i
NBS SPECIAL PUBLICATION 415

U.S. DEPARTMENT OF COMMERCE / National Bureau of Standards

Biomaterials
NATIONAL BUREAU OF STANDARDS

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Biomaterials

Proceedings of a Symposium

held in conjunction with the

Ninth Annual Meeting of the

Association for the Advancement

of Medical Instrumentation,

New Orleans, April 19-20, 1974

Edited by

Emanuel Horowitz and

John L. Torgesen

Institute for Materials Research

National Bureau of Standards

Washington, D.C. 20234

U.S. DEPARTMENT OF COMMERCE,

NATIONAL BUREAU OF STANDARDS, Richard W. Roberts, D/rector

Issued May 1 975
Library of Congress Cataloging in Publication Data

Symposium on Biomatcrials, New Orleans, 1974. Biomaterials.

(National Bureau of Standarils Special Publication: 415)

Papers presented at the symposium held in conjunction with

the 9th annua] meeting of the Association for the Advancement

of Medical Instrumentation, New Orleans, Apr. 19-20, 1974. Supt. of Docs. No.: C 13.10:415

1. Biomedical Materials—Congresses. I. Horowitz, Emanuel, Ed.

II. Torgesen. John L., Ed. III. Association for the Advancement

of Medical Instrumentation. IV. Title. V. Series: United States. National Bureau of Standards. Special Publication: 415. QC100.U57 No. 415 [R857.M3] 389'.08s [610'.28] 74-30305

National Bureau of Standards Special Publication 415

Nat. Bur. Stand. (U.S.), Spec. Publ. 415, 109 pages (May 1975)

CODEN: XNBSAV

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WASHINGTON: 1975

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Forward

For many years the National Bureau of Standards has conducted a research program for the

improvement of dental materials. We are now embarking on a program involving other biomaterials.

Our goals are the development of standards, specifications, and data necessary for the creation

of improved synthetic implants, improved from the standpoint of increased durability and compatibility in the specialized environment of the body.

Success in such a program can only come from the interaction of our materials experts with

the medical profession. We are pleased to cooperate with the Association for the Advancement of

Medical Instrumentation, and to have participated in their Symposium on Biomaterials. I welcome

such Joint efforts, and hope that this publication will increase interest in this vital field.

Richard W. Roberts

Director

iii
Preface

The Association for the Advancement of Medical Instrumentation concerns itself with virtually

the complete spectrum of medical measurements and medical devices — the instrumentation

involved, the techniques employed, safety, efficacy, and education. Within this broad scope of

interest and concern are biomaterials including those used in the fabrication of synthetic implants

which are employed to assist or replace body parts. We recognize that biomaterials research is broadly interdisciplinary in nature, requiring medical, surgical, engineering, and materials experts

for its advancement. It is necessary that opportunities be provided for these experts to come

together to discuss problems, progress, and the potential that biomaterials have in their applications to advances in health care. We welcome, therefore, a collaboration with the National Bureau of Standards and its materials

scientists in organizing and conducting technical sessions at our annual meetings to provide the

forums for interchange of ideas between the medical, materials, and engineering communities.

The present Symposium on Biomaterials was a feature of the Ninth Annual Meeting of AAMI,

held in April, 1974 and is the second in a series organized by NBS on this topic.

We endorse the cooperation of the National Bureau of Standards in publishing this volume of

the Proceedings of the Symposium as one in its series of Special Publications. We feel that this

form of publication will contribute to a wide circulation of the volume and we pledge the full cooperation of AAMI in the execution of this worthwhile effort.

Harold Laufman, M.D.

President, 1974

The Association for the Advancement

of Medical Instrumentation

iv
Contents

Page

Foreword iii Preface iv Abstract vi Introduction vii Plasma-formed Polymers for Biomedical Applications. Part I. Synthesis and Fundamental

Studies 1 K. G. Mayhan, A. W. Hahn, M. R. Havens, and B. W. Peace

Plasma-formed polymers for Biomedical Applications. Part II. Biocompatibility and Applications 13

A. W. Hahn, K. G. Mayhan, J. R. Easley, and C. W. Sanders

Interfacial Behavior of Ceramic Implants 19

L. L. Hench, H. A. Paschall, W. C. Allen, and G. Piotrowski

Soft Tissue Response to a Series of Dense Ceramic Materials and Two Clinically Used

Biomaterials 37

W. C. Richardson, S. F. Hulbert, J. J. Klawitter, and B. W. Sauer

Engineering and Biological Studies of MetaUic Implant Materials 45

N. D. Greene, C. Onkelinx, L. J. Richelle, and P. A. Ward

Materials Characterization of Implantable Porous Electrodes 55

R. B. Beard, J. F. DeRosa, S. F. Dubin, L. Sturm, R. M. Koerner, and A. Miller

Properties of Fibrous Biomaterials with Statistically Dispersed Orientation 63

E. Y. Robinson

A Simple In Vitro Test for Screening the Blood Compatibility of Materials 75

H. Kambic, T. Komai, R. J. Kiraly, and Y. Nose.

Conformation of Adsorbed Blood Proteins 83

B. W. Morrisey, L. E. Smith, C. A. Fenstermaker, R. R. Stromberg, and W. H. Grant

A Nerve Implant Prosthesis for Facilitating Peripheral Nerve Regeneration

Part I. Development of the Prosthetic Device and System of Repair 91

W. E. Kuhn and J. L. Hall

Part II. Development of the Prosthetic Device and System of Repair Neuroanatomical

Aspects 99

J. L. Hall and W. E. Kuhn

V
Abstract

This volume is based on papers presented at the Symposium on Biomaterials, held in con- junction with the Ninth Annual Meeting of the Association for the Advancement of Medical Instrumentation, New Orleans, La., April 19-20,1974. It provides a review of special topics in biomaterials

research selected to focus attention on some noteworthy achievements. The topics covered include

plasma-polymerized polymers and their application in biomedicine; biocompatibility of ceramic

materials and their application both as inert coatings for synthetic implants and as porous materials for bone repair; the selection of metallic implant materials through engineering and medical

considerations; in vitro testing of thromboresistance; adsorption of blood proteins on synthetic

substrates; a prosthesis for nerve regeneration; and properties of fibrous biomaterials.

Key words: Biocompatibility; biomaterials; blood protein; ceramic implants; implantable electrodes; metallic implants; nerve prosthesis; synthetic implants; thromboresistance.

vi
Introduction

Papers published in this Special PubHcation of the National Bureau of Standards were presented orally at a Symposium on Biomaterials, held in conjunction with the Ninth Annual Meeting

of the Association for the Advancement of Medical Instrumentation, New Orleans, La., April

19-20, 1974.

The aim of the Symposium and this pubhcation is to review some of the important advances

that have been made by the medical profession, materials scientists and engineers, and clinicians

in the field of materials used for synthetic implants designed to restore body function and activities

to patients.

Special topics of general interest were selected for the Symposium in order to focus attention on some noteworthy achievements. Thus, there is reported the plasma-generated synthesis of

polymers whose properties are vastly different from those of polymers synthesized by conventional

methods and which offer great potential for use in synthetic implants.

An intercomparison of the biocompatibility of certain ceramic materials points to the potential

of these inert materials to withstand the body environment. A paper on the application of ceramic

materials both as ceramic coatings achieved by a unique flame-spray technique and as porous

material for bone repair involving bone ingrowth represents an important contribution to biomaterials research.

Metals and alloys are discussed from the standpoint of their selection for biomedical application based on engineering and biological considerations, the latter emphasizing systemic effects

of metal corrosion products.

Thrombogenicity is the subject of two papers, the first describing a reliable in-vitro test system for thromboresistance of materials, the second a report on fundamental studies of adsorption of blood proteins on substitutes, which studies are attempting to elucidate the mechanisms

of blood clotting.

The fabrication and behavior of implantable porous electrodes are described, work which

attempts to decrease polarization losses in galvanic cells used to power pacemakers.

The properties of fibrous biomaterials as dependent on orientation of fibers and layers are discussed as applied to the development of composite biomaterials whose properties approach

those of bone.

The final paper discusses the development and application of a prosthesis which is very

successfully used in the regeneration of a severed nerve.

In including this symposium in the program of its Ninth Annual Meeting, the Association for

the Advancement of Medical Instrumentation is performing an important service to the biomaterials

community, as well as to the general public. The Symposium provides a forum for the discussion

of progress in research on synthetic implant materials to assist in the development of standards

and specification for these materials and improved methods of analysis, characterization and

testing leading to improved materials for use by the surgeon to benefit the patient.

Emanuel Horowitz

John L. Torgesen

Institute for Materials Research

National Bureau of Standards

vii

NATIONAL BUREAU OF STANDARDS SPECIAL PUBLICATION 415,

Biomaterials. Proceedings of a Symposium held in conjunction with the Ninth Annual Meeting of the Association for the Advancement of Medical Instrumentation, New Orleans, La., April 19-20, 1974

(Issued May 1975).

Plasma Formed Polymers for Biomedical Application

Part I. Synthesis and Fundamental Studies

K. G. Mayhan

Graduate Center for Materials Research and Chemical Engineering Department

University of Missouri - Rolia, Rolla, Mo. 65401

A. W. Hahn

John M. Dalton Research Center

University of Missouri, Columbia, Mo. 34929

M. R. Havens

Graduate Center for Materials Research and Chemical Engineering Department

University of Missouri - Rolla, Rolla, Mo. 65401

and

B. W. Peace

E. I. du Pont de Nemours, Wilmington, Del. 19898

The deposition of polymeric coatings through Rf plasma techniques is a unique process from

several points of view. By introducing monomeric gases into an inert gas Rf plasma, the monomers

are converted to uhrathin continuous polymer hims. These films can be made to vary in thickness

from less than 1 ;u,m to 10 /xm or more. These materials are insoluble to common organic solvents and are not attacked by strong mineral acids. In addition, the monomeric materials in the reactor are

distilled gases and are therefore quite pure in comparison to monomers used for ordinary linear poly- merizations. As a result, the polymer coatings which are formed do not contain residual catalysts or

activists and are in a highly purified state. We have utilized the plasma polymerization process to produce polymer films intended for biomedical application. Plasma coatings have been formed from

a wide variety of monomers. These coatings have been applied to various nonmetallic, metalhc and

organic substrates. In each instance it was found that continuous adherent films could be deposited

after suitable substrate preparation and reactor operation parameters were established. It has been

our experience that a general cleaning procedure for all substrates cannot be dictated and that the preparative steps involved are dictated by the environment to which the final product will be subjected.

Key words: Plasma-formed polymers; polymer films; polymer permeability; polymerization.

1. Introduction

During a recent NSF biomaterials planning

session held in Washington, D.C., some basic

definitions were laid down to guide research efforts

in the biomaterials area. The definition adopted was:

"Biomaterials are defined as any synthetic or modified

natural materials which come into contact with body tissues

or fluids. Biomaterials research is aimed toward providing the biomedical community with the best possible materials needed

to aid in the restoration of bodily function and/or appearance.

This research is directed towards obtaining knowledge which

includes any information which may be expected to be useful

in the development, synthesis, formulation and characterization of the biomaterials."

Plasma polymers are capable of being an integral

part of the materials applications encompassed by

this definition.

Research into plasma deposited polymers has

grown recently to cover a wide range of topics. The

Subcommission on Plasma Chemistry of the Inter- national Union of Pure and Applied Chemistry

recently released a tabulation showing some 150

investigators, located in Australia, Austria, Belgium,

Canada, Czechoslovakia, Denmark, France, Germany, Italy, Japan, The Netherlands, New Zealand,

Switzerland, and the United Kingdom, who are

actively working on some phase of plasma chemistry

1
or plasma physics. Also, some 100 investigators in

the United States are pursuin*;; similar endeavors.

The research topics range from very fundamental

mechanistic considerations to plasma technology,

processing and applications. This work encompasses almost every Held of engineering and science.

To date, little has been reported on the application

of plasma deposited polymers as a biomaterial.

This paf)er summarizes the authors' bioappHcations work which includes three main areas of

plasma deposited polymers: (1) synthesis and

characterization, (2) polymerization mechanisms

and (3) applications.

When discussing biomedical applications, it is important to understand the differences between

conventional linear polymers and those polymers

formed in a plasma from the same monomers. For

instance, conventional linear polymers, such as those derived from styrene, methyl methacrylate,

vinyl acetate, and vinyl chloride, are produced by

either free radical or ionic mechanisms. The re- sulting polymers consist of coiled chains having

molecular weight distributions as low as two and

as high as 10. When these polymers are placed in

good thermodynamic solvents, they begin to swell

and dissolve. Linear polymers can be melted and

extruded and can also be made to decompose at fairly low temperatures. Outside of the normal

branching found in most hnear polymers, the repeat

group is predictable.

By comparison, plasma polymers from the same

monomers are formed by a multitude of mechanisms,

all occurring simultaneously. The plasma deposited

polymers are randomly branched, highly cross- linked, intertwined networks. The polymer resulting

from formulation in the plasma is no longer soluble

in good thermodynamic organic solvents.

Plasma deposited polymers are not even swollen.

In fact, plasma polymers have been refluxed in

boiling organic solvents for up to 7 days with only a 1-2 percent weight loss. Plasma deposited polymers

do not melt, therefore cannot be extruded, but only

decompose at comparatively higher temperatures.

Also, plasma deposited polymers have shown

considerably higher chemical resistance to reducing

or oxidizing environments than conventional

polymers. This resistance could be taken, at times,

as inertness. Chiefly for the last two properties,

minimal solubihty and maximum chemical resistance, plasma deposited polymers deserve consideration as biomaterials.

2. Historical

A plasma was first defined by Langmuir in the

early 1930's as a state a gas achieves when it is excited to the point of ionization. Plasma is more

correctly defined as the region in which the active

species are actually formed. Faraday and Crookes

are perhaps the first to report observing a plasma

glow. De la Rue and Muller reported a relationship

between pressure, the type of glow and the resist- ance of the gas in the discharge. Crookes stated in

1879 that a plasma was a fourth state of matter.

C. T. R. Wilson reported in the early 1900's that

the electrical conductivity of plasmas was due to ionization of gas molecules by collisions with themselves. In 1910 Dewar and Jones reported the first silent discharge from a DC discharge in an evacuated Wood's tube. The silent discharge operated at much lower temperatures and would prove to be

much more useful in chemical synthesis [1].' Shortly after plasmas were observed, fundamental studies of their behavior were carried out

primarily by physicists with little regard for the

chemical aspects or possible applications of these

plasmas. De Wilde [2] and A. Thenard [3] in 1874

reported simultaneously that acetylene reacted

readily in a silent discharge leaving no gaseous

residue and forming a hard, brittle solid which was

insoluble in common solvents. This was the first report of plasma polymerization; however, no effort

was made to study the phenomenon beyond this

point. At this point in the history of chemistry,

polymers were considered to be undesirable byproducts and were of no interest.

The first and the oldest form of a plasma is the

disruptive discharge, the arc and spark. These discharges produce very high temperatures since they

operate with a high voltage source at either atmospheric or higher pressures. They are not useful in

organic chemical synthesis since they operate at several thousand degrees Kelvin.

The other type of plasma is the silent plasma

which operates in a vacuum and over a large

temperature range which may be controlled by the

power input to the plasma generator. There are

five important types of the silent or nondisruptive

discharge. The first is the glow discharge which is established in low pressure gases by low frequency

rf applied across electrodes sealed into the reaction

tube. Second is the electrodeless discharge which

is formed by passing radio frequency current through

a coil surrounding the tube. Third, the point

discharges are established between conductive

points operating at a high voltage, but arranged so

as not to cause an arc. Fourth is the corona discharge

which is similar to the point discharge except a

fine wire is used in place of a conductive point. The

fifth type is a variation of the corona and is widely

used to produce ozone and thus is known as an

ozonizer.

Both the glow and electrodeless discharge can

operate at rather high power levels and remain at essentially ambient temperatures. For this reason

they are usually chosen for organic synthesis

studies.

Plasma Mechanisms. The initial discharge oc- curs as a result of the free electrons being acceler-

' Figures in brackets indicate the literature references at the end of the paper.

2
ated in an electric field until they can cause the gas

molecules to ionize. The electrons that are released

as a result of this ionization are further accelerated

and cause subsequent ionization. This type of chain

reaction wiU soon cause the gas to become conductive; the current or flux in the gas rises and a discharge is started. An equihbrium is estabhshed

almost immediately where rate of ion formation

is equal to the rate of recombination of the species.

For an example of these reactions, let us consider

the possible reactions of molecule AB which can

be given by the following equations:

AB + e^A+B- (1)

AB + e^A+B (2)

AB + e^A+ + B- + e (3)

AB + e^A- + B+ + e (4)

AB + e^AB+ + 2e (5)

A+ + e^A (6)

A- + B+-»AB (7)

AB+ + e^AB (8)

Equations (1-5) represent the formation of active

ionic species while (6—8) represent recombination

reactions to form neutral species.

Free radicals are another type of active species

formed and are of special importance in polymerization reactions. Radicals are formed when a molecule

dissociates into two species in the discharge as given in eqs (9-12).

AB^A+B- (9)

AB+^A- + B+ (10)

RH->R+H- (11)

CzHsH^CzHs+H- (12)

Equation (12) is an example of an organic radical

that has been observed by spectroscopy techniques.

The radicals can propagate polymer chain growth

until they combine to form a neutral species. This can occur in the gaseous phase by two mechanisms:

1. Disproportionation

2C2H5-^ C2H4 + C2H6 (13)

2. Combination

2C2H5-^C4Hio (14)

Chemical Synthesis in Plasmas-Organic. The

early work in cold plasmas dealt with the trans- formation of a small organic molecule to different

molecules. For example. Coats [4] identified 25

different hydrocarbon compounds after passing

7i-hexane through a microwave discharge plasma.

Two free radical reactions, disproportionation and

combination, have already been discussed. Some

additional free radical reactions wiU be given now. Hydrogen abstraction occurs via the free radical

route and an example given by eq (15) was reported

by Cher.

J0'-CH3+CH3 -^CH4+- CH2^ (15)

Here a methyl radical removes a hydrogen from

toluene to form another free radical. Tomkinson [5]

reported halogen abstraction could occur in like manner. A reaction which can lead to polymerization was

reported by Brinton and is given in eqs (16-18).

;0'-CH3 + C2H4-^ • C3H7-J9- (16)

je'-CgHy +C^H, -CsHii-F (17)

C„H2„ + i- + /iC2H4^ C(C2H4)„C- (18)

Brinton's reactions terminated before polymer was

formed; however, there seems to be no reason why

eq (18) could not occur.

Investigations involving polymei" plasma chemistry reported to date fall into two classifications:

treatment of conventional polymer films in a plasma

[6, 7] and actual polymerization of monomers using

the plasma as an initiator [8—11].

Hanson et al. [121, reported that when several

polymers (polyethylenes) were treated in a mild

oxygen plasma, the bulk properties of the polymers were unaltered as the penetration was restricted to the first few hundred Angstroms of the films. In

other cases surface ablation of the polymers was

reported. With alkane polymers a highly oxidized

surface was created which had vastly improved

wettability characteristics. Hanson and Schonhorn

[13] reported that treatment of polymer films of

polyethylene and teflon in an oxygen plasma in- creased their adhesion properties. Weininger [14,

15] studied the effect of active nitrogen on polyethylene and polypropylene and found increased

cross-hnking and increased unsaturation resulting

from the treatment. HoUahan, Stafford, Balk and

Payne [16] reported that amino acid groups could

be attached to polymer surfaces by treating the

films with nitrogen and hydrogen plasmas. Bazzarree

and Lin [17] treated fiberglas filaments with various

vinyl type monomers (vinyltriacetoxysilane and

vinyltriethoxysilane) in an attempt to increase their

tensile strengths.

Shaw polymerized a sihcone oil on the surface

of small electronic devices and produced an im3
permeable, insulating, thin film. He used a medium

power four megahertz capacitively-coupled plasma

[18|. Ozawa [19] employed plasma polymerization

teciiniques using styrene and tetrafluoroethylene to make thin film insulators for micro capacitors.

He used an alternating current (low frequency)

electrode discharge plasma in a long glass tube. He

found a more uniform coating was formed when the

substrates to be coated were placed away from the

electrodes in the plasma glow. Mearns [20] pubhshed

a review which includes the use of plasma polymerization in the production of thin insulating films.

Some work with vinyl type monomers has been

reported. Goodman [21] described the formation of

uniform thin (0.1 to 2 ^tm) polymer films with a plasma from a high voltage alternating current

discharge between two parallel plates. Some of the

monomers used were styrene, methyl methacrylate,

tetrafluoroethylene, monochlorotrifluoroethylene,

and difluoroethylene as well as some nonvinyl

compounds such as benzene, toluene, and chlorodifluoromethane. Williams and Hayes [22] also

reported the flow discharge polymerization of

several vinyl type monomers. Denaro, Owens, and

Crawshaw [23] investigated the flow discharge

polymerization of styrene. This is the more comprehensive plasma polymerization study reported.

The initial work summarized above led to many

investigations both basic and applied. The following

investigations represent only those which interface

the work of the authors and, thus, are not to be

considered as a comprehensive list. Several books

[24-26] and one general review [27] related to plasma chemistry have been published. Suhr [28]

and Stille, [30-31] among others, have investigated

some basic organic synthesis using plasmas. Other

investigations have looked into oxidation [32] and

decomposition [33-36] of various materials in

plasmas. A number of investigators have looked

into the polymerization aspects of plasma chemistry

[37-44]. Others have looked at specifics in the

polymeric films which are usually produced [45,

46]. Some specific applications of these films have

recently begun to appear in the literature [47-49].

The authors and others [50] have been investigating

acceptability of the plasma-deposited polymeric

films to biomedical applications.

3. Plasma Reaction Procedures

Before discussing some of the specifics of biomaterials work with the plasma polymers, the actual

physical and chemical processes involved in their

formation should be described in general terms.

Figures 1 and 2 show block diagrams of the reactor

system. Figure 3 is a photograph showing one of our more refined reactor systems where basic studies are performed. Figure 4 shows a low cost reactor

which was designed and constructed in our laboratories to be utilized in the applications research.

RF COIL REACTOR

LEAK

VALVE

TO ARGON

PURGE

MONOMER

INLET

hooQUANTITATIVE BACKSTREAMING COLD TRAP COLD TRAP

"O" RING JOINT 1

DIFFUSION PUMP

ROUGHING PUMP

Figure 1. Schematic of reactor system.

RF TRANSMITTER

MATCH

BOX

POWER

METER

TO COIL

Figure 2. Schematic of power input to load coil.

The reactors can all be characterized simply as glass

tubes wrapped with a copper coil which serves to direct the incident rf power into the tube volume.

One end of the reactor contains the inlet system

which is one or more leak valves. The other end is connected to a cold trap and then to a vacuum

system. The electronics for the rf power are a

straightforward application of conventional ham

radio parts.

The first step in the process involves consideration of the type of substrate to be either coated,

cleaned, or converted. Substrate preparation and

cleaning procedures used vary considerably with

the type of material to be coated. The cleaning

procedures may vary from rinsing with residue free

solvents to chemical etching followed by thorough

distilled water rinses. In some cases we have

resorted to ESCA and Auger spectroscopy to determine the effectiveness of overall surface preparation prior to deposition of plasma polymer films with

regard to resulting film adhesion. An example would

be a glass substrate. The structure of the cleaned

glass substrate and subsequent adherence of the

plasma polymer is a function of the chemical composition of the glass. Each material (metals, ceramic,

polymer) presents its own unique problems which

must be worked out for each application. After the

samples to be coated have been cleaned, care is taken not to contaminate the surface through improper storage or handling.

4
Figure 3. Research reactor.

Figure 4. Applications reactor.

The substrate is then loaded into the reactor

either for further cleaning or to be coated. The

reactor is then secured and evacuated to a vacuum

of below 10"'* torr. A prepurified gas (Argon, Helium,

Nitrogen, etc.) is bled into the system to achieve a desired pressure. The rf power is tuned up and

followed by initiation of the plasma by means of a Tesla coil. By utihzing the initial gas at a higher

power level than for polymerization, it is possible,

in many instances, to obtain additional substrate

cleaning through the interactions of the ionized gas

and the substrate. Similar treatments have also been

effective in removing impurities from the surfaces

of implant devices. This interaction step can be a prelude to sterilization. At this point the monomeric

gas of choice can be introduced simultaneously

with the inert gas flow or the inert gas can be

discontinued and simply replaced with monomeric

gas. The pressure and power are adjusted to their

final values, respectively, by setting the leak valves

and tuning and matching the rf transmitter and load.

The coating process then proceeds until the

desired thickness of material is obtained. Relatively

pure polymer films with thicknesses of 0.5 to 10 fim can be formed effectively by this means. These

films, as will be noted in part 2, have little or no

tissue reactions. Any tissue reactions which may

occur can be assigned to the polymer structure

itself and not to the impurities present in most

conventional polymer systems. This assignment can be made because it is inherent to this coating

process that all monomers enter the reactor as gases. Thus, the gases can be considered to be

essentially purified by their being distilled away

from any initiators, catalysts, or inhibitors which

normally appear as contaminants in the product.

Also, no sensitizers, initiators or catalysts are needed in the reaction.

The total time for deposition can vary from 15

minutes to 12 hours. The rate of polymer deposition

is largely dependent on the power input to the load

coil, the effective concentration of the monomer gas

or gases and reactivity of the particular monomer

involved. After polymerization, the reactor is re- turned to atmospheric pressure by slowly introducing the desired terminating gas.

Recent work in the authors' laboratories has

shown that an alternate embodiment of the above

procedure can also lead to useful polymer films.

After proper cleaning, the substrate is coated with a

thin film of a nonvolatile low molecular weight

polymer or a nonvolatile high molecular weight

monomer where both are generally in the form of a

liquid. After evacuation of the reaction vessel, a non- polymerizing gas is introduced into and excited in

the rf field through proper adjustment of the reaction parameters. The liquid film is converted to a

solid insoluble polymer having properties similar

to those contained in the gas phase deposition.

Several interesting systems have been developed

through this special synthesis technique.

Other related work has been completed recently

which involves hnear polymer film interactions

with plasmas. Both plasmas of polymerizable and

nonpolymerizable gases were used. In both casts,

the result is that the surface of the hnear polymer

film is converted to material similar to that obtained

in the previously discussed hquid phase-plasma

interactions.
4. Generalized Results

In the authors' laboratories, the tubular type

inductively coupled reactors have been found to be

superior to the capacitively coupled plate reactors

in the applications investigated. Figure 5 (Plasma

ethylene) shows the type of polymer distribution

obtained for one reactor through more refined design

techniques. The bell-shaped curve may be further

flattened allowing more uniform polymer deposition.

2 0

10 20 30

DISTANCE FROM INLET (cm!

Figure 5. Polymer formation distribution.

It should be noted that equal deposition around the

circumference of the reactor is observed. Kinetic

studies have revealed that the rate of polymer

deposition at any point in the reactor is constant

with time with all other reaction parameters being

constant. Figure 6 illustrates the temperature rise

in the reactor as a function both of power input and

thermocouple position in the reactor. General

experience shows that acceptable polymer films can be obtained outside of the reactor volume cir- cumscribed by the coil with little or no temperature

rise while operating at fairly high power levels. It has also been shown that monomeric conversion can be controlled by adjusting power inputs and

monomer concentration.

22

O CENTERED IN COIL • 35 cm. FROM COIL

6 8 10 12 POWER (WATTS)

14 16 18 20

Figure 6. Temperature rise as a function of power input and

reactor position.

At first glance, it would appear that plasma

polymerization products possess a much greater

thermal stability than corresponding Unear poly- mers. Indeed, this is true after a fashion. Figure 7

shows thermograms for linear and plasma polystyrene heated in air to 700 °C. The plasma polystyrene still retains 40 percent of its initial weight

whereas all of the linear polystyrene has decomposed below 400 °C. However, a closer look at these

thermograms indicated that the plasma polymer

actually began to lose weight initially at a considerably lower temperature than the linear polymer.

Therefore, any statements about thermal stability

must take into account the temperature range under

consideration. The four curves of plasma polystyrene in figure 7 show the effect of polymer film

thickness. It becomes apparent that the rate of

escape of the decomposition products from the

thicker films is a diffusion controlled process. Work

currently underway is designed to determine the

nature of the apparent instability of these materials.

Ways and means are being sought to stabilize the

plasma polymer films through quenching techniques in order to flash off any low molecular weight

volatiles.

In one aspect of application, some work has been

completed recently in investigating plasma deposited polymers for use as membranes. The pur100

80

u

a> a. 60

LiJ

X

o

LlJ 5

40

20

\

^\

\

\

\

\

STD, PS

• 22 WATT- 12 HR. 22 WATT- 9 HR. 22 WATT- 6 HR.

• 22 WATT-3 HR.

\ \ \

\ \ \

\ \ \

\ \ \

200 400 600

TEMPERATURE {°C)

Figure 7. Thermogram of plasma polystyrene.

6
Table 1. Permeability of gases through uncoated and coated films

G3S

Permeability

/cm3,STP)cmX

\ cm Hg-cm'^-s /

Uncoated 1 mil (25 /xm)

polyethylene film

1 mil (25 fJLm) polyethylene film coated with 1-2 micrometers

polypropylene

Helium 6.290 3.33 Neon 2.070 0.72 Oxygen 3.250 i!oo

Argon 0.375 C)

Carbon dioxide 9.290 1.93

^ For conversion to SI units (mol - cm -cm"^ - Pa"' -s"'), multiply values by 3.3 X 10-". Permeation rates were too slow for satisfactory measurements.

pose of this work was to show that adherent plasma

polymer films could be deposited on smooth

polymer membrane surfaces thereby altering the

membrane's permeability characteristics. Unplasticized linear polyethylene membranes 1-mil (25 fMm)

thick, were chosen at the start of the experiments.

On these membranes was deposited a 1-2 micrometer film of adherent polypropylene.

The permeability of the coated and uncoated

membranes to He, Ne, Ar, O2, and CO2 was measured. The results are tabulated in table 1. In comparing the data for coated and uncoated membranes,

it is plain that the thin (2 /u,m) film of plasma deposited polypropylene considerably reduced the

permeability constants. These permeability constants are plotted against the square root of the

permeating gas' molecular weight in figure 8. Both

oxygen and carbon dioxide show a definite dissolution in the uncoated membrane but appear to approach Graham's law behavior with the coated

membrane. The slight change in slope between

helium and neon in both coated and uncoated cases

indicates that some pores have been altered or plugged in the coating process.

12.0

9.0

0,0

O Uncoated • Coated

a.

1 0

0

K 6.0

1- He

-Jm

<

LUS •

(T 3.0

hi

o

COa

O

Ne

O

O2

Ar ,0

0.0

Figure 8.

2.0 4.0 6.0 8.0

Permeability of uncoated and coated 1-mil polyethylene

The significance of this data must be evaluated

by taking into account the two terms of the permeabiHty constant. This constant is the product of the

diffusion constant and solubihty constant, P = DS.

The large reduction in the permeability cannot be

explained wholly by the reduction in diffusion

through pores which have been altered or plugged

by the plasma deposited film. It is known that

plasma deposited films are highly cross-linked into a tight three-dimensional network. For example,

plasma polypropylene samples have been refluxed

in aliphatic and aromatic hydrocarbons, chlorinated

hydrocarbons, esters, and ketones for up to a week

without appreciable changes in solubility. Thus,

the plasma deposited film reduces the solubility of

O2 and CO2 which allows diffusion to be the predominant transport mechanism, although it is also

significantly reduced. The result is to approach

Graham's law.

For these reasons, plasma deposited polymer films

may have certain advantages as membrane materials. It appears possible that by tailoring thickness

and/or composition of the film, a membrane can be

made to very nearly approach Graham's law behavior. Behavior like this could be important in

those applications where either a definite ratio of

two gases is desired for a particular transfer

application or where a specific diffusion rate is required. However, greater film thicknesses could

conceivably block gas permeation across the

membrane.

A complimentary membrane study involved

taking the same membrane material and subjecting

it to a high wattage argon plasma. The investigation

was aimed at alterations of the permeability constant by interaction of the membrane surface with

the nonreactive plasma as opposed to plasma

polymer depositions. Results of this work for carbon

dioxide and helium permeation are shown in

figure 9. PermeabiUty was reduced but not as

significantly as was the case with plasma deposited

polymer films.

Since the aforementioned properties of plasma

deposited polymers are markedly different from

7
9

o

X 8 -

<

a:

Q.

I I 1 1 0 50 100 150

REACTOR PRESSURE- MICRONS

Figure 9. Permeability of gases through plasma treated 1 -mil polyethylene.

those of their generic polymers, an investigation

into the more fundamental property of density was

initiated. The densities of nine rf plasma produced

polymers were determined in a gradient density

column. Generally, these polymers were found to be

more dense than conventional polymers produced

from the same monomer species via common free

radical, ionic, or condensation polymerizations.

The results of the density determinations are given

in table 2. A density range is given for each polymer

because the density was found to vary with the

monomer concentration, the power level and the

position within the reactor. For comparison, the

densities of both the single crystal and the amor- phous conventional linear polymer are reported.

Low angle x-ray diffraction patterns of plasma

polystyrene and polyethylene indicated that these

materials were amorphous. Since the generic polymers of styrene and ethylene are highly crystalline,

this indicates that substantial differences in structure and morphology are present. This density data

has been correlated with refractive index measure- ments of the plasma deposited films. Further

probing of these properties has resulted in preliminary data from elemental analysis that indicates

oxygen and sometimes nitrogen are incorporated

into the plasma deposited films. For example,

plasma deposited polyethylene was found to contain

from 20 to 30 weight percent oxygen. In addition,

the density, refractive index and chemical composition have all been found to vary as functions of both

reactor design and operation. Current investigations are aimed at determining the source of these elements and the method of their incorporation into

the polymer.

5. Current Problems and

Future Work

The early work of the authors involving the

coating of bioelectrodes and various implantable

substrates went well and yielded promising results.

It appears that this early work stayed within the

then undefined tolerance limits of substrate preparation, reaction parameters and coating techniques.

Reports of dissimilar results froni other laboratories

for the same monomers regarding both changes in

the polymers after formation and elucidation of

different reaction mechanisms were disturbing since

these phenomena had not been observed in our

laboratories. However, these problems and many

more began to surface once a reaUstic approach was

Table 2. Polymer densities

Starting monomer

Density range of Crystal density [3] Amorphous density [3]

plasma polymer of linear polymer of linear polymer

(g/cm^) (g/cm^) (g/cm^)

1.332-1.408 1.126 1.04 1.141-1.231 1.014 0.852

Vinyl chloride 1.463-1.499 1.44-1.522 1.39 Vinyl fluoride 1.341-1.395 1.44 1.331-1.364

1.402-1.436 1.11 syndiotactic

1.54 isotactic Allene 1.195-1.203

Methylvinyl-dichlorosilane 1.435-1.454

1.212

8
Figure 10. Electronmicrographs of plasma deposited polymers.

taken to the biocompatibility and use applications

of these polymers. After several changes in reactor

design, more sophistication in electronics and

modifications to both the vacuum system and

process procedures, it became apparent that ex- periments were being run out of control. Results

which were favorable during one experiment were

not favorable during another experiment, although

both were run under presumably identical conditions. Most of the problems involved adhesion of the

thin polymer films to the substrates. An extreme

example of the differences in adhesion is shown in

figure 10. Both films were deposited on glass. In

micrograph A, a 2 /nm plasma polypropylene film

is adherent and so smooth that an artifact had to be

introduced onto the surface in order to locate the

surface in the SEM. Micrograph B shows cracking

of another 2 plasma polypropylene deposited

under the same conditions but in another run. A

series of studies involving both the changing of

reaction conditions and the effects of position within

the reactor upon adhesion led to the conclusion that

more reproducible results were obtained with lower

power inputs and where the substrates were positioned toward the trailing edge of the plasma field. These findings only partially eliminated the adhesion problem.

It finally became necessary to curtail all routine

experiments and launch a full scale investigation

(basic studies have been in progress for 2 years)

into the cause and effect of the various parameters

upon the final properties of the polymer films.

Elemental analyses of the plasma polymer films

confirmed that the polymer structure contained

elements not present in the original monomer.

Specifically, oxygen and nitrogen were contained

in the final polymer structure. Figure 11 gives some

typical results. This investigation is continuing in

conjunction with a more fundamental study. Additional studies are underway to determine whether

Elemental Analysis of Plasma & Conventional Polymers

Plasma Polymer Conventional Polymer

X C % H 1 N I C I H _ I 0

Polyethylene, 10 80^m 76.16 8.28 15.3 85,71 11,29

10 V, 60/tM 77.12 8.37 11,2

10 W. lO^H 73.96 8.36 17,7

10 W, 20^„ 76.72 8.20 15,1

Polystyrene- 10 W, 'iO^h 52,20 1.68 21,5 18.63 92,31 7,69

26 H. lO/iM 55.69 5.26 21,7 11,35

51.32 1.67 21,0 19.97

Polyacetonitrile 55.81 3.95 11,8 28.39 58,51 7,31

Polyvinyl Chloride, 17 W, 25^h 16.10 3.11 0,16 38,10 1,80

Polyacrylonitrile, 17 W, 27^m 18.60 1.73 25,5 21,22 67,92 5,17

Polyvinyl Flouride, 16 W- 25^m 62,88 5.82 1,73 52,17 6,52

Figure 11. Elemental analysis of plasma polymers.

12 (= I I I I I I H

' ' I I I I I 0 20 40 60 80

TIME (Hours)

Figure 12. Weight gain of plasma polyethylene as a result of

atmospheric exposure.

9
or not the source of these elements is contamination

of the monomers, perpetual leaks in the system,

or whether the polymer is reacting with atmospheric

gases after formation. A quartz crystal microbalance

was set up in the reactor shown in figure 3 to follow

the. rate of polymerization and also to follow gas

adsorption after polymerization. Figure 12 shows

the weight gained by a plasma polyethylene sample

deposited on one of the quartz crystals. It is evident

that the polymer film is indeed reacting with atmospheric gases.

In conjunction with the quantitative data from

this quartz crystal, qualitative data on the species

involved was obtained by depositing the plasma

polymer films on sodium chloride crystals for IR

spectroscopy. Figures 13 and 14 show the IR spectra

of such a film and the changes which it has undergone. Preliminary identification of the new absorption peaks indicated the presence of both nitrogenbearing compounds and cyclic oxygen compounds.

It should be noted that this particular film was not

adherent and after 7 days exposure to the atmos3000 2000 1500 1000

WAVENUMBER CM"'

Figure 13. Infrared absorption, atmospheric quenched

polyethylene.

Infrared Absorption Spectra

Atmospheric Ouenched Polyethylene

Spectrum Mo, Elapsed Time Comments

2 30 MiN.

3 90 MiN,

4 5 Days Mew absorption peaks

evident at 3900-i|100,

3100-3300- 2^100-2500,

2000-2400, 1800-2000,

700-1350 wave numbers.

Peaks at 3200-3700, 1500-

1800 MORE intense,

Coating losing adhesion,

5 6 Days Same as Spectrum ^.

6 7 Days Coating peeled - additional

spectra not possible.

Figure 14. Absorption data associated with figure 13.

10
phere had completely peeled from the sodium

chloride crystal. Elemental analyses of the peeled

film showed the presence of both oxygen and nitrogen. Another experiment was run in which the

finished plasma polymer film was allowed to soak

in an ethylene atmosphere in the reactor for 8

hours prior to being removed and exposed to the

atmosphere. This film has not shown a tendency

to peel. Looking at its spectra in figures 15 and 16,

it is obvious that the hydroxyl and carbonyl regions

have become more intense during observation for a period of 61 days. However, the new adsorption

peak previously noted in figures 13 and 14 have

not appeared.

Additional experiments are underway to determine how the elemental composition of the polymer

changes as a function of length of time that it is exposed to the atmosphere. These experiments

are being run whereby the plasma films are being

soaked or quenched with reactive and nonreactive

gases, as well as being subject to thermal treat- ments. Studies are also being conducted to follow

the decay, of the active species by means of electron

spin resonance spectroscopy over a similar period

of time.

100

80

60

(0s

inz

<

I2 40 o

u

0.

20 —

l-t = l5 Minutes 3-t= 5 Doys

5-t«22 Days

9-t«6l Days

C= 75.81 % Wt. (Analysis After 61 Days)

H>8.ll %

N = 0.75 %

0« 133 %

4000 3000 2000

WAVENUMBER CM'

1500 1000

Figure 15. Infrared absorption curves, ethylene quenched polyethylene

Infrared Absorption Spectra

Ethylene Quenched Polyethylene

Spectrum No,

1

2

3

Elapsed Time

20 MIN,

1 Day

5 Days

22 Days

25 Days

26 Days

60 Days

61 Days

Comments

Absorption peaks at 3200-3700

1600-1800 and 800-1300 more

intense. No other peaks detected. Adhesion uneffected,

3200-3700, 1600-1800 and

800-1300 MORE intense

No change,

No change,

3200-3700- 1600-1800 and

800-1300 MORE intense.

No change. Test concluded.

Figure 16. Absorption data associated with figure 15.

11
6. Summary

Plasma polymerization constitutes a unique technique f(jr forming thin polymer films on a variety

of substrates. In most instances the polymers

formed by these techniques are more uniform and

more stable than conventional polymer films. There

is a great deal to be investigated concerning the

mechanism by which these polymers are formed.

In addition, the kinetic and thermodynamic variables must be understood before one can hope to deposit adherent films on substrates intended for

biomedical apphcations. Prehminary results have

indicated that these polymers are potential candidates for biomedical purposes.

We wish to acknowledge the advice and assist- ance of Dr. Theodore Wydeven, NASA Laboratories,

Mofifett Field, California. We also wish to acknowledge The National Science Foundation who partially

funded this work.

7. References

[1] McTaggart. F. K.. Plasma Chemistry in Electrical Dis- charges, (Elsevier Publishing Company, New York, N.Y.,

1967).

[2] de Wilde, P., Ber. 7,4658(1874).

[3] Thenard, A.. Comp. Rend. 78, 219 (1874).

[4] Coats. A. D., U.S. Dept. Comm. Office Tech. Serv. A. D., 419-618 (1962).

[5] Tomkinson, D. M., Calvin, J. P.. and Pritchard, H. 0.. Free

radical dissociation of halogenated aromatics in a plasma,

J. Phys. Chem. 68 , 541 (1964).

[6] Ogurtsoua, N. N., and Podmosleineskll, J., Zh. Prikl. Spektrosk. 8, 1043 (1968).

[7] Sacher. E., J. Polym. Sci. Part A, 6, 1813 (1968).

[8] Sakurada, I., Macromolecules 1, 265 (1968).

[9] Bashara, N. M., and Doyty, C. T., Appl. Phys. 35 , 3498

(1964).

[10] Kronick, P. L., and Jesch, K. F.. J. Polym. Sci. Part A, 7,

767 (1963).

[II] Bradley, A., and Hammes, J. P., J. Electrochem. Soc. 110,

15(1963).

[12] Hanson, R. H., Pascale, J. V., de Benedictis, T., and Rentzepis, P. M., J. Polym. Sci. Part A 3, 2205 (1965).

[13] Hanson. R. H.,' and Schonhorn, H., Polym. Letters 4, 203

(1966).

[14] Weininger, Nature 186, 546 (1960).

[15] Weininger, J. L., The reaction of active nitrogen with

polyolefins, J. Phys. Chem. 65, 941 (1961).

[16] Hollahan, J. R.. Stafford, B. B.. Balk, R. D.. and Payne,

S.T.,J. Appl. Polym. Sci. 13,807(1969).

[17] Bazzarree, D. F., and Lin, L. J., Quarterly report to Naval

Ordnance Systems Command. No. N00017-69-C-4424,

(1969).

[18] Shaw, W., Personal Communication, Western Electric Co.,

Kansas City, Mo., (1970).

[19] Ozawa, P. J., IEEE Trans. PMP-5, 112 (1969).

[20] Mearns, A. M., Thin solid films 3, 201 (1969).

[21] Goodman, J., J. Polym. Sci. 44, .552 (1960).

[22] Williams, T.. and Hayes, M. W., Nature 209, 769 (1966).

[23] Denaro, A. R., Owens, P. A., and Crawshaw, A., Eur. Polym. J. 4, 93 (1968).

[24] Venugopalan, M., Reactions Under Plasma Conditions,

(Two Volumes), (Wiley-Interscience, New York, N.Y.,

1971).

[25] Nasser, E.. Fundamentals of Gaseous Ionization and Plasma

Electronics, (Wiley-Interscience, New York, N.Y., 1971).

[26] Gould, R. F. (Ed.), Chemical reactions in electrical dis- charges, (American Chemical Society, Washington, D.C..

1969).

[27] Kolotyrkin, V. M., Gilman, A. B., and Tsapuk, A. K., Russian

Chem. Rev. 36, No. 8, 579 (1967).

[28] Suhr, H., Naturwissenschaften 55, 168 (1968).

[29] Swift, F., Sung, R. L.. Doyle, J., and Stille. J. K.. J. Org.

Chem. 30, 3114 (1965).

[30] Stille. J. K., Sung, R. L., and Vander Kool, J.. J. Org. Chem.

30, 3116 (1965).

[31] Stille. J. K., and Rix, C. E.. J. Org. Chem. 31, 1591 (1966).

[32] Lawton, E. L., J. Polym. Sci. A-1, 10, 1857 (1972).

[33] Taki, K., Bull. Chem. Soc. of Japan 43, 1574 (1970).

[34] Taki, K., Bull. Chem. Soc. of Japan 43, 1578(1970).

[35] Taki, K., Bull. Chem. Soc. of Japan 43, 1580 (1970).

[36] Simionescu, C, Asandei, N., Denes, F., Sandulouici, M.. and Popa, G., Eur. Polym. J. 5, 427 (1969).

[37] Yasuda. H., and Lamaze, C. E., J. Appl. Polym. Sci. 15,

2277 (1971).

[38] Westwood, A. R., Eur. Polym. J. 7, 363 (1971).

[39] Westwood, A. R., Eur. Polym. J. 7, 377 (1971).

[40] Brown, K. C, Eur. Polym. J. 8, 117 (1972).

[41] Thompson, L. F., and Mayhan, K. G.. J. Appl. Polym. Sci. 16, 2291 (1972).

[42] Thompson, L. F.. and Mayhan, K. G.. J. Appl. Polym. Sci. 16,2317(1972).

[43] Brown, K. C. and Copsey, M. J., Eur. Polym. J. 8, 129

(1972).

[44] Kronick, P. L., Jesch, K. F., and Bloor, J. E., J. Polym. Sci. A-1, 7, 767 (1969).

[45] Tkachuk, B. V., et al, Vysokomol. soyed., A9-9, 2018

(1967),Transl. K. A. Allen.

[46] Wendel, C. T., and Wiley, M. H., J. Polym. Sci. A-1. 10,

1069 (1972).

[47] Thompson, L. F., and Smolinsky, G., J. Appl. Polym. Sci. 16, 1179 (1972).

[48] Bradley, A., and Fales, J. D., Chem. Tech., p. 232 (April

1973).

[49] Final Tech., Report F-C2260, The Franklin Institute Research Lab., (NIH) (1969).

12
NATIONAL BUREAU OF STANDARDS SPECIAL PUBLICATION 415,

Biomaterials, Proceedings of a Symposium held in conjunction witii tiie Nintii Annual Meeting of the Association for the Advancement

of Medical Instrumentation, New Orleans, La., April 19-20, 1974

(Issued May 1975).

Plasma-Formed Polymers for Biomedical Applications

Part II. Biocompatibility and Applications*

A. W. Hahn

John M. Dalton Research Center

University of Missouri, Columbia, Mo. 65201

K. G. Mayhan

Graduate Center for Materials Research and Chemical Engineering Department

University of Missouri — RoUa, Rolla, Mo. 65401

and

J. R. Easley and C. W. Sanders

John M. Dalton Research Center

University of Missouri, Columbia, Mo. 65201

Different polymer films generated by rf plasma techniques and deposited on glasses of varying

chemical composition, on implant alloys, and on formed prosthetic polymers have been implanted

in New Zealand white rabbits and various canine species and have shown minimum tissue reactions

after periods of time up to six months. It has further been found that the substrates upon which the plasma polymers are formed are more detrimental to cell cultures than the polymers themselves. These

findings, along with other implant work, indicate that plasma-formed polymers will play a definite

role as biocompatible materials in the future.

Key words: Biocompatibility; inflammatory response; plasma polymers; tissue reaction.

1. Introduction

In Part I of this joint paper [1],' Mayhan et al. have reviewed the fundamentals of the synthesis

and deposition of the plasma polymerized films on

various surfaces. Our studies were undertaken to explore the compatibility of these polymers with

biological tissues and some selected applications

to the fields of biology and medicine.

2. Compatibility Studies

Our investigation of compatibility with biological

tissues has taken two forms. First, the polymers,

coated on a suitable substrate, were exposed to tissue culture. To date we have conducted these

tissue culture exposures for the following plasma

polymerized materials: polyvinyl chloride, polyethylene, polyallene, polytetrafluroethylene, poly-

*Supported in part by grants from the National Science Foundation {#NSF GK

38957) and the National Aeronautics and Space Administration (#NASA NCR 26-

004-099).

' Figures in braci^ets indicate the literature references at the end of the paper.

chlortrifluroethylene, polystyrene, polyacyrilonitrile,

and polyvinyl fluoride.

The procedure used for studying these materials

was to obtain various thicknesses of coatings on

both glass and commercially available polystyrene cover slips especially made for tissue culture. These

5 X 10 mm shps were coated, on one side only,

under varying reactor conditions. The materials were then brought to our laboratory at University

of Missouri — Columbia where the glass cover sHps were exposed to autoclaving at standard conditions

for sterilization and the polystyrene cover slips

were exposed to ethylene oxide sterilization under

appropriate conditions.

Control shdes of both glass and polystyrene were

also sterihzed under identical conditions. Each of

the study materials was then placed in Leighton

tubes which had been previously sterilized, and 5 ml

of cells and a standard tissue culture media were added to each tube. The cells used were the HEP-2

cell fine which has been well developed and standardized for tissue culture studies. Approximately

10** cells were added to each tube. These were then

placed in a carbon dioxide incubator at 37 °C and

13
observed at both 24 and 48 h for cell attachment to the particular polymer coated (or control) sUde, for

attachment to the Leighton tube itself and for

metabohsm (i.e., or utilization) of the media in the

tube. The utilization of the media would determine

whether or not the cells were undergoing rapid

growth, free from any products which might inhibit

this growth.

Figure la is a low power photomicrograph of the

growth of these cells at 24 h in the control tube.

This particular tube (all tubes were done in triplicate) is an uncoated glass shde. Note the stellate

appearance of the cell types indicating good

attachment and growth at 24 h. Figure lb is a photomicrograph of the same tube at 48 h, again showing

a high concentration of the cells and the beginnings

of the development of a monolayer of cells over the

slide. During this time it was also noted that there

was change in the indicator to a more acid state

indicating utilization of energy products within the

media. The uncoated polystyrene control slides

showed the growth similar to those on the glass

sUdes.

Figure 1. Low power {X40) photomicrographs of HEP-2 cell growth in the control tube; (a) 24 hours post-inoculation;

(6) 48 hours post-inoculation.

Figure 2. Low power (X 40) photomicrograph of HEP-2 cell growth on a poly-TFE coated cover slip. This represents a

4 + toxic reaction.

Figure 2 is a photomicrograph of a polymer that

is considered to have some form of toxic byproducts

inhibiting both growth of the cells within the media

and their attachment to the slide. This is considered

to be a 4+ toxic reaction. There is a very sparse

number of cells throughout the media and all of

these have a rounded appearance. This rounded

appearance is indicative of a resting phase of the

tissue culture cells showing that there is some form

of growth inhibitor within the media.

Earlier preliminary studies showed that none of

the polymers that were subjected to this form of

screening markedly inhibited the growth of the

HEP-2 cells. Both the 24 and 48 h examinations

of the cells showed that the monolayers appeared

normal when compared with controls. An exception

to this rule was that polyallene material did show

a shght inhibition of growth.

With these encouraging preliminary data, these

studies were repeated on a new batch of polymers

that had been synthesized in Dr. Mayhan's laboratory by slightly different techniques. Five different

batches (TFE-3 and 4 in table 1) of polytetrafluroethylene were examined. It was found that at both 24 and 48 h there were few or no attached

cells, and there was little, if any, metabolism in

the tubes. Only one of these forms of poly-TFE

showed itself to be less toxic at 24 h, but at 48 h it appeared that there were fewer cells attached than

earlier. On the other hand, other forms of polymers

examined, with the exception of polyvinylHuoride,

showed themselves to allow normal attachment

and metabolism when compared with the controls

at both 24 and 48 h. AU of these data are summarized in table 1. These recent results were disturbing. At the

present time it is felt that the post-synthesis

treatment of, particularly, poly-TFE is of a high

order of significance. It would appear, although not

yet established conclusively, that a large number of

14
Table 1. Summary of tissue culture results

Material

Slide growth Tube ^1 ^ W 1 1

1

Vjijiiiii icii La 24 h. 48 h. 24 h. 48 h.

Control" A 0 0 0 0

Control—

B

0 0 0 0

TFE(l) 0 0 0 J TFE(2) 0 0 0 ] TFE(3) J 4 4 4 4 TFE(4) 3 1 3 4 PVC 0 0 0 0

PE(1) 0 0 0 0

PE(2) 0 0 0 0

PolyaOene 1 1 0 0

C-TFE 0 0

Polystyrene(l) — 0 — 0 Polymer peeled.

Polystyrene(2) 0 0

Acrilonitrile 0 0 0 0 Polymer surface

crazed. PVF 3 0

Code: 0; normal cell attachment to surface

1; approximately half of cells attached

2; approximately one-quarter of cells attached

3; most cells not attached

4; no cell attachment —; slow early growth

Polymer abbreviations:

TFE; Polytetrafluorethylene

PVC; Polyvinylchloride PE; Polyethylene

C-TFE; Polychlortrifluorethylene

free radicals are present in the plasma synthesized

polymer. These probably are responsible for the

inhibition of growth in the tissue culture medium.

Dr. Mayhan and his group at University of Missouri —

RoUa are presently pursuing these findings and

attempting to neutralize the active sites that may

be contributing to the tissue culture dilemma.

It is interesting to note that in the same reactor

and under nearly identical polymerization conditions, the polychlorotrifluorethylene, the polystyrenes and polyacyrilonitriles had httle adverse

effect on the rephcation and attachment of the

HEP—2 cells in tissue culture.

3. In Vivo Studies

Immature New Zealand white rabbits were used

to evaluate the tissue reaction to these polymeric

films. Various types of polymeric films were deposited on 1 X2 cm section of flat Silastic sheeting of medical grade quahty. Silastic was used

because of its universal use and known tissue

compatibility.

After anesthetizing the animals with Halothaneoxygen, the implants were sutured with 3/0 silk to the paravertebral muscle fascia 2 cm from the

dorsal midhne between the last rib and the sacrum. Two implants were placed in a retroperitoneal

position without sutures through a small paracostal incision made posterior to the kidney. Sterile

technique was used in all surgical procedures. Most

animals received four implants, two on either side

of the midhne approximately 5 cm apart. The

experiments are summarized in table 2. At necropsy the tissues and attached implants

were recovered and fixed in 10 percent neutral

buffered formalin. Thin tissue sections were embedded in paraffin and cut at 6 fim and stained

with hematoxylin and eosin.

Necropsy revealed no systemic effects to the

implants, and local gross tissue reaction was

limited to encapsulation by connective tissue.

The tissue reaction to the implants was deter2 Dow Corning, Midland. Mich.

^Certain commercial materials may be identified in this publication in order to adequately specify the experimental procedure. In no case does such identification imply recommendation or endorsement by the National Bureau of Standards, nor does

it imply that the material are the best available for the purpose.

Table 2

Implant Number of animals

Number of implantsLength of study (days)

Reaction

Connective

tissue

Inflammatory

Polyethylene

Polyvinyl chloride. TFE

Allene. SiaOene

Polypropalene

Control

4

1 144

3

12

403-432

432

296-418

403-418

296-352

209-230

296-432

2.0

1.0 2.8

1.5 2.0

2.9

3.3

1.0 0

1.3 1.5 0.3

2.2

1.4

The average connective tissue reaction was graded on a scale of 1-5 as follows: (1) Grade 1: « V4 low power field (Ipf) (lOOX) in diameter; (2) Grade 2: > V4 « V2 Ipf; (3) Grade 3: > V2 « »/4

Ipf: (4) Grade 4: > "A « l Ipf; and (5) Grade 5: > 1 Ipf. The inflammatory reaction was graded on a scale of 0-3 as follows: (1) Grade 0: minimal, « 1 cell/high power field (400X); (2) Grade 1: mild, > 1 « 3 cells/hpf: (3) Grade 2: moderate, > 3 « 10

ceUs/hpf; and (4) Grade 3: severe, > 10 cells/hpf. The grade was based on the average of 10 repre- sentative high power fields.

15
mined by measuring the connective tissue and

inflammatory cell response histologically. The

results are summarized in table 1. Figures 3a and

3b are photomicrographs of typical responses. In

all the experiments the connective tissue reaction

was less than that produced by the control (uncoated) strips, and the inflammatory cell reaction

to about two-thirds of the polymer strips was less

than that of the control.

Figure 3. Low and high power photomicrographs of a typical

polymer implant capsule given grade one connective tissue reaction and grade one inflammatory cell reaction; (a) 100 X

magnification of the capsule surrounding polymer implant.

The capsule is composed basically of acellular connective

tissue matrix with minimal inflammatory cell reaction.

The dense layer in the upper part of the field is the tissue in direct contact with the polymer implant: (b) 430x magnification of the dense layer of the same capsule. The

tissue is composed basically of acellular connective tissue matrix, but a few large inflammatory cells are noted.Applications

The results of our in vitro and in vivo studies

have shown that, while certain polymeric forms can

be rejected because of cytotoxicity, a number are acceptable as candidates for implant apphcations.

Certainly, the in vivo studies to date have revealed

a paucity of problems. With these encouraging

results, our attention was then directed to actual

applications for these types of unique polymers.

The biomedical scientist can immediately

conceive of numerous potential uses for this type

of material, ranging from the coatings of vascular

protheses (assuming antithrombogenic properties can be engineered), the protection of implantable

electronic devices and the coating of various types

of in vivo sensing devices. Efforts of our laboratories

have concentrated on the coating of implantable

oxygen sensing electrodes in order to obviate the

problems inherent in the use of polarography; i.e.,

long-term instability of 02-sensing surfaces and

poisoning of the electrode surfaces.

Principally through the efforts of our colleagues,

R. E. Barr and V. G. Murphy [2], our laboratory has

undertaken a group of studies aimed at a better

understanding of the electrochemical phenomena

taking place at these sensing surfaces, in vivo,

and at the use of the conformal polymerized coatings

to obviate certain of the problems encountered.

A particular problem with 02-sensing polarographic electrodes which utilize O2 reduction on

noble metal surfaces has been with the poisoning

of these surfaces by macromolecules, particularly

those associated with the sulfur containing amino

acids. It was felt that if these surfaces could be

protected with a membrane that was permeable to oxygen, but impermeable to the large macromolecules, a far more stable sensing electrode for

modestly long-term totally implantable use might

be possible.

Figure 4 is a low power scanning electron micrograph of the particular form of oxygen sensing

electrode that we are presently using. This is a very

thin platinum wire which has been coated with

glass. It is the flat surface that is coated with the

polymer film.

Figure 4. Low power {X300) scanning electron micrograph

(SEM) of an uncoated platinum-iridium oxygen electrode used in the studies.

The sensing surface is the disc in the center of the photograph.

16
60 60

(a)

10

10

(b)

10 8 6

TIME (min)

10 8 6

TIME (min)

Figure 5. Current generated from Pt-Ir oxygen electrodes during stirring; (a) uncoated (Note the high noise level of the current);

(b) coated (poly-ethylene) electrode of the same design.

While the magnitude of the current has decreased, the noise level is now quite acceptable.

We have found that by coating this sensing

surface with a uniform polymer membrane that is tough and firmly applied, we can almost negate the

starting artifact that is present around each of

these electrodes. Figures 5a and 5b show a recording of the currents generated by polargraphic

oxygen electrodes with both uncoated and coated

sensing surfaces. Note that the uncoated sensing

surface (fig. 4a) when placed in a media which is stirred, is strongly sensitive to stirring artifact;

in other words, the platinum surface is seeing

random concentrations of local oxygen. With the

polymer film on the surface, it may be noted in

figure 5b that, while the overall current is cut down

in magnitude, there is almost no noise present in

this particular recording. This is due to the uniform

oxygen boundary layer that the oxygen reducing

surface of the platinum sees rather than being

bombarded by random concentrations of oxygen

even in a highly turbulent medium. The useful Ufe

of Oa-sensing electrodes is prolonged to days rather

than hours by the use of the film coatings.

While there is certainly a large amount of work

remaining to be done in the area of precoating

preparation of these electrodes and of their durability

in long-term tissue applications, there is much to be gained by the application of a uniform film to the

sensing surfaces to allow for protection of the

sensing surface from tissue macromolecules and

other poisoning effects. It is our opinion at this time

that the potential for oxygen sensing in tissues,

insofar as long periods of time is concerned, is greatly enhanced by the apphcability of these

particular forms of membranes.

4. References

[1] Mayhan, K. G., Havens, M. R., Peace, B. W., Hahn, A. W.;

Plasma Formed Polymers and Biomedical Applications —

Part I: Synthesis and Fundamental Studies, this volume,

page 1.

[2] Mayhan, K. G., Hahn, A. W., Barr, R. B.; A New Approach

to Coating Oxygen Electrodes. Proc. 25th A.C.E.M.B.

(Annual Conference on Engineering in Medicine and

Biology), Miami Beach, Fla., p. 85, 1972.

17

NATIONAL BUREAU OF STANDARDS SPECIAL PUBLICATION 415: Biomaterials. Proceedings of a Symposium held in conjunction with the Ninth Annual Meeting of tiie Association for the Advancement of Medical Instrumentation. New Orleans. La.. April 19-20. 1974

(Issued May 1975).

Interfacial Behavior of Ceramic Implants

L. L. Hench

Department of Materials Science and Engineering

University of Florida, Gainesville, Fla. 32611

H. A. Paschall

College of Medicine

J. Hillis Miller Health Center and Chief of Orthopaedics

Veterans Administration Hospital, Gainesville, Fla. 32610

W. C. Allen

College of Medicine

Department of Orthopedic Surgerv

J. Hillis Miller Health Center, Gainesville, Fla. 32611

and

G. Piotrowski

Department of Mechanical Engineering,

University of Florida, Gainesville, Fla. 32611

Recent studies of bioglass, bioglass-ceramic. and alumina implants have produced an understanding of the chemical nature of interfacial tissue reactions to bioceramics. Significant differences between hard and soft tissue reactions are due to the influence of surface chemical reactivity on the

ultrastructurai histology as revealed by transmission electron microscopy. Modern surf ace characterization of the implants correlates with the histological reactions. Applications of the results to a variety of orthopaedic prostheses show promise in animal experiments. Biomechanical analyses of interfacial bonding between bioceramic prostheses and tissues are presented.

Key words: Bioglass-ceramic: ceramic implants: flame spray coating: hip prosthesis: segmental bone

replacement.

1. Introduction

Two years ago a review article on ceramics in

medicine by one of the authors reported that such

materials are seldom used for chronic applications

in medicine [IJ.' However, pioneering efforts of the

last few years have changed this situation. Dr. P. Boutin of Pav, France recently reported over 500

successful cases of total hip replacements using

high density, high purity AI2O3 components over a 4 year period [2J. Application of the alumina

ceramic joint was based upon studies of very low

wear rates of alumina in contact with alumina when

both ball and socket are finely polished to a 1 /xm

surface finish. Biocompatibifity of the high purity

(99.3%) alumina was first established in dogs and

humans using both acicular and button configura-

' Figures in bracl^ets indicate llie literature references at tlie end of tiie paper.

tions of the ceramic. Only a thin fibrous membrane

several cell layers deep was present at the implanttissue interface. Similar results of tissue response

and low rates of wear of dense high purity AI.Os

implants have been demonstrated bv a West German

research team [3-5]. These efforts have led to a systematic development of a total hip prostheses

that is all AI2O3 ceramic. This is in contrast to the

Boutin prostheses that utihzes a Ti metal femoral

stem pressure fitted onto an alumina ball for the

femoral portion of the hip prostheses.

A combined West German and Austrian research

team reports total replacement of vertebrae,

humerus, and radius in humans with high density,

high purity AI2O3 ceramics [6J. Successful use of the

implants for times in excess of 1 year are recorded.

With application of ceramics in humans commencing, it becomes imperative to understand

fully interfacial reactions between ceramic materials

19
and bone. The studies reported by Hulbert and

Klawitter et al. [7-10], and those cited above have

demonstrated that very close apposition of mineralized bone occurs at the implant interface. However,

if dense implants are used, taking advantage of the

40,000 psi (0.28 GPa) tensile strength of high

density alumina, a thin fibrous membrane is always

observed. The presence of the membrane causes one to reflect on the possibility of implant loosening

during long term application with attendant stress

locahzation and bone resorbtion currently characteristic of long term metallic implants. Will the

membrane remain firmly attached to the ceramic

interface or will interfacial fluids develop with

stress-corrosion fatigue at the ceramic grain

boundaries? Can compositional, microstructural,

or surface modifications be made to the alumina

ceramics to avoid fibrous membrane formation

without degrading strength? Answers to these and

other fundamental questions are only beginning to be obtained. Alumina with 200 /xm pore diameters

and 50 percent porosity appear to be hmited to 10,000 psi (0.07 GPa) tensile strength. Will 10,000

psi tensile be strong enough at the interface to transmit the loads required in weight bearing

apphcations? Successful load bearing segmental

bone replacements tested in monkeys in our laboratory using a bioglass-ceramic of less than 10,000

psi tensile strength suggest that mechanical design

hmits for such materials are presently not understood [11]. In other words, we do not know how

strong a ceramic prosthesis must be to withstand

chronic applications in humans. The data in animals

is only beginning to be achieved in spite of much

recent effort [1-15]. Much additional effort directed

toward understanding interfacial reactions, strength

requirements, and strength deterioration [16-19]

must be completed before large scale chronic

human apphcations seem warranted.

Alternative directions towards achieving satisfactory ceramic prostheses also continue to show

promise. The research of Graves et al., at the University of Dayton has demonstrated considerable

control over osteogenesis in monkeys using variable

P2O5 content in calcium-aluminate ceramics [12].

Strength enhancement in this system also seems

quite probable.

The efforts of our laboratories have concentrated on the investigation of specially designed bioglass

and bioglass-ceramic compositions [20-25]. This

research will be reviewed in the remainder of this

paper. Special emphasis will be placed upon the

results of combining the osteogenic features of the

surface active bioglass system with the high

mechanical strength of 316L surgical stainless steel.

The results indicate that flame spray coatings of

bioglass on stainless steel or high density AI2O3 may

provide a solution to many of the interfacial tissue

and strength problems described above.

2. Bioglass-Steel Flame Spray

Concept

Previous studies in this laboratory have resulted

in the development of a nonporous glass and glassceramic composition which is compatible with bone

on an ultrastructural scale [20-25]. The bioglass

and bioglass-ceramic contains Ca, Na, and P ions

(table 1) which are slowly leached from the surface

of the implant material during the healing process.

Table 1. Bioglass compositions developed for orthopaedic

prostheses

Code number

Weight percentage

Si02 NajO CaO CaF2 P2O5 B2O3

45S5

45S5F

45B,5S5

45.0

43.0

30.0

24.5 23.0

24.5

24.5

12.0 24.5

16.0

6.0

6.0

6.0 15.0

The surface ion release leaves behind a polymerizing

SiOrrich gel [22, 23, 26, 27]. The surface gel

appears similar in composition to that preceding

hydroxyapatite mineralization in normal bone

growth [28]. As a consequence of the surface

activity, collagen bonding and bone minerahzation

are initiated at the implant surface [20-25]. The

bone growth from the implant-bone interface results

in an interfacial chemical bonding which is suffi- ciently strong that the bone or the implant will

fracture before the interface does [11].

Although the tensile strength of the bioglassceramic material is in the range of 10,000 psi when

crystalhzed in a controlled manner [29], there are numerous prostheses apphcations where even

greater strength or impact resistance may be

desirable. This is especially true for plates, nails,

pins, and femoral head shafts where a small cross- sectional area of the implant is required. Stainless

steel devices satisfy the strength requirement for

these applications. However, a fibrous tissue

reaction at the steel interface results in loosening

of the implant with time and corrosion of the implant

produces metallic particles which become incorporated in the neighboring tissues.

An investigation was conducted to combine the

superior strength and impact resistance of surgical

stainless steel with the biocompatibility of Na, Ca,

P containing bioglasses. This combination is achieved by applying the bioglass as an intimate

coating on the surface of stainless steel prostheses.

The coating is appUed by a flame spraying method

developed in collaboration with the ceramics group

at nTRI.3 This paper presents the development of

^ Illinois Institute of Technology Research Institute. Ceramics Division, Chicago,

111., S. W. Freiman and E. J. Onesto, Investigators.

20
the bioglass flame spray coating process, the design

of a flame spray coated partial hip prosthesis,

surgical results, and a histological comparison of

tissue responses adjacent to the partial hip prostheses with those of a standard surgical stainless

steel screw.Flame Spray Process

Flame spraying has been in use for well over half

a century in the apphcation of metallic coatings

while efforts to apply ceramic coatings, with emphasis on refractory crystalHne oxides, began about 20

years ago. The flame spray process consists of

passing a material in powder form through a flame

where it is heated to a temperature at which it becomes plastic and is readily deformed. The

molten material is then impinged onto a substrate

at high velocity with consequent deformation and

adherence of the particles. The combination of

heat and impact which is available in flame spraying

permits rapid particle-to-particle sintering at temperatures and times where normal sintering

would not be obtained. Consequently, the substrate

does not reach a high temperature. For this reason,

substrates can be flame spray coated without

deformation and loss of tolerance. Diffusion

processes between the substrate and the coating

are also retarded. Both of these factors are important in producing a high quality coated medical

prosthesis.

The spray efficiency and the quahty of adherence

of a flame sprayed coating depends on many factors,

but of particular importance are the thermophysical

properties of the sprayed material, and the surface

conditions, thermal conductivity and thermal

expansion of the substrate. For a given material

and substrate, spray conditions are readily optimized empirically as to spray distance, gas flows,

powder feed rate, etc.

3. Flame Spray Equipment

The flame spray apparatus consists of an oxyhydrogen torch*'^ in conjunction with a powder feed

hopper, shown schematically in figure 1. The oxygen

line to the torch passes through the feed hopper

which is equipped with a variable speed screw-feed

mechanism to regulate powder flow. The powder is entrained in the oxygen stream and passes through

the torch. The powder laden oxygen is combined

with the hydrogen at the torch, the powder passing

through the resulting oxyhydrogen flame. Since the

theoretical flame temperature is approximately

3000 °C, the powder is rapidly heated and becomes

molten.

VotiiiH tD=

* Continental Coatings Corporation. Certain commercial materials and instruments may be identified in this publication

in order to adequately specify the experimental procedure. In no case does such identi- fication imply recommendation or endorsement by the National Bureau of Standards,

nor does it imply that the equipment or instruments identified are necessarily the best

available for the purpose.

Figure 1. Schematic representation offlame spray apparatus.

4. Materials Factors in Flame

Spraying Glass

When the material being sprayed is a glass, a special situation exists. Glasses differ markedly

from metals and most crystalline ceramics in that

they display high optical transparency and relatively

high viscosity above the softening temperature.

These differences are very significant to the flame

spraying process, as was observed in previous work

at IITRI [30].

As the particles are projected through the flame,

they are heated by convection from the hot gases

and by radiation from neighboring particles. The

lack of the radiation component to heat transfer in

the case of transparent glasses results in the evident

low temperature of the glass particles relative tc spray particles of opaque metals and ceramics. On

the other hand, too high a particle temperature may

destroy the glass, particularly compositions containing phosphate such as the bioglasses. This is due to the high vapor pressure of the phosphate

groups. Also, the low thermal emittance of the

particles imparts a slow coohng rate, hence more

deformation and flow after impact (provided that

the viscosity is sufficiently low).

The long softening range of glasses is advantageous where the substrate is hot or where the

thermal conductivity is low because flow can continue for a longer period, resulting in better adherence and a tighter structure. A glass with high

viscosity will not flow even when well over the

softening temperature. The more rapid the quench

rate of the particle, the lower the viscosity must be.

It has been found that a low viscosity is required

for sintering as well as for flow of glass particles.

Because of the smoothness of the molten glass

surface the free energy is small, and the material

must flow in order to sinter. This situation is not

observed with metal and ceramic particles at high

21
temperature where surface defects give rise to a high free energy for sintering [31].

The relationship between softening range, quench

rate and viscosity of a glass important to flame

spraying is shown in figure 2. The volume decrease

associated with the cooling of the molten droplets

is shown as dark curves in the figure. If the liquid

does not crystallize upon cooling to T,,,-, it can retain

its random, liquid-like structure through a supercooled hquid regime, B, and solidify into a glassy

material within the transformation range, Tg. The

transformation temperature of a given flame sprayed

glass is dependent upon its quench rate. A rapid

quench produces a high Ty and a less dense glass, curve C. Likewise, a slower quench rate results in

high densities and lower transformation temperatures, curves D and E.

Figure 2. Volume-temperature relations for a liquid (A), supercooled liquid (B), glasses (C, D, E), and crystal (F), all shown in solid lines.

Viscosity-temperature dependence for two glasses shown as dotted curves.

If a glass has a high viscosity before the glass

transformation occurs, such as the dashed curve #1, the flame sprayed particles cannot flow and

create a uniform glassy coating on the substrate.

However, when the viscosity of the glass is low (curve

#2), spreading and interparticle flow occurs prior to solidification of the glass. This condition must be

achieved by adjusting the glass composition. There

is a hmit on compositional adjustment, however,

since crystallization at Tm will occur if sufficient

three-dimensional network formers are not present

in the molten glass particles. Because of the large

volume change associated with crystallization

(curve F) it is very difficult to obtain a smooth,

coherent crystalhne coating.

To supplement the low viscosity, a low surface

tension of the glass is also required for good wetting

of the substrate. Particle flow and sintering after

impact will be a function of both viscosity and

interfacial tension.

It is most important that the coefficient of thermal

expansion of the glass sprayed on a relatively cold

substrate be nearly the same as the substrate or

slightly lower. A lower expansion will provide

thermal shock resistance and a minimum of tensile

stress estabUshed within the glass when the coating

shrinks on cooling. However, as shown in figure 3,

if there is large expansion mismatch between the

glass and the substrate, tensile stresses wiU develop

at the interface when the glass cools below its transformation temperature. Cracking will result

if the stresses become too large [32].

Figure 3. Development of interfacial stresses between a flame

spray coating and a prosthesis due to thermal expansion

mismatch.

5. Development of Bioglass Composition for Flame Spraying

Initial flame spraying experiments were con- ducted using flat plates of 316L surgical stainless

steel as the substrate for the 45S5 bioglass (table 1)

originally developed for bulk glass-ceramic prostheses. All substrates were grit blasted with 20

mesh SiC at a pressure of 80 psi to completely

roughen the surface before spraying. The devices were then thoroughly scrubbed in a detergent

solution and rinsed first in water and then in

isopropyl alcohol to remove any grit, blasting debris,

dirt and oils picked up in handling the devices.

A variety of experimental variables were ex- amined in an effort to achieve a high adherence

coating of the 45S5 bioglass [33]. None succeeded.

It was concluded that the major reason for the lack

of success in flame spraying the 45S5 glass was its very high viscosity at the spraying temperature.

This high viscosity prevented the sprayed material

from flowing and forming an adherent glassy layer.

The high viscosity of the glass is probably due to its relatively high silica content. Previous studies had

shown that glasses which form the best flame

sprayed layers are those that are low «5%) in

SiO2[30].

This interpretation was borne out by an analysis

of the thermal expansion coefficient of the 45S5

bioglass and the dilatometric softening point. As

shown in figure .4 (determined on 2-inch specimens

with an Orton Automatic Recording Dilatometer),

the expansion of the 316L stainless steel substrate

and the 45S5 bioglass are well matched. However,

the softening of the glass under stress, revealed by

the hump in the curve, occurs at a relatively high

temperature (530 °C) for this composition. Thus,

the glass viscosity is too high for a successful flame

sprayed coating.

22
Two approaches were taken to decrease the glass

viscosity; substituting B2O3 for Si02 in the glass

and CaF for CaO (table 1). Both changes decreased

the viscosity as shown by the lower softening points

in figure 4. The decrease in softening point to 470 °C

was the same for both the fluorine and boron containing bioglasses. The magnitude of the thermal

expansion coefficient was also the same for both

compositional series. However, the higher temperature fluidity of the fluorine-containing bioglass

45S5F was less under the flame spraying conditions

and produced a smooth, uniform glassy coating on

the stainless steel with excellent adherence.

Since it is required that implant devices be

virtually encapsulated in a glass coating, it was

necessary to flame spray each device individually.

Samples were hand held in forceps modified to grip them with minimum contact. In this manner.

front, back, and interior surfaces (in the case of

cylinders) could be uniformly preheated and coated

with a well fused, glassy appearing coating.

Coatings on the prostheses were obtained by

preheating the substrate above the softening point

of the glass, then initiating the powder feed through

the torch and allowing the glass particles to flow,

rather than quench, as they struck the surface.

Altering the bioglass composition in order to obtain suitable flame spray behavior does not yield a unilateral change in properties. Consequently,

evaluation of solubility characteristics and in-vivo

responses of the new compositions were necessary

before it was justified to produce and test coated

prostheses. Previous publications show that the

in-vivo behavior of both the fluorine and boron

containing compositions were suitable for prostheses

applications [21-25].

0 100 200 300 400 500 600 700

Figure 4. Thermal expansion and softening points of 316L stainless steel and

three bioglasses.

23
6. Bioglass Flame Spray Application

to Prostheses

Several configurations of prostheses and test samples have been flame coated with the 45S5F

and 45B15S5 glass compositions discussed above.

Figure 5 shows an array of flame spray coated

316L stainless steel prostheses and test samples.

Several bone bridges, rat cortex inserts and a tissue

culture blank are shown. Both the 45S5F and

45B15S5 bioglasses have been used for flame

spraying such samples.

Figure 5. A variety of flame spray coated prostheses, iSBi^SS

composition. On the left are segmental bone replacements for

rats; at center is a tissue culture substrate, and at right are

rat bone chips.

A femoral head replacement for a Rhesus monkey

was selected as the test model for a major prosthetic

implant made of 316L stainless steel and flame

sprayed with a bioglass coating. The objective of

such a composite design is to combine the strength

of the stainless steel core with the biocompatible

surface properties of the glassy coating.

A prosthetic femoral head replacement is comprised of three major segments: the head, which

replaces the head of the femur, the stem, which

anchors the device in the shaft of the femur, and

the junction between the head and the stem. These

three parts are pointed out in figure 6, which is a photograph of the completed prosthesis. The head

ideally should be an exact copy of the femoral

head, but this is not practical. It was therefore

decided to manufacture the head of the prosthesis

as a sphere, with a diameter closely matched to the

mean diameter of the femoral head.

The stem of the prosthesis anchors the device by

fitting snugly into the marrow cavity of the femur.

The stem needs to be quite long, so that bending

loads may be resisted with a minimum of contact

stress developed at the extremes of the stem. The

stem should not be circular in cross section, so

that the prosthesis cannot rotate about the long

axis of the stem. The stem must also be designed

to allow itself to be driven into the medullary canal

without splitting the femoral cortex. The junction

between the head and the stem must not interfere

Figure 6. A prosthetic femoral head replacement for a Rhesus

monkey is labeled to show the three segments of the device.

The stainless steel prosthesis has been flame sprayed with a glass-ceramic coating.

with muscle attachments at the greater and lesser

trochanters. The junction also prevents the prosthesis from migrating distally along the shaft of

the femur.

7. Design of the Prosthesis

The geometric constraints imposed by the monkey's anatomy are illustrated in the x-ray (fig. 7). The head appears to be a sphere of about 14 mm

(0.56 in). The inner diameter of the marrow cavity

is about 6 mm (0.24 in) at a point about halfway

down the femur. The proximal end, being narrower,

is to be reamed out to an inner diameter of about

6 mm as well. Since the smallest outer diameter at about the level of the isthmus, is over 10 mm (0.40

in), the proximal femoral cortex is not greatly

weakened by the reaming of the bone.

To prevent rotation of the stem within the marrow

cavity, the stem was designed with a square cross

section, 5 mm (0.20 in) square. Since a diagonal of

the cross section is 7.1 mm (0.28 in) long, the stem

must be driven into the marrow cavity and cut four

grooves, triangular in shape and approximately 0.5

mm (0.02 in) deep, into the endosteal surface. The

end of the stem was cut off square, essentially

providing four cutting corners with zero rake angle.

24
Figure 7. The design for the prosthesis was evolved directly on

this x-ray of a Rhesus monkey pelvis.

Note that the design of the device does not interfere with muscular attachments to the lesser and greater trochanters.

The length of the stem was made as great as possible. A length of 43 mm (1.69 in) was selected; a longer length would have no effect in terms of

anchoring the prosthesis since the distal portion of

the stem would have protruded into the flared region

of the distal femur and not be in contact with any

cortical bone.

The junction between the head and stem was

designed so that both the greater and lesser tro- chanter would be left undisturbed. Thus, neither

the major flexors nor abductors of the hip would be

affected.

The thickness of the ceramic coating of the prosthesis is about 0.5 mm (0.02 in). Therefore, the metal

portion of the device is 0.5 mm smaller all around,

as shown by figure 8. Note that the complete device

fits entirely within the original volume of the femur.

Thus the range of motion available after implantation is not compromised by the design of the implant.

8. Mechanical Factors

13.5

3.0 4.0

All dimensions in mm.

Figure 8. The final design of the metal portion of the prosthesis

compensates for the ceramic coating, which is of the order of

0.5 mm thick.

The thickness of the stem and junction was reduced to 3.0 mm to reduce the danger of

splitting the femur during insertion.

Once the geometry of the implant was established, a series of calculations were performed to compare the strength and rigidity of the intact

monkey femur and the prosthesis. The calculations are based on the following assumptions. (1) The

femur is a hollow cylinder of constant cross- sectional geometry, with an outer diameter of 10 mm

(0.40 in) and an inner diameter of 5 mm (0.20 in).

Use of this model to determine the mechanical

parameters of the femur assures that the values of

strength and rigidity will be lower than the actual

values. (2) The modulus of elasticity of the bone

was estimated at 2 X 10^ kgf/cm^ (2.5 X 10« psi),

while the ultimate tensile strength was estimated

to be about 1000 kgf/cm^ (15,000 psi). (3) The effect

of the ceramic coating on the strength and rigidity

of the prosthesis were ignored, which is justified

considering the small thickness and low modulus

of the coating.

The rigidity of the stem of the prosthesis, defined

as the product of the modulus of elasticity and the area moment of inertia, was computed to be of the

order of 4500 kgf-cm^ (1530 Ib-in^), while the rigidity

of the femoral shaft is approximately 8250 kgf-cm^

(2760 lb-in2).

Thus, the shaft of the monkey femur is approximately twice as rigid as the stem of the prosthesis

in bending. Bending loads applied to the femur will

therefore be primarily carried by the bone rather

than the implant. This tends to inhibit disuse

osteoporosis in the femur.

Based on a yield strength of 5300 kgf/cm^ (75,000

psi) for 316L stainless steel, the strength of the

stem of the prosthesis is approximately 57 kgf-cm

(50 in-lb). The femur, however, can withstand a bending load of 132 kgf-cm (112 in-lb) before

fracture. Thus the bone is also twice as strong as the implant. The superior strength and rigidity of

the bone, despite the lower modulus of elasticity

and failure stress, arises from the fact that the bony

25
material is distributed much farther away from the

axis of the femoral shaft than the implant material.

It is generally accepted that a load of about three

times body weight acts on the head of a human

femur during single-legged stance. If this same

proportion is applied to the Rhesus monkey, a bending load of about 21 kgf-cm (18 in-lb) is seen by the prosthesis at the upper end of the stem. The

maximum contact stresses developed by the junction bearing on the bone just above the lesser

trochanter are of the order of 70 kgf/cm^ (1000 psi),

far short of the failure stresses of either bone or

steel.

If contact between the stem and the femur is assumed to occur in a strip 0.5 mm (0.02 in) wide

along each edge of the stem, a torque of 29 kgf-cm

(25 in-lb) is required to raise the contact stress to 700 kgf/cn# (10,000 psi). The muscles around the

hip are not capable of producing such torsional

loads. Thus this mode of failure would come about

only through accidental application of very large

external forces.

Figure 6 shows the device after coating with the

45S5F bioglass. Most of the surface is covered with

a clear transparent glass which appears mottled

and gray because of the oxidized steel surface

underneath. A portion of the end of the femoral

shaft shown in figure 6 is more transparent and

therefore gives the illusion of the coating being

missing.

9. Surgical Procedure

The bioglass flame sprayed monkey hip prosthesis was placed in two stump tail monkeys for a period of 14 months. The monkeys were adult

females taken from the colony of stump tails at the

animal farm and operated on under aseptic surgical

conditions.

After the monkeys were anesthetized, their right

hips were shaved and prepped in the usual manner,

and a posterior lateral approach used to expose

the hip. Through this approach, the gluteus maximus

was partially split with the majority of the muscle

retracted toward the midline. The gluteus medius

and minimus muscles were then defined and re- tracted superiorly. This allows visualization of the

sciatic nerve, which is located just posterior to the

greater trochanter. The nerve was protected and

the short rotators of the hip cut through an arc of

180° to allow exposure of the femoral head and the

labrium cut transversely to allow dislocation posteriorly. Following good hemostosis, the base of the

neck was osteotomized, and using a small Hoke

osteotome the greater trochanter was notched so

as to accomodate the neck of the prosthesis. The

femur shaft was reamed using a twist drill and the

prosthesis introduced into the femur. With proper

notching and reaming, the prosthesis was seen to

fit snugly. This is shown in figure 9. The prosthesis

was then reduced into the acetabulum, and the

labrum and capsule repaired using chromic sutures.

The short rotators were sewed to their insertion

making no particular effort to reposition them

anatomically. The wound was then closed in layers

using absorbable sutures on the skin.

Figure 9. Bioglass flame spray coated femoral stem replacement as inserted in femur at time of operation.

Although the procedure is analagous to the human

procedure, there are two notable differences. These

are: (1) the close proximity of the sciatic nerve

to the posterior capsule of the hip, and (2) the shape

and more proximal position of the greater trochanter

relative to the base of the neck of the femur.

10. Histological Evaluation

The objective of this part of the study was to evaluate the histological response of the soft tissue

reactions about one of the bioglass flame spray

coated partial hip prosthesis. The histology of the

soft tissue removed from a corroded area about a

metallic bone screw on a fixation device is presented for comparison.

Soft tissue used to demonstrate the effects of

corrosion upon the soft tissues was taken from an

area of black discoloration of the soft tissues

beneath a bone plate and adjacent to the bone sur- rounding a screw which demonstrated corrosion

with metallic scarring between the screw and the

plate (fig. lOA). The internal fixation device had

been used in a 58-year old woman who had undergone an osteotomy of the femur for degenerative

arthritis of the hip. The implant had been in place

for approximately 18 months at the time of removal.

The second specimen was soft tissue removed

from the bioglass flame spray hip prosthesis described above at the location shown in figure lOB.

The prosthesis had been experimentally inserted

as a replacement device in a stumptailed monkey,

and had been in place 14 months at the time of

26
Figure 10. Schematic of the location of soft tissues for histology

evaluation.

A. Tissue adjacent to a stainless steel screw.

B. Tissue adjacent to the bioglass flame sprayed partial hip prosthesis.

sacrifice. This soft tissue showed no evidence of

discoloration and appeared healthy and normal in

all respects.

In each instance the tissue was taken fresh from

the operating room and placed immediately in cold

buffered gluteraldehyde. Tissues were fixed for 2

hours in the gluteraldehyde, then washed with

additional buffer and fixed an additional hour in

1 percent osmium tetroxide buffered to a pH of

7.4. The tissues were then dehydrated in graded

alcohols and then embedded in Epon 812^.

Sections one micrometer in thickness were

taken for routine histology and orientation for

sectioning for electron microscopy. The blocks were then further trimmed for electron microscopy

and the blocks were sectioned on an MT-2 Porter

Blum microtome using a diamond knife, and the

sections collected on carbon-stabilized collodianShell Chemical Company.

coated grids. The sections for electron microscopy were stained with lead citrate and uranyl acetate.

Figure 11a is a photomicrograph of soft tissue

surrounding the corroded metal screw. This is normal dense collagenous connective tissue and

the cells which are seen are fibroblasts. However,

one notes in the fibroblasts that there are dense

collections of dark staining material which one may assume to be metal particles.

Figure lib is a photomicrograph of soft tissues

surrounding the bioglass coated implant. This tissue

is also dense collagenous connective tissue and one

sees numerous fibroblast and fibrocyte-like cells

surrounded by a dense collagen matrix. These

cells do not demonstrate any abnormal inclusion

at this magnification.

Figure 11. a. Light micrograph of reaction to metallic implant. Note black particle in cytoplasm offibroblasts. (1,300 X)

b. Light micrograph of area surrounding bioglass-coated implant.

(1,300 X)

27
Figures 12 through 15 are representative electron

micrographs of the soft tissue seen in figure 11a.

With the resolution afforded by electron imaging, one can now easily discern that the dark staining

inclusions within the fibroblasts are made up of

aggregations of metallic particles. These particles are enclosed in smooth-walled vesicles. These

particles are very densely packed in some of the

vesicles so that one cannot distinguish the individual

particles but in others there is less dense packing

and the individual metal particles are easily visualized.

Figures 16 through 19 are representative electron

micrographs of the soft tissue removed from the

bioglass-coated implant. In this tissue the cellular

morphology is undisturbed. The cells are all typical

of fibroblasts actively producing collagen or fibrocytes which are mature and which are enclosed in a very dense collagen matrix. In none of the sections

of this tissue were there any inclusions within the

cytoplasm containing inorganic crystalline material.

From these observations one can conclude that

the material removed from the surface of the

metallic implant by corrosion is taken into the

fibroblasts of the soft tissue surrounding the

implant. These metallic particles are apparently

Figure 12. Fibroblast-like cell containing aggregates of

metallic particles in cytoplasmic vacuoles. (23 ,000 X)

Inset — High magnification electron micrograph to demonstrate metallic particles within membrane-bound vacuoles. (55,000 x)

Figure 13. Multiple smooth-walled vacuoles containing metallic

particles.

Note density of packing in some vacuoles. (43,750 X)

Figure 14. Fibroblast containing two dense aggregations of

metallic particles. (23,900 X)

/f V

/

CM

Figure 15. Metallic particles from implant enclosed in smoothwalled vacuoles. (23,200 X).

28
Figure 16. Representative field from tissue surrounding

bioglass-coated implant.

Note absence of cytoplasmic vacuoles containing inorganic material. (12,600 X)

Figure 19. Fibroblast with normal-appearing cytoplasm.

Note abundant rough endoplasmic reticulum without cytoplastic inclusions of inorganic

material. (33,000 X)

Figure 17. Mature fibrocyte with large nucleus and scant cytoplasm.

Cell is completely surrounded by mature collagen fibers cut both longitudinally and in cross-section. (21,700 >0

insoluble and are removed as foreign bodies and

packeted in the fibroblasts. It is noteworthy that

in the soft tissue surrounding the bioglass coated

implant, no similar inorganic particles were noted

within the cytoplasm of the cells. This would suggest that any elemental constituents released from

the ceramic are totally soluble in body fluids and

are released into the circulation rather than re- tained within the cells surrounding the implant.

Another obvious conclusion is that the stainless

steel implant was protected from corrosion and that no metaUic particles were released to be phagocytized by the cells of the surrounding soft tissues.

This report demonstrates a histological comparison between soft tissues removed from a metallic

implant and about a bioglass coated stainless

steel implant. The evidence supports the conclusion

that by coating the stainless steel with a low re- activity biocompatible glass coating, the stainless

steel implant is protected from corrosion. Complete

mechanical stability of the bioglass flame spray

implant after 14 months also demonstrates the

potential of using biocompatible coatings to provide

orthopaedic device stabilization without resorting

to polymer injection. No fibrous capsule was

observed at the bioglass coating-bone interface

and the implant could not be forcibly extracted

from the femoral shaft.

Key to Symbols

Figure 18. Mature fibrocyte.

Smooth-walled vesicles do not contain any inorganic material. (29,0(X) X )

c = collagen

CM = cell membrane

ER = endoplasmic reticulum

FB = fibroblast

MP = metallic particles

N = nucleus

n = nucleolus

SM = smooth membrane

V = vacuole

29
11. Development of a Segmental

Bone Replacement: Rats

Application of the bone-bonding bioglass-ceramic

material as a segmental replacement for bone has

involved three stages of development. The first study involves the use of small bioglass-ceramic

washers, ~2 mm thick, as a replacement for a total cross section of rat femur. In the second study,

larger segments of bioglass-ceramic, ~7 mm long,

were used to replace total cross sections of monkey

femora. In the third program bioglass flame spray

coated stainless steel segments, —10 mm, were

evaluated. Comparisons of the interfacial strengths

developed for the three types of implants, the

animal models, and the strength data of the opposite

bones of the femoral pairs follow.

The rats were anesthetized with sodium pentabarbital. After a longitudinal incision was made in

the right thigh, the femur was exposed by separating

the muscles along a fascial plane. The femoral

diaphysis was cleaned of soft tissue at the midpoint and osteotomized with a circular dental saw. A Kirschner wire was inserted into the medullary

canal of the distal half of the femur, drilled through

the condyles, and bent to a right angle with a Vs

in (3 mm) leg at the distal end. The fracture was

reduced after placement of a bioglass-ceramic

washer over the wire and against the osteotomy

surface by inserting the proximal end of the wire

into the medullary canal of the proximal fragment

and drilHng it through the greater trochanter. The

bent end of the wire was placed so it ran subcutaneously down the anterior surface of the tibia

and between the lateral and medial condyles of the

femur. Apposition of the fracture site was achieved

by pulUng on the proximal end of the wire, which

protruded from the bone (fig. 20, x-ray). The skin

was closed with metal wound staples after apposing

the muscles over the fracture site with two silk

sutures.

Figure 20. A typical animal specimen from the initial series of

rat femoral segmental replacement experiments.

Sixteen rats received the bioglass-ceramic

washers, which completely separated the two

femoral fragments. Thirteen other rats were subjected to the same operative procedure, but did

not receive a bioglass-ceramic implant and thus

served as controls for the fracture heahng process.

Animals in both groups were sacrificed at varying

time intervals, ranging from 6 to 28 weeks. The

specific durations are fisted in table 2. At sacrifice.

Table 2. Summary of the results of the mechanical testing of pairs of rat femurs

In all cases the right femur was osteotomized

Duration

With implant Without implant

Animal

Number kgf-cm kgf-cm

Animal

Number

Tu

kgf cm

Tr

kgf - cm

43 4.8 0.1 0.96 40 5.1 3.9 0.13 44 4.4 0.5 0.80 41 5.2 1.8 0.49 45 nonunion 48 5.1 1.6 0.52 46 5.3 0.6 0.80 49 6.6 1.8 0.57 47 nonunion 50 nonunion

8 weeks 13 nonunion 14 4.2 1.8 0.40

16 4.4 0.6 0.76 15 6.6 1.4 0.65 18 nonunion 19 4.6 3.8 0.10 20 nonunion 23 nonunion

25 nonunion 26 5.0 4.5 0.05

4 nonunion 9 4.7 3.2 0.19 5 fibrous-union 11 4.6 1.4 0.53

7 fibrous-union 12 5.0 4.2 0.09 10 nonunion

28 weeks ; 33 4.2 3.9 0.04 34 5.3 1.8 0.49

30
both femurs were removed from the animals and

all soft tissue removed from them. Any callus

bridging a washer was carefully removed, so that

the bioglass-ceramic washer completely separated

the two femoral fragments. The left femurs, which

were not operated on served as additional controls.

A block of metal-plastic potting compound was

cast around each specimen end, so that the bones

could be gripped securely for the subsequent

torsional test. The configuration of the torsional

test specimens is shown in figure 21. All femurs were loaded to failure in torsion in the Rapid Loading Torsional Tester, which is described in detail

by Frankel and Burstein [34]. The instrumentation

yielded the fracture torque for each specimen.

Cross-sections of the reconstructed femurs were

used to calculate shear stresses at failure, using

the computer program developed by Piotrowski

and Wilcox [35].

Figure 21. Arrangement of bone test specimen, in center, with

ends embedded to facilitate gripping for torsional testing.

The torques required to fracture all of the rat femurs are summarized in table 2. The percent

side-to-side difference of the fracture torques is defined by

_ Tl-Tr

where Tl and Tr are the torques required to fracture

the left and right femurs of a given animal, respectively. Values for d-T are also tabulated in table 2. Two studies of the normal side-to-side differences

in the fracture torques of paired rat femurs show

that there is no significant left-right bias (average

value of c?x — ~ 0.015 for the 27 pairs, with a standard deviation of 0.110). One should expect that

about 30 percent of aU rat femur pairs normally

wiU exhibit a side-to-side difference in fracture

torque greater than 0.110. This finding is consistent

with normal side-to-side differences found for canine

fibulas and monkey tibias as well [12].

Figure 22 summarizes the side-to-side differences

for the animals receiving bioglass-ceramic implants

and the control animals. Superposed over the diagram is the distribution of the side-to-side differ- ences in 27 normal rat femur pairs. A positive value

of d-r indicates that the right femur, which was

fractured, is weaker than the left femur, which

serves as a control.

; In the group without implants, two nonunions

occurred, as listed in table 2. The relative strengths

of the healed bones compared to the intact controls

as indicated in figure 22, where a histogram of the

d^s is plotted. In the group with the bioglassceramic washer implants, 10 femurs failed to unite

by the time the animals were sacrificed.

For both groups, the right femurs, by and large,

failed to recover structural integrity. The large

NUMBER OF OCCURRENCES 12

DISTRIBUTION FOR NORMAL PAIRS

(N:2J)

WITHOUT

IMPLANT

I

i

1 1 1 II II

I I 1 1 1 1 1 11

1 M I I I I II

WITH IMPLANT DURATION

6 WEEKS

8 WEEKS

17 WEEKS

28 WEEKS

NON UNIONS

NON UNIONS

-.2

Figure 22. Bar graph showing the percent side to side differences, dr, in the fracture torques of both

normal pairs of rat jemurs, pairs with sham operations (without implant), and pairs with implants

in the right femur.

31
degree of scatter in the relative strength, as well

as the large number of nonunions, appear to be

due to the fact that the fixation used affords no

resistance to torsional loading of the fractured

femur. Fewer nonunions were observed with the

control group because the rough bony surfaces

acting on each other provide more friction than the

bone-implant interface.

The maximum shear stresses experienced at fracture were calculated for 12 left femurs. The

distribution of these stresses is presented in

figure 23. The mean value for the computed fracture

stress was 1040 kgf/cm^, with a standard deviation

of 397 kgf/cm2.

3000

SHEAR STRESS

KGF/C m'

Figure 23. Distribution of fracture shear stresses for normal

rat femora.

All but one of the right femurs containing the

implants failed at the bone-implant interface. The

femurs of animal 33, shown in figure 24, clearly

show that fracture occurred away from the interface,

and in fact backed away from the interface after

approaching it. Shear stresses for bone 33R were

computed to be about 515 kgf/cm^ at the fracture

site, and 150 kgf/cm^ at the bone-bioglass interface.

Thus the interface was subjected to a stress equal

to 29 percent of the fracture stress of the same bone.

The results of this study were rather inconclusive

due to the small number of specimens. However,

Figure 24. Femurs of animal number 33 after mechanical

testing.

Note that the right femur did not fracture at the surface of the cerainic washer.

the striking success of animal 33 stimulated further

interest in the determination of the strength of

the bone-bioglass bond. A subsequent series of

rat implants failed to generate a technique for rigid

fixation in both bending and rotational modes. The

rat studies were then abandoned in favor of segmental replacement study involving monkey femurs.

12. Development of a Segmental Bone

Replacement: Monkeys

Segments of monkey femurs were surgically

removed and replaced by either a sohd bioglassceramic implant, or a flame-sprayed metallic implant, as shown in the accompanying x-rays (figs.

25 and 26). The compositions involved are listed

in table 1. All but two of the implants were gas

sterilized and conditioned in a buff^ered solution

for 24 h prior to implantation. Fixation of the

implant is provided by a 3 mm Schneider nail

which spans the length of the femur. Periodic

x-rays were used to evaluate the post operative

progress.

32
Figure 25. Development and remodeling of a normal callus

at a 45S5 bioglass-ceramic segmental bone replacement in a monkey femur.

In one monkey (fig. 27) paralysis of the right

linnb occurred when the sciatic nerve was engaged

by the drill during the reaming operation. Of the

20 other monkeys that were implanted, four animals

died within 1 week post op. One monkey fractured

both bones of the lower right leg and succumbed

to the subsequent complications. One monkey

escaped. Of the remaining 14 monkeys eight re- ceived bulk bioglass-ceramic implants and six

received bioglass flame-sprayed steel implants.

One of the animals with a bulk bioglass-ceramic

implant was sacrificed for histologic studies.

The progress of the repair processes is illustrated

by the accompanying x-ray sequences. Animal

2B18 received a bulk bioglass-ceramic implant

(fig. 25). The callus is beginning to mineralize at 4 weeks, is fully established at 16 weeks, and is

Figure 26. Development and remodeling of a callus at the extremity of a 45S5F bioglass flame sprayed 316L stainless

steel segmental bone replacement.

Figure 27. Minimal development of callus at a 45S5 bioglassceramic segmental bone replacement on a paralyzed monkey

femur.

33
to a large extent resorbed by 38 weeks. The larger

flame-sprayed implant in animal 3A14 (fig. 26)

prevents total encapsulation by the callus (13 weeks)

but also appears to form a contiguous structure

with the remodeled callus (37 weeks). Animal 2B3

sustained complete paralysis of the right leg during

the implantation, and shows a minimal amount of

reparative activity at 16 weeks. At 36 weeks little evidence of fracture healing was seen (fig. 27).

The animals were sacrificed at periods between

23 and 50 weeks. All major long bones of the lower

extremities were removed, the Schneider nails

withdrawn, and the bones tested mechanically to failure in a rapid loading torsion tester [34]. The

bones were kept physiologically moist during sacrifice and test. After the torsional tests were completed, the bones were reassembled, embedded in a clear epoxy potting compound, and sectioned.

The geometry of these sections was transmitted

to the SCADS computer program [35, 36] which

was used to determine the torsional stresses generated both at the fracture sites and at the implant

surfaces.

torques required to fracture the left and right bones

respectively. Pairs of tibias were used as controls

for this portion of the experiment. The group of

tibias exhibits a mean difference of 0.010, indicating that there is no left-right bias, and a standard

deviation of 0.07 [37]. The group of femurs containing a bulk bioglass-ceramic implant showed an

average difference of 0.29; this difference is quite

significant, in that it estabUshes that the structural

integrity of the femurs which were operated on

has not recovered fully by approximately 40 weeks.

The femurs containing the flame-sprayed implants

failed at extremely low torques, resulting in an

average difference of 0.66. The torques indicate

that the flame-sprayed implant system probably

would not be able to sustain severe physiological

torques. However, the monkeys used the legs with

the flame-spray coated prostheses without difficulty or any noticeable physiological deficit throughout

the post-operative period. Close examination of the

failure surfaces, which were always near the implant

bone interface, showed that the glass coating was

still adhering to the bone but that it was torn off

the metal substrate during the mechanical testing.

Table 3. Mechanical properties of monkey femurs

In all cases, the right femur contains the implant.

Animal Fracture Computed stress Number torques right femur

Right Left Fracture Interface

(kgf-cm) (kgf/ cm^)

Bulk bioglass-ceramic

2B2 9.0 24.0 0.45 * 547

2B5 18.0 30.0 0.25 515 822

2B6 22.0 24.5 0.05 1162 661

2B13 11.0 25.0 0.39 649 615

2B14 33.0 39.0 0.08 1174 955

2B18 10.0 27.0 0.46 * 101 2B19 14.0 30.0 0.36 859 1191

mean 0.29 872 849

Flame-sprayed metal

2F7 3.0 24.0 0.78 * 150

2F8 4.0 23.0 0.70 * 52 3A13 5.0 18.0 0.57 * 132 3A14 3.5 15.0 0.62 * 150

2G14 5.0 23.0 0.64 * 534

mean 0.66 204

*Failure occured at bone-implant interface.

The results of the mechanical testing are summarized in table 3. The structural integrity of the

experimental bone is characterized by the percent

difference between it and its opposite, control,

bone. This difference, d^, is defined by the equation

c?T = Tl — Tr/Tl + Tr, where Tl and Tr are the

12

10

FRACTURE STRESSES NORMAL BONES —

;

(LEFT FEMURS & TIBIAS) S=1560 kgf/cm2

<r = 300 kgf/cm^

FRACTURE STRESSES BIOGLASS-CERAMIC IMPLANTINTERFACIAL STRESS, BIOGLASS-CERAMIC

S= 848.8 cr= 234.0

400 800

-INTERFACIAL STRESS FLAME- SPRAYED

S = 203.8 <r= 189.1

1600 2000 2400

STRESS, KGF/CM^

FRACTURE STRESS,

RIGHT FEMURS WITH

BIOGLASS-CERAMIC S= 871.8 (r= 298.9

0 400 800 1200 1600 2000 2400

STRESS, KGF/CM^

Figure 28. Distribution of computed shear stresses of normal

bones at fracture, bone with implants at fracture, and shear

stresses developed at bone-implant interfaces in monkeys.

34
Figure 28 shows the results of the stress analyses

of the various bones involved. The upper histogram

shows the distribution of fracture stresses of

normal bones, i.e., left femurs and tibias in a bold

outhne. The average stress required to fracture a normal monkey bone is about 1560 kgf/cm^ with

a standard deviation of 300 kgf/cm^ No difference

was noted in the strength of bone from left femurs

and from tibias. Stresses from three groups of bone

implant systems are summarized in figure 28: bones

with flame sprayed stainless steel implants, bones

with bioglass-ceramic implants which broke during

the mechanical testing, and bones with bioglassceramic implants which fracture through the bone

rather than through the implant. The average

stress at which the bones containing flame sprayed

implants failed was 203 kgf/cm^. As noted before

this is really the stress required to fracture the

glass from the metal substrate (see fig. 29). Two

Figure 29. Fractured end of the flame spray coated segmental

implant showing the exposed metal surface fracture site (grey areas).

bulk bioglass-ceramic implants failed through the

implant rather than through the bone (fig. 30), and

these two are shown as crosshatched blocks in

figure 28. This low strength shows that at the present

state of the art the bulk material is not of adequate

strength to rehably support physiological loads to

failure. Five femurs containing bulk bioglassceramic implants failed away from the bone implant

interface. The lower histogram of figure 28 shows

the distribution of the fracture stresses calculated

in those, bones, with the corresponding region in

the upper histogram denoting the interfacial stresses

Figure 30. Fallure of bone bioglass-ceramic implant test speci- men (2B18) failing within the implant.

Note the implant bone interface withstands the fracture stress applied.

developed at the surface between the bioglassceramie and the newly healed bone. The average

stresses in these two distributions are very close

to each other, and are not significantly diff"erent

from each other.

13. Discussion

The structural strength of the monkey femurs

was compromised considerably by the presence of

the implant. During the torsion tests, fracture

occurred at various sites, but appeared to be

limited by either the strength of the implant material or the strength of the healing bone near the

implant. The fact that the fracture stresses of the

right femurs containing bioglass-ceramic implants

was considerably lower than the average fracture

stress for normal monkey bones was probably due

to the fact that the Schneider nails used for fixa- tion were not removed from the animal until just

prior to test. This means that a substantial portion

of the bending loads to which the bone-ceramic system was subjected was transmitted across the implant site by the nail itself. Thus the bone was protected from full physiologic stress and is therefore probably weaker than normal monkey bone.

It should be kept in mind, however, that normal

axial compression forces were able to act on the

implant-bone interface. This axial compression is probably responsible for the in-vivo implant failures.

The distribution of the interfacial stresses

calculated for the bulk bioglass-ceramic implants

indicates that the strength of the interface is nearly

as great as that of the newly healing bone in the

same vicinity. The data in figure 28 also indicates

that the strength of the interface is at least 75

percent of the strength of normal monkey bone.

Due to the fact that failure did not occur preferentially at the interfaces in our study, it must be

kept in mind that these figures represent only a lower hmit on the strength of the bond developed

between the bone and the bioglass-ceramic.

35
The authors gratefully acknowledge the able

assistance of H. Parkhurst, A. W. Smith, and J. McVey of the University of Florida; S. W. Freiman

and E. J. Onesto of IITRI, the financial support

of U.S. Army Medical R&D Command Contract

No. DADA 17-70-C-OOl and cooperation of the

Gainesville, Florida Veterans Administration

Hospital.

14. References

[1] Hench, L. L., Ceramics, glasses and composites in medicine,

Med. Instr. 1, 136-144 (1973).

[2] Boutin, P., Direct attachment of ceramic joints, Symposium

on Materials for Reconstructive Surgery, (Clemson

University, April 1974).

[3] Griss, P., Heimke, G., von Andrian-Werburg, H., Krempien,

B., Reipa, S.. Hartung, J., and Lauterbach, J., Morphological and biomedical aspects of AI2O3 ceramic joint

replacement; experimental results and design considerations for human endoprostheses. Symposium on Materials

for Reconstructive Surgery, (Clemson University, April

1974).

[4] Griss, P., von Andrian-Werburg, H., Krempien, B., and

Heimke, G., Biological activity and histocompatibility of dense Al203/MgO ceramic implants in rats, J. Biomed.

Res. Symp., No. 4, pp. 453-462, (Interscience, New York,

N.Y. 1973).

[5] Griss, P., Krempien, B., von Andrian-Werburg, H.. Heimke,

G., and Fleiner, R., Experimentelle Utersuchung zur Geivebsuertraglichkeit oxidkeramischer (AI2O3) Abri- ebteilchen. Arch, fiir orthopadische und Unfall-Chirurgie

76, 270-279 (1973).

[6] Englehardt, A., Salzer, M., Zeibig, A., and Locke, H., Experiences with AI2O3 implantation in humans to bridge

resection defects. Symposium on Materials for Reconstructive Surgery, (Clemson University, April 1974).

[7] Klawitter, J. J., and Hulbert, S. F., Applications of porous ceramics for the attachment of load bearing internal orthopedic applications, J. Biomed. Res. Symp., No. 2, Part 1, pp. 161-229 (Interscience, New York, N.Y. 1972).

[8] Hulbert, S. F., Cooke, F. W., Klawitter, J. J., Leonard, R. B., Sauer, B. W., Moyle, D. D., and Skinner, H. B., Attach- ment of prostheses to the musculoskeletal system by

tissue ingrowth and mechanical interlocking, Biomed.

Res. Symp. No. 4, pp. 1-23, (Interscience, New York,

N.Y., 1973).

[9] Hulbert, S. F., Young, F. A., Mathews, R. S., Klawitter, J. J., Talbert, C. P., and Stelling, F. H., Potential of ceramic

materials as permanently implanted skeletal prostheses,

J. Biomed. Mat. Res. 4, 433-455 (1970).

[10] Hulbert, S. F., Klawitter, J. J., and Leonard, R. B., Compatibihty of bioceramics with the physiological environ- ment. Ceramics in Severe Environments, W. W. Kriegel

and H. Palmour, IH (Eds.), pp. 417-434 (Plenum Press,

New York, N.Y., 1971).

[11] Piotrowski, G., Hench, L. L., Allen, W. C, and Miller, G., Mechanical studies of the bone-bioglass interfacial bond.

Symposium on Materials for Reconstructive Surgery,

(Clemson University, April 1974).

[12] Graves, G. A., The influence of compositional variations on bone ingrowth of implanted porous calcium aluminate

ceramics. Symposium on Materials for Reconstructive

Surgery, (Clemson University, April 1974).

[13] Graves, G. A., Hentrich, R. L., Stein, H. G., and Bajpai,

P. K., Resorbable ceramic implants, J. Biomed, Res. Symp. No. 2, Part 1, pp. 91-115 (Interscience, New York,

N.Y. 1972).

[14] Bhaskar, S. N., Brady, G. M., Getter, L., Grower, M. P.,

and Driskell, T. D., Biodegradable ceramic implants in bone. Oral Surg. 32, 336 (1971).

[15] Bhaskar, S. N., Cutright, D. E., Knapp, M. J., Beasley, J. D., Perez, B., and Driskell, T. D. Tissue reaction to intrabony

ceramic implants. Oral Surg. 31, 282 (1971).

[16] Schnittgrund, G. D., Kenner, G. H., and Brown, S. D., In vivo and in vitro changes in strength of orthopedic

calcium aluminates, J. Biomed. Res. Symp. No. 4, pp. 435-452 (Interscience, New York, N.Y., 1973).

[17] Kenner, G. H., Brown, S. D., Pasco, W. D., and Lovell, J. E., Bull. Amer. Ceramic Soc. 52 , 432 (1973).

[18] Frakes, J. T., Brown, S. D., and Kenner, G. H., Bull, Amer.

Ceramic Soc. in press, (1973).

[19] Jecman, R. M., Eggerding, C. L., Brown, S. D., and Schnitt- grund, G. D., J. Biomed. Mat. Res. 7, 369 (1973).

[20] Hench, L. L., Splinter, R. J., Allen, W. C, and Greenlee,

T. K., Mechanisms of interfacial bonding between ceramics and bone, J. Biomed. Res. Symp. No. 2, Part 1, pp. 117-143 (Interscience, New York, N.Y., 1972).

[21] Hench, L. L., and Paschall, H. A., Direct chemical bond of bioactive glass-ceramic materials to bone and muscle,

J. Biomed. Res. Symp. No. 4, pp. 25-42 (Interscience,

New York, N.Y., 1973).

[22] Clark, A. E., Hench, L. L., and Paschall, H. A., The influence

of surface chemistry on implant interface hisotlogy: A

theoretical basis for implant materials selection. Symposium on Materials for Reconstructive Surgery, (Clemson

University, April 1974).

[23] Hench, L. L., and Paschall, H. A., Histo-chemical responses

at a biomaterials interface, J. Biomed. Res. Symp. on

Prostheses and Tissue: The Interface Problem, (Clemson

University, 1973).

[24] Beckham, C. A., Greenlee, T. K., Jr., and Crebo, A. R., Bone formation at a ceramic implant interface, Calc.

Tis. Res. 8, No. 2, 165 (1971).

[25] Greenlee, T. K., Jr., Beckham, C. A., Crebo, A. R., and

Malmborg, J. C, Tissue responses at the interface of a ceramic, J. Biomed. Mat. Res. 6, 244 (1972).

[26| Clark, A. E., and Hench, L. L., Eff'ect of P^+, 6^+, and F"

on corrosion of invert soda-lime-silica glasses. Submitted

to J. Amer. Ceramic Soc. (1974).

[27] Clark, A. E., Pantano, C. G., and Hench, L. L., Auger

spectroscopic analysis of invert glass corrosion films. Submitted to J. Amer. Ceramic Soc. (1974).

[28] Carlisle, E. M., Silicon — a possible factor in bone calcification. Science 167, 279-280 (1970).

[29] Housefield, L. G., Mechanical property control of a bioglassceramic system. Masters Thesis, University of Florida

(1972).

[30] Bhton, J. L., Ceramic coatings for cementitious and metallic surfaces, IITRI Final Report, Project No. B8009, (Inter- national Lead Zinc Research Organization, 1964).

[32] Oel, H. L., The relationship between free energy and kinetics

in sintering processes. Kinetics of High-Temperature

Processes, W. D. Kingery (Ed.), (John Wiley and Sons,

Inc., New York, N.Y., 1959).

[32] Kingery, W. D., Introduction to Ceramics, pp. 628-647

(John Wiley and Sons, Inc., New York, N.Y., 1960).

[33] Hench, L. L., Onesto, E. J., and Freiman, S. W., Development of a bioglass flame spray coated hip prosthesis, U.S. Army Medical Research and Development Command

Contract No. DADA 17-70-C-OOOl, Report No. 3, 180 pp.

[34] Burstein, A. H., and Frankel, V. H., A standard test for laboratory animal bone, J. Biomech. 4, 155 (1971).

[35] Piotrowski, G., and Wilcox, G. A., The STRESS program: a computer program for the analysis of stresses in long

bones, J. Biomech. 4, 497 (1971).

[36] Piotrowski, G., and Kellman, G. I., A stress calculator for

arbitrarily drawn sections — the S.C.A.D.S. computer

program, U.S. Army Medical Research and Development

Command Contract No. DADA 17-70-C-OOOl, Report

No. 4, 126 pp. (1973).

[37] Miller, G., and Piotrowski, G., VariabiHty of the torsional strength of paired bones, J. Biomech., 7 , 247 (1974).

36
NATIONAL BUREAU OF STANDARDS SPECIAL PUBLICATION 415,

Biomaterials, Proceedings of a Symposium held in conjunction with the Ninth Annual Meeting of the Association for the Advancement of Medical Instrumentation, New Orleans, La.. April 19-20, 1974

(Issued May 1975).

Soft Tissue Response to a Series of Dense Ceramic

Materials and Two Clinically Used Biomaterials

W. C. Richardson, Jr.

Tulane University, New Orleans, La. 70112

S. F. Hulbert

Medical Center, University of Alabama, Birmingham, Ala. 35294

and

J. J. Klawitter, and B. W. Sauer

Clemson University, Clemson, S.C. 29631

Disc-shaped implants of spinel, alumina, muUite, zircon, a cast Co-Cr-Mo alloy, and ultra-high

molecular weight polyethylene (UHMWPE) were implanted in the parasipinaUs musle of twelve adult,

male. White New Zealand rabbits. Prior to implantation the implants were characterized with respect

to size and shape, weight and surface roughness. After periods of 1 month, 2 months, and 4 months

the rabbits were sacrificed and the tissue specimens were retrieved with the implants stiU intact. Histological examination of the tissues surrounding the implants along with changes in the size and

shape, weight, and surface roughness of the implants were used as criteria for evaluating these materials

for implant purposes. No surface degradation of any of the materials was detected using scanning electron microscopy.

Fibrous tissue seems to adhere to the UHMWPE implants more than any other material used in this study. Large amounts of fibrous tissue were also found to adhere to the cast Co-Cr-Mo alloy implants.

The histological results indicated that within the limits of this investigation, the biocompatibility

of the ceramic materials used in this study compares favorably with the clinically used cast Co-Cr-Mo

alloy implants and the UHMWPE implants.

Key words: Biocompatibility, ceramic implants, histological evaluation; implant characterization.

1. Introduction

Ceramic materials are beginning to find a place

in the field of biomaterials as new demands are made

for materials that can withstand the highly corrosive

environment of the human body [1-26].^ The

physical and chemical properties of ceramics have

recently led researchers to consider these materials

for use in the replacement of bone [1, 3, 4, 6, 9, 10,

12-16, 19-26] and teeth [17-19].

Ceramic materials have certain advantages and

disadvantages as compared to other materials for

implantation. The major advantage of ceramics is that they are highly chemically inert in the physiological environment. They are, however, characteristically brittle, notch sensitive materials. Thus, a

careful consideration of their mechanical properties

wiU be essential to their applications as biomaterials.

' Figures in brackets indicate the literature references at the end of the paper.

Fixation of orthopedic appliances to bone is a

definite problem to the orthopedic surgeon. Some

investigators have approached this problem by

developing a mechanical interlocking between the

implant and the tissue [3, 6, 10, 12-14, 17, 19, 20,

22, 24, 26]. The mechanical interlocking was

achieved by tissue growth into the interconnecting

pores of porous ceramic materials. Klawitter [10]

reported that a minimum pore size of approximately

100 fxm is necessary for bone ingrowth.

Hench et al. [21], reported achieving direct chemical bonding of a glass-ceramic material with bone

and soft tissues. They accompUshed this through the

development of a series of surface-active glasses

and glass-ceramics.

2. Compatibility Testing

One of the problems of conducting a biocompatibihty experiment is that of methodology. Thus far

there appears to be no universally accepted method

for evaluating tissue compatibility [27-42].

37
There are two main techniques which have been

used to evaluate candidate materials for implantation, in vitro and in vivo studies. In vitro studies

incorporate the use of tissue culture or simulated

body fluids to test the biocompatibility of candidate

materials and they have advantages in that they are considerably less expensive. While in vitro studies serve as an excellent tool for screening materials,

they do not provide conclusive information because

the total dynamic physiological environment is not

simulated [36, 39, 40].

In vivo biocompatibility studies, most generally,

consist of implanting materials of interest into

research animals for varying periods of time and

subsequently performing a histological evaluation

of the adjacent tissues.

There are some inherent problems associated

with determining compatibility of biomaterials

[44]. When implants are placed into the body, the

material is subjected to a barage of rejection and

healing mechanisms. As the body reacts to the

foreign material, certain histological processes

take place which are observable in the tissues

surrounding the implant as well as in special organs

that are far removed from the implant. The combined activity of the many cell types, tissues, and

organ systems produce such a complex sequence

of interactions that it is often difficult to explain

certain histological observations that occur when an

implant is placed in the body [44]. It is essential

for the researchers in these types of studies to have

a basic understanding of the nature and processes

of foreign-body reactions and their characteristic

signs. A paper by Hulbert et al. [44], provides a discussion of phagocytosis, the immune response,

abscesses, neoplasms, poisons, inflammation and

normal wound healing and relates these phenomena

to their importance in evaluating the biocompatibility of implant materials.

Many investigators attempt to monitor the biocompatibility of an implant by designing a device

of the type desired and to test it in the exact location

of the intended application [7, 17, 18, 24, 25, 37, 45].

In an attempt to reduce the number of experimental

variables, other researchers have implanted sample

materials in various shapes and forms [5, 6, 9-12,

15, 27-30, 33, 34]. Wood et al. [29], reported that an increased tissue response which typically occurs near the ends of rod-shaped implants does not occur around disc-shaped implants.

Autian [38] mentioned that a material implanted

in the body may initiate a response in two general

ways: (1) the effects of the body on the material

and (2) the effects of the material on the body. In

view of this fact, it seems that a properly designed

compatibility experiment would seek to not only

investigate the histological reactions, but would

also attempt to monitor changes which may occur

to the material. Indeed, material characterization

is an extremely important but often neglected

portion of biocompatibility studies.

3. IVIaterials and Methods

3.1. Scope of the Research

Disc-shaped implants of spinel ^ (MgO AI2O3),

alumina 2 (AI2O3), muUite 2 (3Al2Qi 2Si02), zircon 2 (ZrSi04), a cast Co-Cr-Mo aUoy,^ and ultra-high

molecular weight polyethylene, were implanted iri the paraspinalis muscle of twelve adult, male,

White New Zealand rabbits. Prior to implantation

the implants were characterized with respect to size and shape, weight and surface roughness.

After periods of 1 month, 2 months, and 4 months

the rabbits were sacrificed and the tissue specimens were retrieved with the implants stiU intact.

Histological examination of the tissues surrounding

the implants along with changes in the size and

shape, weight and surface roughness of the implants were used as criteria for evaluating these

materials for implant purposes.

3.2. Materials Selection

SpineP (MgO AI2O3), alumina^ (AI2O3), mullite^

(3AI2O3 2Si02), and zircon 2 (ZrSi04), were chosen

for this experiment because they represent a group

of ceramic materials which are often selected for use in severe environments and which were available from commercial suppliers.^

In order to facilitate quality comparisons of these

ceramic materials with clinically used biomaterials,

this experiment included the implantation of an ultra-high molecular weight polyethylene

(UHMWPE)4 and a cast cobah-chromiummolybdenum alloy (meeting the ASTM Standard

Specification for Surgical Implants, designation

number F75-67) as experimental controls.

3.3. Materials Characterization

Each implant was characterized with respect to size and shape, surface roughness, and weight

before and after implantation. Scanning electron

microscopy was also used before and after implantation. Chemical analysis, x-ray diffraction, and

reflected light microscopy were used to characterize the materials before implantation.

2 Supplied by American Lava Corporation, Chattanooga, Tenn.

3 Zimaloy, Zimmer Manufacturing Company, Warsaw, Ind. *RCH 1000, Hoechst, Supplied by Zimmer Manufacturing Co., Warsaw, Ind.

^ Certain commercial materials and instruments may be identified in this publication in order to adequately specify the experimental procedure. In no case does such

identification imply recommendation or endorsement by the National Bureau of Standards, nor does it imply that the equipment or instruments identified are necessarily

the best available for the purpose.

38
Size and Shape

The disc-shaped implant was chosen for this

experiment for reasons demonstrated by Wood

et al. [29]. The diameter and thickness of each

implant was determined.

3.4. Surface Roughness

Variation in surface roughness from 4 microinches to 25 microinches (0.1-0.6 /xm) was shown

not to influence the degree of tissue reaction in a prehminary experiment [48].

Attempts were made to achieve a uniform surface

roughness on all implants. The ceramics were

pohshed on a 30 /xm diamond grinding wheel with

a water lubricant. The UHMWPE implants were

finished by hand with 320 grit (approximately equal

to 30 jU,m) SiC grinding paper and with a water

lubricant. These implants were finished by hand to reduce the degree of feathering along the edges of

the implant which occurred using the grinding

wheel. The Zimaloy implants were polished on a 320 grit SiC grinding wheel. The surface roughness

of the implants was measured using a profilometer ^ before and after implantation.

3.5. Weighing the Implants

Prior to weighing the implants, they were washed

in an ultrasonic cleaner in an effort to remove any

particulate matter and any organic film which may

have been present. The following procedure was

used:

Benzene wash 15 minutes

Acetone wash 15 minutes

Absolute Ethanol wash 15 minutes

Dry and desiccate in a Bell Jar 8 hours

From this point on, the implants were only

handled with a pair of forceps.

Each implant was characterized with respect to size and shape, surface roughness, and weight prior

to being implanted. After the implantation period,

the implants were to be characterized in the same

fashion to determine any changes jn these characteristics due to implantation. A separate experiment

was conducted in order to determine the possibility

of any changes which may result from cleaning

and/or sterilization procedures. The results showed

that the sterilization and/or cleaning procedures do

not significantly affect the size or shape, the surface

roughness or the weight of the implants.

Chemical Composition and Crystal Phase State of the

Implants

The chemical composition of the implants was

"Brush Electronics Co., Cleveland, Ohio. Model BL- 110-57-2041.

determined by spectrochemical analysis^ of the

materials. X-ray diffraction* was used to determine

the crystal phase state of the ceramic implants.

Microstructure of the Implants

Photomicrographs and the average crystal size

data for the four ceramic materials were acquired

from the supplier of these materials.

Metallographic analysis was performed to determine the microstructure of the cast Co-Cr-Mo

alloy (Zimaloy) used in this study.

3.6. Surgical Procedure

Three groups of four male White New Zealand

rabbits approximately 6-8 months old received six

implants and a sham operation (an incision without an implant) in the paraspinalis muscle of each

rabbit. The first group of four rabbits received

implants for 4 months. The second group were 2- month implants and the third group were 1-month

term implants. The relative positions of six implants

and the sham were chosen completely at random.

This process was carried out for each of the 12

rabbits.

Sodium pentobarbital (approximately 60.0 milligram per kilogram of body weight) was used for

induction and maintenance of surgical anesthesia.

A large region of hair was chpped to reveal the skin on the back. The skin was then scrubbed with

alternating providone-iodine surgical scrub sponges

and alcohol sponges approximately three times. A

film of providone-iodine surgical preparatory solution was sprayed on the skin and left for surgery.

Sterile technique was used throughout surgery.

A midline skin incision approximately 7 cm long

was made with an electrocautery unit. The subcutaneous facia was dissected out of the implantation

region. Incisions were made with a scalpel approximately 2 cm into the paraspinalis along the longitudinal axis of the muscle and approximately 1 cm

long. The implants were placed into the incision

using a pair of polypropylene forceps. The muscle

tissue was closed with a single horizontal mattress

stitch of 000 medium chromic gut. The facia was

closed in the same manner. The sham operation

was performed in the same fashion except that the

forceps were inserted into the muscle without an

implant. The subcutaneous tissue was closed with

a simple continuous suture pattern of 000 chromic

gut. The skin was closed with interrupted 000

monofilament nylon suture material.

3.7. Post Operative Care

Following surgery, a clean bandage was secured

'Spectrochemical Laboratories, Inc., Pittsburgh, Pa. 'Norelco Type 42202 Diffractor, North American Philips Co.. Inc., Mount Vernon,

N.Y.

39
with adhesive tape over the midportion of the

rabbit. Procain PeniciUin-G (100,000 units) and

dihydrostreptomycin (0.125 g) were administered

twice a day for 5 days post-operatively.

3.8. Tissue Preparation

The rabbits were sacrificed using a 1 cm^ intra- venous injection of Barb-Euthol. After sacrifice,

tissue specimens measuring approximately 2.5 cm

by 1.5 cm by 0.5 cm were retrieved with the implants

in situ and mounted on a piece of balsa wood before

fixation in 10 percent formalin. After at least 2 days

in 10 percent formalin, the implants were carefuUy

dissected out of the tissue specimens by making

an incision along the midportion of the implant

parallel to the longitudinal axis of the muscle fibers.

The tissue specimens were infiltrated and embedded

in paraffin blocks to facilitate sectioning.

Sections 8 )U,m thick were taken from the tissue

blocks using a rotary type microtome. These sec- tions were mounted on glass slides using an albumin

fixative and were dried overnight before staining.

The sections were stained using a normal hematoxylin and eosin staining technique and were

covered with microscopic cover glasses for histological examination under a light microscope.

3.9. Post-Implantation Characterization of

Materials

Post-implantation characterization of the implants

was performed using the same technique as in the

pre-implantation characterization.

4. Results and Discussion of Results

4.1. Results of Materials Characterization

Geometric Analysis

The diameter and thickness of each implant was

determined before and after implantation. All implants had an approximate diameter of 1 cm and

an approximate thickness of 1 mm. When consideration is given to the diameter and thickness of the

implants and the sensitivity of the instrument used,

the minimum percent change detectable for the

thickness is 1 percent and the minimum percent

change detectable for the diameter is 0.25 percent.

A summary of the data for each material is listed in

table 1. The UHMWPE implants were the only implants

to have a noticeable increase in diameter. The

increase in thickness can probably be attributed to the fact that some tissue remained attached to the

surfaces of the implants after washing the implants

in successive washings of benzene, acetone and

Table 1. Summary of geometric analysis

IVle3.n percent IVlean percent

IVlatcna.! name change in change in diameter thickness

U.U U.U MuUite (3Al203 2Si02) .0 .0 Zircon (ZrSi04) .0 .0 Spinel (MgO AI2O3) .0 .0 UHMWPE +0.3 + 1.0 0.0 + 1.0

absolute alcohol. One of the UHMWPE implants

was processed through the normal H & E staining

procedure in order to identify the tissue which had

adhered to these implants. These UHMWPE implants permitted the transmission of fight and thus,

the histological observation was made with relative ease. These histological results revealed that a thin

layer (approximately one cell layer thick) of young

fibrous tissue was firmly attached to these implants.

4.2. Weight Determinations

The weight of the implants used in this study was

determined before and after implantation. When

consideration is given to the weights of the implants

and the sensitivity of the instrument used, the

minimum percent change detectable for the weights

is 0.01 percent. A summary of the data for each

material is given in table 2. The spinel implants were the only implants to have an appreciable change in weight.

The weight changes observed here were biased

as mentioned earlier by the fact that some tissues

continued to adhere to each sample after the

washing sequence.

Table 2. Summary of weight changes

Mean weight before Mean percent

Material name implantation change in (grams) weight

Alumina (Al^Os) 0.2994 -0.02

MuOite (3Al203 2Si02) .2452 + 0.01 Zircon (ZrSi04) .3211 0.00 Spinel (MgO AI2O3) .2044 -0.38

UHMWPE .0707 + 0.05 Cast Co-Cr-Mo alloy .6869 -0.01

4.3. Surface Roughness Determinations

The surface roughness of each implant was

determined before and after implantation. When

consideration is given to the sensitivity of the

profilometer, the minimum percent change detecta40
ble for the surface roughness is 1 percent. A sum- mary of the results is presented in table 3.

Table 3. Summary of changes in surface roughness

Mean surface Mean percent

roughness change in

IVlalcricll Ilalllc before s ui lace implantation roughness

(microinches) ^

Alumina (Al-iOa) 14.9 13 MuUite (3Al203 2Si02) 15.8 21 Zircon (ZrSi04) 10.6 75

Spinel (MgO AI2O3) 15.1 37

UHMWPE 27.4 6 Cast Co-Cr-Mo alloy 6.1 259

" 1 microinch= 0.025 micrometers

All of the materials showed an overall increase in

surface roughness. The cast Co-Cr-Mo aUoy implants showed a considerably larger increase than

any of the ceramic materials or the UHMWPE

implants. The pre-implantation measurements for

the cast Co-Cr-Mo alloy implants indicated a surface roughness of approximately 6 microinches

which is considerably less than the surface rough- ness of the other implants before implantation. The

tissue which had adhered to the cast Co-Cr-Mo

alloy implants, along with the low before implantation measurements, are probably responsible for the

large percent increase in surface roughness.

4.4. Scanning Electron Microscopy

One representative sample of each material was

observed under a scanning electron microscope^

(SEM) before and after implantation.

After removing the implants from the tissue

blocks, all of the materials had some tissue which

had remained intact. Two specimens of each

material (from the 4-month implant group) were

chosen to determine an approximate percentage

of the surface covered with tissue. An eyepiece

objective with grid markings was used to observe

the implants under reflected light. Three areas of

the surface of each implant were chosen at random

and the number of grid blocks falling on the surface

with and without tissue was counted and percentages calculated.

The SEM analysis provided supportive evidence

in agreement with the data presented in table 4. Very little difference was seen in the SEM photomicrogiaphs of the four ceramic materials before

and after implantation. However, the UHMWPE

appeared to be entirely covered with tissue.

^Mark 2A Slero Scan. Kent Cambridge Scientific Inc.. Morton Grove, III.

Table 4. Approximate percentages of the surface ofthe implants

covered with tissue

Approximate

percentage of the Material surface covered

with tissue

Q0

Mullite (3Al203 2Si02) 1115 Spinel (MgO AI2O3) 8

UHMWPE 96

Cast Co-Cr-Mo alloy 72

The cast Co-Cr-Mo alloy implants had relatively

large amounts of tissue which had remained intact.

This is in agreement with observations recorded by

Homsy, et al. [49].

4.5. Spectrochemical Analysis

Spectrochemical analysis of the ceramic materials

used in this study is given in table 5.

Table 5. Semiquantitative spectragraphic analysis " (weight

percent)

Constituents Alumina Mullite Zircon Spinel

AI2O3 Major Major 3.0 20.0 Zr02 *0.01 "0.01 Major 0.01 MgO .15 10.0 2.0 Major

Si02 .35 10.0 25.0 0.10 CaO .05 0.06 1.0 .07 Ti02 .02 .02 0.25 .05 Fe203 .15 .15 .25 .15 CuO .10 .05 .01 .02 CoO .07 C) C) C)

BaO C) C) .02 (") NiO .005 .003 .002 .003 PbO .002 .002 .002 Cr203 .001 C)

ZnO .001 .001 .001 B2O3 .001 .001 C) C^) Sn02 C) C) {') C)

MnO .001 .001 .001 .001

"Data obtained from Spectrochemical Laboratories, Inc., Pittsburg, Pa. Not detected. Zr Interference.

4.6. Crystal Phase State of the Ceramic

Materials

The identity of the ceramic materials used in

this study was confirmed using x-ray diffraction. A

representative specimen of each of the ceramic

implants was crushed, ground and a powder x-ray

diffraction analysis performed using Cu Ka radiation. No unreacted raw materials were observed in

any of the specimen analyses.

41
Microstructure and Average Crystal Size of the Materials A summary of these results appears in table 7.

Photomicrographs of the ceramic materials were

provided by the American Lava Corporation. The

average crystal size of the ceramic materials is given in table 6.

Table 6. Average crystal size of the ceramic materials "

Material Average crystal size

Alumina (AI2O3) typic2Llly 2 ju.m Mullite (3AI2O3 2Si02) 6 ixm

Spinel (MgO AI2O3) 6 /xm

Zicron (ZrSi04) typically =S 2 /lim

" Data obtained from American Lava Corporation, Chattanooga, Tenn.

5. Results of Histological Evaluation

Three groups of rabbits received implants for

4 months, 2 months, and 1 month, respectively.

One of the rabbits (from the 2 month implant group)

died 2 weeks post-operatively. The cause of death

was a chronic respiratory infection not related to the implant study. All other rabbits remained healthy

the full duration of their implant periods.

Upon sacrifice the implant sites were observed

macroscopicaUy. All of the implant sites appeared

to be healthy and normal with no incidence of

excessive swelling or redness.

The histological sections were evaluated using the

following criteria:

slight: thin fibrous tissue capsule with uniform

thickness and no inflammatory response

evident.

moderate: thin fibrous tissue capsule with some

locahzed thickening and with httle or no

inflammatory response evident.

severe: uniform thick fibrous tissue capsule

associated with a chronic inflammatory

response.

All of the ceramic materials indicated a tissue

compatibiUty that compared favorably with the

compatibiUty of the clinically used cast Co-Cr-Mo

alloy and the UHMWPE. The spinel implants

seemed to have a slightly better compatibihty than

any of the other materials used in this study. The

mulhte implants indicated a tissue compatibility

slightly less than any other material.

Many of the histological sections indicated a much

thicker fibrous membrane on one side of the implant site with a very thin membrane on the other

side of the implant site. This could probably be

attributed to mechanical factors rather than poor

compatibihty. Wood et al., has described a "clubbing" effect which occurs around rod-shaped

implants [29]. This "clubbing" effect is the increased

tissue response around the ends of a rod-shaped

implant. The effect occurred around many of the implants used in this study. Since this type of increased

tissue response is thought to be due to mechanical

irritation, these regions of "clubbing" were not con- sidered in the histological evaluation.

There are some inherent advantages and disadvantages in this method of evaluation. The removal

of the implants before histological procedures allows

the investigator to use a paraffin embedding technique. This technique is an easier and quicker

technique than the process of embedding the tissue

and implant in poly(methylmethacrylate) [10].

Another advantage in this technique is that it allows

pre- and post-implantation evaluation of the implants.

The disadvantages in paraffin embedding are twofold. Firstly, the actual tissue-implant interface

cannot be observed. Secondly, the criteria for histological evaluation in this study considered the thick- ness of the encapsulating membrane. Since part

of this membrane was removed along with the implant, it seems that the histological evaluation may

have been adversely affected.

Table 7. Summary ofhistological evaluation

Material Fraction with Fraction with Fraction with a slight a moderate a severe response response response

Alumina (AI2O3) 5/12 6/12 1/12

Mullite (3Al203 2Si02) 4/12 8/12 0

Spinel (MgO AI2O3) 9/12 2/12 1/12

Zircon (ZrSi04) 8/12 2/12 2/12

UHMWPE 5/12 7/12 0 Cast Co-Cr-Mo alloy 8/12 3/12 1/12

Note. All of the sham operations indicated a normal wound healing process.

42
6. Overall Discussion of Results and

Recommendations

Changes in weight, size and shape, and surface

roughness of the implants have been monitored on

a pre- and post-implantation basis. Table 8 serves

to summarize the results. The spinel implants were

the only implants to exhibit an appreciable weight

loss. These implants had the best tissue response

of all materials used in this study. No obvious

changes were noted in the surface of these spinel

implants using scanning electron microscopy.

Within the hmits of this investigation, the biocompatibihty of the ceramic materials compares

favorably with the clinically used cast Co-Cr-Mo

implants and the UHMWPE implants.

One of the major disadvantages in this technique

of evaluation is that the tissue-implant interface

cannot be observed histologically. Future biocompatibility studies could possibly become more

significant by implanting two series of implants. One

group could be used to evaluate the material on a

pre- and post-implantation basis as described in this

experiment. Another group could be sectioned with

the implant in situ in order to observe the actual

tissue-implant interface using the technique described by Klawitter [10].

Another recommendation is that future studies

include the implantation of a negative control. This

would further validate the sensitivity of the implant

study.

Table 8. Overall summary of results

Alumina

(AI2O3)

Mullite

(3AI2O3 2Si02)

Zircon

(ZrSi04)

Spinel

(MgO AI2O3)

UHMWPE Cast Co-Cr-Mo

alloy

Mean percent change in weight -0.02 -fO.Ol 0.00 -0.38 + 0.05 -0.01

Mean percent change in diameter 0.0 0.0 0.0 0.0 + 0.3 0.0

0.0 0.0 0.0 0.0 + 1.0 + 1.0

Mean percent change in Surface Roughness + 13 + 21 + 75 + 37 + 6 + 259

Fraction with slight response 5/12 4/12 8/12 9/12 5/12 8/12

Fraction with moderate response 6/12 8/12 2/12 2/12 7/12 3/12

Fraction with severe response 1/12 0 2/12 1/12 0 1/12

7. Conclusions

Within the limits of this investigation, the following conclusions can be drawn:

1. The biocompatibility of the ceramic materials

used in this study compares favorably with the

clinically used cast Co-Cr-Mo alloy implants

and the UHMWPE implants.

2. Fibrous tissue seems to adhere more to the

UHMWPE implants than to any other material

used in this study.

3. Fibrous tissue seems to adhere more to the

cast Co-Cr-Mo alloy than to any ceramic

material used in this study.

4. No surface degradation of any of the materials

was detected using scanning electron microscopy.

8. References

[1] Ross, G. A., and Smith, L., Cerosium. Symposium on Use of Ceramics in Surgical Implants (Clemson University, 1969).

[2] Helmer, J. D., and Driskell, T. D., Research on Bioceramics,

Symposium on Use of Ceramics in Surgical Implants

(Clemson University, 1969).

[3] Hentrich, R. L., Graves, G. A., Stein, H. G., and Bajpae, P. K., An evaluation of inert and resorbable ceramics for future

clinical orthopedic applications, J. Biomed. Mater. Res. 5,

pp. 25-51 (1971).

[4] Welsh, R. P., and Macnab, Ian., Ceramics in surgery, J. Biomed. Mater. Res. 5, pp. 231-249 (November 1971).

[5] Davis, S. D., Gibbons, D. F., Martin, R. L., Levitt, S. F., Smith, J., and Harrington, R. W., Biocompatibility of ceramic implants in soft tissue, J. Biomed. Mater. Res. 6, pp. 425-449 (September 1972).

[6] Hulbert, S. F., Klawitter, J. J., Talbert, C. D., and Fitts, C. T., Materials of construction for artificial bone segments,

Research in Dental and Medical Materials, pp. 19-67

(Plenum Press, 1969).

[7] Hulbert, S. F., Morrison, S. J., and Klawitter, J. J., Compatibility of porous ceramics with soft tissue; application

to tracheal prostheses, J. Biomed. Mater. Res. Symposium,

No. 2, pp. 269-279 (1971).

[8] Hulbert, S. F., Klawitter, J. J., and Leonard, R. B., Compatibility of bioceramics with the physiological environ- ment. Ceramics in Severe Environments — Materials

Science Research, 5 (Plenum Pubhshing Corp., New

York, N.Y.)

[9] Rhinelander, R. W., Rouweyha, M., and Milner, J. C,

Microvascular and histogenic responses to implantation

of a porous ceramic into bone, J. Biomed. Mater. Res. 5,

pp. 81-112(1971).

[10] Klawitter, J. J., A basic investigation of bone growth into a porous material, Ph.D. Thesis, Clemson University (1970).

[11] Morrison, S. J., Tissue reaction to three ceramics of porous and nonporous structures, M.S. Thesis, Clemson University (1971).

43
[12] Bowman, L., Characterization of tissue growth into pellets

and partial sections of porous ceramics implanted in bone, M.S. Thesis, Clemson University (1971). [31

[13J Hulbert. S. F., Klawittcr, J. J., Sauer, B. W., and Matthews,

J. R., Characterization of tissue ingrowth into porous

bioceramics. Technical Report No. 2, Office of Naval [32

Research Contract No. N00014-71-A-0339-0001 (May

1972). [33

[14] Talbert, C. D., A basic investigation into the potential of ceramic materials as permanently implantable skeletal prostheses, M. S. Thesis, Clemson University (1969). [34

[15] Welsh, R. P., Pilhar, R. M., and Macnab, Ian., Surgical

implants. Journal of Bone and Joint Surgery 53-A, No. 5, [35 pp. 963-977 (July 1971).

[16] Greenlee, T. K., Jr., Beckham, C. A., Crebo, A. R., and Mal- morg, J. C, Glass ceramic bone implants, H. Biomed. [36 Mater. Res. 6, pp. 235-244 (1972).

[17] Hamner, James E., Ill, Reed, O. M., and GreuUch, R. C,

Ceramic root implantation in baboons, J. Biomed. Mater. [37 Res. Symposium, No. 3, pp. 1-13 (1972).

[18] Bartels, John C, Dental ceramics. Journal of Prosthetic

Dentistry 1 1, No 3, pp. 537-551 (1961).

[19] Monroe, E. A., Votava, Ward, Bass, D. B., and McMullen, [38; James, New calcium phosphate ceramic material for bone

and tooth implants. Journal of Dental Research 50, No. 4 [39

(1971).

[20] Benson, J., Elemental carbon as a biomaterial, J. Biomed.

Mater. Res. Symposium 5, No. 2, Part 1, pp. 41^7 (1971).

[21] Hench, L. L., and PaschaU, H. A., Direct chemical bond of bioactive glass-ceramic materials to bone and muscle, [40

Symposium on Materials and Design Considerations for the Attachment of Prostheses to the Musculo-Skeletal

System, Clemson University (1972).

[22] Klawitter, J. J., and Hulbert, S. F., Application of porous [41 ceramics for attachment of load bearing internal orthopedic

appliances, J. Biomed. Mater. Res. 5, pp. 161-231 (1971).

[23] Peterson, C. D., Miles, J. S., Solomons, C, Predecki, P. K., [42

and Stephen, J. S., Union between bone and implants of open pore ceramic and stainless steel, a histological study.

Journal of Bone and Joint Surgery 51-A, pp. 805 (1969).

[24] Topazian, R. G., Hamner, W. B., Talbert, C. D., and Hulbert, S. F., The use of ceramics in augmentation and re- [43

placement of portions of the mandible, J. Biomed. Mater.

Res. Symposium, No. 2, pp. 311-332 (1972). [44

[25] Leonard, R. B., Sauer. B. W., Hulbert, S. F., and Per-Lee,

J. H., Use of porous ceramics to obliterate mastoid

cavities. Symposium on Materials and Design Considera- [45 lions for the Attachnfent of Prostheses to the Musculoskeletal System, Clemson University (1972).

[26] Stanitski, C. L., and Mooney, Vert., Osseous attachment to [46 vitreous carbons. Symposium on Materials and Design

Considerations for the Attachment of Prostheses to the Musculo-Skeletal System, Clemson University (1972).

[27] Laing, P. G., Ferguson, A. B., and Hodge, E. S., Tissue [47

reaction in rabbit muscle exposed to metalhc implants, J. Biomed. Mater. Res. 1 , pp. 135-149 (1967).

[28] Ferguson, A. B., Laing, P. G., and Hodge, E. S., The ionization of metal implants in living tissues. Journal of Bone [48

and Joint Surgery 42-A, No. 1, pp. 77-90 (January 1960).

[29] Wood, N. K., Kaminski, E. J., and Oglesby, R. J., The signifi- cance of implant shape in experimental testing of biologi- cal materials: disc vs. rod, J. Biomed. Mater. Res. 4, [49 pp. 1-12(1970).

[30] Stinson, N. E., The tissue reaction induced in rats and guineapigs by polymethylmethacrylate (acrylic) and stainless

steel (18/8/Mo), British Journal of Experimental Pathology

45, No. 1 (1964). Cohen, Jonathan, Assay of foreign-body reaction. The Journal

of Bone and Joint Surgery 41-A, No. 1, pp. 152-166

(January 1959).

Brown, D. E., Tissue reaction to plastic and metal implants.

Arch Otolaryng. 88, pp. 283-287 (September 1968). Fitzpatrick, Barry, A comparative study of some implant

materials, Australian Dental Journal 13, pp. 422-434

(December 1968).

Roydhouse, R. H., Implant testing of polymerizing materials,

J. Biomed. Mater. Res. 2, pp. 265-277 (1968). Johnsson-Hegyeh, R. I. E., and HegyeU, A. F., Interaction

of blood and tissue cells with foreign surfaces, J. Biomed.

Res. 3, pp. 115-144(1969).

Johnsson, R. I., and Hegyeli, A. F., Tissue culture techniques

for screening of prosthetic materials. Annals of the New

York Academy of Sciences 146, pp. 66-76 (1968). Homsy, C. A., McDonald, K. E., Akers, W. W., Short, C,

and Freeman, B. S., Surgical suture-canine tissue inter- action for six common suture types, J. Biomed. Mater.

Res. 2, pp. 215-230 (1968).

Autian, John, Toxicologic aspects of implants, J. Biomed.

Mater. Res. 1, pp. 433-449 (1967).

Rosenbluth, S. A., Weddington, G. R., Guess, W. L., and

Autian, J., Tissued culture method for screening toxicity of plastic materials to be used in medical practice. Journal

of Pharmaceutical Sciences 54, No. 1, pp. 156-159

(January 1965).

Homsy, C. A., Ansevin, K. D., O'Brannon, W., Thompson,

5. C, Hodge, R., and Estrella, M. E., Rapid in vitro screening of polymers for biocompatibility, J. Macromol. Sci.

Chem. A4, No. 3, pp. 615-634 (May 1970).

Cohen, Jonathan, Tissue reactions to metals — the influence

of surface finish. The Journal of Bone and Joint Surgery

43-A, No. 5, pp. 687-699 (July 1961).

Evarts, C. M., Steffee, A. D., and McCormack, L. J., In- vestigation of canine tissue reaction to TEE-fluorocarbon

resin, to high-density polyethylene and to vitallium. Journal of Surgical Research 10, No. 2, pp. 91-94

(February 1970).

Ludwigson, D. C, Today's prosthetic metals. Journal of Metals 16, pp. 226-231 (1964).

Hulbert, S. F., Morrison, S. J., and Klawitter, J. J., Problems

associated with determining compatibility of biomaterials,

Symposium-Workshop on Research Techniques in Bio- materials Evaluation, Clemson University (March 1971).

Richbourg, Henry Leroy, Evaluation of a metal-ceramic

composite hip prosthesis, M.S. Thesis, Clemson University

(1973). Laing, P. G., Ferguson, A. B., and Hodge, E. S., Spectrechemical determination of trace metals in normal striated muscle in the rabbit. The Journal of Bone and Joint Surgery

41-A, No. 4, pp. 737-744 (June 1959). Ferguson, A. B., Akahoshi, Y., Laing, P. G., and Hodge, E. S., Trace metal ion concentration in the liver, kidney, spleen,

and lung of normal rabbits. The Journal of Bone and

Joint Surgery 44-A, No. 2, pp. 317-322 (March 1962).

Hulbert, S. F., Klawitter, J. J., and Richardson, W. C, Jr., Soft tissue response to a dense ceramic and varying surface

characteristics, 3rd Annual Meeting of the Biomedical

Engineering Society, Baltimore, Maryland (April 7, 1972).

Homsy, Charles A., Stanley, Rufus F., Anderson, M. Sidney,

and King, Joe W., Reduction of tissue and bone adhesion

to cobalt, alloy fixation appUances, J. Biomed. Mater. Res. 6, pp. 451-464 (1972).

44
NATIONAL BUREAU OF STANDARDS SPECIAL PUBLICATION 415,

Biomaterials, Proceedings of a Symposium held in conjunction with the Ninth AnnueJ Meeting of the Association for the Advancement

of Medical Instrumentation, New Orleans, La., April 19-20, 1974

(Issued May 1975).

Engineering and Biological Studies of Metallic

Implant Materials

Norbert D. Greene, Claude Onkelinx, Leon J. Richelle, and Peter A. Ward

Schools of Engineering, Dental Medicine, Medicine, and the Institute of Materials Science, The University of

Connecticut, Storrs, Conn., Box U-136 06268

The aim of this investigation is the development of improved alloys for short term (0.5 to 5 years)

orthopaedic implants. The program is interdisciplinary in nature — simultaneous studies of corrosion,

inflammatory response, and systemic effects of iron, nickel, cobalt, titanium, and tantalum base alloys are being determined. Corrosion tests under in vitro and in vivo conditions have been performed via linear polarization and other electrochemical methods.

The inflammatory responses of various implant alloys are being determined by both in vivo and

in vitro experiments. In vitro chemotactic assays on rabbit neutrophilic granulocytes and mononuclear

cells are performed in the presence and absence of appropriate metal ion concentrations. Corrosion rate data described above are employed to select the proper concentrations.

Systemic effects of metallic corrosion products have been determined via radioisotope and analy

tical techniques. Radioactive metallic salts at appropriate concentrations are injected intravenously in

rats of known age and sex. Following this, the concentration of metallic products is determined as a function of time in various biological samples (plasma, urine, feces, etc). These data permit the establish- ment of models which can predict the distribution of individual elements released from continuously

corroding metal implants.

Keywords: Corrosion; inflammatory response; metallic implants; systemic eff^ects.

1. Introduction

A successful orthopaedic implant must possess

sufficient mechanical strength to resist the stresses

it encounters in service and it should be compatible

with its host environment. A yield strength of 100,000

psi (.07 GPa), tensile strength of 150,000 psi (0.1 GPa)

and an elongation of 10 percent probably represent

minimal characteristics. High yield and tensile

strengths are needed to prevent bending and

breaking due to tensile overload and fatigue,

while adequate elongation is needed to avoid

brittleness. In addition, alloys should be easy to fabricate into complex shapes and consequently,

those which can be strengthened by heat treatment

after machining hold special interest. To be compatible, an alloy should corrode very slowly when

implanted and its corrosion products should not

affect adjacent tissue (inflammatory response) and

produce no adverse reactions in the rest of the

organism (systemic effects). It is especially important that the alloy resists crevice corrosion since

many implants are multicomponent devices containing shielded areas. None of the current implant alloys meet the above

requirements. Specifically, Type 316 stainless

steel lacks sufficient strength and corrosion resistance, especially crevice corrosion. Cast Vitalhum ' although much more inert and consequently more

compatible with adjacent tissue, tends to exhibit

brittle behavior. Also, the fact that it can be

fabricated only by casting is a limitation. Haynes 25

("Ductile Vitalhum") has the strength of cast Vitallium, together with sufficient ductihty to prevent

brittleness and to permit fabrication by conventional

methods. Its corrosion resistance, although not as good as Vitallium, is superior to stainless steel.

Ideally, its strength should be somewhat higher.

The actual biological compatibility of this alloy is not known since the manufacturer has not distinguished it from cast Vitallium in his products.

Titanium is very corrosion-resistant and may be

the least toxic although this has not been confirmed

by quantitative clinical tests. Unfortunately, the

strength of pure titanium is essentially identical to stainless steel.

Summarizing, all of the presently employed implant alloys are inadequate. Alternate materials

are needed since only small improvements can be

^ Certain commercial materials and instruments may be identified in this publication in order to adequately specify the experimental procedure. In no case does such

identification imply recommendation or endorsement by the National Bureau of Standards, nor does it imply that the equipment or instruments identified are necessarily the best available for the purpose.

45
accomplished with current implant alloys. Cold

working is utilized by most manufacturers to increase the yield strength of Type 316 stainless

steel, but this is done at the expense of decreasing

ductility and it cannot be conveniently applied to all devices. Coating stainless steel implants to improve

their corrosion resistance is limited by two factors.

First, orthopaedic implant surfaces are subjected

to high tensile and abrasive forces. Coating failures

would most likely occur at the most critical areas

(e.g., at bone plate-screw interfaces). More importantly, even if an ideal coating were applied to stainless steel, it would not alter its basically in- ferior mechanical properties. Similarly, there is no way to significantly increase the ductihty of Vitallium

or the strength of titanium. The only approach

capable of producing a significant improvement is a modification of alloy composition (i.e., substitution).

Historically, the development of implant alloys

has utilized empirical in vivo implantation tests. The American Society for Testing and Materials

(ASTM) is now standardizing such a test for the purposes of evaluating new implant materials [1].

^ Cylindrical specimens are inserted in the muscles

and bones of rabbits, rats and/or dogs and histological examinations made at various intervals up to 104 weeks. However, this test is insufficiently

sensitive to detect significant differences between

any of the present implant alloys, as shown by the

studies of Laing et al. [2]. Current implant materials,

in the absence of crevices, corrode so slowly that

histologically observable tissue reaction does not

occur. Thus, crevice-free samples of present orthopaedic alloys are sufficiently inert to render conventional implantation methods obsolete. The solution to this is the implantation of large numbers of

multicomponent appliance samples for extended

periods (e.g., 5 years or more). This of course would

be prohibitively expensive and time-consuming

except for more promising candidate materials.

Sensitive, accurate screening techniques for selecting potential alloys are therefore required for future

implant alloy development. In the future, as implants remain in place for longer periods, the possibilities of local inflammatory complications and toxic

systemic effects become increasingly important.

resemble wrought Vitallium on the basis of composition and corrosion resistance. Both possess markedly superior strength characteristics. Elgiloy is used as lead wires in cardiac pacemakers because

of its high strength, resistance to fatigue fracture

and excellent tissue compatibihty. There is Uttle

doubt that these alloys could function as improved

orthopaedic implants.

Although high yield and tensile strengths together

with good ductility are desirable for screws and bone

plates, there may be optimum limits for these

mechanical properties. There is some evidence to indicate that less ductility is required as tensile

and yield strength are increased. Also, it is Ukely

that beyond a certain tensile strength level it is impossible to torsionally fracture bone screws

during conventional orthopaedic procedures.

If so, then higher mechanical properties could

be "traded off' for better corrosion resistance, lower

toxicity or other features.

Since there are so many potentially useful

implant alloys, selection can be made on the basis

of minimum long term inflammatory and systemic

effects. To accomplish this, three major coordinated

research studies are used in this program. Corrosion measurements are being used to determine

the rate of metal solution into adjacent tissue. These

data are then utilized to determine inflammatory

and systemic effects. Also, optimum mechanical

properties for bone screws are being determined

by laboratory and in vivo experiments. The interactions between these different studies is schematically illustrated in figure 1. The corrosion

rates of most current and potential implant alloys are so low that loss of mechanical sti'ength is not an important consideration. Consequently, the

corrosion rates of most implanted metals are meaningless, per se. Corrosion rate only has

biological significance, and this can be described

in terms of inflammatory and systemic effects.

As noted in the diagram, there is feedback between

these two general biological effects and the alloy

corrosion rate. These effects are influenced by

both the specific metallic elements and their

rate of discharge (corrosion rate) under in vivo

conditions.

2. Research Approach

New, improved implant alloys together with rapid,

sensitive methods for predicting their in vivo

performances are urgently needed. Actually, a survey of present commercial alloys shows that

many materials with properties superior to these

now used for implants already exist. This is especially true of the superalloys — high strength, corrosion

resistant alloys developed for jet turbine applications. For example, Elgiloy and MP-35N closely

ORTHOPAEDIC SURGERY MECHANICAL

PROPERTIES

INFLAMMATORY

RESPONSE

IMPLANT

ALLOY SYSTEMIC

EFFECTS

CORROSION

RESISTANCE

^ Figures in brackets indicate the literature references at the end of this paper. Figure 1. Schematic diagram showing multidisciplinary ap- proach to the selection of new implant alloys.

46
Although mechanical properties can be used to qualitatively select potentially useful alloys for

orthopaedic applications, the optimum properties can only be defined by empirical mechanical tests with actual apphances (e.g., bone screw and plates).

Figure 1 illustrates this aspect by showing that

mechanical properties can be used to select suitable

alloys but the selection process can be improved by

feedback dialog between engineers and orthopaedic

surgeons. It follows that the selection of new implant

alloys requires an interdisciplinary research team,

competent in the areas of metallurgy, corrosion,

orthopaedic surgery, inflammatory response and

systemic effects.

3. Corrosion Studies

3.1. Background

There are two general features important in

evaluating the corrosion characteristics of a

material. These are the rate and type of corrosive

attack. For implant applications, the corrosion

rate determines the amount of foreign metal

compounds released into surrounding tissues.

To avoid significant local reaction, an implant

must corrode at a very slow rate. In general, rates

two or three orders of magnitude lower than those

commonly employed in conventional industrial

practice are necessary [3]. The measurement of

such very low rates introduces experimental diffi- culties since conventional weight loss methods

which are usually employed are insufficiently sensitive. Fortunately, electrochemical methods [3-8J,

particularly hnear polarization techniques and

charging curve analyses [3], are ideally suited for

the measurements of ultra low rates. Briefly, these

methods involve the measurement of the potentialcurrent relationships of a corroding specimen near

its corrosion potential. Over a potential range of

10 mv or less, there is a hnear relationship between

voltage and apphed current. The slope of this linear

region can be used to calculate the instantaneous

corrosion rate of a metal or alloy. The accuracy of

these techniques have been verified under in vitro

and in vivo conditions. Corrosion rates, especially

those of alloys protected by passive films, usually

vary with time. Thus, the ability to measure corrosion

as a function of time is very valuable in comparing

implant performance.

The evaluation of resistance to localized corrosion

such as crevice attack can also be accomplished

by electrochemical techniques. The initiation of

crevice attack is readily detected by monitoring

the corrosion rate of a sample via hnear polarization

techniques or charging curve analyses [9]. Also,

it has been shown that characteristic changes in

the corrosion potential also accompany initiation

of localized corrosion [9]. Locahzed attack, because

it is concentrated at isolated points, can be detected by microscopic (hght and electron) techniques. Such methods are also very useful for

examining specimens after in vivo exposure.

3.2. Procedures

Pretreatments of implants, such as sterihzation

procedures, cause marked changes in the corrosion

resistance of a given material under both in vitro

and in vivo environments [8, 10]. Also, the surface

finish of an implant (e.g., pohshed versus grit

blasted) has a major influence on corrosion characteristics [11]. Thus, in screening possible implant

alloys it is important that they be tested with a representative surface finish after exposure to typical chnical sterilization techniques. Where

possible, highly pohshed surfaces which have been

exposed to high temperature steam or dry heat

sterilization for long periods have been employed

since these conditions yield surfaces with the

greatest corrosion resistance [10].

The corrosion rates of various implant alloys are being measured as a function of time using linear

and transient linear polarization methods together

with charging curve analyses [3]. Also, when apphcable, passive current-time measurements [10] are used to monitor corrosion rates. Passive current

measurements are simpler than hnear polarization

techniques and they have the added advantage of

permitting continuous monitoring of corrosion

rate during the test period. Corrosion rates are monitored for periods ranging between 100 hours

and several months depending on the behavior of

a particular system. The corrosion rate of most

metals decreases exponentially with time [3, 10]

in a regular fashion. Therefore, the relative resist- ances of various materials can be compared on

the basis of short time exposures. However, the

most promising materials will be subjected to long

term tests in both in vitro and in vivo environments.

Resistance to crevice corrosion can be evaluated

with the above electrochemical measurements.

Since the corrosion potential must be monitored

during hnear polarization measurements, potential

changes which accompany the initiation of crevice

corrosion can be detected. Thus, the experimental

procedures used to determine corrosion-time characteristics are also capable of detecting locahzed

attack.

Isotonic sahne solutions at average body temperature can be used to quahtatively measure the

corrosion rate of implant alloys [8]. Although the

absolute value of the corrosion rates measured

under these conditions may not correspond to in

vivo environments, they are sufficiently accurate to determine if a potential alloy has a sufficiently low

corrosion rate to be considered for further testing.

The most promising materials will be subjected to in vivo tests utihzing specially modified electrochemical apparatus and procedures [4, 8]. Rats,

rabbits, dogs, and monkeys wiU be employed. Of

special interest are comparisons of different in vivo

47
environments. Previous studies [8J suggest that they

are very similar, but additional data are needed.

3.3. Results

Since crevice corrosion is the most common form

of corrosion damage observed on surgical implants,

we have devoted special attention to this aspect of

the program. A new technique for quantitatively

evaluating the crevice corrosion susceptibility of

alloys has been developed [11 J and is presently

being refined further. The method employs a paper

covered metal sample immersed in a suitable electrolyte. The presence of the paper simulates the

restrictive convection conditions within a crevice

without the resistance drop ("IR") errors normally

associated with shielded areas. Crevice corrosion

susceptibility is directly related to the change in

dissolution currents observed in the presence of a paper substrate.

Figure 2 illustrates the construction of the paper

shielded electrode used in these measurements, and

figure 3 shows the effect of paper shielding on Type

304 stainless steel — a material which is very sus- ceptible to crevice attack. Photographic paper with

its gelatin coating provides a very effective convective diffusion barrier and its influence on alloy

dissolution rate is very pronounced. Titanium, which

is almost completely resistant to crevice attack,

shows the opposite effect — paper shielding decreases

dissolution current (fig. 4). This decrease in current

is probably due to the increase in pH caused by the

accumulation of sulfate ions. Additional metals

and alloys are now being compared using this

procedure in several different electrolytes.

Passive currents versus time are being determined

in isotonic saline at 37 °C using newly developed

mini potentiostats [12]. These devices, based on

solid state operational amplifiers, are ideally suited

for multiple, long time measurements because of

THREADED ^MOUNTING ROD

Figure 3. Polarization curves of Type 304 stainless steel in sulfuric acid solution.

Increased current density in the active region for the specimen covered with photo- graphic paper illustrates the tendency for crevice corrosion.

o

<

u.s 1 1 1 1 1—1 1 1 1

1

1 1 1 1 1 1 1

0.6 *

TITANIUM

H; SAT O: 22*C

SS Hj SO,: BARE SPECIMEN M f

0.4

• 5N H.SO,; PHOTOGBAPHIC PAPEB T \ T

COVERING V •

0.2

5li HjSO,« 2M NaiSO,; BARE W 4

SPECIMEN 1 7

\

0

0.2

0.4

0.6! } ^corr

1 1 1 1 1 1 1 1 1 1 1 1 1 1 1 L 1

CURRENT DENSITY ^LZ/m^

Figure 4. Polarization curves of titanium in sulfuric acid

solution.

No sensitivity to crevice attack in the active region is indicated.

RUBBER BAND

SPECIMEN —

;

TEFLON GASKET

THREADED

HOLE

PAPER COVERING

Figure 2. Construction of paper-shielded electrode used to measure crevice corrosion susceptibility.

their simplicity and low cost. These potentiostats

together with specially designed polarization cells

and a shielded wiring system are being used to collect passive current-time data. MP35N alloy

possesses passive corrosion characteristics similar

to wrought Vitallium (HS-2 alloy). This together with

its outstanding mechanical properties indicates

potential usefulness as an implant alloy. However,

the recent systemic kinetic analyses of nickel distribution (see below) suggest that although nickel base

alloys such as MP35N may appear promising, their use should be approached with caution because of

possible toxic and carcinogenic effects.

48
4. Inflammatory Response

4.1. Background

The inflammatory response is, by definition,

the reaction of the body to injury. In this context,

the most important consideration is the change

in the functional behavior of blood vessels, the acute

inflammatory response being characterized by an

increased vascular permeability with a leakage of

fluids into extravascular sites. At the same time that

salts, water and proteins begin to leak from blood

vessels, a concomitant but independent accumulation of leukocytes (neutrophils or neutrophilic

granulocytes) occurs. It seems clear that the localization of neutrophils at a site outside blood vessels

is primarily a defensive reaction, since these cells

avidly ingest and kill bacteria, providing an important defensive function. This is particularly relevant

if there has been a break in the integument, such as

a laceration or abrasion in the skin. The inflammatory response then progresses, in the case of

wound heahng, to the subacute stage at which point

collagen is rapidly formed by fibroblasts and a scar

is produced. We are primarily concerned at this

point with the acute inflammatory response.

Although it is usually beneficial, in some situations the acute inflammatory response goes out of

control. The clinical results may be rheumatoid

arthritis or glomerulonephritis. One of the ways in

which these comphcations may occur is by the over- production or the inappropriate production of factors

that cause the accumulation of leukocytes.

Why the excessive accumulation of leukocytes at a site outside blood vessels should be deleterious to the tissue is explained by the fact that these cells

contain a battery of proteolytic enzymes, and it has

been demonstrated that elastin, coUagen, basement

membrane, and other structural proteins of tissues are suitable targets. Thus, in glomerulonephritis

there is actual evidence of basement membrane

degradation. In the vascular lesions of serum

sickness we find evidence for dissolution of elastic

fibers. In such situations where neutrophilic

granulocytes appear, it is readily apparent that

unless the inflammatory response is carefully

controlled, tissue injury may result.

The above discussion points out the proinflammatory effects of chemotactic agents [13-16]. If a compound impairs chemotactic responsiveness,

it will have anti-inflammatory effects by preventing

the delivery of cells to extravascular sites. This is the way in which several of the potent anti-inflammatory drugs have their eff"ect. Drugs showing antichemotactic effects include corticosteroids, derivatives of various quinohnes, and other agents that

by clinical evaluation have been shown to be

effective anti-inflammatory agents [17]. The significance of knowing a material has chemosuppressive

effects on leukocytes, aside from the obvious cUnical

apphcation, hes in the fact that such agents may

render an area of tissue susceptible to invasion by

micro-organisms. It seems possible that corrosion

products from implanted metals may affect leukocyte

migration either positively or negatively as described

above.

The possibility that corrosion products from

metalhc prosthetic devices may suppress the

chemotactic responsiveness of leukocytes is of

great interest, since in wound healing the arrival

(by chemotactic means) of neutrophils, assuming

excessive numbers are not dehvered, provides the area with a protective shield of phagocytes, whose

duty is to ingest and enzymatically degrade bacteria.

If, for any reason, inadequate numbers of neutrophils fail to arrive in the vicinity of a surgical

wound, serious trouble may develop in the form of a wound infection. The development of this complication in patients receiving metalhc implants almost

always dictates removal of the prosthesis. The

finding jsf antichemotactic activity in any of the

metal salts that represent corrosion products of

metalhc prosthetic devices is therefore of great

clinical significance and it is this part of the program,

based on prehminary data, that is described below.

Recent advances in the understanding of agents

that affect leukocyte migration have broadened our concepts regarding the control of inflammatory

responses [13-16]. For instance, it is now recognized

that several of the complement proteins of plasma

(and extracellular fluids) are substrates of endoge- nous enzymes (tissue proteases, plasmin and en- zymes derived from the complement sequence), the

products of which are the chemotactic factors for

leukocytes. The generation of any one of these

factors will set the stage for leukocyte accumulation,

a hallmark of the inflammatory response. Other

agents such as bacterial products are directly

chemotactic for leukocytes and require no other

proteins for the expression of their biological activity, [18]. If this happens in the context of bone or soft

tissues, the outcome wiU be an inflammatory re- sponse that may be inimical to the affected tissue.

Conversely, agents that suppress the chemotactic

responsiveness of leukocytes are potentially antiinflammatory agents.

4.2. Procedures

In vitro chemotactic assays on rabbit neutrophihc granulocytes and mononuclear cells have been

carried out using modified Boyden chambers [19]

that employ micropore fihers [20]. In this assay,

the leukocyte suspension is placed on one side of

the filter, the chemotactic factor on the other side.

After a 3-5 hour period of incubation at 37 °C, the

chambers are disassembled and the filters fixed,

stained, dehydrated and cleared in xylene, then

mounted on a glass shde. By conventional hght

microscopy the number of cells that have migrated

in five high power fields (selected at random) is determined. This reflects the chemotactic activity

49
of the material added to the compartment opposite

that containing the leukocyte suspension.

The cells used in this bioassay are rabbit

peritoneal cells, the neutrophils collected from exu- dates 4 hours after the installation of 0.1 percent

glycogen in saline into the peritoneal cavity, the

mononuclear cells collected 3-4 days after the intraperitoneal injection of mineral oil [20, 21]. For

the neutrophil system, micropore filters with pore

size 0.65 micrometer are used, while the mononuclear cell system involves the use of filters with

5 micrometer pore size[20, 21].

The following protocol is used for evaluation of a metal compound:

(a) Intrinsic chemotactic activity — test material

by itself in medium 199 (see below) is added to the

lower compartment (opposite the cell suspension).

(b) Effects on cell function — test material is added

to upper compartment containing cell suspension.

If the material suppresses chemotactic function,

the dose that causes 50 percent inhibition is determined.

(c) Effects of metaUic derivates on the chemotactic

factor — metal salt is added to the lower compartment containing the chemotactic factor. Of interest are those elements contained in potentially promising alloys. These include chromium, cobalt,

molybdenum, nickel, tintalum, tatanium, and

tungsten. These are added in the form of an

appropriate compound (e.g. chloride).

In all of these studies the chemotactic factor to be employed is the soluble, low molecular weight

factor produced by E. coli grown in a simple,

defined, protein free culture medium (medium 199).

If any of the metal derivatives are found to depress

chemotactic function of granulocytes, the complement-derived chemotactic factors C3a, C5a, and

C567 [22, 23] will also be tested to determine if the suppressive effect is specific for a given chemotactic factor or if it affects a variety of chemotactic

factors.

4.3. Results

The data in tables 1 through 3 provide prehminary evidence for the effects of metal salts on

chemotactic activity of neutrophilic leukocytes.

The findings can be summarized by the following

statements: (1) Several of the metal salts have a pronounced antichemotactic effect. None of those

studied have a prochemotactic effect. (2) Not all salts are antichemotactic. (3) The mechanism by

which the salts exert an antichemotactic function

seems to be by a direct, and in some cases irreversible, effect on the leukocytes rather than on the

chemotactic factor per se. In table 1 a series of metal salts at varying concentrations have been mixed with the neutrophils

obtained from glycogen-induced peritoneal exudates

of rabbits. The cell suspensions containing the salts were then placed in the test chamber and a standard

chernotactic factor a culture filtrate of the bacterium

(E. coli) deposited on the opposite side of the filter. Numbers of cells responding chemotactically to the bacterial factor were determined by the standard

method. By extrapolation, the dose of metal salt

causing 50 percent inhibition of cell migration was

determined. The gold salt was the most antichemotactic agent found, inhibiting 50 percent of the cell

responsiveness at a concentration of 2.3 X 10 "^M.

Table 1. Inhibitory effects of various metal salts on chemotactic

function of neutrophils

Salt tested Concentration (M) giving 50%

inhibition of chemotaxis

HAuCU-SHzO 2.3 X 10-" RhCla 3 X 10-" CuCl, •2H2O 3 X 10-" CrCl3-6H20 3 X 10-" MnCl2 •4H2O 5 X 10-" AICI3 • 6H2O 10-3 CuCl 10-3 FeCl2-4H20 10-3 FeCla • 6H2O 10-3 WO3

ZnCl2

M0O3 > 10-3 C0CI2 • 6H2O

NiCl2 • H2O

Table 2. Direct inhibitory effect of metal salts on neutrophil

chemotactic function

Salt tested

Inhibition of response'*

chemotactic Percent

Salt mixed

with factors

Salt mixed

with cells

HAuCl4-3H20 86 100 RhClj 41 71 CuCl2-2H20 52 32

CrCl3-6H20 5 22 MnCU-ffliO 5 80

''Salts added to cell suspension or to chemotactic factor in a concentration of 5 X 10 -"M.

It is interesting to note that gold salts have been

used for years to treat patients with rheumatoid

arthritis. The mechanism for this action has never been explained. It is quite possible that the antichemotactic, and the anti-inflammatory, effect of

gold salts shown in table 1 is the basis for this

clinical effect.

As described in table 1, several other metal salts

also have an antichemotactic effect, whereas the

last five in the list have no demonstrable effect on chemotaxis in doses as high as 10"^M.

In table 2 salts (at 5 X lO-^M) have been mixed

either with leukocytes or with the chemotactic

(Bacterial) factor in the chemotaxis chamber and

the resulting chemotactic response of the cells

assessed. Five different salts were tested, each of

which was previously determined to have an anti50
chemotactic effect (table 1). As shown in four of

the five cases (cupric chloride excepted) the metal

salts had an accentuated effect if mixed with the

cell suspension, in contrast to their effects when

mixed with the chemotactic factor. These findings

indicate that in most cases the antichemotactic

effect of the metal salts is the result of a direct

action on the cells. These findings were further explored (table 3)

where washing experiments were done. Briefly,

cells were treated with various metal salts and then

one-half of each suspension was diluted, centrifuged,

and the cells suspended in culture medium lacking

the particular metal salt. It is evident that in several

cases, particularly with the gold and the manganese

salts, that substantial inhibitory effects persisted

when the salts were removed. These data indicate

that in some instances the inhibitory effect on

chemotactic responsiveness of leukocytes is irreversible whereas in other cases this is not the case. The data confirm those in table 3 and indicate that

the inhibitory effect of the salt is the result of a

direct action on the ceDs.

Table 3. Irreversible inhibitory effect of metal salts on neutro- phil chemotactic function

Salt used^

Inhibition of chemotactic

response. Percent

Cells washedCells not washed

HAuCU-H^O

RhCla

10020

1546

75

10090

100

50

100

CuCl2-2H20

CrCl3-6H20

MnCl4-4H20

''At a concentration of 5 X 10 "^M.

5. Systemic Effects

5.1. Background

The study of the systemic effects of any foreign

molecules introduced into the body has two fundamental aspects. The first component of such a study attempts to determine the amount of the

active form of the molecule which is present, at any given time, in various parts of the body. The

second component of the study deals with the interactions between the active form of the molecule

and some receptor sites. In general, such interaction

wiU modify one or more physiological or biochemical

phenomena, thus leading possibly to the development of a pathological condition.

It is clear that when metallic elements are re- leased into the body by the corrosion of implanted

materials, their distribution in time and space is an essential step before conducting any logical

study of their potential toxicity. In discussing the

hmitations of current techniques for the evaluation

of new biomaterials, Hegyeh [24] insists on "the

lack of willingness to select, adapt and actually

conduct objective tests of high predictive value

for human patient response". She also suggests

that "any implanted device should be considered

as a potentially hazardous bulk pharmaceutical."

For the study of systemic side effects she notes

the shortcomings of in vitro methods that result

from the fact that physiological conditions in the

body are not adequately duplicated. Moreover,

some in vitro conditions such as gas atmosphere,

acidity, evaporation, temperature, etc. are not

always controlled as they are in vivo. These parameters influence the interactions between implants

and the biological environment. In the past, experiments have been carried out to study the distribution of trace elements in the body. For instance,

A. B. Ferguson et al. [25, 26] measured the trace

metal ion concentrations in various organs of the

rabbit. Using those data as a basehne, they were

then able to investigate the release of trace ions

from embedded metal implants and their subsequent uptake by various organs. These tedious

experiments (6 to 16 weeks) did not allow estabhshing in detail how these ions are distributed in the

body. Differences were noted from animal to animal,

between various metals, etc. Some organs, like

spleen and liver, appeared to pick up more material

than other tissues. Sunderman [27] has produced a

bibliography that describes how a variety of elements (beryUium, cadmium, chromium, cobalt,

iron, lead, nickel, selenium, zinc) find their way

into various organs when administered via different

routes (dietary, intramuscular, subcutaneous,

intravenous, intraperitoneal, etc.). Though these

papers are specifically oriented towards carcinogenicity, their content clearly indicates that the

systematic approach that we employ for studying

the distribution of trace elements released by

metalhc implants, would provide much needed

information.

In toxicology and pharmacology, kinetic analysis

[28] is, by now, a classical method to study "drug

transference." This term refers to all the phenomena

which result in a given drug or molecule being

transported between various points in the organism

or being transformed from one chemical form into

another. Both classes of phenomena are grouped in one general category — transference — because the

mathematical analysis of these phenomena is based

on a number of common concepts proper to kinetic

analysis [29]. Such methods have been used to study the metabohsm of calcium, an element

normally present in the body, and" radioactive

strontium, a fallout product resulting from the

explosion of nuclear devices [30—33].

The specific aims of this part of the program are:

1. Estabhsh the distribution throughout the

body of trace elements released by implanted

biomaterials by making use of appropriate

kinetic models. Single intravenous injections

of radioactive isotopes will be used for this

purpose.

51
2. Predict, using these kinetic models, the time course of the concentration of elements in

various parts of the body, when they are released from implants, taking into account

the in vivo rate of corrosion as determined by

the studies described above.

3. Analyze, chemically and radiochemically,

those organs and tissues which appeared to retain potentially toxic amounts of the elements

for sufficiently long periods of time, after in

vivo implantation of biomaterials.

4. Study possible interactions between the elements retained in toxic amounts and biological

receptor sites in order to determine the

precise nature of any pathological process.

5.2. Procedures

The rationale of the methodology is apparent

from the above list of specific aims. In a first step,

rats of known age and sex are injected intravenously

a single dose of a radioactive tracer of the element

under study. By sampling blood at regular intervals,

the rate at which the serum loses the injected dose

is determined. Over the same time period, urine

and feces are collected. If most of the loss is found

in urine, feces, or both, then the kidneys and/or

liver will be considered as likely target organs. If

the difference between what is lost from the blood

and what is found in both urine and feces is considerable, the skeleton is then sampled and analyzed, since it acts as a sink for many metals.

Once a model is estabhshed, it can be used to predict the distribution of the element when,

instead of a single intravenous injection, the element

is continuously released, at a known rate, by the

corrosion of the implanted material. It is also possible to check these predictions, either by continuous

perfusion of animals (acute experiments), or by.

implanting samples of the material containing a radioactive isotope of the element under study

(chronic experiments).

6. Results

Nickel was the first metal chosen for study in

terms of whole body kinetics. The reasons for this

choice were as follows: (a) nickel is a major component of many alloys used in implants; (b) corro- sion characteristics of nickel containing alloys

have been determined both in vivo and in vitro,

[8, 10]; (c) its potential toxicity is well established

particularly in relation to carcinogenesis [35, 36].

Rabbit data were kindly made available to us by Dr. Sunderman (Department of Laboratory

Medicine, Health Center, University of Connecticut). The experiments were conducted as follows:

18 albino rabbits were administered radioactive

nickel chloride (^^NiClz) in a single intravenous

injection of 0.24 mg Ni/kg body weight (equivalent

1.43 mCi/kg). At intervals ranging from 1 hour to 7 days after the injection, blood samples were

collected and the concentration of radioactive nickel

in the serum was determined. The animals were

kept in metabolism cages in order to collect urine

in which radioactive nickel was subsequently

measured. Similar measurements were also performed in the bile collected from two animals

under anesthesia.

The results indicate that nickel is rapidly cleared

from the plasma during the first two days after an

intravenous injection, but disappears at a much

slower rate from 3 to 7 days (fig. 5). An average of

78 percent of the administered ^^Ni (i.e. 0.187

/Ag/kg of rabbit) was recovered in urine during the

first 24 hours after the injection. The biliary excretion was estimated at approximately 9 percent

(4.6 fxglkg body weight) of the administered dose.

No data were available for the net elimination of

^^Ni in the feces.

8-

( hours)

Figure 5. Time-course of ^^Nilll) concentrations in serum of

rabbits following a single i.v. injection o/^^NiCU. (Each point

represents the mean of measurements in at least 3 rabbits.)

Experiments with rats were conducted as follows.

Female Wistar rats, 10 weeks of age, were administered 100 /LtCi of ^^Ni (II) in a single intravenous

injection. The injection was performed under ether

anesthesia into the left subclavial vein. Immediately

thereafter, the animals were placed in metabolism

cages in order to coUect feces and urine separately.

Blood samples were withdrawn from the tip of the

tail with the help of capillary tubes at a series of

time intervals between 2 hours and 8 days after

the injection. Plasma was separated by centrifugation and weighed before radioactive nickel determination. Feces and urine were collected separately

for a period of 3 days starting at injection time.

They were dried in crucibles then ashed in a muffle

furnace (24 hours at 300 °C, followed by 24 hours

at 600 °C). The ash was dissolved in 6 N HCl over

a hot plate and the solution transferred into a 500

52
ml volumetric flask. While bringing the solution

to volume, the pH was adjusted to 3 with NaOH.

Radioactive nickel was measured in all the samples

with a TRI—CARB Packard scintillation counter

(scintillating mixture: Aquasol, New England

Nuclear Co.). The count rate was corrected for

quenching using an internal standard.

As was the case with the rabbit, the concentration

of ^^Ni in the plasma decreased very rapidly during

the first 2 days after the injection (fig. 6). Thereafter

the decrease took place at a much slower rate. After

5 days, most of the radioactive nickel had been

eliminated in the urine and to a lesser extent in

the feces.

As demonstrated in figure 5 for the rabbit, and

in figure 6 for the rat, the disappearance curve

of ^^Ni (II) from the plasma can be approximated

very closely by a sum of two exponential terms:

S= Aie-°.' + ^2e-°2' (1)

in which S stands for the concentration of nickel

in the plasma (in /Ag/liter) and t is time measured

from the time of injection (hours).

This fact strongly suggests that nickel dilutes

0 50 100 150 200

( hours)

Figure 6. Time-course of *^Ni(II) concentrations in plasma of

rats following a single i.v. injection of ^^NiCU. (Each point

represents the mean of measurements in at least 15 rats.)

itself within a distribution volume made up of two

compartments and is ehminated from at least one compartment according to first order kinetics

(within the range of concentrations observed in

these experiments). Compartment I, which includes

the plasma is the central compartment in which

^^Ni (II) is injected at time ^ = 0 and from which

elimination takes place. Compartment II represents

a hypothetical volume related to the first compartment by an exchange process. We assume that at time t — 0, ^^Ni (II) is rapidly and uniformly distributed within the first compartment.

In the rabbit experiment, the time course of

^^Ni (II) concentration in the plasma is very well

approximated by the following equation:

S ^ 1165 e-0 0925< + 4 g-0.0084(

(2)

(units: /iig/liter; amount injected: 0.24 mg/kg

body weight)

whereas in the rat experiment the following equation

was determined:

S = 226.3 e-« iio3r+ 0.572 e-o oi4< (3)

(units: /Ag/Uter; amount injected: 17.0 /ig

per animal)

These equations were obtained by fitting expression

(1) to the average values of nickel concentration

in plasma as a function of time for each group of

animals. The degree of fit is demonstrated in

figures 5 and 6. A mathematical analysis of the two-compartment

model, based upon the data reported above, can be worked out in terms of differential equations.

This allows a complete quantitation of nickel

kinetics in the body. In particular, the distribution

volume of each compartment as well as the clear- ances of nickel exchange and excretion can be

evaluated.

Furthermore, once the parameters of the model

are known, it is possible to predict the time-course

of nickel concentration in both compartments when

nickel is continuously released into the plasma at a constant rate. From what is known about nickel

corrosion, one can estimate that in humans who

have implants made of a nickel-containing alloy,

the rate of nickel release from the device can range

between 5 and 500 mg/year per individual. This

corresponds approximately to a range of 0.81-0.0081

ixgjh per kg of body weight on the basis of an aver- age body weight of 70 kg for humans. The following

table gives the estimated steady state values of

nickel concentration in plasma resulting from 3

different input rates within that range.

Infusion rate

/Lig/h per kg of body weight

Steady state Ni

concentration

/Ag/Iiter

Rabbit Rat

0.81 45 19.9 0.081 4.5 1.99 0.0081 0.45 0.20

These figures have to be compared with the normal

range of nickel concentration in human plasma. In

a recent paper, McNeely et al. [38] reported a value

of 2.6 /Ag/Hter (±0.08) for a presumably normal

population sampled from the area of Hartford

(Conn.), while in Sudsbury, Ontario, the site of the

largest open-pit nickel mines in North America,

the plasma level was 4.6 ju,g/liter (±1.4). It is quite

53
clear from the table that an implant can constitute a continuous source of nickel resulting in far higher

plasma levels that those due to the heaviest

atmospheric pollution.

Although there is no direct evidence as yet that

these high levels of plasma nickel are necessarily

toxic, one should certainly exercise caution with

regard to the potential hazards associated with its use in implants. Indeed our results reinforce the

opinion voice by Hueper [35] according to which

"the evidence on hand indicates that metal implants

which contain nickel and which remain over long

periods in human tissues might create delayed

potential cancer hazards to their recipients."

PREDICTED CURVES FOR THE CASE OF A COIITISUOUS INFUSION OF

NICKEL INTO TNE 8L0OO AT A RATE OF 1 ug/hr/kg OF BOD» UT. rabbit

rat

50 100

(hours)

150 200

Figure 7. Computer-plotted curves for the predicted concentrations o/^Ni(II) in plasma or serum of rats and rabbits, assuming a continuous i.v. infusion o/^NiCl2 at a rate of 1 fj.g Ni/kgh.

We can add filially that the steady state under a continuous input is reached very quickly, as

illustrated in figure 7 which pictures the transient curve from the time at which the continuous infusion

is started. As can be seen from the figures, it takes

about 7-8 hours to achieve a plasma concentration

half the maximum value.

This research is supported by the National

Science Foundation and the Institute of Materials

Science, University of Connecticut.

7. References

[1] Recommended Practice for Experimental , Testing for Biological Compatibility of Metals for Surgical Implants,

ASTM Committee F-4 on Surgical Implants, ASTM. » PhUadelphia, Pa.

[2] Laing, P. G., Ferguson, A. B., Jr., and Hodge, E. S., J. Biomed. Mat. Res. 1, 135 (1967).

[3] Jones,-D. A., and Greene, N. D., Corrosion 22, 198 (1966).

[4] Colangelo, V. J., Greene, N. D., Kettelkamp, D. B., Axelander, H., and Campbell, C. J., J. Biomed. Mater. Res. 1,

405 (1%7).

[5] Skold, R. v.. and Larson, T. E., Corrosion 13, 139t (1957).

[6] Stern, M., Corrosion 14, 440t (1958).

[7] Stern, M., and Weisert, E. D., Proc. ASTM 59, 1280 (1959).

[8] Revie, R. W., and Greene, N. D., J. Biomed. Mater. Res. 3,

465 (1969).

[9] Jones, D. A., and Greene, N. D., Corrosion 25, 367 (1969).

[10] Revie, R. W., and Greene, N. D., Corr. Sci. 9 , 755, 763

(1969).

[11] Picard, R. J., and Greene, N. D., Electrochemical Reactions

at Paper Shielded Metal Surfaces to be published.

[12] Greene, N. D., Moebus, G. A., and Baldwin, M. H., The

Mini Potentiostat: A Versatile Power Source for Electro- chemical Studies, accepted publication in Corrosion.

[13] KeUer, H. O., and Sorkin, E., Separatum Experientia 24,

641 (1968).

[14] Mergenhagen, S. E., Snyderman, R., Gewurz, H., and Shin, H. S., Significance of complement to the mechanism of

action of endotoxin, Current Topics in Microbiology and

Immunology 50, 37 (1969).

[15] Tempel, T. R., Synderman, R., Jordan, H., and Mergenhagen, S. E., J. Periodont. 41, 71 (1970).

[16] Ward, P. A., Arth. & Rheum. 13, 181 (1970).

[17] Ward, P. A., J. Exp. Med. 124 , 209 (1966).

[18] Ward, P. A., Lepow, I. H., and Newman, L. J., Am. J. Path. 52 , 725 (1968).

[19] Boyden, S., J. Exp. Med. 115, 453 (1962).

[20] Ward, P. A., Cochrane, C. G., and MuUer-Eberhard, H. J.,

J. Exp. Med. 122, 327 (1965).

[21] Ward, P. A., J. Exp. Med. 128, 1201 (1968).

[22] Shin, H. S., Snyderman, R., Friedman, E., MeUors, A., and

Mayer, M. M., Science 162 , 361 (1968).

[23] Taubman, S. B., Goldschmidt, P. R., and Lepow, I. H., Fed. Proc. 29,434(1970).

[24] Hegyeli, R. J., Limitations of Current Techniques for the Evaluation of Biohazards and Biocompatibility of New

Candidate Materials, Medical AppUcations of Plastics, H. P. Gregor, Ed. (Wiley, New York, N.Y., 1971) pp. 1-14.

[25] Ferguson, A. B., Akahoshi, Y., Laing, P. G. Hodge, E. S.,

J. Bone and Joint Surg. 44-A, 317 (1962).

[26] Ferguson, A. B., Akahoshi, Y., Laing, P. G., Hodge E. S.,

J. Bone and Joint Surg. 44-A, 323 (1962).

[27] Sunderman, F. W., Jr., Food and Cosmetics Toxicology 9,

105 (1971).

[28] Ariens, E. J., Molecular Pheirmacology (Academic Press,

New York, N.Y., 1964).

[29] Riggs, D. S., The Mathematical Approach to Physiological

Problems, (Williams and Wilkins, Baltimore, 1963).

[30] Berman, M., and Weiss, M. F., Users Manual for SAAM,

PubUcation No. 1703, U.S. Public Health Service.

[31] Aubert, J. P., Bronner, F., Richelle. L. J., J. Clin. Invest. 42 , 885 (1963).

[32] Bronner, F., Aubert, J. P., Richelle, L. J., Saville, P. D., Nicholas, J. A., Cobb, J. R.. J. CUn. Invest. 42, 1095 (1963).

[33] Richelle, L. J., Bronner, F., in Proc. First Europ. Symp.

Calcif, Tissues, H. J. J. Blackwood, Ed. (Pergamon Press,

London, 1964).

[34] Moore, C. V., Dubach, R., in Mineral, Metabolism, and

Advanced Treatise, Vol. 3, C. Comar, F. Bronntr, eds. (Academic Press, New York, N.Y. 1963).

[35] Heuper, W. E., Carcinogenic hazards from arsenic and

metal containing drugs. Potential Carcinogenic Hazards

from Drugs, R. Truhart, Ed., (Springer-Verlag, Berlin,

1967), pp. 7^104.

[36] Sunderman, F. W., Jr., The current status of nickel carcinogenesis, Am. CUn. Lab. Sci. (in press).

[37] Van Soestberger, M., and Sunderman, F. W., Jr. ^Nicomplexes in serum and urine after injection of ^NiCl2

(in press).

[38] McNeely, M. D., Nechay, M. W., and Sunderman, F. W., Jr., Clinical Chemistry 18, 992 (1972).

54
NATIONAL BUREAU OF STANDARDS SPECIAL PUBLICATION 415,

Biomaterials, Proceedings of a Symposium held in conjunction with the Ninth Annual Meeting of the Association for the Advancement of Medical Instrumentation, New Orleans, La., April 19-20. 1974

(Issued May 1975).

Materials Characterization of Implantable Porous Electrodes

R. B. Beard, J. F. DeRosa, S. F. Dubin, L. Sturm, and R. M. Koerner

Drexel University, Philadelphia, Pa. 19104

and

A. Miller

Temple University, School of Dentistry, Philadelphia, Pa. 19122

Porous platinum and palladium black electrodes have been used as cathodes for reducing body

oxygen in implantable hybrid ceOs supplying energy to pacemakers. The power-generating capabilities of a cell are greatly decreased under load, i.e., increased current density, due to polarization at the electrode interface. A greatly increased surface area of the porous electrodes, i.e., number of sites for the electrode reaction, permits a greater current density with lower overpotential or polarization.

Similarly, at the pacemaker stimulating electrode and at electrodes used in making electrical impedance

measurements there is polarization and a consequent loss of energy in a charge transfer at the electrode

interface. Porous electrodes in these cases have also greatly reduced the overpotential. The physical

electrochemical properties of the porous electrodes have been characterized by specific adsorption,

i.e., BET measurements; scanning electromicrographs; galvanostatic and potentiostatic measurements; and electrical impedance measurements. Histopathological studies have been made in order to deter- mine the biocompatibility of the tissue-electrode interface.

Key words: Implantable hybrid cells; pacemakers; palladium black; platinum black; polarization; porous electrodes.

1. Introduction

Implantable hybrid ceUs consisting of catalytic

cathodes where dissolved oxygen is reduced and

sacrificial anodes are galvanically oxidized are capable of powering pacemakers fl-10], ^ carotidsinus, bladder and electrophrenic stimulators [11].

The implantable hybrid cell has a long shelf life,

costs no more than conventional batteries and has

a potential for lasting 10 years or longer. The power

generating capabilities of the hybrid cell depend

upon the cathode electrode potential and current

drain. Under zero current drain the electrodes in

the body develop an open circuit voltage. When

current is drawn through the load, the electrode

potential decreases from its open circuit value due

to the overpotential developed at the electrode.

Higher power outputs are produced by electrodes

which can maintain higher potentials for given

current drains. In-vivo and in-vitro studies have

demonstrated that the power of the cell is decreased

due to the overpotential at the catalytic cathode.

There is sufficient oxygen available in the subcutaneous tissue according to Roy et al. [12], to depolarize for current loads in excess of 200 /xA

which will produce sufficient power to drive pace-

' Figures in braclcets indicate the literature reference at end of this paper.

makers. At Drexel, implanted hybrid cells powering

commercial pacemakers pacing dogs use porous

cathodes which have operated with current loads of

610 ^l^ or 25 /i-A/cm [12, 13].

In-vivo and in-vitro studies on platinum and

palladium black cathodes consisting of hydrophobic

binders of teflon and hydrophihc binders of polyvinyl

chloride and sihcone rubber have demonstrated

considerable overpotential with a resulting power

loss [8, 13].

In order to increase the number of available sites

for reducing the oxygen and to obtain a more

biologically compatible interface [14], platinum

and palladium porous cathodes were fabricated

using the techniques of powder metallurgy [15, 16].

The polarization or overpotential at the various types

of cathodes has been studied using the galvanostatic

technique. In-vivo studies have been made implanting the electrodes under load subcutaneously in

the lateral abdomen and thorax of dogs. Material

and biological compatibility characterization studies

have been made on the various types of cathodes.

The surface areas of the palladium and platinum

powders and compacts have been determined using

the Brunauer-Emmett-Teller (BET) gas adsorption

technique and the scanning electron microscope.

The compatibility of the electrodes have been

determined by histopathologic evaluation of the

55
tissues surrounding the porous metal implants

[15, 161. Besides the problem of decreasing the electrode

polarization (i.e., overpotential) of the power

source, there is also the problem of a significant

waste of stimulation energy at the cardiac pacemaker electrode surface due to polarization [17, 18,

19, 20]. A decrease in the electrode surface in- creases the current density, lowering the threshold

voltage for stimulation, but increasing the polarization voltage [20]. There is therefore an optimum

size electrode and pulse shape [17, 42] from the

standpoint of both stimulation threshold and pacemaker energy.

Although a variety of electrodes have been investigated for delivering an electrical stimulus to the heart [17, 18, 20, 21, 22, 23], noble metals such

as platinum and platinum alloys are commonly used

and recommended because of their inertness to body tissue [17]. Very high thresholds due to infection have been reported [23]. The current

threshold is determined by the minimum current

density at the electrode required for stimulation of

the surrounding myocardium. Fluctuations which

occur in the current threshold immediately after

implantation are attributed to the process of tissue

formation. When the myocardial tissue reaction is complete the threshold stabilizes. Infection around

the electrode with tissue degeneration results in a further increase in threshold current [23]. The above

investigations [17, 19, 21, 22, 23] have been limited

to electrode materials currently used in cardiac

pacemaker stimulating electrodes. The polarization that occurs at the smooth platinum electrodes

used as pacemaker stimulating electrodes is much

greater than in the case of platinum black electrodes.

The porous stimulating electrode on the other

hand can be made small with the electrode polarization being reduced without effecting the current

density just outside the electrode surface area. The

rise in threshold is attributed not to an increase in

the electrical impedance of the scar tissue surrounding the electrodes, but rather due to an increase in

the size of the electrode surface due to the con- ductive scar tissue with a resulting decrease in

current density [20]. Porous ceramics with 45 to 100 fjim pore size have shown thin fibrous encapsulating layers of 4-6 cells in thickness with a good

blood supply or no observable encapsulation as contrasted to thick encapsulation around impervious

discs [24, 25]. Thus, porous stimulating electrodes

should have lower electrode polarization and a lower threshold of stimulation due to tissue in- growth. Tissue ingrowth wiU also help to fix the

stimulating electrode in place. One of the principal

complications remaining with the transvenous

technique has been reported to be electrode displacement in the right ventricular inflow tract [26].

Electrode polarization is also a problem in impedance measurements on biological materials,

and corrections for this polarization must be made

if excessive measurement errors are to be avoided

[27J. Corrections are not only large but uncertain

at sufficiently low alternating current frequencies

where polarization becomes large [28, 29]. At

present, electrodes in our laboratory and elsewhere

[27] are coated with platinum black by electroplating from a Kohlrausch solution. This normally

reduces electrode polarization to a relatively low

level for frequencies above 100 kHz. However, for

measurements on physiological fluids at lower

frequencies, corrections for polarization must be

made [18, 19, 28, 29, 30, 31] or different electrode

configurations used. For example, the four-electrode

nuU technique [28] has been applied to impedance

measurements over the range of 10 Hz to 1 kHz.

Here voltage across the sample is applied by one

pair of electrodes and another pair of noncurrentcarrying electrodes is used to measure the voltage across the sample. The electrodes have a platinum

black coating to further decrease polarization.

Platinum black electrodes deteriorate with use in

biological fluids which is believed due to macromolecules effecting the platinum black surface [27]. If macromolecules are blocking the sites of the

platinum black interface and decreasing the ease

of charge transfer, this would account for the increased polarization when the electrodes are used

in protein solutions. Again, the development of

porous electrodes could be appHed to improve

biological impedance measurements.

DeRosa and Beard [30, 40] at Drexel have considered polarization in terms of rate determining steps.

The effects of electrode d.c. potential, pH and oxygen coverage on the mechanism of cathodic reduction of oxygen and their influence on the cathode

polarization has been studied [30]. Anodic polarization under the variation of the above parameters

has also been studied. The anodic as weU as the

cathodic reactions are important in a.c. polarization

in electrical impedance measurements and at bipolar pacemaker electrodes.

2. Experimental Procedures

2.1. Modified Transient Galvanostatic Method

for Studying Polarization [33]

When using the galvanostatic technique to study

the variation of electrode potential as a function of

current density, the cell current is changed suddenly

from zero to a given value, and the resulting variation of potential with time is recorded. The electrode

under test and a counter electrode are placed in a Ringer's solution which has an ionic concentration

approximating that of subcutaneous tissue fluid.

The potential of test electrodes with reference to a silver-silver chloride electrode is measured as shown in figure 1. The cell current is changed from

zero to a given current density by a current pulse

of approximately 1 ms. width and 500 /xA amplitude.

The electrode potential is measured with respect

56
Resistance

Oscilloscope

Counter

Electrode

Reference Electrode

Saline

Figure 1. Transient galvanostatic technique.

to the reference electrode. Since the current through

the reference electrode is very small, there is negligible overpotential and voltage drop in the solution

between it and the electrode under test. When a stimulus such as a step current is apphed to an electrode interface, most of the initial

current is used to charge up the double layer while

later on the constant current density is used to promote electronation or a charge transfer reaction

[33]. By recording the variation of the potential

with time between a nonpolarizable reference

electrode and the test electrode, the kinetic parameters of the electrode reaction can be obtained

[33, 34].

2.2. The Brunauer, Emmett, and Teller (BET)

Gas Adsorption Method

The Brunauer, Emmett, and Teller (BET) Gas

Adsorption Method has been used to measure the

surface area of porous compacts and powder [16,

35-38]. The utilization of the technique requires one to obtain an experimental gas adsorption isotherm from which the volume corresponding to monomolecular gas coverage is estimated. Once

this volume is known along with the average area

of the molecules of the adsorbate, the surface area

of the specimen can be calculated.

The experimental setup which was used is similar

to Kreiger [38] and is basically a combined gas

buret and mercury manometer. To obtain more

accurate pressure measurements, a thermocouple

was also used. Before testing began, the specimens

were outgassed at 0.5 ^tmHg at 200 °C for approximately 16 hours. Helium was used for dead space

determination and nitrogen for the adsorbate during

the actual measurement portion of the tests. All

tests were made with the sample container immersed in a Uquid nitrogen bath. The resulting

pressure versus volume data results in a Type II Brunauer isotherm [39] in which the "knee" of

the curve (fig. 2a) is generally accepted as being

monomolecular gas coverage of the surface involved.

For both convenience and greater accuracy, this

data can be plotted as shown in figure 2(b) with the

resulting straight line slope and intercept giving

monomolecular coverage directly, i.e.,i;m=l/(6+m).

PRESSURE

(b)

Figure 2. (a) Type II Brunauer isotherm, (b) BET free-surface linear plot.

57
The product of Vm, the cross-sectional area of the

absorbate and Avogadro's number corrected to standard temperature and pressure gives the

surface area, Sbet2.3. Scanning Electronmicrographs

A scanning electron-microscope was used to study the loose powders and compacts. The preparation of biological materials for the scanning electron

microscope was as follows: Upon removal from the

animal, the sample was washed in saline to remove

mucous and fixed in glutaraldehyde. The sample

was then using a succession of ethyl alcohol-water

baths, i.e., 58, 90, 95, and 100 percent alcohol. The

alcohol was dehydrated then replaced with amyl

acetate. The samples were then freeze-dried in

conventional manner, coated with gold, and ex- amined in the scanning electron microscope.

2.4. Impedance Measurements

Electrode polarization, which greatly influences

electrical impedance measurements, can be studied

both from time and frequency domain measure- ments [19, 21, 33, 40]. Plastic cells of poly(tetrafluoroethylene) and poly(methyl methacrylate), consisting of thin-walled measuring chambers

surrounded by circulating water for temperature

control, were used for studying a variety of porous

electrodes. For time domain measurements a constant current pulse source is connected to the

electrodes and the amplitude of the current pulse

is increased in steps. The potential across the electrodes and current are measured at each level. The

impedance measurements were made using a Wayne Kerr B221 Admittance Bridge which measures the parallel conduciance and capacitance.

The amplitude of the applied sinusoidal voltage is only 25 millivolts RMS which insures the electrical

impedances are measured in the linear range, i.e.,

not in the range where impedance is a function of

the amplitude of the appHed signal. Measurements

were made at 25 °C over the frequency range from

200 Hz to 20 kHz. While cell temperature was controlled and the pH and oxygen concentration was

monitored with an oxygen meter.

3. Materials

3.1. Fabrication of Porous Electrodes

Following powder metallurgy techniques, the

loose powder is placed in a hardened metal die and

is compressed under high pressure. Under pressure,

the particles become cold welded together and the

magnitude of the compacting pressure greatly

influences the density and strength of the resulting

compact. The required compacting pressure depends upon the ductility of the base metal from

which the powder is formed, e.g., aluminum compacts at 10,000-30,000 psi (0.07-0.2 GPa), ferrous

powders at 30,000-60,000 psi (0.2-0.4 GPa), and

tungsten at 80,000-150,000 psi (0.5-1.0 GPa).

Since no published information was available

regarding compacting pressure versus density or strength, a number of platinum and palladium

powder electrodes * were prepared at different

compacting pressures. These pressures were 10,000,

30,000 and 50,000 psi, which resulted in compacts

with gradually increasing density. The green

strength of the compacts was sufficient to withstand

sterilization in a steam autoclave and they were

therefore left unsintered.

The stimulating electrodes have been fabricated

under a variety of techniques. Some electrodes

were fabricated from platinum- and palladium-black

powders with an isostatic compaction technique,

then sintered. In the fabrication of large pore sizes,

poly(methyl methacrylate) (PMMA) powder was

blended with the palladium blaclc and compacted

under a hot press technique [41].

4. Results

In-vivo results for the porous electrodes used as power source electrodes demonstrated that there

was a six-fold increase in power density for these

electrodes [15]. Figure 3 shows the comparison

between a porous palladium electrode and others

with plastic binders. Porous Pd electrodes implanted in dogs over a period of 146 days under loads

of 70 fjL A/cm^ had no adverse tissue reactions [43,

44].

Figure 4 illustrates that the surface area of

platinum and palladium is of the order of lO'* times

greater than foil. The large surface area, with a large increase in sites, accounts for the marked

decrease in overpotential.

Figure 5 demonstrates the decrease in polarization of the porous stimulating electrodes which

are compared to the standard platinum-iridium tips.

The potential difference versus time across the

platinum-iridium electrode has, as shown by its initial jump, a considerable potential across its electrodes and, as shown, in the curve of figure 5, a large overpotential across the interface. The amplitude of the 1 ms. current pulse was held to 200 fxA,

which gave a current density of 200 /uA/cm^.

' Type JSM-2, Japan Electron Optics Laboratory Company, Ltd.

' Certain commercial materials may be identified in this publication in order to adequately specify the experimental procedure. In no case does such identification imply recommendation or endorsement by the National Bureau of Standards, nor does

it imply that the material are the best available for the purpose. 'The metal powders were obtained from Matthey Bishop, Inc., Malvern, Pa.

58
20

10 20 30 40 30 60

Time In days

Figure 3. Power density versus time. An implanted porous palladium compact and platinum black with poly (tetrafluoroethylene) and poly (vinyl chloride) binders

POTENTIAL vs. TIME

PLATINUM

o PALLADIUM

10COMPACTION

30 50

PRESSURE (ksi)

80

60

40

20

Platinum - Iridium

954 Pd

Si P^WA

0.2 0.4 0.6 0.8 1.0 1.2

Milliseconds

RE 4. Surface area of platinum and palladium compacts FIGURE 5. Potential difference versus time across a platinumfor various compacting pressures. iridium electrode and a porous palladium electrode.

59
Figure 6. Scanning electronmicrograph of porous palladium

electrode; magnification 100; 97 percent palladium 3 percent

PMMA; sintered at 450 °C.

The scanning electromicrographs of figures 6 and

7 demonstrate that the pore sizes are irregular and

vary in size, with some pores 70 (xm long. Figure 7

shows tissue ingrowth in the larger pores. More

recent electrodes have been fabricated with 100

fxm pore sizes and are in the process of being

evaluated. It appears that porous palladium and

platinum structures must, like ceramics, have pores

greater than 45 jttm for marked tissue ingrowth

[25]. Pore sizes above 100 /Am produce a weakened

structure so that there is a balance between strength

and tissue ingrowth.

Additional linear alternating current electrode

polarization impedance measurements are being

made to determine the effect of various background

ligands on the electrode polarization impedance.

Ahrland [46] has shown that the tendency of metal

ions (acceptors) to form stable complexes depends

on the relative coordinating affinities of nonmetal

groups (donor atoms). Impedance measurements

with various Ugands such as iodide, bromide, and

chloride appear to influence the electrode polarization in a predictable manner.

This study was made possible by the support of

the Biomaterials Program of N.S.F. Grant No.

GH 33748.

5. References

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Figure 7. Enlarged scanning electronmicrograph of figure 6

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[3] Reynolds, L. W., and Konikoff, J. J., Utilization of Bioelectric Potentials, First Interim Report, 1963, Ames

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60
[14] DeRosa, J. F., Beard, R. B., and Koerner, R. M., Improved

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i

61

NATIONAL BUREAU OF STANDARDS SPECIAL PUBLICATION 415,

Biomaterials, Proceedings of a Symposium held in conjunction with the Ninth Annual Meeting of the Association for the Advancement of Medical Instrumentation, New Orleans, La., April 19-20, 1974

(Issued May 1975).

Properties of Fibrous Biomaterials With Statistically Dispersed Orientation

E. Y. Robinson

Prototype Development Associates, Santa Ana, Calif. 92705

Many factors influence the interaction between bone or soft tissue and implanted synthetic bio- materials, e.g., biocompatibility, implant configuration, functional requirements, bone and tissue structure, and relative mechanical properties. Of the many active factors, one aspect is considered

here: the theoretical consequences of the fibrous-lamellar structure of bone, and of the degree of

fibrous orientation present in the individual lamellae. This orientation is known to be statistically

dispersed about certain preferred directions in each layer, with possible large orientation changes

from layer to layer. Analysis of this type of structure is presented with graphical illustration of the

effects of orientation and of statistical dispersion of orientation on conventional engineering material

parameters.

New biomaterials are being evolved which combine fiber reinforcement with polymer resin matrices (e.g., graphite fibers/polyethylene matrix). Such materials may be tailored to yield certain

specific properties by controlling fiber orientation and quantity. Materials of this type are of interest

in implants which may be required to behave in a similar fashion to adjacent bone tissue (as in bone

sphnts, hip prostheses, etc.). The analysis presented here provides a rapid and convenient basis for calculating the effect of controlled and dispersed fibrous orientation on material properties.

The methods described lay a base for first approximations and show certain directions which

should be followed in further investigation. Graphical results include examples for both fibrous bone

tissue models and synthetic types of fiber-reinforced biomaterials.

Key words: Biomaterial properties; biomechanics; composites; fiber orientation; fibrous biomaterials.

List of Symbols

(r(i) Stresses

e(i) Strains

El Principal Young's modulus

E2 Secondary Young's modulus

^ = E2IE1 ; anisotropy parameter

Vl Primary Poisson ratio

Gi Shear modulus in 1-2 plane direction;

elastic constants without subscripts

denote the isotropic asymptotes

Qiij) Stress-strain matrix

U{i) Coefficients of the multiple-angle formulation for coordinate transformation of

Qiij)

6 Angle of orientation of the natural longitudinal direction measured positive

counter-clockwise from the reference

coordinate axis; variable of integration

fid) Frequency distribution function of orientation angle

Qiij) Stress-strain (stiffness) matrix of the

multilayer laminate, analogous to the

single layer properties defined above

* This paper presents the results of one phase of research carried out at the Jet Propulsion Laboratory. California Institute of Technology, under Contract No. NAS

7-100, sponsored by the National Aeronautics and Space Administration.

Qiij) Isotropic stress-strain (stiffness) matrix

a Intended orientation; integration limit

for e ±8 Bounds of orientation dispersion

1. Bone and Soft Tissue as Layered

Fibrous Materials

The structure of bone tissue is highly fibrous and

the fiber networks are often found in aggregated

layers or lamellae. The nature of the lamellar bone

structure varies within the musculoskeletal system

and also within localized regions. Enlow in Evans

[1]^ gave a concise description of several characteristic bone structures, two of which are reproduced here in figure 1. Various mechanisms of bone development described by Enlow result in generally striated or at least localized layering of bone. In some cases this

oriented lamellar structure develops by successive

layering around a nucleus, such as an osteon. In

other cases the process seems to be a progressive

densification, as in cancellous compaction. In still other cases the bone exhibits regular strata with

^Figures in brackets indicate the literature references at the end of this paper.

63
Figure 2b. Schematic of one type of alternating lamellar

aggregation.

The lamellar structure is clearly evident, as well

as a change of character in adjacent series of layers.

A schematic of lamellar orientation shown in figure

2b shows two types of lamellar arrangement. The

properties of the layered structure must depend

not only on the lamellar arrangement but also on

the details of fibrous orientation within each lamella.

Evidence for these types of arrangement has been

found in bone tissue. LocaUzed aggregations of

Figure 1. Schematic layered bone structures.

(a) Zones B and C are surface plates; Zone A shows cancellous compaction.

(b) Laminar bone with stratified vascular canals. (After Enlow in Evans, [IJ).

Thus, in one case the material may be laced with

highly oriented, dense (and probably strong and

stiff) unidirectional structural elements such as osteons, while in another the properties reflect a system of pores which may be highly organized,

"en echelon." Each of these structures imparts a

particular and pecuUar behavior to gross material

properties.

Many published studies are devoted to elucidating

the fibrous sfrxicture not only of bone but soft tissue as well, for example Walton in Buckles [2], Currey

[3], Dempster [4], and others. Direct studies of

single osteons have been pursued for some time,

notably by Ascenzi et al; see for example in Evans

[1]. Figure 2a, taken from Ascenzi in Evans [1], shows an osteon segment after fracture testing.

* m

Figure 3a. Fibrous mats and oriented structure of mandibular

tissue in cortical plates.

64
Figure 3b. Fibrous mats and oriented structures in mandibular tissue.

Successively magnified views of dentin.

Type 1 are shown in figure 2b, in osteons, while

Type 2 structure is suggested in the SEM photograph of figure 3a. Implications of bone structure

orientations have been considered by Evans and

Vincentelli [5] , Bonfield [6, 7] , Cooke [8] , Currey [9]

,

and others. The Type 2 lamellar arrangement, illus- trated in figure 2b, is shown in an electron micrograph of mandibular cortical tissue in figure 3a.

Figure 3b shows a sequence of progressive

magnifications in dentin. Here the effect of orientation must be attributed to the parallel ranks of

tubules (see Lees and RolUns [10]), while fibrous

structures around the tubules appear to be quite

random, without obvious preferential orientation.

However, the highly organized geometry strongly

suggests some corresponding fiber orientation.

65
Figure 4. Fiber patterns in femoral cartilage [After Millington et al., in Kenedi, [11])

Studies of fiber arrangements in soft tissue have

yielded many revealing and instructive configurations. Figure 4, from Millington et al., in Kenedi [11],

illustrates changes in fiber organization through

the thickness of femoral cartilage. Adjacent layering

and change in predominant orientation direction

between adjacent layers of dermal collagen may

be seen in figure 5. The cross-linking betw^een adjacent layers may

not be great, allowing a "scissor" action between

such layers. On the other hand, if there is fairly

intimate binding between layers, the "scissoring"

is inhibited and a reaction stress is induced in each

layer. If the fibers are adversely oriented, for example in a highly transverse direction, the induced

stress may be intolerable and splitting results, as shown in figure 5b. Figure 6 is a schematic representation of a distended carotid artery, showing

relative compaction of the inner collagenous layers

and an apparently random arrangement on the

outer boundary.

While the bone tissue, characterized by high

mineral content, may be analyzed in somewhat

simpler terms (linear elastic) than soft tissue, there

are evidently many microstructural aspects which

are similar, and a properly conceived analytical

approach could shed additional light on the behavior of both tissue types.

2. Peculiar Behavior of Anisotropic

Materials

Before estabhshing a simplified fiber-reinforced

laminate model for tissue, it is instructive to re- view briefly some of the pecuhar properties which

accompany material anisotropy. In engineering

structures these properties tend to be of secondary

importance and are often avoidable by design. In

the case of natural tissue these effects could be of

importance and are therefore of interest here.

Consider a single layer of material containing

perfectly ahgned fibers bonded together by rela66
tively weak cross-link bonds and/or a binder matrix

which permeates the layer. The fiber longitudinal

and transverse directions are the "natural" directions of the layer, representing axes of elastic

symmetry. If forces are applied along these directions, the material behavior is described as orthotropic. If forces act at some arbitrary angle to a

natural direction, the material behavior is generally

anisotropic. The mathematical implications are shown in figure 7.

Figure 6. Schematic representation of distended carotid artery.

ORTHOTROPIC

Y,T

ANISOTROPIC

T Y L

e

L

"

1 E

L

E T

0

."L \

1 0

0 0

1

~

1 "Y

'x ^X 'l

S ."x 1 ^x E Y

"2

'^XY 'l '2

1

LT

XY

Figure 7. In-plane stress-strain relation for orthotropic and

anisotropic material.

PLATE TUBE NO

LOAD

AXIAL LOAD

SHEAR

STRAIN TORSIONAL

STRAIN

Figure 8. Illustration of shear coupling.

Figure 5. Dermal collagen under load (After Millington et al.,

in Kenedi, [11]).

(a) "Scissor" action, (b) Transverse splitting.

The nonzero terms of the bottom matrix mean

that normal or transverse stress wiU cause shear

strain. This effect is illustrated in figure 8.

67
Figure 8 shows the shear distortion of an anisotropic sheet subject to axial tension. Similar behavior of a tubular section is manifested by rotation

accompanying axial tension. The shear distortion

of a single + 6 anisotropic ply is prevented by

attaching a second layer oriented at — 6. This is referred to as balanced orientation. However, such

a two-layer laminate will tend to twist when axial

force is apphed, as illustrated in figure 9.

NO LOAD

NO SHEAR

Figure 9. Illustration of twist coupling.

Combining this two-layer laminate with another

two-layer laminate having opposite twist tendencies

to form a "symmetric" and balanced laminate will

result in conventional, uncoupled deformation under

applied tensile or compressive in-plane forces, and

the material system will be orthotropic. However,

if this symmetric, balanced laminate is subjected

to bending, as illustrated in figure 10, the twisting

tendencies reappear. Generally speaking, it is not

possible to simultaneously eliminate both loadtwist and bend-twist coupling.

3. Defining a Simplified Analytical

Model

It is clear from even a general perusal of published

results that fiber-orientation in layered tissue is not

precise. While certain directions predominate,

there is dispersion of orientation about these directions. Cox [12] provides a useful and general analysis

applicable to a broad class of fiber networks.

For the present purpose the model will be

limited to systems of two-dimensional layers.

Applied or induced stresses in the transverse

"thickness" directions are considered negligible.

This assumption can lead to overestimation of

elastic modulus of three-dimensional materials.

The effect of fibrous structure in a lamella is approximated by considering the layer to be a homogeneous, orthotropic continuum. The orthotropic properties of this equivalent layer are derived by analytical consideration of individual

fiber properties and matrix or cross-linking

properties.

All layers in the laminate are assumed to strcdn

contiguously; that is, the strain is uniform through

the transverse thickness direction. If "scissoring"

and interfacial slippage occur between layers, this

assumption will tend toward overestimation of

modulus and underestimation of Poisson's ratio.

Orientation dispersion of fibers is assumed to be

uniform but bounded symmetrically about a predominant, preferred direction a: that is, in the

range of a — d to a + 8 where 8 is the extent of

orientation dispersion. An illustration of dispersedorientation networks is shown in figure 11. The

network, as it might appear to the eye, is shown

above while the dispersion angle is depicted in the

lower portion of figure 11.

Figure 10. Illustration of bend-twist coupling.

As the number of layers increases and the orientation patterns become more dispersed, these

coupling effects '^end to diminish. They should,

however, be recognized and may in fact play a

significant role in the microscopic behavior of single

fibers such as helically configured collagen fibrils.

DISPERSION ANGLE 8

0° 40° 120° 180°

Figure 11. Schematic illustration of dispersed orientation networks.

In the analysis which is carried out in the following section, two material types are considered and

analyzed by the same method (Robinson, [13]). The

68
Table 1. Basic unidirectional lamellar properties used in analyses

Longitudinal Transverse in-plane Shear modulus Major

Material elastic modulus. elastic modulus. GN/m2(10«psi)' Poisson

GN/m2(10«psi) GN/m2(10«psi) ratio

Carbon/polyethylene 69.0 (10) 0.69 (0.1) 0.69 (0.1) 0.3

Bone tissue 47.5 (7) 5.1 (0.75) 3.45 (0.5) 0.1

first is a synthetic fiber- reinforced, polymer-matrix

material, a type which is becoming of great interest

in biomedical and prosthetic application; see, for

example, Giltrow [14], Musikant [15], and Sclippa

[16]. One of the attractions of the material is the

prospect of being able to tailor material properties to meet specific and individual objectives. This is achieved, in principle, by varying the quantity of

fiber reinforcement and the degree and nature of

fiber orientation. In reality, many additional,

subtle factors can influence resultant properties of a fiber-composite material. The material considered

here is a carbon fiber-reinforced polyethylene.

The second material is a hypothetical bone-tissue

model built up of idealized fiber-reinforced lamellae.

The basic orthotropic properties ascribed to an

ideal unidirectional lamella of each material are shown in table 1. The data in this table are determined in part from published data on collagen,

hydroxyapatite and highly oriented tissue, and in

part estimated, since complete orthotropic characterization of natural tissues could not be found in a fairly persistent literature search.

4. Formulation and Analysis

In current engineering structural application of

fiber-reinforced composites, the material is usually

in the form of a multilayer laminate containing

parallel, unidirectional fibers in each layer or ply

(lamina or lamella). Consequently the analysis for

such materials is referred to as laminate theory or laminate analysis. Published work in this field was

strongly influenced within the last decade by substantial U.S. Air Force research and development

sponsorship directed toward improved lightweight

materials for aircraft and space applications. The

introduction of carbon and graphite fibers, boron

filaments, new single-crystal whiskers and other

new fiber types is a direct result of this sponsorship.

Also, for many years U.S. Forest Products Laboratory personnel have investigated theoretical and

experimental behavior of nonisotropic materials

related to plywood and fiberglass analysis. They

are an excellent source of useful analysis and copi- ous experimental data.

Lamination theory is covered in the Structural

Design Guide for Advanced Composite Materials

[17] and Ashton et al. [18]. Some elements of laminate analysis are given here to provide a useful

description of the equations appropriate to partly

random layer orientation. A two-dimension generalized Hooke's Law is used, with material stiffness

elements Q{ij) defined by the matrix equation

cT{i)=[Qiij)][e{j)].

(See Glossary for definitions of symbols and terms

used in this article.)

The 3X3 stiffness elements of Q{ij) depend on

the orientation of the principal elastic axes (natural

directions of each layer or ply) relative to the api

phed forces. When these directions coincide (i.e., 6 = 0), the matrix elements are the natural direction

orthotropic properties given by

0(11)= 0(22)=^ (1)

C(12)=5T^ <?(66) = G,.

When the natural direction is oriented at 6, the

transformed stiffness coefficients of each layer,

Qiij), are given by.

<?'(11) cos (2^) + C/3 cos m

Q'{22) cos (2^) + C/3 cos m

<?'(12) = u,-u. cos m

(?'(66) cos m

<?'(16) = + 1/2 U2 sin (26) + Us sin m

Q'{26) - + 1/2 f/2 sin (26) - Us sin m

(2)

C/i, U4, U5 are invariants, independent of orientation. The coefficients U{i) are given in terms of the

natural direction orthotropic parameters by

69
Ux = 1/8 [3(?(11) + 3(?(22) + 2^(12)

+ 4(2(66)]

t/2 = l/8 [4C>(11) -4(3(22)]

£/3= 1/8 [(2(11) +(3(22) -2(3(12) -4(3(66)]! (3)

f/4 = 1/8 [(3(11) + (3(22) + 6(3(12) - 4(3(66)]

f/5= 1/8 [(3(11) +(3(22) - 2(3(12) + 4(3(66)]

C/5= [^(l)-^(4)]

In this formulation, the problem is limited to twodimensional elastic behavior.

The stiffness properties of the two-dimensional

multilayer laminate are calculated by summing the

stiffness contributions of each layer. Note that for

balanced orientations of similar layers (at + 6 and - 6) the (>'(16) and (3 '(26) terms will cancel. These

are shear couphng terms discussed earlier.

In order to account for orientation dispersion we

consider the case of a large number of hypothetical

layers in a laminate of unit thickness, with orientations distributed from — S to +8. The laminate

stiffness elements are obtained by summing the

stiffness contributions from each hypothetical

ply. For large n and small angular interval, we can

pass from the sum to the integral:

Q{ij) = r Q'{ij)f{d)dd (4)

where fi9) is the normahzed frequency distribution

of orientation, and Q' (ij) are given by eq (2).

4.1. Quasi-Isotropic Limits

If the dispersion angle extends from — 7r/2 to + 77/2, the material becomes apparently isotropic.

Therefore, the in-plane isotropic elastic limits of a composite result when the orientation distribution

is uniform, /(0) = 1/(28), and unbounded:

1 f^l'^ Qiij) = - Q'{ij)dd.

TT J -77/2

The terms from eq (1) including functions of 6 will

drop and this equation shows that the "invariant"

terms £/i, f/4, and f/5 are the unbounded 6 averages

of the Q' {ij) for symmetrical, uniform probabihty

distributions. The isotropic stiffness matrix is,

then.

. Qiij) =

isotropic

f/4 0

f/4 f/l 0

f/l0 0 7

The quasi-isotropic engineering parameters are

f/4 v =

f/l (5)

It is instructive to examine the invariants in terms

of the conventional engineering constants of the

individual orthotropic layers:

8

3 + 4|i + ^ (3 + 2t.i-4i.?^

1-M

[1-4^

E,

'l + 6i.i + 4t^f I^J

8 1 J

E,

"1+4^

E, 'l-2^i-4^?|-)'

8 1 -^^\ J

f/4=^ (6)

f/..=

^

This form exhibits the contribution of element

Poisson ratio and anistropy to the invariant parameters. The theoretical value for the isotropic

Poisson ratio, given by U4IU1, shows a fundamental

trend to values around 1/3 which is only slightly

altered by varying the anisotropy parameters. This

tendency to Poisson ratios around 1/3 is seen to be

a structural consequence of combining large numbers of randomly oriented orthotropic lamaUae

into a quasi-isotropic laminate. For layers of random

unbonded filaments = E2 = Gi = 0) acting as a contiguous network, eq (5) predicts v = 1/3. This

result sheds hght on the well-known fact that nearly

all macroscopically isotropic materials exhibit

Poisson ratios in the vicinity of 0.3. The isotropic

sheet may therefore be modeled as an ensemble of

random orthotropic layers. For highly anisotropic

constituent layers (small ^), the quasi-isotropic

two-dimensional parameters can be estimated by

E = Eil3, G = £1/8, V = 1/3. Cox [12] gives a classic

discussion of both two- and three-dimensional fiber

networks and reinforced systems.

If the basic layer properties are known, that is,

£1, E2, Gl, and vl, the quasi-isotropic limits may

be computed with equations (1), (3), (5), or (6).

4.2. Off-angle Lamallae Containing Dispersed

Fiber Orientation

In the bone tissue model the concern is with

random partial orientation of individual layers

distributed about an intended nominal direction, as well as the effect of orientation dispersion of the

constituent filaments. The orientation dispersion

is assumed to be distributed in the interval — 8 to + 8 about the nominal predominat direction angle a.

70
The stiflfness elements for unit thickness laminate

are given by

Q(ij)=r"'Qiij)fie)dd (7)

For a uniform distribution, /(^) = 1/28, the integration gives

Q{ll) = Ui + U2Cos (2a)

sin 2§

28

+ Us COS (4a) (

sin 48

(2(22) = f/i-C/2 COS (2a)

sin 28' ~28"

+ C/3C0S (4a)

sin 48

48

(?{12) = U4-U3Cos (4a)

(2(66) = t/5 - f/3 cos (4a)

sin 48

48

>(8)

^2 . . / sin 2 8

,2(16) =ysm (2a) ( —

I rr • /J \ /sin 48

+ C/3sm (4a) (^-^

^(26) =1 sin (2a)(^-i|f

t/3 sin (4a)

sin 48

48

As the dispersion limit 8 approaches zero, the ideal

anisotropic results of eq (2) are obtained. When 8

reaches 7r/2, the isotropic limits result. Equations

(8) are identical in form (homologous) to eq (2) if modified values are used for U-z and t/3:

Uo = U,

sin (28)

28

sin (48)

48

Equations (8) provide a new generahzed form of

the constitutive stiffness elements given by eq (2). The homologous formulation allows simple alteration

of existing computation procedures to incorporate

effects of uniform, bounded dispersion and for

analysis of partially oriented muhilayer laminates

(lamellar materials).

The laminated material description is greatly

simplified if the orientation at + a is balanced by

orientation at —a. In such cases, the material is said to be orthotropic, instead of generally anisotropic, and the conventional engineering constants

may be computed by

(2(12) -_Qi22)

(9)

(2(22) (2(11)

£. = (2(ll)[l-i^?] E2 = iE, G = (2(66)

using the Q{ij) values from eq (8).

5. Results of Numerical

Computations

Figures 12 to 14 show the results obtained with

the equations of Section 4 applied to the idealized

material given in table 1 of section 3. Two types of

structural arrangements were analyzed. In the first,

the multilayer laminate was assumed to contain

perfectly oriented (no dispersion) unidirectional

fibers in each layer, with the orientation of alter- nate layers changing from + ^ to — ^ with respect to the laminate principal direction. In the second case

the fibers within the lamellae were assumed to have

dispersed orientation, ranging uniformally from — 8

to + 8. Figures 12a and 12b show the computed

results for these two structure types for a carbon

fiber/polyethylene matrix composite. The quantity

and type of carbon fiber reinforcement was chosen

to yield the particular set of orthotropic layer properties shown in table 1. The properties of these

fibers offer a very broad range of reinforcement

and property control with polymeric matrix composites. Figures 13a and 13b illustrate the analogous

results for a hypothetical fibrous bone-tissue model.

The two structure types are combined to show the

range of orthotropic properties which might be

observed in a lamellar structure, wherein each

lamella contains orientation dispersion and the

preferred layer orientation alternates from layer to layer. The graphical result for such a fibrous bone

tissue model is shown in figure 14.

71
10

3

Qo

b

zo

rE2

\'"'

/

A

\

\

/ \ /

/ \

/ \ \ /

\ /

/

/

\ \

\

/

/

\

'\

1

/

/ \

t

30 60

ALTERNATE LAYER ORIENTATION, ±0

90

- 4

O

t—a

z

- 2 O

oa.

- 1

_ 0

Figure 12a. Calculated in-plane properties of a carbon fiber- polyethylene matrix composite.

8 3

on6 2 z?

.

/

\

/ \ — >

i h

V

\

1 \

\

\

1 \

-\

1

1

1 \ /

/

30 60 ALTERNATE LAYER ORIENTATION ±890

Figure 13a. Calculated in-plane orthotropic properties of a

fibrous bone tissue model.

±9 layering, no orientation dispersion.

t6 layering, no orientation dispersion.

10

3

o

O

5

i/i 6 4 zo

>-

m \El\

i— \\

1 N

/

1

\^ —

N

1

(- \ \

/

o

I—

z

2 OoQJ 0

E

5

LU

B 4

ZDQo

5 3 6

O

>-

1

0

El

7^

1 -1—

\

\

/ \

1 —/

^E2

f

0.6

0.4 g

Zo

0.2 2

30 60 DISPERSION ANGLE (8) , deg

90 30 60 DISPERSION ANGLE ( S )

, deg

90

Figure 12b. Calculated in-plane properties of a carbon fiber- polyethylene matrix composite.

Bounded uniform orientation dispersion (preferred orientation is 0°)

Figure 13b. Calculated in-plane orthotropic properties of a

fibrous bone tissue model.

Bounded uniform orientation dispersion ^preferred orientation is 0°)

72
8^

2

6 3 2

O ,

0.

m

m V

N p

V

1

\ \ V

f

0.8

0.6

g

I—

z

0.4 O

0.2

30 60 AITFRNATF LAYER ORIENTATION -19

90

Figure 14. Calculated in-plane orthotropic bone properties

showing range of values for ±9 layers having up to ±15 degrees

orientation dispersion.

6. Discussion of Results

The analytical results presented in section 5

illustrate the ranges of material property variations

implied by the ideahzed fibrous structures considered. It should be noted that the resultant

orthotropic properties are interrelated. Experimental data on real bone tissue samples can present

combinations of properties which are not describable by the type of model considered here. This

should not be surprising in view of the photomicrographs presented in section 1 showing twisted,

kinked, irregular fibers, with varying degrees of

cross-Unking. The layering is also irregular in

thickness and physical arrangement, and even the

highly lamellar osteons may curve and change

direction rather sharply within the cortex. The

range of properties shown in figure 14 appears to be typical, more or less, of much bone-tissue data

pubUshed in the hterature. However, the applicability of this model is limited, and many of its underlying assumptions are probably too ideahzed.

Nevertheless, a basis is provided for calculating

detailed interaction between synthetic materials

used in prostheses and adjacent bone tissue.

Examination of the equations in section 4 and

results in section 5 reveals that direct, axial loads

in a laminate will be borne primarily by the stiffest

lamaUae; however, the transverse stresses, induced

by differential contraction according to individual

layer Poisson ratios, may be much more equally

shared if the lamellae act contiguously. The price

for transverse load sharing when axial properties

of the layers differ sharply will be induced internal

stress and a tendency for the interface to delaminate

in efforts to relieve these induced stresses.

The material properties selected for the basic

unidirectional lamella are, to some degree, based

on pubhshed data for collagen and hydroxyapatite

(or similar crystalHne materials). The analysis

provides a logical connection between the properties

of a lamellar material and the nature and degree of

orientation dispersion which is present. Conversely,

if the dispersion is characterized a priori and the

properties measured, it is possible to use the

analytical model in reverse to infer the maximal

properties of constituent reinforcements.

Another potentially important feature of the

analytical results is the behavior of Poisson's ratio.

In certain regions of orientation dispersion, as shown

in figure 14, the value of this parameter becomes

very high and is sensitive to orientation. Thus, as orientation of the fibrous tissue changes under the

action of apphed forces, the instantaneous elastic

properties also change in correspondence with the

instantaneous orientation. Poisson ratio for the

model analyzed would be expected to increase with

applied tension loads if the dispersion angle was

less than about 35 degrees. If the angle exceeded

35 degrees, the Poisson ratio would be expected to decrease. This could be a significant factor in bone

remodeling dynamics, since a high (and possibly

increasing) Poisson ratio would cause the fibrous

network to close up under tension and to open up

under compression. In a flexural loading situation

(as in the femur or tibia) the tensile side would

inhibit the flow of nutrients into and through the

tissue fine-structure, while on the compressive side

the "opening" of the structure should encourage

the flow of nutrients. This type of structural response

may affect localized resorption and bio-electric

behavior.

This is consistent with some observed cases

of bone remodehng in which apposition occurs on

the compressive side and resorption on the tensile

side of a flexuraUy loaded bone. This direction of

remodeling acts to reduce flexure arising from

eccentric compressive loads until equilibrium is reached.

7. Conclusions and Recommendations

It is a well-known fact that natural tissue is comprised of fibrous networks of various types,

including systems of well-defined lamellae (laminates). The mathematical approach to analyzing

fiber-reinforced laminates has been presented and

extended to include the effect of bounded uniform

orientation dispersion. The model analysis was here

confined to hnear, elastic, small-deflection, twodimensional conditions.

The resultant equations (sec. 4) may be easily

used to infer the orthotropic engineering elastic

properties of a fairly arbitrary class of lamellar

materials containing dispersed orientation. The

same analytical procedure is also applicable to the

new types of synthetic biomaterials which in73
corporate fiber reinforcements in a polymer matrix

(e.g., carbon fiber/ polyethylene matrix). These

materials are convenient to mold and, by selection

of reinforcing fiber quantity and length, the resultant

properties may be tailored over a fairly broad range.

This property range can be estimated, at least to a first approximation, by the methods developed

here.

Finally, further extension of this work should

respond to the explicit limitation described above.

Methods for three-dimensional analysis with non- Unear material behavior will yield more faithful

simulations of real tissue behavior and will be

applicable to soft tissue. Analytical development

should be guided by experimental determination of

actual orientation and orientation dispersions as well as the degree of interlamellar contiguity and

individual fiber configurations. Basic properties of

unidirectional bone-tissue microstructural elements

should be measured, and procedures should be

developed for the rehable experimental characterization of bone tissue samples as orthotropic

(or anisotropic) lamellar materials.

8. References

[IJ Evans, F. G., Ed., Studies on the Anatomy and Function of Bone and Joints, pp. 92-112, 121-141 (Springer Verlag,

New York, N.Y., 1966).

[2] Buckles, R. G., Ed., Formation and Structure of Calcified Tissue, Advances in Bioengineering, (American Institute of Chemical Engineers Symposium, Series No. 114, 67,

1971).

[3] Currey, J. D., Three analogies to explain the mechanical

properties of bone, Biorheology 2, 1-10 (1964).

[4] Dempster, W. T. and Liddicoat, R. T., Compact bone as a

non-isotropic material, American J. Anatomy 91, 331

(1952).

[5] Evans, F. G. and VincentelU, R., Relation of collagen fiber orientation to some mechanical properties of human

cortical bone, J. of Biomech. 2, 63 (1969).

[6) Bonfield, W., Mechanism of deformation and fracture in bone. Composites 2, No. 3, 173-175 (June 1971).

[7] Bonfield, W. and Clark, E. A., Elastic deformations of compact bone, J. Mat. Sci. 8, 1590 (1973).

[8J Cooke, F. W., et al.. The fracture mechanics of bone —

another look at composite modeling, J. Biomedical Mat. Research Symposium: Materials and Design Considerations for the Attachment of Prostheses to the Musculoskeletal System (Interscience, New York, N.Y., 1973).

[9] Currey, J. D., The relationship between the stiffness and

the mineral content of bone, J. Biomech. 2,477-480(1969).

[10] Lees, S. and Rollins, F. R., Jr., Anisotropy in hard dental

tissues, J. Biomech. 5, 557 (1972).

[11] Kenedi, R. M., Ed., Advances in Biomedical Engineering,

pp. 162-168, 202-212 (Academic Press, New York, N.Y.,

1971).

[12] Cox, H. L., The elasticity and strength of paper and other

fibrous materials, Brit. J. Appl. Phys. 3, 72-79 (1952).

[13] Robinson, E. Y., On the Elastic Properties of Fiber Composite Laminates with Statistically Dispersed Orientations,

(Proc. of the 28th Annual Technical Conference, Reinf.

Plast. and Composites, Soc. of the Plastics Industries,

New York, N.Y., 1973).

[14] Giltrow, J. P. and Lancaster, J. K., Friction and wear of polymers reinforced with carbon fibers, Nature 214, 1106

(1967).

[15] Musikant, S., Quartz and Graphite Filament Reinforced

Composites for Orthopedic Surgical Application, J. Biomed. Mat. Research Symposium; Medical AppUcation

of Plastics (Interscience, New York. N. Y., 1972).

[16] Sclippa, E., and Piekarski, K., Carbon fiber reinforced

polyethylene for possible orthopedic uses, J. Biomed.

Mat. Research 7,59(1973).

[17] Structural Design Guide for Advanced Composite Materials,

Second Edition Prepared under AF Material Laboratory

Contract No. F33615-69-C-1368 by Los Angeles Division

of North American Rockwell Corp., Los Angeles, Ca.,

January, 1971.

[18] Ashton, J. E., Halpin, J. C. and Petit, P. H., Primer on Composite materials (Technomic Publishing Co., Stamford,

Conn., 1969).

[19] Haut, C. H., and Little, R. W., A Constitutive Equation for CoUagen Fibers, J. Biomech 5, 423 (1972).

[20] Lukes, R. S. and Katz, J. L., Transformation of the Viscoelastic Functions of Calcified Tissues and Interfacial Biomaterials into a Common Representation (Fifth

Annual Clemson Biomaterials Symposium, April, 1973).

74
NATIONAL BUREAU OF STANDARDS SPECIAL PUBLICATION 415,

Biomaterials, Proceedings of a Symposium held in conjunction with the Ninth Annual Meeting of the Association for the Advancement of Medical Instrumentation, New Orleans, La., April 19-20, 1974

(Issued May 1975).

A Simple /w VitroTest for Screening the Blood Compatibility of Materials

H. Kambic, T. Komai, R. J. Kiraly, and Y. Nos^

Department of Artificial Organs

The Cleveland Clinic Foundation, 9500 Euclid Avenue,

Cleveland, Ohio 44106

An in vitro blood compatibility test was developed to evaluate thromboresistant properties of materials. This method is CcJled the closed-cell kinetic blood coagulation test. A closed cell system

eliminates any air-blood interface. The blood is withdrawn directly from the animal into the cell, minimizing the exposure to foreign surfaces other than the one being studied and eliminating the use of anticoagulants through the process. The technique includes the simultaneous blood filling of eight cells with the test materials, and

eight cells lined with a control material. As a control material we have selected silicone rubber, which

has reasonably good thromboresistant properties, is widely accepted, and commercially available. The cells are opened at different predetermined times, and the clot formation is then measured by two complimentary methods: weighing the clot and colorimetry of the unclotted blood. The two methods

correlate and can differentiate between red and white thrombus.

The results are presented as clot formation curves versus time for the material under test and for the control. The variability of blood properties makes this control curve essential. Detailed analyses of the curves wiU offer a new approach to the understanding of the mechanism of thrombus formation on various types of materials. Results will be presented for tests conducted on Hydron®, segmented polyurethane, purified natural rubber, as well as chemiceilly treated tissue.

Key words: Blood coagulation; Blood compatibility; Blood platelet consumption; kinetic clotting test.

1. Introduction

A large number of materials is available for use

in intravascular devices. Since the processes

occurring when materials are brought in contact

with blood are not well known, a screening test is required in order to establish their material blood

compatibility.

The advantages of an in vitro screening test are obvious: low cost, independence of flow dynamics,

and short time. However, questions have been

raised as to their reliability as well as their correla

tion with the actual behavior of the materials once

implanted in vivo.

A simple technique designed in our laboratory

has been reported [5].* Further refinement has

resulted in a more reliable and reproducible test

[11]. This present report describes the results

obtained studying candidate materials with the

technique.

' Figures in brackets indicate the literature references at the end of the paper.

2. Materials and Methods

The technique employed is the kinetic clotting

test, as described by Nose et al. [11]. Basically, the

test consists of the simultaneous filling with donor

blood of eight cells containing the candidate

material and eight cells containing the control

material (Silastic 372® V'^ The cells are sequentially opened at predetermined intervals to follow

the kinetics of clot formation.

The test cells consist basically of two sheets of

the material under study, enclosed within an acrylic

chamber (fig. 1). Blood is drawn in between the

material sheets by aspirating saline from the acrylic

chamber as shown in figure 2. The blood comes

directly from the donor dog, contacting only the venous puncture needle and a short segment of

' Dow Chemical Co., Midland, Mich.

= Certain commercial materials and instruments may be identified in this publication

in order to adequately specify the experimental procedure. In no case does such

identification imply recommendation or endorsement by the National Bureau of Standards, nor does it imply that the equipment or instruments identified are necessarily the best available for the purpose.

75
r

1 1

CELL FILLING TECHNIQUE

TWO ML OF SALINE IS WITHDRAWN FROM

AROUND TEST SAMPLES DRAWING BLOOD

BETWEEN THE TWO SAMPLE SHEETS

Figure 2. The cell filling technique. Two ml of saline are with- drawn from around the test sample, resulting in blood being

drawn between the test sample sheets.

Afterward, they are placed in a slowly rotating

device to prevent sedimentation.

Two cells, one with Silastic and one with the test material, are opened every twenty minutes over a two-hour period. The clot formation is ascertained

using two techniques. The first is the direct measure- ment of the amount of clot formed. This is done by

determining the dry weight of the clot, normalized

by the amount of blood in the cell.

Normalized clot weight =

ISgauge

Needle

Figure 1. Closed cell kinetic test chamber; (a) expanded view

showing test material and filling needle; (b) assembled cell.

flexible Silastic tubing. No anticoagulants are used

throughout.

All test cells are weighed before and after filling to determine the amount of blood within each cell.

Dry clot weight (gm)

Total blood weight (gm)

The actual values of normalized clot weight are

plotted versus time. The difference between the

clot weights obtained with the test material and the

Silastic control is also plotted as a function of time;

the area under the difference curve provides a quantitative direct index for comparing different

materials. The more negative the number obtained

with one particular material, the better its blood

compatibility.

Second, the amount of unclotted blood which

remains within the test cell is measured. The

content of the cell is immersed in 100 ml of distilled

water, where the unclotted blood hemolyzes. The

resulting hemoglobin concentration is measured

spectrophotometricaUy at 5400 A. The hemoglobin

readings are normalized by the initial amount of

blood, and plotted as functions of time. The differ- ences between the amounts of unclotted hemoglobin

obtained with the test material and Silastic control are also plotted, and the area under the curve

76
provides an indirect index where positive numbers

predict better blood compatibility than Silastic.

The test materials and Silastic control sheets

were cut 2V2-in in diameter and soaked in saline

overnight prior to use. All test materials were

prepared at one time and stored in covered glass

jars. The materials used were:

1. Hydron® (2 — type E hydrophilic polymer, polyhydroxyethyl-methacrylate).

2. Purified Natural Rubber— double centrifuged,

prevulcanized latex; ^50 percent solids exposed

to a 4 percent formaldehyde solution after

curing and then saline rinsed.

3. Segmented Polyether-polyurethane Ethicon ^—

prepared according to the method of Boretos

[!]• 4. Silastic 372, MDX4-4156- vulcanized at 150

°C, coated with NaHCOs, and washed in saline

several days before use (control material).

5. Bovine Pericardium — exposed to 4 percent

glutaraldehyde solution (24 h) and saline rinsed.

Healthy 20-25 kg dogs were fasted overnight and

anesthetized intravenously with 300 mg of sodium

thiamylal.^ Six donor dogs were used in total; one dog was used for each test.

3. Results

Figure 3a shows the weight of clot formed by

Hydron and that formed by Silastic. The difference

in the clot weights is shown in figure 3b. The direct

index was —3.02 min. The comparison of unclotted hemoglobin is shown in figure 3c. The

difference in the amounts of unclotted hemoglobin

between Hydron and Silastic (fig. 3d) yields an

indirect index of 16.14 min, figures 4, 5, and 6a, b,

c, and d show the results for segmented polyurethane, glutaraldehyde treated pericardium and

formaldehyde treated natural rubber, respectively.

AU results are presented as mean values; the bars

represent standard errors of the means. The clotting

indices for all materials tested are outlined in

table 1.

Hydro Med Sciences, Inc.. New Brunswick, N.J.

^ Hygienic Dental Manufacturing Company, Akron, Ohio.

* Ethnor Medical Products, Somerville, N.J.

^ Suritai (R). Parke Davis, Detroit, Mich.

Tabll 1. 120 Minute kinetic clotting modulus

Direct clotting Indirect clotting

Material modulus modulus

(min)

Hydron ® -3.02 +16.4

4 percent glutaraldehyde

treated pericardium -3.33 +24.7

Segmented polyurethane -3.72 +11.2

4 percent formaldehyde

treated natural rubber +0.66 +6.8

(Silastic 372 Control)

77
.16 n

.14 -

.12 -

.10 -

o- .08

.06

.04

.02

AVERAGE CLOT WEIGHT VS TIME

FOR HYDRON AND SILASTIC

O HYDRON • SILASTIC

I 20 —

r~ 40

I 60

MINUTES

80 100 120

AVERAGE OPTICAL DENSITY VS TIME

FOR HYDRON AND SILASTIC

3 0 -1

2 6

2 2

1.8

o

t

IT)

Cid

1.0

06

0,2

O HYDRON • SILASTIC (c)

I 20 40 60

MINUTES

—I

80 100

AVERAGE DIFFERENCE IN Hb CONTENT

OF UNCLOTTED BLOOD IN CONTACT

WITH HYDRON AND SILASTIC

n

120

AVERAGE CLOT WEIGHT INDUCED

BY HYDRON COMPARED TO SILASTIC

.02

SILASTIC

-.06

~l

20

I 40

1

60

MINUTES

—I

80 100 120

JO -J

1

60 MINUTES

—I

80

Figure 3. Comparative blood compatibility of Hydron and

Silastic.

(a) average clot weight as a function of time; (b) average clot weight induced; (c) average

optical density (hemoglobin content of unclotted blood) as a function of time; (d) average difference in hemoglobin content of unclotted blood.

78
AVERAGE CLOT WEIGHT VS TIME FOR

SEGMENTED POLYURETHANE AND SILASTIC

AVERAGE OPTICAL DENSITY VS TIME FOR

SEGMENTED POLYURETHANE AND SILASTIC

AVERAGE CLOT WEIGHT INDUCED BY

SEGMENTED POLYURETHANE COMPARED TO SILASTIC

I

1 1 1 1 1

1

I I 1 I I I I 0 20 40 60 80 100 120 0 20 40 60 80 KX) 120

MINUTES MINUTES

Figure 4. Comparative blood compatibility of segmented polyurethane and Silastic.

(a) average clot weight as a function of time; (b) average clot weight induced; (c) average

optical density (hemoglobin content of unclotted blood) as a function of time; (d) average difference in hemoglobin content of unclotted blood.

79
AVERAGE CLOT WEIGHT INDUCED BY

4% GLUTARALDEHYDE TREATED

PERICARDIUM COMPARED TO SILASTIC

AVERAGE DIFFERENCE IN Hb CONTENT OF

UNCLOTTED BLOOD IN CONTACT WITH 4%

GLUTARALDEHYDE TREATED PERICARDIUM AND SILASTIC

.02 -I

.02 -

.04 -

.06 -

.08

AREA = -3.33 min

1 20

SILASTIC

-.0975 1 (b)

I 40

1 60

MINUTES

I 80 100 120

.44 -I

.40

.36

.32

.28

24

.20 -

.16 -

.12

.08

.04

0

-.04

SILASTIC (d)

20 40 60 MINUTES

80 100 120

Figure 5. Comparative blood compatibility of 4 percent glutaraldehyde-treated pericardium and Silastic.

(a) average clot weight as a function of time; (b) average clot weight induced; (c) average

optical density (hemoglobin content of unclotted blood) as a function of time; (d) average difiference in hemoglobin content of unclotted blood.

80
AVERAGE CLOT WEIGHT INDUCED

BY 4% FORMALDEHYDE TREATED

NATURAL RUBBER AND SILASTIC

05 -I

(b)

20 40 60 80

MINUTES

100 120

AVERAGE DIFFERENCE IN Hb CONTENT OF UNCLOTTEC

BLOOD IN CONTACT WITH 4% FORMALDEHYDE

TREATED NATURAL RUBBER AND SILASTIC

40

.30 -

,20 -

01 .10

-.10 -

-.20 -I

lASTIC

AREA 6.8 min

(d)

20 40 60

MINUTES

80 100 120

Figure 6. Comparative blood compatibility of4 percent formaldehyde-treated natural rubber and Silastic.

(a) average clot weight as a function of time; (b) average clot weight induced; (c) average

optica] density (hemoglobin content of unclotted blood) as a function of time; (d) average difference in hemoglobin content of unclotted blood.

81
4. Discussion

In vivo screening tests for determining the thrombogenicity of materials have been proposed [4, 6]. Previous in vitro systems, even though simpler and

less expensive, have not been readily adapted

because of some experimental drawbacks such as

blood-air interface, exposure to materials other than

the one under study, and the need for anticoagulants

[5, 7]. The kinetic closed cell test described herein

minimizes these drawbacks [11]. Furthermore, the

kinetics of the coagulation process can be followed;

thereby determining the beginning, slope, and the

end point of the reaction. This is not possible with

in vivo tests. Moreover, in vivo testing involves

variables which are extremely difficult to control,

such as the blood flow patterns on the surface of

the material analyzed.

The test uses material in sheet form; the need for

special configurations is eliminated. Also, the ratio

of material surface to the amount of blood is increased over previous methods [7]. The same cells can be utilized for analyzing other parameters of

material blood compatibility. Preliminary data (fig. 7)

illustrates the decrease in platelet counts induced

by the listed materials, showing a significant dif- ference between pericardium and the other three.

The variability of the blood properties requires

that all tests be run with a simultaneous control.

Silastic was selected as a control material, since it

PLATELET CONSUMPTION

100

80

O

Lu 40

60 -

20 -

Si

SILASTIC

1281

47. FORMALDEHYDE

TREATED NATURAL RUBBER

(11)

SEGMENTED

POLYURETHANE

(111 4% GLUTARALDEHYDE

TREATED PERICARDIUM

(6)

Figure 7. Average platelet consumption by tested biomaterials.

Numbers in parentheses indicate the number of individual tests made on each.

is readily available, and widely used for blood

contacting devices. Results from individual tests can be affected by uncontrolled variables; therefore, a minimum of 10 tests is desirable for statistical

evaluation of the data.

The results presented here show that this test can discriminate between different materials. If

these in vitro results correspond with experimental

and clinical results obtained by using the same

materials in blood contacting applications remains

to be demonstrated.

These materials can be rated quantitatively as

to the rate of clot formation on their surface. Natural

rubber proved to be the least desirable material

with a direct modulus of 0.66 min. The blood

compatibiUty of natural rubber is considered

to be inferior to the other materials tested [2, 3,

8, 9, 10]. The direct clot modulus of polyurethane,

Hydron, and glutaraldehyde treated pericardium

was similar. The indirect modulus for the glutaraldehyde treated pericardium, however, was twice

that of the polyurethane, suggesting a higher percentage of white clot formation.

The authors wish to thank T. Agishi, I. Koshino,

and C. Carse for their technical assistance. Dr. J. Urzua provided advice in the preparation of the

manuscript.

5. References

[1] Boretos, J. W., and Pierce, W. S., Segmented polyurethane:

A polyether polymer, J. Biomed. Mater. Res. 2, 121-130

(1968).

[2] Boretos, J. W., Detmer, D. E., and Donachy, J. H., Segmented polyurethane, II. Two years experience, J. Biomed.

Mater. Res. 5, 373-387 (1971).

[3] Brack, S. D., Rabin, S., Fergusin, R. J., Evaluation of bio- compatible materials, Biomat. Medical Devices and

Artificial Organs 1, 191-222 (1973).

[4] Gott, V. L., Keopke, D. E., Daggett, R. L., Zamstorff, W.,

and Young, W. P., The coating of intravascular prosthesis

with colloidal graphite— The relative importance of

electrical charge in the prevention of clot formation.

Surgery 50, 382 (196).

[5] Imai, Y., and Nose, Y., A new method for evaluation of antithrombogenicitv of materials, J. Biomed. Mater. Res. 6, 165 (1972).

[6] Kusserow, B., Larrow, R., Nichols, J., Analysis and meas- urement of the effects of materials of blood. Proceedings

of the Artificial Heart Program Conference (R. J. Hegyeli,

Ed.), (National Institutes of HeaUh, Bethesda, Md.,

June 9-13, 1969), U.S. Government Printing Office,

Washington, D.C., p. 233.

[7] Lee, R., and White, P., A clinical study of the coagulation

time of blood. Am. J. Med. Sci. 145,495 (1913).

[8] Leininger, R. I., Polymers as surgical implants, CRC

Critical Reviews in Bioengineering (October 1972).

[9] Levowitz, B. S., La Guerre, J. N., Calem, W. S., Gould,

F. E., Scherrer, J., and Schoenfield, H., Biologic compatibility and applications of Hydron, Trans. Amer. See.

Artif. Int. Organs 14, 82-87 (1968).

[10] Nose, Y., Tajima, K., Ogawa, H., Klain, M., von Bally, K., and Effler, D. B., Cardiac prosthesis utilizing biological

material, J. Thor. Cardiov. Surg. 62 , 714-724 (1971).

[11] Nose, Y., Kambie, H., Kiraly, R., Komai, T., and Urzua,

J., In vitro test of compatibility of blood with various

natural and synthetic surfaces in: P. Didisheim, T. Shimamoto, and H. Yamazaki, Editors, Platelets, Thrombosis, and Inhibitors, Seminar Proceedings, Honolulu,

Hawaii, Dec. 19-21, 1973 (Stuttgart-New York: F. K. Schattauer Verlag, 1974), pp. 87-95.

82
NATIONAL BUREAU OF STANDARDS SPECIAL PUBLICATION 415,

Biomaterials, Proceedings of a Symposium held in conjunction with the Ninth Annual Meeting of the Association for the Advancement

of Medical Instrumentation, New Orleans, La, April 19-20, 1974

(Issued May 1975).

Conformation of Adsorbed Blood Proteins

B. W. Morrissey, L. E. Smith, C. A. Fenstermaker,

R. R. Stromberg, and W. H. Grant

National Bureau of Standards, Institute for Materials Research, Washington, D.C. 20234

The likelihood that surface-induced blood coagulation results from specific interactions between

proteins and materials has led to a study of the conformation of adsorbed blood proteins. Infrared

difference spectroscopy was used to determine the bound fraction, i.e., the fraction of carbonyl groups

of an adsorbed molecule directly interacting with a silica surface, and ellipsometry was used to measure

the average extension (thickness) of adsorbed protein films. In situ measurements were made on serum

albumin, fibrinogen, and prothrombin as a function of the amount adsorbed, time of adsorption, and

surface energy. The bound fraction results obtained for serum albumin and prothrombin indicate that the internal bonding of these globular proteins is sufficient to prevent changes in conformation while adsorbed, even

at low surface population. The bound fraction of fibrinogen increases with increasing adsorbance,

suggesting possible interfacial aggregation. The conformation of all three proteins was found by both

I. R. difference spectroscopy and ellipsometry to be independent of the time of adsorption. In addition, the eUipsometiic studies show that while the adsorbance of fibrinogen and prothrombin does not vary

for a number of surfaces, their extensions increase with decreasing surface energy. Studies of cross-linked and denatured serum albumin have shown that changes in conformation

concomitant with adsorption of the native protein, if they occur, are small.

Key words: Adsorbed protein conformation; blood protein; protein conformation.

1. Introduction

Proteins and enzymes localized at interfaces

play an important role in many biological phenomena. Matrix insolubilized enzymes are utilized

clinically and commercially [1],' glyco-proteins on

the ceU surface can promote and control both

cellular aggregation [2] and cellular growth [3], the

cytochrome enzyme system for oxidative phosphorylation is localized on the mitochondrion

membrane [4], while the deleterious effect of materials on blood resulting in surface-induced coagulation has long been recognized [5-7]. The possible

effects of a given surface on a protein mixture

include permanent or reversible adsorption with or without denaturation, preferential adsorption of

specific proteins, and changes in the protein microenvironment. The details of these interactions can

best be understood by determining the conformation of the adsorbed molecule and the conformational changes concomitant with adsorption.

Various techniques have been utilized for the

determination of the conformation of adsorbed proteins. Adsorbed molecular area and solution di- mensions have been compared by Bull [8] for egg

albumin using film balance techniques, and by

'Figures in brackets indicate the literature references at the end of the paper.

Brash and Lyman [9] for numerous serum proteins

using infrared internal reflection spectroscopy.

Perturbations in the transmission fluorescence

spectrum arising upon adsorption of trypsin and

chymotrypsin were analyzed by Katchalski [10] to demonstrate conformational changes, while

Loeb and Harrick [11] induced fluorescence of

adsorbed serum albumin using internal reflection

techniques and showed that the protein is apparently

native when adsorbed. Kochwa and coworkers [12]

applied potentiometric titrations to show that yGglobuhn apparently unfolds upon adsorption to a polystyrene latex under conditions of low surface

coverage and acquires an antigenicity similar to that of heat denatured material. EUipsometry has

been used quaHtatively by Vroman [13] to study

sequential, competitive adsorption of blood proteins.

As indicated by these studies, protein adsorption

is a complex process and each technique provides

only partial answers. As a result, a consistent model

of protein adsorption at the solid-solution interface

does not exist at present.

In general, an adsorbed polymer is attached to the surface at a number of locations along the

chain. Adsorbed segments can occur singly or in runs with attached portions separated by loops of

unattached segments which extend away from the

surface into the solution as shown in figure 1. The

number and arrangements of the attached portions and the size and distribution of the unattached

83
Figure 1. Schematic representation of a random coil polymer adsorbed with few

attachments and large loops.

loops and tails define the conformation of the

adsorbed polymer molecule [14]. In the case of

proteins, numerous internal constraints limit the

accessible conformations, both in solution and on

the surface, and make the prediction of the adsorbed conformation more difficult.

In this paper we review some of our studies [15]

of the direct measurement of both the number of

protein carbonyl group-surface attachments made

using infrared difference spectroscopy, and the

average dimensions of the adsorbed protein layer

determined ellipsometrically for a number of

blood proteins and materials. These results are viewed as providing essential information on the

conformation and conformational changes of the

adsorbing molecule. These studies have been

carried out in situ on individual proteins as a

function of the amount adsorbed, time of adsorption, and surface energy of the adsorbent. Additional information on blood-protein-material in- teractions is essential to an understanding of

surface-induced coagulation and the development

of a rational approach to the selection of thromboresistant materials.

2. Techniques

The interaction of the chromophores of an

adsorbed molecule with a surface frequently results

in a shift of their characteristic spectral absorption

bands. For the protein studies presented here, a

shift of — 20 cm of the amide I band for free and

bound carbonyl groups was typically observed.

Since there is a significant overlap of these two

bands, difference techniques must be utilized.

From the data obtained, one determines the

bound fraction, i.e., the fraction of carbonyl groups

of an adsorbed protein directly interacting with

the surface. The number of carbonyl attachments

may then be calculated using the known amino

acid composition of the protein.

Fontana and Thomas [16] originally described a method whereby a polymer is adsorbed from

solution on a high surface-area powder, the sus- pension centrifuged, and an infrared difference

spectrum recorded for the resulting gel. To prevent

the introduction of uncertainties in polymer concentration, this method has been modified [17] to permit a direct analysis of the suspension. This

modification has been used by a number of investigators [18-20] to study the conformation of synthetic polymers adsorbed from organic solvents.

For the determination of the bound fraction for an adsorbed protein, it is necessary to use D2O

solutions of deuterated proteins to obtain a window

in the 1650 cm"i region. The difference spectrum

resulting from a protein solution-silica suspension in

the sample beam versus the same protein solution

in the reference beam (fig. 2) was resolved using a Beer's Law analysis of the optical density of the

unbound peak.

I. R. Dl FFERENCE MEASUREMENTS

REFERENCE

BEAM

PROTON IN OjO BUFFER

SAMPLE

Mil

% T

FREE SEGMENTS

IN SOLUTION

1655 cm

FREE SEGMENTS IN

SOLUTION AND ON

ADSORBED PROTEIN

BEAM

PROTEIN IN OjO BUFFER i

Cab-O-Sil

BOUND SEGMENTS ON ADSORBED PROTEIN

% T

% T —

1655 1625 cm"

1655 1630 cm

Figure 2. Schematic of infrared difference spectroscopy measurements showing idealized signals and actual resulting

difference spectrum.

The technique of ellipsometry has been described

in detail [21], and its application to the study of

polymer adsorption from solution in situ has been

84
THE COMPONENT PARTS OF AN ELLIPSOMETER

POLARIZED LIGHT

Figure 3. Schematic showing arrangement of ellipsometer components used in protein studies.

discussed [22]. Briefly, ellipsometry is an optical

technique in which changes in the state of polarization of light upon reflection from a surface (fig. 3) are analyzed to characterize the surface and to determine the thickness and refractive index of a

thin film on the surface. The thickness of a protein

film on a surface is related to the extension and

distribution of loops as shown in figure 1. This

average extension of the adsorbed molecule normal

to the surface is therefore a measure of the dimensions of an undisturbed protein layer on the surface.

Calculations of the average extension utilize an

iterative solution [23] of the Drude equations [24].

Additionally, one calculates the amount of protein

adsorbed using the experimentally measured film

refractive index and the refractive index increment

for the protein at the wavelength of light of the

measurements. These calculations are, however,

based upon the model of a homogeneous film of

constant refractive index and discrete boundaries.

Since an adsorbed protein film would probably be an inhomogeneous film with refractive index decreasing

with distance from the surface, the calculated

values reported represent an upper bound for the

thickness. Appropriate inhomogeneous film models

[25] may be utilized to relate the average extension

to actual molecular dimensions.

3. Experimental Detail

The proteins selected for study are either major

constituents of blood plasma or are implicated as being important in the clotting process.

3.1. Ellipsometry

Human serum albumin (4X crystallized) and

bovine prothrombin (Cohn III—2) were obtained from

Nutritional Biochemicals (NBC)^ and used without

^Certain commercial materials and instruments are identified here and elsewhere

in this pubUcation in order to adequately specify the experimental procedure. In no

case does such identification imply recommendation or endorsement by the National Bureau of Standards, nor does it imply that the equipment or instruments identified are necessarily the best available for the purpose.

further treatment. Human fibrinogen (NBC) was

purified by the procedure of Batt, et al. [26], a modification of a procedure developed by Laki [27].

Commercial ferrotype plate consisted of chromium electroplated on either brass or stainless

steel. Platinum was 99.95 percent pure and rolled

to a mirror finish. The quartz was fused silica. These

materials were cleaned before use with hot 50:50

HN03:H2S04 followed by three rinses in boiling

distilled water. They were then heated for 5 minutes

at 500° in a muffle furnace and placed while still warm into the adsorption cell filled with buffer.

Low temperature laminar (LTL) type pyrolytic

graphite was kindly supplied by Dr. Jack Bokros

of General Atomic Corp. Commercial polyethylene

sheet was cleaned in ethanol and cast under vacuum

at 180° between plate glass sheets. Several samples

were prepared from NBS Standard Reference

Material 1475 linear polyethylene which has a

density of 0.978 g/cm^, also by casting against glass

under vacuum. AU results on these two types of

polyethylene were identical. The carbon and polyethylene samples were cleaned before use by re- fluxing with ethanol for two weeks in a Soxhlet

extractor followed by a similar washing with distilled water.

AU solutions were made up in pH 7.4 phosphate

buffer and concentrations were determined by

UV spectroscopy. AU adsorption measurements

were made in a ceU thermostatted at 37±0.1°C.

3.2. Infrared Difference Spectroscopy

Bovine serum albumin (4X crystaUized), bovine

fibrinogen (Cohn fraction I, 60% clottable), and

bovine prothrombin (fraction ni-2) were obtained

from NBC. The serum albumin was deuterated

[28], lyophilized, and stored in vacuo at 4 °C. The

prothrombin was dissolved in O.lM D2O phosphate

buffer pD 7.4, dialyzed overnight against buffer,

and filtered through a weU washed 0.8 fxin pore size

filter just prior to use. Fibrinogen was purified by

the method of Laki [27] with the foUowing modifi85
cations: The ammonium sulfate precipitate was

resuspended in the pD 7.4 phosphate buffer and

dialyzed overnight against the same buffer. The

resuhing solution was filtered and used immediately. These protein solutions were prepared

using O.lM D2O phosphate buffer which was standardized using NBS standard reference materials

2186-1 and 2186-11. Adjustments of pD were made

with DCl or NaOD. Ionic strength was adjusted

with KCl.

The mild heat denaturation of serum albumin

was carried out at 66 °C [28]. The quantitative,

limited, reduction of disulfide bonds of serum

albumin was performed by the method of Sela et

al. [29], using a reaction time of 1.5 hours and

omitting the urea. Serum albumin was cross linked using diethyl

malonimidate dihydrochloride. This reagent was

prepared from malononitrile and HCl gas in anhydrous ethanol-dioxane according to McElvain and

Schroder [30]. The reaction of diethyl malonoimidate dihydrochloride with serum albumin generally followed Button et al. [31], using a 1.5 percent

protein solution and a reaction time of 1 hour.

The adsorbent used for all experiments was a fumed, nonporous silica with a nominal particle

size of 0.012 fim (Cab-O-Sil M-5, Cabot Corp.).

The adsorbent was heated in vacuo at 110 °C over- night and cooled over silica gel just prior to use. The

surface area, as determined by BET N2 analysis,

was 204 ±20 m^/g.

The adsorption isotherms were determined at 23.5 ±0.5 °C. Known amounts of silica and protein

solution were shaken for three hours and the silica

then removed by filtration through a 0.8 fxm filter. The amount of adsorbed protein was determined

from the protein concentration changes measured

spectrophotometrical at 280 nm. Changes in the

ratio of silica to protein solution volume produced

no effect on the protein adsorption isotherms. For

studies of adsorbance as a function of time, the

quantities of silica and protein solution were suitably scaled up and aliquots withdrawn and filtered

with time.

For bound fraction studies, 0.029 grams of silica

per ml protein solution was found generally to give

stable suspensions suitable for infrared analysis.

The difference spectrum of the protein-silica suspension versus the same protein solution was obtained

for the region 1750-1550 cm^^ using matched 0.1

mm CaF2 cells at 10 X ordinate expansion. The

extinction coefficient of the adsorbed carbonyl

group was obtained by scanning successive dilutions

of the reference solution against the protein-silica

suspensions and noting the dilution at which the

peak due to the unbound chromophores disappeared. Each reported bound fraction value represents the average of from two to four infrared scans on the same protein suspension.

4. Results and Discussion

4.1. Number of Attachments

The isotherms of serum albumin and prothrombin

are given in figure 4 along with values of the bound

fraction. The bound fraction of fibrinogen was found

in general to depend on the equilibrium concentration (and hence the adsorbance). The results of a

set of four experiments on fibrinogen are shown

in figure 5. In each experiment, purified protein

was utilized to prepare solutions covering a range

of concentrations. Portions of the fibrinogen isotherm were inaccessible for bound fraction measure- ments using the technique described above because

of the large adsorbance of this protein coupled

with the limited solubility of fibrinogen in the

D2O — phosphate buffer which prevented the

preparation of highly concentrated solutions for

the adsorption experiments.

For serum albumin and prothrombin, no systematic change in the bound fraction was apparent

over the measured portion of each isotherm. The

mean values with their computed standard deviations were 0.11 ±0.02 and 0.11 ±0.03 for serum

albumin and prothrombin respectively. Results

for fibrinogen show an increase in the bound fraction with increasing amount adsorbed. The dashed

Une in figure 5 represents a least-squares linear

fit to the data. The unstable nature of the fibrinogensilica suspension is at least partially responsible

for the large scatter in the bound fraction values,

and, in conjunction with the high adsorbance,

prevented measurements of the bound fraction at larger adsorbance values.

The number of carbonyl contacts made per

molecule with the silica surface can be calculated

using the bound fractions reported and the total

number of carbonyl groups per protein obtained

from the amino acid compositions. The bound

fraction data reported then translate into a minimum

of 77 contacts with the silica surface for serum

albumin, 80 for prothrombin, and 176—703 for

fibrinogen, as shown in table 1. If the adsorbed

protein molecules were spread nearly flat on the

surface, large values of the bound fraction would

be expected. Adsorption of poly(ethylene o-phthalate) [19], poly(methyl methacrylate) [18], polystyrene [20], and poly(4- vinyl pyridine) [17] on Cab—O—Sil gave bound fractions of 0.37, 0.35, 0.24,

and 0.35 respectively. In each of these cases, the

synthetic molecules appear to lie close to the sur- face in a presumably spread condition. Since bound

fractions of 0.11 were obtained for prothrombin

and serum albumin, the adsorbed protein molecules

were still relatively unchanged by the surface,

despite the large number of attachments and great

interaction.

86
600 n

500 -

CONCENTRATION, mg/ml

Figure 4. Adsorption isotherms for bovine serum albumin and prothrombin at pD

7.4 on silica.

Values of bound fraction given for isotherm points.

The bound fraction data can yield information on changes in the conformation of the adsorbed

protein molecule as a result of increasing competition for surface sites. A decrease in the bound

fraction values with increasing surface population

would be expected if changes in conformation were

to occur. Such changes in conformation of synthetic

random coil polymers have been observed [17,

18, 22]. It has been argued [33, 34] that an adsorljed

serum albumin molecule should initially unfold,

thereby increasing the value of p, the bound fraction,

which would presumably then decrease with

increased surface population. Our results for serum

albumin and prothrombin, figure 4, show that the

bound fraction remains constant, indicating that

the conformation does not change as the surface

population increases. The relatively low value of

the bound fraction for these proteins as compared

to the synthetic polymers, and the constant value

of the bound fraction over the isotherms, strongly

suggests that the internal bonding and disulfide

0 24-|-

0 20-

Q.

16-

1—o

<

cr 0

Li_ 12- Q

§0

O

CD

.08-

0 .04-

0 .00--

0 40 80 1 20 160 200 240 280 320 360 ADSORBANCE, mg/gm

Figure 5. Bound fraction versus adsorbance of bovine fibrino- gen for four different experiments-.

Each symbol represents a separate run. Adsorption at pD 7.4 on sihea. [Reprinted from Ref. 15, J. Colloid Interface Sci.J

Table 1. Summary of adsorption and bound fraction data for proteins

Protein

Maximum

adsorbance

mg/g

Maximum

adsorbance

mg/m^

Equil. Cone, for Maximum adsorbance

mg/ml

Bound

fraction

No. Carbonyl

contacts per molecule

with surface

180 0.88 11.0 0.11 77

Prothrombin 440 2.16 4.0 0.11 80

Fibrinogen 890 4.36 1.0 0.05-0.20 176-703

87
crosslinks of the globular proteins are sufficient

to preserve the basic tertiary structure of the

molecules.

The adsorption behavior of fibrinogen (fig. 5),

which shows a direct dependence of bound fraction on adsorbance, does not necessarily represent a conformational change. As noted above, previous

experience with synthetic polymers has shown that

conformational changes resulting from increases

in surface population are accompanied by decreases

in the bound fraction. The bovine fibrinogen molecule is rodlike with a length of between 400-600 A

[35] and can easily interact with more than one

silica particle. It has been postulated [36] that the

solubility of fibrinogen is due to electrostatic re- pulsion between regions of high negative charge

density localized in fibrinopeptides A and B. If

adsorption and interaction with the hydrophilic

surface were to effectively neutralize this repulsion,

fibrinogen molecules adsorbed to different silica

particles could interact like fibrin, thereby in- creasing the surface attachments of each fibrinogen

molecule. Such an interfacial aggregation [34] could

result in an increase in bound fraction values with

increasing adsorbance, without significantly hindering further adsorption to the remaining available

surface.

Our data for maximum adsorbance can be con- verted to amounts/unit area using the measured

surface area. These results are presented in table

1 and agree quite favorably with those found for

adsorption of human blood proteins on quartz using

the ellipsometric techniques. Results obtained by

MacRitchie [34] for adsorption of bovine serum

albumin on a silica surface (also 200 m^g) at pH 7.5 were only one third as great as that found in the present study, while adsorption on flat polymer surfaces

determined using internal reflection techniques [9]

gave generally higher amounts. The reasons for

these differences are not obvious.

The possibility that the conformation of the

adsorbed protein might vary as a function of the

time of adsorption was tested by determining the

bound fraction for each protein periodically fol- lowing the formation of a suspension. Figure 6 illus- trates some typical results obtained for serum

albumin, prothrombin, and fibrinogen where the

bound fraction values are listed over each curve

at the time of measurement. The rates of adsorption cannot be compared since the initial conditions

are different for each experiment. It is clear that

there is no variation in the bound fraction for any

of the proteins over the time period investigated at

E

256

252-

248

206

Prothrombin, C=0.24

-a

' EEF-

^ 202

<

CD rii a:

'-^ 132

Q

<

Fibrinogen, C=0.069

Serum Albumin, C = 0.98

75 100 125 150 175 200 225

TIME, min

Figure 6. Adsorbance versus time for bovine serum albumin, prothrombin, and

fibrinogen.

Values of bound fraction are given over each curve at the time measured. Adsorption at pD 7.4 on silica. [Reprinted from ref. [15], J. Colloid Interface Sci.J

88
these conditions. This indicates that molecular

rearrangements do not occur following adsorption

at the concentrations studied. Although the adsorption rates for prothrombin and fibrinogen were

extremely high as a result of the conditions re- quired for bound fraction measurements, the

relatively slow adsorption of serum albumin allows

measurements of the bound fraction as a function

of time which shows that the conformation is not

dependent upon the surface population, as was

found in the equihbrium adsorption studies (fig. 4). The shght decrease in adsorbance of prothrombin

with time in figure 6 is possibly caused by a dissociation at the low final solution concentration.

Prothrombin has been reported [37] to associate

reversibly at low ionic strengths.

The conformational changes of blood proteins

upon adsorption are extremely important to any

understandting of surface induced coagulation. A

priori, one does not know how much denaturation,

if any, occurs upon adsorption of native proteins.

The effects of any such denaturation are necessarily included in our values of the bound fraction.

The presence of an infrared shift in the difference

spectrum for the bound carbonyl group does not

necessarily mean that protein denaturation has oc- curred upon adsorption. The shift may occur for

adsorbed protein, even if the tertiary structure has

not changed. It is possible, however, with the aid

of certain assumptions to gain some insight regarding the amount of distortion of the molecule that occurs upon adsorption.

The observed bound fraction for serum albumin

can be considered to be the sum of two contributions: the first due to contacts permitted by the

native tertiary structure with the surface and the

second resulting from surface induced rearrangements that occur as a consequence of adsorption.

These contributions can be separated by determining the bound fraction of a highly cross-Unked,

and therefore rigid, material. Cross-linking of serum

albumin with diethyl malonimidate dihydrochloride

[32] produces a protein which is resistant to denaturation, has the same electrophoretic mobility,

and most important, retains the same antigenic

activity as native serum albumin. We prepared

such a cross-finked material which had 45% of its free amino groups blocked [38]. The average result

of six experiments shown in table 2 gave a bound

fraction which was experimentally indistinguishable

from that of the native protein. We conclude that

surface induced rearrangements, if present in

native serum albumin, are small.

As a further means of assessing the conformational change of native serum albumin upon adsorption, we have denatured the protein under

mild conditions and determined the bound fraction

of the denatured material when adsorbed. Table 2

fists the results for heat denaturation at 66 °C and

thioglycoUic acid reduction. It is clear that the

bound fraction for a given protein is strongly dependent on its tertiary structure. Heating, which

causes the irreversible aggregation of serum al- bumin [39], leads to a bound fraction one third that

of native material. Mild reduction, where one disulfide bond was broken as determined by a p-chloromercuribenzoate titration [30], leads to a bound

fraction of about IV2 times that for native protein.

Table 2. Bound fraction of modified serum albumin

Sample Bound fraction

Native (control) 0.11

0.10

0.04

0.19

Cross-linked

Heated at 66 °C (2-60 min)

These studies also shed fight on the conformation

of native adsorbed serum albumin. If the native

protein were unfolded by an amount equivalent

to the breaking of one disulfide bond, an increase

of 55 contacts (tables 2 and 1) would resuh. On

the other hand, pretreatment of the serum albumin

in a manner which produces aggregates resulted

in a decrease of 50 surface contacts, implying that no aggregation occurs upon adsorption of the native

serum albumin. These results for denatured serum

albumin show that bound fraction is sensitive to what in our view are small conformational changes.

3.2. Molecular Extension

Two aspects of the eUipsometric results should

be considered, viz. the time dependence of the

adsorption process and the long time or equilibrium

value.

In general, no change in extension was observed

during the time period from 10 minutes to 24 hours.

However, there was a small, but definite increase

in adsorbance. The absence of change in extension

with adsorption time is a general feature found for

all three proteins on aU of the surfaces examined

to date. Rates could not be inferred from these

results as the eUipsometer cell was not stirred for

these experiments. Extension measurements made

at adsorption times shorter than 20 minutes generally showed poor reproducibifity due to refractive

index gradients present in the cell before temperature equihbrium is reached.

This behavior can be interpreted as indicating a

lack of significant molecular rearrangement of

the protein molecule once adsorption has taken

place. In the early time periods, additional molecules

are still arriving at the surface as evidenced by

the increasing adsorbance, but the surface pressure

of these molecules is not sufficient to cause the

molecules already adsorbed to change their con- formation. Nor, apparently, do the late arrivals

adopt a significantly more extended conformation

89
than the majority of the adsorbed molecules. It should be noted that a random coil homopolymer

has been shown to undergo a significant change in

extension with the adsorption time indicating

more conformational freedom [22].

Two general features are noted in the equilibrium

adsorbance and extension values of the three proteins adsorbed on the various surfaces studied. First,

in spite of the varied nature of the surfaces, the

measured adsorbances covered only a small range.

Typical results were from 2-4 mg/m^ and the

variations in adsorbance showed no relation to the

surface energy. If surface energy is estimated by

the critical surface tension for wetting, then the

surface energies ranged from the high energy

metals, chromium and platinum, to the low surface

energy of polyethylene. There was, however, a correlation of surface energy with extension — the

extension of a given protein increased with decreasing surface energy. Typical extensions vary

from approximately 200 A on the high energy sur- face to 5-600 A on the lowest energy surface. This

would be in accord with flat, tightly held con- formations on the higher energy surfaces and

looser, more extended conformations on the lower

energy surfaces.

Secondly, the measured extensions cannot be

construed as molecular dimensions of the adsorbed

protein. The model assumed for the calculation of

this extension is that of a film of homogeneous

refractive index with discrete boundaries. Obviously,

any large adsorbed molecule will show some refractive index inhomogeneities normal to the surface

and this will affect the calculated extensions directly.

This problem has been considered by McCrackin

and Colson [2.5] who have shown that extensions

calculated assuming a homogeneous film may be

converted to root-mean-square segment to surface

distances if the actual refractive index distribution

is known. For example, the homogeneous film thick- nesses have to be divided by factors of from 1.5 to 1.7 if the refractive index distributions are linear,

gaussian, or exponential. As we have no estimate

of the refractive index distribution of an adsorbed

protein, we have no way of making such a conversion, but the homogeneous extensions may perhaps

be considered an upper hmit approximation to molecular dimensions. The relative values of the

extension on the different surfaces may properly be

compared assuming that the refractive index

distribution does not change markedly from one

surface to another.

This work was supported in part by the Biomaterials Program of the Division of Blood Diseases

and Resources of the National Heart and Lung

Institute.

4. References

[1] Mosbach, K., Sci. Amer. 224 (3), 26 (1971).

[2] Pessac, B., and Defendi, V., Science 175 , 898 (1972).

[3] Schnebli, H. P., and Burger, M. M., Proc. Nat. Acad. Sci. USA, 69,3825 (1972).

[4] Green, D. E., and Young, J. H., Amer. Sci. 59, 92 (1971).

[5] Hulbert, S. F., King, F. M., and Klawitter, J. J., J. Biomed.

Mater. Res. Symp. No. 2 (Part 1), 69 (1971).

[6J Bruck, S. D., Biomaterials, Med. Devices and Artificial Organs 1, 79 (1973).

[7] Brash, J. L., and Lyman, D. J., in The Chemistry of Bio- surfaces — M. L. Hair, Ed. (Marcel Deicicer, Inc., New

York. N.Y. 1971), p. 177.

[81 Bull. H. B., J. Colloid Interface Sci. 41 , 305 (1972).

[9J Brash, J. L., and Lyman, D. J., J. Biomed. Mater. Kes. 3,

175 (1969).

[101 Gable, D., Steinberg, I. Z., and Katschalski, E., Biochem.

10, 4661 (1971).

[Ill Loeb, G. I., and Harrick, N. J., Anal. Chem. 45,687 (1973).

[12] Kochwa, S., Brownell, M., Rosenfield, R. E., Wasserman,

L. R., J. Immun. 99, 981 (1967).

[13] Vroman, L., and Adams, A. L., Thrombos, Diathes.

Haemorrh. 18, 510 (1967).

[14] DiMarzio, E. A., and Rubin, R. J., J. Chem. Phys. 55,

4318 (1971).

[15] A more complete account of part of this work has appeared

in J. Colloid Interface Sci. 46, 152 (1974).

[16] Fontana, B. J., and Thomas, J. R., J. Phys. Chem. 65,480

(1961).

[17] Thies, C, Peyser, P., and Ullman, R., Proceedings of the

4th International Congress on Surface Activity, Brussels,

1964, Vol. 2, Gordon and Breach, N.Y. 1967, p. 1041.

[18] Thies, C, J. Phys. Chem. 70, 3783 (1966).

[19] Peyser, P., Tutas, D. J., and Stromberg, R. R., J. Poly. Sci. A-1, 5,651 (1967).

[20] Herd, J. M., Hopkins, A. J., and Howard, G. J., J. Poly. Sci. Part C, 34, 211 (1971).

[21] McCrackin, F. L., Passaglia, E., Stromberg, R. R., and Stein- berg, H. L., J. Res. Nat. Bur. Stand. (U.S.), 67A (Phys.

and Chem.), No. 4, 363-377 (July-Aug. 1964).

[22] Stromberg, R. R., Tutas, D. J., and Passaglia, E., J. Phys.

Chem. 69, 3955 (1965).

[23 1 McCrackin, F. L., A Fortran Program for Analysis of ElUpsometer Measurements, Nat. Bur. Stand. (U.S.), Tech. Note 479, 83 pages, (April 1969).

[24] Drude, P., Ann. Physik 272, 532 (1889); 272, 865 (1889);

2 75 , 481 (1890).

[25] McCrackin, F. L., and Colson, J. P., in Ellipsometry in the Measurement of Surfaces and Thin Films (Symposium

Proceedings), E. Passaglia, R. R. Stromberg, and J. Kruger, Ed., Nat. Bur. Stand. (U.S.), Misc. Publ. 256,

365 pages (Sept. 1964), p. 61.

[26] Batt, C. W., et al., J. Biol. Chem. 245, 4857 (1970).

[27] Laki, K., Arch. Biochem. Biophys. 32,317 (1951).

[28] Susi, H., Timasheff, S. N., and Stevens, L., J. Biol. Chem.

242 , 5460 (1967).

[29] Warner, R. C, and Levy, M., J. Amer. Chem. Soc. 80,

5735 (1958).

[30] Sela, M., White, F. H., and Anfinsen, C. B., Biochem.

Biophys. Acta 31,417 (1959).

[31] McElvain, S. M., and Schroeder, J. P., J. Amer. Chem.

Soc. 71,40 (1949).

[32] Button, A., Adams, M., and Singer, S. J., Biochem. Biophys.

Res. Comm. 23 , 730 (1966).

[33] Baier, R. E., Loeb, G. I., and Wallace, G. T., Fed. Proc. 30, 1523 (1971).

[34] MacRitchie, F., J. Colloid Interface Sci. 38,484 (1972).

[35] Laki, K., Fibrinogen, p. 4, (Marcel Dekker, Inc., New York,

N.Y., 1968).

[36] Laki, K., Op. Cit. p. 6. [37] Tishkoff, G. H., WilUams, L. C, and Brown, D. M., J. Biol. Chem. 243, 4151 (1968).

[38] Shick, A., and Singer, S. J., J. BioL Chem. 236, 2477 (1961).

[39] Neet, K. E., and Putnam, F. W., J. BioL Chem. 2 1 0, 2883

(1965).

90
NATIONAL BUREAU OF STANDARDS SPECIAL PUBLICATION 415,

Biomaterials, Proceedings of a Symposium held in conjunction with the Ninth Annual Meeting of the Association for the Advancement of Medical Instrumentation, New Orleans, La., April 19-20, 1974

(Issued May 1975).

A Nerve Implant Prosthesis for Facilitating Peripheral Nerve Regeneration

Part I. Development of the Prosthetic Device and System of Repair

W. E. Kuhn

Department of Materials Science and Metallurgical Engineering

and

James L. Hall

Department of Anatomy, University of Cincinnati, Cincinnati, Ohio 45221

The design rationale and requirements of a system of sutureless nerve repair employing a special

thin walled porous stainless steel tube and a vacuum technique for applying the tube and approximating the nerve ends wiU be outlined. A description of the tubular prosthesis, its fabrication, the

surgical instruments and neurorrhaphical procedures will be presented.

A slurry dipping process has been developed for the production of uniform "green" tubes in quantity. These are sintered in a hydrogen atmosphere to impart the strength required to retain their tubular configuration under the forces imposed by the surgical procedures and the vacuum pressure. Placing the porous tube on the nerve and drawing the nerve ends into approximation is both simple

and rapid.

Key words: Neurorrhaphical procedures; peripheral nerve repair; tubular prosthesis.

1. Introduction

The problem of peripheral nerve regeneration

has been studied intensively by many investigators

and has been found to depend on manv factors.

The most important of these are the patient's age,

type of injury, length of tissue, and finally, the

technique of nerve repair. A repair is satisfactory

when it is executed with as little trauma as possible

and when anatomic approximation is achieved. It

is difficult to fulfill the requirements without magnification, microsurgical instruments, and the finest

of sutures [1, 2]'. Despite the advances that have

have been made to date, the reestablishment of

normal function in the damaged nerve is considered

poor [3]. According to Hakstian, "One remaining frontier

in which a successful breakthrough has" not yet

occurred is that of peripheral nerve injury [4]. This

is not surprising considering the complex structure

of the peripheral nerve over its entire length and

the unique manner of heahng of this tissue compared to others. Whereas most other tissue, excepting tissue of the central nervous system, survive

both distally and proximally of the injury site, nerve

tissue (axons) degenerates peripherally [4].

*This research was supported by the National Scieiiee Foundation under Grant GH-33873.

'Figures in bracitets indicate the hterature references at the end of this paper.

If results are to improve, in addition to gaining

greater understanding of the physiology and the

pathology of neural tissue and processes of nerve

regeneration, it will be necessary to attain greater

sophistication in the surgical manipulation of nerve

tissue. It is the latter goal that the present research

is concerned with. More specifically, it is concerned

with the repair of small nerves, that is, nerves 2 mm

in diameter and less such as cranial, digital and

facial nerves. Digital nerves are the most fre- quently severed peripheral nerves, and though

small in size are of critical importance because of

their properties of discriminatory touch [5]. Injury

to facial nerves, though less common, is of tragic

consequence because of loss of communication

through expression, interference with speech and

daily activities as eating, drinking, washing and

shaving, and the impairment to social and economic

welfare produced by psychological effects [6, 7, 8]. Of particular interest to us in the current study

has been the use of other tissues or materials to form a channel through which the nerve fibers

may grow and bridge the gap between the cut ends.

Severed nerves have been drawn through blood

vessels [9] or through decalcified bone tubes.

Nerve fibers have also been wrapped in sheets of

rubber, cargyl membrane, fascia lata or various

other materials in an effort to reduce the ingrowth

of scar tissue between the ends of the severed

fibers. The methods employing organic materials

91
have been relatively unsuccessful, largely because

phagocytes have removed the nerve covering prematurely. A tubulation method has found some

success in neurorrhaphy, whereby sheets of Silastic are wrapped about the anastomosis [10, 11, 12, 13,

14.]. The success of this method for facilitating

regeneration of nerves is questionable; further it suffers from the drawback of requiring the use of

sutures. Methods employing adhesives, especially

the cyanoacrylates, as a sutureless technique to anastomose severed nerves, though highly attractive, has so far been ineffective [15, 16].

2. Principles and Practice of Current

Nerve Repair Techniques

In this new system of nerve repair it was con- sidered essential to incorporate the applicable

general principles employed in the three types of

neurorrhaphy currently practiced, that is, epineurorrhaphy, perineurorrhaphy, and nerve grafts.

These principles may be summarized as follows

[17, 18, 19].

• Magnification to permit easier, more atraumatic identification, isolation, and preparation

of nerve ends.

• Correct anatomic alignment which is mandatory.

• Use of a minimum of fine sutures to achieve

approximation or eliminate sutures entirely.

• Tension across a suture Une lends to the proliferation of connective tissue within the nerve and a decrease in axon regeneration.

• Ideal nerve repair occurs only when there is no tension at the repair site.

• Axons pass through two suture lines without

tension more easily than through one suture

line under tension.

• Fibrin adherance is of sufficient strength to maintain the approximation if the repair is accompHshed without tension.

• Epineurium contributes to scar formation and

should be resected.

• Fasicular grafts can survive without an estabUshed blood supply as a free graft, whereas

larger grafts undergo central necrosis.

The surgeon using the epineural technique has

little or no control over alignment and hence can

never be fully confident that correct anatomical

approximation has been achieved. Correct tension

is essential for successful epineural nerve repairs.

If the repair is too loose, a gap results, which will

be filled with scar tissue. If the repair is too tight,

buckling of the fascicles occurs. According to Bora, even if tension is exactly right, the chances of

correct orientation of a fascicular level is only one

in four [20]. To overcome the problems of epineural

repair, such as gap, overriding, buckUng and

straddUng of fascicles, the fascicular suture technique has been devised [4], whereby two sutures

are placed in the perineurium on opposite sides of

each fascicle and the severed ends approximated.

The use of sutures is minimized by techniques

developed by MiUessi, which depends on the use

of grafts to eliminate tension [18, 19].

Nerve injuries are commonly repaired by approximating the severed nerve stumps with sutures

through the epineurium, and through the perineurium where the fascicular suture technique is used. The very act of suturing destroys nerve fibers.

Only the larger nerves can be sutured without

serious damage to axons, and even these require

a highly skilled surgeon. Thus, ideally, complete

elimination of sutures is desired.

The most critical factor of technique in neurography is the accurate approximation of the cut

ends of the fascicles. Even though other factors

may be ideal, the neurography wiU fail if a significant number of fascicles do not match up. This is because motor axons regenerating through the

sheaths of sensory nerves and vice versa will give no return of sensory or motor function. The fascicles

of the larger nerves, such as the median and ulnar

nerve, are large enough in cross sectional area

that the cut ends of individual fascicles can be

matched up by using the epineural vascular pattern

for rotational orientation and microsurgical suture

techniques. The fascicles of the small diameter

nerves however, are too small to be matched by

suturing. The best that can be done is to approximate the nerve ends using recognizable vascular

patterns for correct anatomical ahgnment. If the

peripheral nerve is cleanly and squarely severed,

a primary repair can be made with good approximation of the fascicles. However if a segmental

deficit exists, the Hkehhood of surgical orientation

or matching of fascicles becomes a matter of chance.

This is attributed to the twisting and plexiform course of the fascicles and axons within the nerve

trunk [3]. It is apparent then, except in the instance

of a clean cut followed by a primary repair, that a

repair of nerves, successful in aU other respects,

will depend on the chances that the fascicles wiU

match up, unless some natural mechanism exists

for directing the axons to sheaths of their own. In

this regard Peacock reports evidence of rotation

and reorientation into good anatomical alignment

prior to a secondary suture of a nerve previously

joined at a single point of junction, even when the

ends were 180° apart when the suture was first placed [21].

92
The matching fascicles becomes less of a problem with the smaller nerves, since there are fewer

fascicles, often no more than one, to contend with.

Thus the chief hazard to the successful repair of

fine nerves is the trauma produced by sutures and

manipulation of the nerve and invasion of the field

of regeneration by epineural and extraneural connective tissue. Many attempts have been made,

almost entirely with larger nerves, to shield the

region of anastomosis from connective tissue and

facilitate longitudinal directional growth of the

epineural and nervous tissue.

A most recent and promising approach by MiUessi

dispenses with wrappings altogether, ehminates

tension in the field of regeneration, and uses a minimum of sutures [19]. This approach has been

extended experimentally to the sciatic nerves of

rabbits to the complete elimination of sutures. This

technique depends on the use of a graft to eliminate

tension and the natural adhesive properties of the nerve ends to hold them in a position of close

anatomical approximation. The strength of the

union can amount to several grams after the first 15 to 30 minutes, sufficient to prevent separation

of the ends and allow closure of the incision [22].

3. Design Requirements for Tubular

Nerve Prosthesis

Consideration and understanding of anatomical

and functional features, pathophysiology of nerve

degeneration and regeneration, and the principles

behind current techniques of nerve repair, suggested the following design requirements for a porous thin-walled tube that could, with proper

surgical procedures, facilitate nerve repair and

regeneration in small diameter nerves:

• The tube material must be compatible with its biological environment and be nontoxic. No re- sidual substances must remain after processing

that will react deleteriously with the nerve tissue

or the biological processes contributing to healing.

• The technique of application must facihtate

approximation of nerve ends quickly with a minimum of manual manipulation. A method of vacuum

approximation has been developed which successfully accomphshes the approximation procedure.

• The tube design must incorporate means to remove the tube without damage to the nerve after

healing has progressed sufficiently to permit its removal.

• The tube must possess a combination of porosity

and wall thickness that imparts to the tube a

sufficient degree of permeability to provide adequate

diffusion and flow of intraneural and extraneural

fluids.

• The tube must possess a combination of porosity

and wall thickness that minimizes the surface area

of the tube consistent with enough strength to

withstand physical insults. Minimum surface area minimizes the release of toxic substances to the

surrounding tissue.

• The pore size of the tube walls must be such as

to prevent ingrowth of extraneural connective

tissue and, thereby, shield the field of regeneration.

• The pore size of the tube must be such as to discourage the ingrowth of intraneural connective

tissue and, thereby, avoid the occurrence of adhesions between the inner wall of the porous tube

and epineural tissue.

• The tube must possess strength to rigidly support

and shield the joined nerve against buckling and

crushing by outside as well as internal forces.

• The method of fabricating tubes must satisfy

the requirements of the models selected for evaluating the efficacy of this method of nerve repair.

This method of fabrication must be capable of producing many tubes simultaneously having a uniform

size and configuration to a wide range of specifications determined by transverse dimensions and

shape of the nerve. A standard size is essential for

biostatistical comparison studies of nerve repair

parameters.

4. IVIethod of Approach and Organization of Research Effort

The general approach taken was to first undertake

an exploratory program encompassing all essential

activities that would lead to a system of neuror- rhaphy employing porous thin-walled tubes. The

main development activities, outlined in figure 1 are summarized in table 1.

Table 1. Major development activities

^./'Development of comminution

process for metal coating

^ti^ slurries

• Development — Development of coating

of tube process for making green

Fabrication ^^^^ tubes Process ^^^^^^^Development of sintering process for thin walled tubes of controlled

porosity and pore size.

• Development of Surgical Instruments for Applying Porous

Tubes.

• In-vivo testing of porous tube, surgical instruments and surgi- cal procedures-. • Study of post operative functional behavior of rats and post

sacrificial histological studies of anastomosed nerves.

Parallel development and study of the four main

activities shown in table 1 permitted earlier production of tubes and hence, an earlier start on surgical and histological procedures than otherwise

93
MATERIALS

DCSICM

'I9CAL PEVONSIBI

CONCEPT THCORY

CnOUP RESPONSIBLE FOR iurriVITIES TYPE Of FLOW

(^^^ WEOICAl.

[ 1 MATERIALS

BIO-WECHAMICS /\ LIFE 9CIENC£S

1 I MANAGERIAL

^ IKFORWATION

<:^^^^^mm rROSTHETK; DCVICES

SURGICAL HARDWARE

Figure 1. Activity chart.

would be possible. Further, the feedback of information, especially from surgical experiments,

hastened the solution of many problems related to the design of the surgical instruments and prosthetic device.

5. Concept of Method of Nerve Repair

and Evaluation Model

Essentially, the method employs a porous thinwalled tube similar to that shown in figure 2. A

small hole is drilled through one side of each tubular

sleeve at the Vs point along its length to serve as a vacuum port. The porous tube is fitted snugly into a matching groove in an adapter at the termination

of the barrel of the "anastomizer" making a "T"

connection with it. (See figs. 3 and 4). The vacuum

port in the porous tube is positioned to match the

opening at the bottom of the groove in the adapter.

Figure 2. Porous tube havine two diametrically opposite longitudinal micro-slots made by a spark discharge machining

technique.

94
Figure 3. Photograph of "vacuum approximator" showing vacuum tube, spring loaded handle for

clamping tube to replaceable tube support.

Squeezing the handle releases the tube from the tip of the applicator.

Figure 4. Photograph showing porous tube clamped to tip of vacuum tube or barrel.

Porous tube fits into the groove at end of a replaceable tube support, which may be removed from the tip of the barrel.

When the vacuum is apphed to the barrel, suction

at the ends of the porous tube first draws the proximal and then the distal nerve end into abutment

within the tube. The nerve is then secured to the

tube, either with an adhesive or mechanically by

surrounding supporting tissue structure. An important feature of the instrument is the small size

of the tube and holding device, thereby allowing access to spaces too small for the application of

sutures.

Presently, experiments are being conducted

with porous tubes 9 mm long and shorter by 1 mm

internal diameter and having a wall thickness

about 100 to 125 /am. The porosity of the tubes

range between 20 and 30 percent. The pores were

elongated circumferentially and longitudinally as a consequence of the platelet shape of the particles

and the nature of the tube forming process as shown

in figure 5. The average length of the pores ranged

from 8.4 to 14 /xm and the width from 2 to 2.5 fxin for those tubes implanted to date.

Figure 5. Photomicrograph showing pore structure of cross section of wall of tube taken from Rat No. 7.

Tube was fabricated from SS 316-L powder milled 192 hours and sintered at 1100°C

for 1 hour in pure hydrogen.

95
6. Device Fabrication

The process steps for tube fabrication are out- lined in figure 6.

METAL POWDEP CORE

(-325 Mesh 316-L POD Stainless Steel)

PCWDER CEARACTERIZATIORCOMMUnjTION

(In EtOH)

POWDEB CMRACTERlZATIOnSLURBY PFEPARATION

SIOTHY CMRACTEF.IZATI0.1-

DfERTSyBSTKATi:

VEHICLE BR BINDER

\ POWDER /

\ i / SLURRY PREPARATIOR

_SLUREY

CHARACTERIZATION

BN COATED ROD

KETAL COATED ROD

OREEH COATING CHARACTERISTICS

SINTERING PROCESS

SINTERED STRUCTURE

ULTRASONIC CLEANING

STERILE PACKAGING

IN VIVO EVALUATION

Figure 6. Flow chart for porous tube fabrication and evaluation.

It has been indicated earher in relation to the

factors contributing to successful nerve repair,

that porosity and permeabiHty of the tube wall are presently considered to be essential properties of

the tube. Thus, powder characteristics and sintered

pore structure have been an important part of the

experimental investigation.

The finest powders commercially available in

grade 316 L stainless steel are nominally —325

mesh (44 fxm and less). Figure 7 shows the appear- ance of the commercial stainless steel powder as revealed with a scanning electron microscope

(SEM). For a tube with a wall thickness of approximately 100 /u,m a much finer powder must be used.

Since ball milling is a common, inexpensive method

of reducing particle size in metal powders, it has

been used to produce fine metal powder in the form

of a coating slurry, a form which facilitates the

pressureless forming technique used to fabricate

the tubes. Figure 8 shows a Megapact high

^Available from Dymatron Inc.. Box 11158. Lexington. Ky. 40511. ^Certain commercial materials and instruments may be identified in this publication in order to adequately specify the experimental procedure. In no case does such

identification imply recommendation or endorsement by the National Bureau of Standards, nor does it imply that the equipment or instruments identified are necessarily the best available for the purpose.

Figure 7. SEM micrograph of as-received 316L powder.

Figure 8. High energy vibratory mill used to mill stainless

steel powder.

energy vibratory mill used to wet grind the stainless

steel powder into a stainless steel slurry of particles

having the size and shape shown in figure 9. The

platelet shape of the particles facilitates the coating

process.

Fabrication from fine powder permits adjustments

to be made to the pore size, wall thickness, and

surface area such that the tube is permeable to body fluids as well as resistant to the ingrowth of

glial and extraneous (scar) tissue which inhibits

regeneration.

A coating of inert material such as boron nitride

is first applied to the core rod to prevent reaction

and alloying between the core rod material and the

stainless steel coating during sintering.

A dipping machine, shown in figure 10, is used to coat short core rods mounted vertically on a boron

96
nitride furnace boat. Coating is achieved by dipping

and then withdrawing the core rods, first in the inert

coating slurry, then, after drying, in the stainless

steel slurry. After thoroughly drying, the double

coated core rods are introduced into a furnace

Figure 9. Stainless steel 316L ball milled for 192 hours in a

rotary ball mill.

chamber and sintered in an atmosphere of hydrogen

at temperatures between 1100 and 1200 °C for about

1 hour. After sintering, the sintered tubes are withdrawn from the core rods, facilitated by the shppery

nature of boron nitride. The tubes are next cut to size, the vacuum port drilled, and sharp edges

resulting from the cutting and drilhng operations

removed. Finally, the tubes are uhrasonically

cleaned to remove the inert coating powder and

stored in absolute alcohol until needed. Typical

tubes are shown in figure 11.

1 M 1 1

j

1 1 ifi lii | III Til

1

1 n 1 |llll| ill|llll| llll|llll{n ill: Ml

14 is 6 : 7 i8 i9 1.0 1.1

Figure 10. Dipping machine used to multiple coat core rods tube fabrication process.

Figure 11. Various porous tubes made by slurry coatingsintering process.

7. Summary

The need for a more reliable method of neurography capable of significantly increasing the

quality and number of neurorrhaphies that result

in useful function was indicated. The general

principles of nerve repair as currently practiced

were reviewed briefly for consideration in the design

of a tubular device to facilitate the repair of severed nerve tissue. The method of approach and organization of a research program to create and achieve

fabrication of a device that attempts to satisfy the

device design and functional requirements was

briefly outlined. The concept of the repair technique

was cdso briefly described. Finally the process

employed to fabricate the tubes was briefly outlined.

Part II of the paper describes some early results

obtained using the method to repair the severed

sciatic nerves of Wistar rats.

97
It should be mentioned that, ahhough the system

described herein was originally designed for the

epineurorrhaphy of small nerves with one or two

fascicles, there is the possibility of employing it for the repair of cleanly cut larger multifascicular nerves without segmental loss. Possibly application can be made to joining individual resected fascicles

in instances of segmental loss necessitating nerve

grafts to replace the deficiency of nerve tissue. The

absence of tension would ensure abutment of the nerve within the porous tube, since no tension force

is acting on them to cause separation of the nerve

endings.

8. References

[1] Kleinert, H. E., Serafin, D., and Daniel, R. K., The place of microsurgery in hand surgery. Orthopedic Clincs of North

America 4, No. 4 (October 1973).

[2] Khodadad, Ghahriman, Microsurgical techniques in repair of peripheral nerves. Surgical CUnics of North America

52, 1157-1166 (1972).

[3] Brown, Paul W., Factors influencing the success of the

surgical repair of peripheral nerves, Surgical Clinics of North America 52, 1137-1155 (1972).

[4] Hakstian, R. W., Perinaural neurrorhaphy, Orthopedic

Clinics of North America 4 (October 1973).

[5] Buncke, Harry J., Digital nerve repairs, Surgical Chnics of North America 52, 1267-1285 (1972).

[6] Smith, James W., Advances in facial nerve repair. Surgical

Chnics of North America52, 1287-1306 (1972) .

[7] Drake, Charles G., A caustic neuroma — repair of facial nerve with autogeneous graft, J. of Neurosurgery 17, 836

(1960).

[8] Campbell, James B., Facial nerve repair: new surgical tech- niques. Trans. Am. Acad. Opthal. Ololaryng. 68, 1068

(1964).

[9] Mayer, E. K., and Kimmel, D. L., The repair of severed motor and sensory spinal nerve roots by the arterial sleeve method of Anastomous, J. Comp. Neural 88, 285-317

(1948).

[10] Carlsson, C. A., and Thulin, C. A., Studies on the bridging

of irreducible gaps in severed ventral roots: an experimental study in cats, J. Neurol. Sci. 12, 163 (1971).

[11] Campbell, James B., Peripheral nerve repair, Clinician Neurosurgery 17, 77 (1969).

[12] Campbell, James B., Andrew, C, Bassett, L., and Bohier,

J., Frozen-irradiated homographs shielded with microfilter sheaths in peripheral nerve surgery, J. of Trauma 3,

303 (1963).

[13] Campbell, J. B., et. al, Microfilter sheaths in peripheral nerve surgery, J. of Trauma 1, 139 (1961).

[14] Bassett, C. A., and Campbell, J. B., Calcification of milhpore

in vivo. Transplantation Bulletin 26, 132 (1960).

[15] Lehman, R. A. W., Hayes, G. J., Leonard, F., Toxicity of

alkyl 2-cyanoacrylates-peripheral nerve. Archives of Surgery 93 , 441 (1966).

[16] Leonard, F., Hemostatic applications of alpha cyanoacrylates: bonding mechanism and physiological degradation of bonds. Chapter 11, Handbood of Biomedical

Plastics, (Pasadena Tech. Press, Pasadena, Ca.).

[17] Kleinert, H. E., and Griffin, J. M., Technique of nerve

anastomosis. Orthopedic Clinics of North America 4

(October 1973).

[18] Millesi, H., Berger, A., and Meissel, G., Fascicular nerve

grafting. Transitions of the Fifth International Congress

of Plastic and Reconstructive Surgery, (London, Butter- worth and Company, Ltd., 1971).

[19] Millesi, H., Meissel, G., and Berger, A., The interfascicular nerve grafting of the median and ulnar nerves, J. Bone

Joint Surg. 54A, 727 (1972).

[20] Bora, W. F., Peripheral nerve repair in cats, J. Bone Joint Surgery 49A, No. 4 (1967).

[21] Peacock, E. F., and Van Winkle, W., Surgery and biology

of wound repair (W. B. Saunders Co., Philadelphia, Pa.,

1970).

[22] Visit with Dr. MiUesi in Vienna, Austria, January 31, 1973.

98
NATIONAL BUREAU OF STANDARDS SPECIAL PUBLICATION 415,

Biomaterials, Proceedings of a Symposium held in conjunction with the Ninth Annual Meeting of the Association for the Advancement of Medical Instrumentation, New Orleans, La., April 19-20, 1974

(Issued May 1975).

A Nerve Implant Prosthesis for Facilitating Peripheral Nerve Regeneration *

Part II. Development of the Prosthetic Device and System of Repair

Neuroanatomical Aspects

James L. Hall

Department of Anatomy, College of Medicine, University of Cincinnati, Cincinnati, Ohio 45219

William E. Kuhn

Department of Materials Science and Metallurgical Engineering, University of Cincinnati, Cincinnati, Ohio 45219

The principal neuroanatomical aspects of peripherical nerve repair are the approximation of the severed ends and the minimizing of trauma during the approximation. The method of using vacuum

to draw the severed ends down a porous tube seems to take these two factors into consideration. The

porous tube provides a shielded environment for the regenerating nerve to grow into the approximated

distal end. The porosity allows for the escape of edematous fluids, prevents the invasion of scar tissue and provides for the flow of nutritional fluids. Histological and statistical results of the procedures are presented. In most instances the axons

of the regenerating nerve grow across the gap and occupy the intact neurilema sheaths of the de- generating distal segments.

Key words: Histological evaluation; nerve axons; surgical procedure.

Any discussion of a method to facilitate peripheral nerve regeneration should first take cognizance

of the fact peripheral nerves are capable of re- generating under ordinary circumstances. Our

prosthesis and the method of application is an

attempt to capitalize and enhance that capability.

Peripheral nerves, unlike their counter-parts

in the central nervous system, are ensheathed in a connective tissue covering called the neurilemma

or sheath of Schwann. This sheath effectively serves

as a conduit or pipeUne through which axons or

fibers of a severed peripheral nerve can grow if no obstacles are encountered enroute to the target

tissue.

If the damage or insult to the nerve has not oc- curred too close to the spinal cord and its accompanying dorsal root ganghon, the nerve cells or neurons in these locations wiU attempt to reconstitute themselves. The degree of success of the re- constitution will depend on the severity of the

injury, the general health of the individual, and,

as mentioned, the proximity of the injury to the cells

themselves. An injury within two or three centimeters of the nerve cells will eUcit greater damage

to the cells and thus result in a poorer chance for

functional recovery than would an injury some

distance away.

* This research was supported by the National Science Foundation under Grant GH-33873.

Our study is centered on nerve separations that occur some distance from the cell -bodies, so we

may assume that a very large percentage, if not

aU, of the cells will recover from the injury.

The axons of the nerve begin to grow very

tenuously from the proximal end of a severed nerve

in about 3—5 days. At first these are very delicate

strands. Such deUcate tendril-hke fibers have been

called fibers of Von Bungner. They are groping or seeking a direction for further growth. It is at this

early stage of regeneration that our prosthesis must

satisfy the principal requirements. The tube provides a guide for the neurilemmal sheath to follow thus

bridging the gap between the proximal intact nerve and the distal severed end. The neurilemma grows

along the long axis of the tube. It is critical, of

course, that the two ends of the nerve be brought

in as close approximation as possible to provide the

optimal situation so that the rapidly proliferating

neurilemma wiU bridge the gap between the two

nerve ends prior to the emerging regenerating axon.

It is also critical that the edematous fluids which

accompany any injury to hving tissue escape through

the pores in the tube. If these fluids are confined

within the tube or any other membrane they will act as a physical barrier to axon growth during the early

phase of regeneration. Such fluids may indeed

crush the axons and prevent regeneration if the

volume is sufficient and no provision is made for

99
release. This undesirable possibility has prompted

many workers to search for either natural or synthetic membranes that could be wrapped around

the nerves and still permit escape of fluids.

Surgical Approach. —The sciatic nerve of the rat has been used for our experimental model. The

nerve has approached through a midline incision

in the posterior thigh, the musculature reflected,

and the nerve cleared of any fascia or excessive

connective tissue. The sciatic nerve in the rat is approximately 1 mm in diameter. The nerve was

exposed for a distance of about 13-15 mm., and was

then severed distal to the branch to the long head

of the biceps femoris. Care was taken to produce

a clean cut, so that the edges of the nerve were

smooth. Usually the two separated ends of the nerve

retracted and a gap of from 5-7 mm developed

between them. The field was cleaned of blood or

debris.

Method. — The method of nerve repair described

in Part I was followed to bring the nerve ends into

apposition. A fluid trap preceded the pressure

gauge and vacuum pump. The pressure differential

between the interior of the tube and ambient

atmospheric pressure was adjusted to give a value

between 50 and 75 mm Hg. This pressure range

sufficed to draw the nerves firmly into the tube.

The anastomizer with the attached tube was

placed in the field, and the two ends of the severed nerve were drawn into the long axis of the tube by

suction. The distance the nerves are drawn into the

tube can be gauged visually but it is not possible for us to determine accurately whether the ends

meet. The nerve will adhere to the tube naturally,

but we are experimenting with various adhesives

to fix them in place. The wounds were sutured and

the animals were given veterinary penicillin at the

site of the wound and also intramuscularly.

Recovery. — The animals were examined frequently during the 120 day survival period. At first the animals dragged the operated limb. The Umb

becomes swollen, sores may develop, and care must

be taken to prevent infection. Usually after 3-4

weeks signs of recovery could be seen. The Umbs

became functional and recovery was fairly complete

by the end of 6 weeks. During the 120 day period a number of the animals regained fuU use of the limb,

and responded to pinching or a pin prick by withdrawal. The rats could stand, walk, run, wash, and

in every way demonstrated a functional recovery.

Histological Results.— The animals were anesthetized with nembutal, and sacrificed by perfusion

with 10 percent formalin through the left ventricle

of the heart, after a wash with warm normal saline

to prevent blood clotting.

Photographs similar to that in figure 1 were taken

of the gross experimental limbs with the tubes

in situ. The nerve was removed and fixed for 10

days in 10 percent formalin. The excess connective

tissue was cleaned off the tubes and the nerves were gently extracted from them. Figure 2 shows a

::':|iiM|iiH|Mii|nn|iiii|iiii|iiililiii|iiii|iiii|iiii{ini|iiii|liii|iHi<Mi:i

2 '3 4 !5 16 7 8 ? • :

Figure 1. Photograph of disected hind limbs of Rat No. 27

after 110 days showing the experimental leg with implanted

tube T on the left and the normal leg on the right.

Nerves N can be observed. Full function was restored to the operated leg with no

indication of atrophy of muscles in the toes, foot or posterior thigh.

Figure 2. Photograph showing sciatic nerve withdrawn from

porous tube takenfrom Rat #15

Before sacrifice rat exhibited withdrawal reflex, could support weight on operated leg, displayed no clawing of toes, and exhibited normal walking gait. Post mortem

examination revealed no discernible atrophy of muscles of foot or thigh.

100
segment of nerve that has been withdrawn from a porous tube.

The nerves were then cut in three sections: —

proximal to the cut, at the level of the cut but with

proximal and distal tissue on either side of the cut,

and distal to the cut. The tissues were blocked in

paraffin and positioned to provide cross sections of

the proximal and distal ends and longitudinal

sections of the area around the cut. The tissues were

impregnated with silver according to the Holmes

modification of the Bielchowsky silver method.

Preliminary histological results seem to indicate

that if the procedure described is done properly

there is a minimal axonal loss in the regenerated

nerve. Comparisons of proximal and distal ends

indicate a diminution of axons so slight as to make

it difficult to distinguish between the two levels.

Longitudinal sections (fig. 2) at the level of the cut

also show no appreciable difference between the

Figure 3. Photomicrograph at the level of sectioning of rat

sciatic nerve.

Note relative similarity between two portions of axons above and below the cut. (Holmes

Silver Impregnation 100 X)

Figure 4. Hieber power magnification at level of cut of regen- erated rat sciatic nerve.

Only a few axons do not show effective repair. (Holmes Silver Impregnation 220 X)

proximal and distal ends. The level of the cut,

however, can be identified. Axonal counts are in

progress but we are not able to report on them in

this paper. However, it would appear that the axonal

difference between the proximal and distal ends will

be less than 20 percent. (See figs. 3 and 4). Further work is needed, but we suggest that this

method of facilitating peripheral nerve regeneration

holds great promise. It may be possible to repair

small nerves where current surgical methods often

produce further axonal loss due to the manipulation

involved. Further, certain cranial nerves inside the

cranial cavity are practically inaccessible for

surgical repair but could be joined with our tube.

In addition, this method may result in the repair

and restoration of many nerves that now must be

sacrificed in order to achieve more significant

surgical objectives.

101

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4. TITLK AND SUBTITLH

Biomaterials (Proceedings of the Symposium on

Biomaterials , 9th Annual Meeting, Association for

Advancement of Medical Instrumentation, New Orleans,

LA. April 19-20, 1974)

3. Recipient's Accession No.

5. Publication Date May 1975

6. Performing Organization Code

7. AUTHOR(S)

E. Horowitz and J. L. Torgesen, Editors

8. Performing Organ. Report No.

9. PERFORMING ORGANIZATION NAME AND ADDRESS

NATIONAL BUREAU OF STANDARDS

DEPARTMENT OF COMMERCE

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15. SUPPLEMENTARY NOTES

Library of Congress Catalog Card Number: 74-30305

16. ABSTRACT (A 200-word or less factual summary of most si^ificant information. If document includes a significant

bibliography or literature survey, mention it here.)

This volume is based on papers presented at the Symposium on Biomaterials, held in conjunction with the Ninth Annual Meeting of the

Association for the Advancement of Medical Instrumentation, New Orleans,

LA. April 19-20, 1974. It provides a review of special topics in biomaterials research selected to focus attention on some noteworthy

achievements. The topics covered include plasma-polymerized polymers

and their application in biomedicine; biocompatibility of ceramic

materials and their application both as inert coatings for synthetic

implants and a? porous materials for bone repair; the selection of

metallic implant materials through engineering and medical consideration

in vitro testing of thromboresistance ; adsorption of blood proteins

on synthetic substrates; a prosthesis for nerve regeneration; and

properties of fibrous biomaterials.

3;

17. KEY WORDS (six to twelve entries; alphabetical order; capitalize only the first letter of the first key word unless a proper name; separated by semicolons)

Key words: Biocompatibility; biomaterials; blood protein; ceramic

implants; implantable electrodes; metallic implants; nerve prosthesis;

synthetic implants;—thromboresistance . 18. AVAILABILITY (xJ Unlimited

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PERIODICALS JOURNAL OF RESEARCH reports National Bureau

of Standards research and development in physics,

mathematics, and chemistry. It is published in two sections, available separately:

• Physics and Chemistry (Section A)

Papers of interest primarily to scientists working in these fields. This section covers a broad range of physical and chemical research, with major emphasis on

standards of physical measurement, fundamental constants, and properties of matter. Issued six times a

year. Annual subscription: Domestic, $17.00; Foreign,

$21.25.

• Mathematical Sciences (Section B)

Studies and compilations designed mainly for the mathematician and theoretical physicist. Topics in mathematical statistics, theory of experiment design, numerical analysis, theoretical physics and chemistry, logical

design and programming of computers and computer

systems. Short numerical tables. Issued quarterly. Annual subscription: Domestic, $9.00; Foreign, $11.25.

DIMENSIONS/NBS (formerly Technical News Bulletin) —This monthly magazine is published to inform

scientists, engineers, businessmen, industry, teachers,

students, and consumers of the latest advances in science and technology, with primary emphasis on the work at NBS. The magazine highlights and reviews such

issues as energy research, fire protection, building tech- nology, metric conversion, pollution abatement, health and safety, and consumer product performance. In addition, it reports the results of Bureau programs in measurement standards and techniques, properties of matter and materials, engineering standards and services, instrumentation, and automatic data processing.

Annual subscription: Domestic, $9.45; Foreign, $11.85.

NONPERIODICALS

Monographs—Major contributions to the technical liter- ature on various subjects related to the Bureau's scientific and technical activities. Handbooks—Recommended codes of engineering and

industrial practice (including safety codes) developed

in cooperation with interested industries, professional

organizations, and regulatory bodies.

Special Publications—Include proceedings of confer- ences sponsored by NBS, NBS annual reports, and other

special publications appropriate to this grouping such

as wall charts, pocket cards, and bibliographies.

Applied Mathematics Series—Mathematical tables,

manuals, and studies of special interest to physicists,

engineers, chemists, biologists, mathematicians, computer programmers, and others engaged in scientific and technical work.

National Standard Reference Data Series—Provides

quantitative data on the physical and chemical properties of materials, compiled from the world's literature and critically evaluated. Developed under a world-wide

program coordinated by NBS. Program under authority

of National Standard Data Act (Public Law 90-396).

NOTE: At present the principal publication outlet for these data is the Journal of Physical and Chemical

Reference Data (JPCRD) published quarterly for NBS

by the American Chemical Society (ACS) and the American Institute of Physics (AlP). Subscriptions, reprints,

and supplements available from ACS, 1155 Sixteenth

St. N. W., Wash. D. C. 20056.

Building Science Series—Disseminates technical infor- mation developed at the Bureau on building materials,

components, systems, and whole structures. The series presents research results, test methods, and performance criteria related to the structural and environmental functions and the durability and safety characteristics of building elements and systems.

Technical Notes—Studies or reports which are complete

in themselves but restrictive in their treatment of a

subject. Analogous to monographs but not so comprehensive in scope or definitive in treatment of the subject area. Often serve as a vehicle for final reports of work performed at NBS under the sponsorship of other government agencies.

Voluntary Product Standards—Developed under pro- cedures published by the Department of Commerce in Part 10, Title 15, of the Code of Federal Regulations.

The purpose of the standards is to establish nationally

recognized requirements for products, and to provide

all concerned interests with a basis for common understanding of the characteristics of the products. NBS

administers this program as a supplement to the activities of the private sector standardizing organizations.

Federal Information Processing Standards Publications

(FIPS PUBS)—Publications in this series collectively

constitute the Federal Information Processing Standards Register. Register serves as the official source of information in the Federal Government regarding stand- ards issued by NBS pursuant to the Federal Property

and Administrative Services Act of 1949 as amended,

Public Law 89-306 (79 Stat. 1127), and as implemented

by Executive Order 11717 (38 FR 12315, dated May 11, 1973) and Part 6 of Title 15 CFR (Code of Federal

Regulations).

Consumer Information Series—Practical information,

based on NBS research and experience, covering areas

of interest to the consumer. Easily understandable

language and illustrations provide useful background

knowledge for shopping in today's technological

marketplace.

NBS Interagency Reports (NBSIR)—A special series of interim or final reports on work performed by NBS for

outside sponsors (both government and non-govern- ment). In general, initial distribution is handled by the sponsor; public distribution is by the National Technical

Information Service (Springfield, Va. 22161) in paper

copy or microfiche form.

Order NBS publications (except NBSIR's and Biblio- graphic Subscription Services) from: Superintendent of Documents, Government Printing Office, Washington,

D.C. 20402. BIBLIOGRAPHIC SUBSCRIPTION SERVICES

The following current-awareness and literature-survey

bibliographies are issued periodically by the Bureau:

Cryogenic Data Onter Current Awareness Service

A literature survey issued biweekly. Annual subscription: Domestic, $20.00; foreign, $25.00.

Liquefied Natural Gas. A literature survey issued quarterly. Annual subscription: $20.00.

Superconducting Devices and Materials. A literature

survey issued quarterly. Annual subscription: $20.00.

Send subscription orders and remittances for the pre- ceding bibliographic services to National Technical

Information Service, Springfield, Va. 22161.

Electromagnetic Metrology Current Awareness Service Issued monthly. Annual subscription: $100.00 (Special rates for multi-subscriptions). Send subscription

order and remittance to Electromagnetics Division,

National Bureau of Standards, Boulder, Colo. 80302.
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