: Metallurgical Studies of Interface

|L Bonding on Implant Alloys

U.S. DEPARTMENT OF COMMERCE

National Bureau of Standards

National Measurement Laboratory

Center for Materials Science

Metallurgy Division

Washington, DC 20234

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October 1982

Prepared for: FDA Contract 22 4795023

Bureau of Medical Devices

ood and Drug Administration

“Iver Spring, MD
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NBSIR 82-2563 (FDA)

METALLURGICAL STUDIES OF INTERFACE

BONDING ON IMPLANT ALLOYS

A. C. Fraker

A. W. Ruff

K. J. Bundy

NOV 4 1982

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with S. A. DeMontigny

P. Sung

A. C. Van Orden

K. M. Speck

U.S. DEPARTMENT OF COMMERCE

National Bureau of Standards

National Measurement Laboratory

Center for Materials Science

Metallurgy Division

Washington, DC 20234

October 1982

Prepared for: FDA Contract 22 4795023

Bureau of Medical Devices

Food and Drug Administration

Silver Spring, MD

U.S. DEPARTMENT OF COMMERCE, Malcolm Baldrige, Secretary

NATIONAL BUREAU OF STANDARDS, Ernest Ambler, Director
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Abstract

A literature review covering articles on stress analyses in total

hip replacement prostheses and on the metal /bone cement interface is

presented. The literature indicated the need for a test which utilized

loading in torsion, and such a test was developed and is described. The

test can be used to determine the influence of various parameters, including surface roughness and passivation and sterilization treatment on

the strength of the metal/bone cement bond. Some preliminary tests were

conducted and results are given. Future work using the mechanical test

developed is discussed. Additional studies were conducted on the surface

preparation of titanium, and data are presented to show that changes in

initial electrochemical behavior and varying degrees of surface roughness

occur depending on whether the metal receives a neutral, alkaline or

acidic washing treatment.

TABLE OF CONTENTS

Page

I. Introduction i

II. Investigation of Stress Analyses and the Metal/Bone Cement

Interface in Total Hip Replacement Prostheses

by Kirk J. Bundy... iiia

III. Surface Preparation and Corrosion Behavior of Titanium Alloys

for Surgical Implants by A. C. Fraker, A. W. Ruff, P. Sung,

A. C. Van Orden, and K. M. Speck 64

IV. Acknowledgments 84
^ \ :
INTRODUCTION

This is a report of a one year effort which was planned and conducted

to investigate the effects of surface treatnents on surgical implant

metals as related to interface problems between metals and bone cement in

joint prosthesis fixations. It has been recognized for some time that

prosthesis loosening is one of the most serious material problems confronting total or partial joint replacement today. While the cement

sometimes is considered to be a grouting material, some adherence occurs

between the bone and cement, and between the cement and the metal. The

importance of this adherence is verified by the fact that the x-ray

radiographic appearance of a line separating either bond is a prelude to

failure of the device. This failure can be due to fatigue and fracture,

or overload fracture which can lead to pain and malfunction of the device.

Recently, at Northwestern University, a study was made of several

factors affecting surgical implant metal/bone cement adhesion. Those

data indicated that surface treatments on metals such as electropolishing

and passivating resulted in the strongest adhesive bond when joined to

the cement at the onset of the dough stage (after 3 minutes). It is

clear from this work that much useful information can be obtained from

studying these interfaces. To date, very little data exists on the

nature of the metal /cement interface and the bond strength. Measurements

of the surface roughness on the metal and the surface.film chemical

composition should be made and correlated with mechanical strength of a

cement bond to the given metal. These measurements can contribute data

for use in standards and tests methods which can provide improved quality

1
assurance against device loosening and more adequately assure safety

and long tern usefulness of the device.

A review of literature and other documented information on surface

treatments and test methods for bond strength has been completed and

is summarized in Part II by Dr. Kirk Bundy, of the Johns Hopkins University,

who worked at NBS on this project. Over 200 published articles pertaining

to bone cement and cemented prostheses were collected and reviewed. In

addition, hundreds of other references which related to other considerations

in prosthesis installation were reviewed. These articles are now part of a

reference file containing over 1000 articles pertaining to skeletal implants.

The literature review indicated that several factors affect the

strength of the bone cement/metal interface bond including the type of

bone cement and the surface conditions of the material. It is clear that

several fracture and fatigue test results on the bone cements are not in

agreement. One study indicates superior fracture toughness and fatigue

behavior for one type of bone cement while the other study shows that a

different bone cement is superior in these qualities. A third study

discussed some differences in the cements but did not indicate differences

in mechanical strengths.

Alternative methods of prosthesis fixation are being considered and

developed. However, there are many patients who always require a cement

fixation of the prosthesis. Metal porous coatings on metal prostheses

are being studied for use in cementless prostheses. Porous metal, porous

polymer, ceramic, and bone cement coatings could be applied to a prostfiesis

in selected areas where the forces on the bone cement/metal interface are

the greatest such as the upper side of the proximal end of the femural

component.
Most manufacturers follow the procedures given in the ASTM F4,

F- 86 , document entitled "Standard Recommended Practice for Surface Preparation and Marking of Metallic Implants." Stems of hip prostheses often

are given a satin finish by dry blasting with AI 2O3 and/or glass beads.

The articulating areas are highly polished. Prostheses are cleaned as

described in F- 86 . There are steps in the procedure where differences

can occur such as washing after different treatments. The final steps

are passivation and sterilization. NBS studies of the surface preparation

and corrosion behavior of titanium and titanium alloys were presented in

a paper at the ASTM Symposium on Titanium for Surgical Implants, May,

1981 and are presented in Part III of this report.

The test methods of ASTM D-14 on adhesives were reviewed (see Part II)

and found not to be appropriate for testing the bone cement/netal interface.

An appropriate test method was developed to the point of application for

measuring bond strength of the cement/netal interface to test various

surface related parameters. The test is described in detail in Part II.

This test will be used in the continuation of the metal/bone cement

interface studies

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Part II

INVESTIGATION OF STRESS ANALYSES AND THE METAL/BONE CEMENT INTERFACE

IN TOTAL HIP REPLACEMENT PROSTHESES

Kirk J. Bundy

TABLE OF CONTENTS

Page

List of tables iv

List of figures v

Chapter I INTRODUCTION 1

Chapter II EFFECT OF ALTERED STRESS DISTRIBUTION ON BONE 3

Background Information 3

Implant Induced Response of Bone to Changes in Stress. 5

Chapter III A REVIEW OF STRESS ANALYSES AND MEASUREMENTS OF TOTAL

HIP IMPLANTS 9

Methods Used to Investigate Stresses in Hip Implants . 11

Problems Examined by Stress Analysis of Total Hip

Replacements 13

Results of Experimental Measurements of Total Hip

Replacements 14

Results of Beam Theory Applied to Total Hip

Replacements 17

Approaches Used for Finite Element Analysis of Total

Hip Replacements 19

Results of Finite Element Method Studies of Total

Hip Implants 22

Calculations of Stresses at the Metal /Bone Cement

Interface 25

Chapter IV CONCEPTS FOR IMPROVEMENT OF STRESS ANALYSES OF TOTAL

HIP IMPLANTS 28

Clinical Observations of Failure Modes of Hip

Implants 28

Joint Loads 30

Chapter V CHARACTERIZATION AND TESTING OF METAL/BONE CEMENT

INTERFACES 34

Interfacial Bond Strength Tests 34

Studies of the Mechanical Behavior of Bone Cement and

Bone Cement/Metal Interfaces 36

Experimental Studies 45

Chapter VI SUMMARY AND RECOMMENDATIONS 53

References 56

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LIST OF TABLES

Page

Table II-l Young's Modulus of Surgical Implant Materials and Bone . 8

Table III-l Various Approaches Employed for Finite Element Analysis

of Total Hip Replacements 20

Table V-1 Shear Strengths of Bone Cement/Material Interfaces for

Various Metals and UHMWPE 42

Table V-2 Experimental Values of Proportionality Constant 6"

between Torque and Twist Angle 50

iv
LIST OF FIGURES

Figure

Figure

Figure

Figure

Figure

III-l Total Hip Replacement Prosthesis and Applied

Joint Force Components

Page

10

III-2 Normal Stress a and Shear Stress i in Bone Cement

n at the Cement/Stem Interface 26

IV- 1 Variation of Joint Force Components with Time

During the Walking Cycle 31

V- 1 Loading Modes Employed for ASTM Standard Tests of

Strength Properties of Adhesives 35

V-2 Length-normalized Torque T/L vs Twist Angle y

clC for Interfaces 7, 8, and 9 46

V

CHAPTER I

INTRODUCTION

This report describes activities pursued in the Metallurgy Division

of the National Bureau of Standards to study the influence of the

metal /bone cement interface on the performance of prostheses. The

rationale for this work is described below.

In recent years the field of surgical implants has become increasingly important for a number of reasons. One reason is that the

average age of the population is increasing and thus greater numbers

of individuals are in need of implants. Another is that devices are

being implanted in younger individuals more often now than previously

was the case. Therefore, these devices will serve within the body's

environment for longer periods of time. A case in point involving a

particularly demanding application is the artificial hip.

As a consequence of these and other trends, there is increasing

interest in the factors which govern the success and failure of orthopedic implants. This subject is a complicated one and involves both

physiological and engineering effects. In the latter category, corrosion of the implant metals and the stress distribution in the implant--

bone cement--bone region have been related to implant failures. The

interfacial regions are particularly prone to loosening and failure.

Another aspect of the matter is that alteration of the stresses in

bone due to the presence of implants can occur which may lead to

adverse physiological consequences.

In order to gain further insight into these problems, the research

described in this report, mainly a literature search, was undertaken.

In Chapter II the effects on bone of alterations of the natural

1
physiological stress distribution caused by the presence of implants

and other means is described. Chapter III is concerned with a review

of stress analyses of total hip replacements. Both experimental and

theoretical studies are examined. Particular attention is paid to

analyses of the stresses at the bone cement/metal interface.

Chapter IV reviews clinical reports of failed bone cement/metal

interfaces and compares them with the results of stress analyses. The

lack of correspondence between the two suggests that torsional forces

which act at the interface could be important in its failure.

Chapter V reviews methodology for testing interfacial bond

strengths. ASTM standards are reviewed and the literature pertaining

to mechanical testing of bone cements and cement/metal interfaces is

discussed. Finally, experimental results of preliminary pilot studies

on measuring torsional strengths of bone cement/metal interfaces are

described. Conclusions and recommendations are given in Chapter VI.

2
Chapter II

EFFECT OF ALTERED STRESS DISTRIBUTION ON BONE

Background Information

One of the main reasons for interest in stress analyses of total

hip replacements is, in addition to insight into failure of the components themselves, to determine the stresses present in the surrounding

bone. Maintenance of an adequate stress level in skeletal structures

is probably the major factor in maintaining the normal balance between

bone formation and resorption processes (Tonino et al . , 1976). This

relationship was first pointed out in 1870 (Wolff, 1870) and has since

become known as Wolff's law which can be stated (Pugh et al., 1973):

... changes in stress of the bone induce changes in its structure

appropriate to support the changed load.

Unfortunately Wolff's law is a qualitative principle and not a

quantitative relationship which can be used to predict a reduction of

bone mass or density which results from a given reduction in bone

stress level. If Wolff's law was a quantitative relationship, much

additional insight into bone's response to stress could be obtained

from stress analyses such as those described in Chapter III.

Various mechanisms to explain Wolff's law have been procosed.

Bone is considered by some researchers to be a piezoelectric material

(Fukada and Yasuda, 1957), (Shamos and Lavine, 1964). It has also

been suggested that physical movement of intercellular fluid over the

charged surfaces of cells, which is caused by mechanical stress,

induces streaming potentials (Dumbleton and Black, 1975). It has been

proposed that the electric fields caused by applied loads (or absent

in the absence of load) act as "signals" which trigger osteoblastic

3
and osteoclastic activity and regulate the rates of bone production

and resorption. Bone in its response to stress has been compared to a

system controlled by a feedback loop (Kummer, 1972). It has been

demonstrated that even in the absence of stress, applied fields of the

appropriate voltages, current levels, and frequencies stimulate bone

growth (Dumbleton and Black, 1975).

In cancellous bone which was caused to remodel by experimental

loading, Pugh, Rose, and Radin (1974) observed significant fracture of

trabeculae on a microscopic scale. The healing of these fractures

produces large amounts of callus which serves to stiffen the bone.

They propose that this mechanism could underlie Wolff's law and could

be responsible for cancellous bone remodeling in response to changes

in stress patterns.

Whatever the mechanism behind Wolff's law, it is clear that

absence or reduction in stresses can lead to adverse consequences to

skeletal tissue. This seems to be true no matter how the stress

reduction comes about: through weightlessness (as with astronauts),

limb paralysis (as in pol iomyel itus) , bed rest, or stress redistribution due to implants.

The time scale required for a stress level reduction to affect

bone can be quite short. The bones of astronauts in weightless

environments undergo demineralization (Kummer, 1972). The effects are

measurable after fourteen days in space (Kazarian and Von Gierke,

1969). On the other hand, as discussed in the next section, a longer

term consequence of reduction in stress, such as that which occurs in

the bone around implants, is resorption.

4
Implant Induced Response of Bone to Changes in Stress

When a load-bearing metal component is introduced into bone

tissue, a reduction of the load transmitted through the adjacent bone

will occur. To a certain degree this is desirable. For example,

fracture fixation plates function by partially taking the load off the

fractured pieces of a bone and rigidly holding them in place so that

healing can occur.

This reduction in load comes about since most implant designs

will produce a situation in which the displacements of the surrounding

bone due to externally applied forces will be constrained to be either

equal to those in the stiffer implant or at least reduced from those

present when an intact bone is loaded. The result is a decrease in

strains in the bone and thus a decline in stress levels also. This

effect is often called "stress shielding" or "stress protection".

Due to the action of Wolff's lav/ stress shielding can have a

number of undesirable consequences for bone tissue as discussed below.

In fracture fixation plates, the need for high stiffness to insure

stability for healing and the need for low stiffness thereafter to

avoid the harmful results of stress shielding pose an unresolved

dilemma in orthopedic surgery.

For an implant of the same material and shape. Woo et al . (1977)

have used finite element calculations to demonstrate that the stresses

in bone diminish as the cross-section increases. Also, it has been

experimentally shown that for implants of the same size and shape, the

stresses diminish as the modulus of elasticity of the material

increases. Woo et al . (1976) conducted 9 to 12 month studies of

implants made from a graphite fiber methyl methacrylate composite

5
resin and from a Co-Cr alloy. They found that stress shielding did

not affect the mechanical properties (ultimate bending strength,

energy absorption capacity, and flexural modulus of elasticity) of

bone as a tissue material but did adversely effect the bone's ability

to bear load due to a change in its structural properties, i.e.,

thinning of the cortex. Significantly less severe effects were noted

for the more flexible nonmetal lie material.

In a somewhat similar study Tonino et al . (1976) have demonstrated

that bone mass per unit length decreases due to both trabeculation of

previously compact areas of bone at the endosteal surface and to

decreased mineralization of newly formed bone in this area. These

results were obtained after 10 to 18 weeks of implantation of stainless

steel or polytrifluoromonochloroethylene in dog femora. Corresponding

reductions were observed in effective ultimate bending strength and

effective Young's modulus (as determined from tests on whole bones).

These effects were more pronounced with the steel. Slatis et al

.

(1980) have observed cortical thinning and development of porosity

near dynamic compression plates and have measured corresponding reductions in torsional loads required to produce fracture.

Observations of disuse osteoporosis beneath plates appear often

in the literature and have also been noted by Dodge and Cady (1972),

Matter et al. (1974), Olerud and Karlstrom (1972), Richon et al

.

(1967) and others. A number of investigators (Cochran, 1969), (Kdbler

and Wiechell, 1971), (Uhthoff and Dubuc, 1971), (Diehl and Mittelmeier,

1974) have suggested that the osteoporosis produced by stress shielding

leads to increased likelihood of fatigue fracture after removal of the

plate. A well recognized complication of the use of plates for internal

6
fixation is a refracture or a fracture through the porotic portion

(Chrisman and Snook, 1968), (Dencker, 1964), (Frankel and Burstein,

1968), (Solheim, 1972).

Since there is a wide disparity between the Young's moduli of the

commonly used implant metals and bone (see Table II-l) and larger,

thicker, more rigid plates are growing in popularity (Moyen et al

.

1978), (Muller et al . 1965) it can be expected that reports of harmful

effects due to stress shielding by fracture fixation plates will

continue to occur in the future. The development of composite materials with elastic properties more closely matched to bone and/or

development of higher strength alloys which can be used in thinner

sections could represent a long term solution to this problem.

Problems with regard to stress protection or shielding are not

confined to fracture fixation plates. One way in which total hip

replacements fail is by fracture of the femoral stem. Lack of bony

support in the area of the calcar femoral e seems to occur often in

cases of stem fracture—always according to Galante (1980) although

Coll is (1977) found other causes related to heavy patient weight,

inadequate valgus positioning, and metallurgical inadequacies.

Whether or not the ultimate result is stem fracture, resorption

in the medial area of the femoral neck is common in total hip replacements. Observed frequencies have been 16.8 percent (Beckenbaugh and

Ilstrup, 1978) and 21 percent (Bocco et al .

,

1977). Inadequate transfer of load from the implant to the calcar due to lack or inadequacy of

physical contact between them has been suggested by several investigators as a biomechanical cause for the resorption (Galante, 1980), (Oh

and Harris, 1978), (Markolf et al . , 1980). These investigators have

7
also mentioned other possible causes however: a relative ischemia

(i.e., a subnormal amount of blood) of the area or tissue reaction to

bone cement wear particles. Another nonbiomechanical cause for resorption has been proposed by Smethurst and Waterhouse (1978). In a study

of failures of metal -on-metal total hip replacements, bone loss was

attributed to corrosion of the components which resulted in release of

Cr and Co ions.

TABLE II-l

YOUNG'S MODULUS OF SURGICAL IMPLANT MATERIALS AND BONE

Material Young's Modulus

(10^ MPa)

Bone 20

316 L Stainless steel 200

Cast Co-Cr alloy 200

Wrought Co-Cr alloy 230

MP35N 230

Pure Ti 100

Ti-6A1-4V 100

PMMA bone cement 2

after Swanson and Freeman (1977)

8
Chapter III

A REVIEW OF STRESS ANALYSES AND MEASUREMENTS OF

TOTAL HIP IMPLANTS

Many studies have been conducted to obtain stress analyses of

total hip replacements for purposes of design improvements or development of understanding of the biomechanical basis of clinical failures.

This work is reviewed in this chapter.

To perform a theoretical stress analysis, the following information must in general be known: (1) the geometry of the structure of

interest, (2) its relevant mechanical properties, in particular, its

elastic properties, (3) the loads applied at its surface or the surface

displacements, and (4) appropriate boundary conditions.

An exact stress analysis of a total hip replacement procedure

which provides information about the stress state at all points in the

metal prosthesis, bone cement, bone itself, and at the interfaces

between them is an exceedingly difficult and perhaps almost intractable

task because of difficulties associated with all four areas mentioned

above.

The complexity of this problem can be appreciated by examining

Figure III-l. The geometry which is involved is irregular and may

require a three dimensional treatment of the problem. There are three

different materials whose elastic properties are important: bone, bone

cement, and the alloy of the femoral component. The properties of the

cement can vary considerably depending upon its handling and various

environmental factors. In the case of bone (at least) the situation

is further complicated by the spatially nonuniform and anisotropic

nature of the tissue and the presence of both cortical and cancellous

9
8

FIG. m-1 Total Hip Replacement Prosthesis and

Applied Joint Force Components

10
bone. The loads in the hip joint which are applied to the prosthesis

have components in all three coordinate directions which vary with time,

and a consensus does not appear in the literature as to what force state

should be used in calculations. As will be discussed later in the

report, the boundary conditions which should be used to model the bone/

bone cement and metal /bone cement interfaces are also uncertain.

It is thus not surprising that a variety of methods have been

used for hip implant stress analysis and that no agreement has yet

emerged as to which is the most valid for this purpose. In what

follows these methods are first very briefly described. Then the

results of representative studies which use these techniques are

discussed. Although these methods are broadly applicable to a great

number of problems dealing with total hip replacements, the topic of

most interest to this report is the stresses which exist at the metal/

bone cement interface.

Methods Used to Investigate Stresses in Hip Implants

The basic principles behind and, where applicable, the limitations

of the various methods used to determine stresses in hip implants are

given below:

Theoretical Methods

1. Finite Element Analysis: The differential equations which govern

the load-displacement behavior are numerically solved at a finite

number of points (called nodal points) in an irregularly shaped

object and the behavior at the other points in the object are

interpolated from this information. Stresses and strains are

derived from this data. Anisotropic elastic properties and

spatial change in elastic properties can be straightforwardly

included in this type of analysis.

11
2 . Beam Theory: In beam theory as applied to problems of concern to

surgical implants (Bartel, 1977), (Barberi et al., 1978), simplifying assumptions are made with respect to deformation (i.e.,

that plane sections remain plane) and the materials used (which

are assumed to be homogeneous, isotropic, and linearly elastic)

so that closed form solutions for stresses and bending moments in

bone, bone cement, and/or the metal stem can be determined along

the length of the prosthesis. Variable geometry and elastic

properties along the length are taken into account with the beam

theory approach. Joint loads which produce torsional stresses

have not yet been examined with this method.

Experimental Methods

1. Strain Gauge Techniques: Strain gauges which consist of several

loops of foil or wires whose electrical resistance is known as a

function of deformation are bonded to the object to be measured

at points of interest. Loads are applied. By monitoring the

change in electrical resistance of the strain gauges with a

Wheatstone bridge circuit, the normal strains at the surface of

the object are determined. Shear strains can be determined

indirectly if the gauges are placed in a "rosette" pattern

(Dieter, 1976). Stresses can be calculated for elastic behavior

assuming that the elastic properties are known.

2. Photoelastic Methods: A model with the geometry of the object of

interest is constructed from a transparent optically bi refri gent

plastic, (i.e., a material with a refractive index which depends

upon the plane of polarization of the light which is incident

upon it). The loads of interest are applied and an optical

12
interference pattern is developed when plane polarized monochromatic light is passed through the object and viewed through an

analyzer (Kolsky, 1963). The stress distribution is inferred

from this fringe pattern.

Problems Examined by Stress Analysis of Total Hip Replacements

Before considering in detail results which have been obtained

with the various stress analysis techniques, a brief overview of the

different classes of problems which have been investigated is given

below.

A number of studies have been concerned with the stresses in the

materials of the total hip implant as they are influenced by various

design parameters e.g., stem length, elastic modulus of the femoral

component and the bone cement, stem shape, presence or absence of a

collar resting upon the medial shoulder of the bone, and presence or

absence of the bone cement layer. Among the studies which fall into

this category are those of Svensson et al . (1977), Cook et al . (1980),

Scholten et al. (1978), Yettram and Wright (1980), Mizrahi et al

.

(1979), and Barberi et al . (1978). Most studies have mainly been

concerned with determination of the stresses in the metallic component

and/or the surrounding bone. Some studies, however, have been concerned

(at least in part) with the stress distribution in the acrylic bone

cement (Svensson et al., 1977), (Yettram and Wright, 1979), and

(Kennedy et al .

,

1979).

Another important class of problems which has been investigated

is that which involves attempts to provide biomechanical explanations

for different types of failures of total hip implants which have been

observed clinically. Topics which have been examined include the

13
desirability of valgus as opposed to varus orientation for the stem,

stress shielding in the calcar region leading to bone resorption and

stem loosening, increases in distal stem stresses due to inadequate

proximal medial support, fracture of the acrylic or interface due to

high stresses resulting in stem loosening, and the importance of the

various force components in causing failure. Some studies in this

category are those by Hampton et al . (1980), Dobbs and Chaplin (1981),

Markolf and Amstutz (1976), Jacob and Muggier (1978), and Jacob and

Muggier (1980).

Finally, it should be mentioned that some studies have been

concerned with comparing a theoretical approach (the finite element

method) with experimental results (Svensson et al .

,

1980), (Rohlmann

et al .

,

1980). Some work has also been done with regard to stress

analysis of intact bones subjected to joint loads (Toridis, 1969),

(Rybicki et al .

,

1972), (Valliapan et al., 1977) and (Piotrowski,

1975).

Results of Experimental Measurements of Total Hip Replacements

Some measurements have been made of the gross response of total

hip replacements to mechanical forces. McCarthy and Wells (1977)

loaded prostheses (which had been implanted in dogs) in the axial

direction and measured the maximum load. They divided this number by

the implant surface area to find the "shear strength" of the bone/bone

cement interface to be 1.49 ± 1.27 MPa. Markolf et al . (1980) measured

displacements of the collar and tip of femoral components implanted in

cadavers due to a 2000 Newton femoral head force.

Studies which have involved more refined measurements have been

concerned with determining the stress distribution in the components

of the total hip replacement. Kennedy et al . (1979) used photoelastic

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techniques to study the stress distribution in bone cement, a factor

which they considered important in regard to loosening of the femoral

component in its bed of cement and for understanding nonuniform load

transfer to the femur which could result in resorption. The models

were constructed from actual hip prostheses, bone material or aluminum

to simulate the femur, and an epoxy with elastic properties similar

to bone cement. They found that a porous coated stem was superior to

a smooth stem with respect to promoting a more uniform stress distribution and that porous coatings prevented loosening of the stem in the

cement.

Mizrahi et al . (1979) used strain gauges to measure the stresses

at the surface of Straight Narrow Stem and Standard Charnley prostheses. They compared their results to calculations using curved beam

theory and found good agreement. The main factor they investigated

was the stresses which result from applied forces which correspond to

different anatomical positions of flexion-extension and abductionadduction. They found that both normal stresses and shear stresses

were maximum at a position corresponding to 20° flexion and 10° abduction and are comparable to the fatigue limit of the material. The

shearing stresses which correspond to twisting or torsion of the

prosthesis in the cement are highest near the proximal end and increase

with the degree of flexion and abduction,

Jacob and Muggier have written two papers (1978) and (1980) in

which strain gauge measurements were made on hip devices implanted in

epoxy models to simulate bone. The aim of this work was to clarify

the causes of stem loosening in the proximal end of the femur. They

found out that a prosthetic hip replacement could greatly diminish the

stresses in cortical bone at the proximal end of the femur such that

15
in some cases the stress level is reduced to 40 percent of that present

in the physiological state. Also, it was shown that a stem which had

loosened in the bone was a very adverse situation which creates high

bending stresses in the prosthesis and could cause stem breakage.

Although the presence of cancellous bone between the cortical bone and

bone cement could exercise a favorable load distribution function, the

possiblity exists that the stresses can be sufficiently high to cause

fracture of trabeculae. Direct bearing of a stem neck collar on the

underlying cortical bone of the calcar did not produce significant

alteration of stresses within either the stem or the bone.

Two studies, (Dobbs and Chaplin, 1981) and (Markolf and Amstutz,

1976), are directly relevant to the main subject of this report.

Markolf and Amstutz used strain gauges to measure stresses at the

surface of metal prostheses potted into 6 cm diameter cylindrical

molds of PMMA. Both cantilever loading, vertical loading, and loading

in which the downward force vector was inclined form the vertical at

angles of -3, +17, and +37° were studied. They found that varus positioning of the stem should be avoided at surgery because it produced

marked increases in stem stress. Varus orientation is when the ball of

the prosthesis at the proximal end is tipped medially and the distal

end of the stem is tipped laterally. They also showed that lack of

adequate support in the proximal medial region, due to acrylic fracture

or poor bone quality creates dangerous stress levels in the stem.

Finally, their finding which is most directly pertinent to this study

was that a prosthesis which is loose in the bone cement causes an

extremely severe rise in stem stress level. At the distal end of the

stem, the stress levels are more than one order of magnitude greater

for stems in which only the distal 1/3 of the stem is fixed in acrylic

16
as compared to a stem in which the total length is rigidly fixed.

They concluded from this that it is very important during hip surgery

to avoid movement of the prosthesis within the acrylic during the time

that the bone cement is setting. It presumably could also be concluded

from this data that materials treatment techniques which would promote

higher strength bonds between the metal and bone cement would be

desi rable.

Dobbs and Chaplin (1981) tested femoral components embedded in

acrylic which was contained in a thin walled cylindrical fixture.

Strain gauges were used to measure the strains in the acrylic, stem,

and fixture. It was demonstrated that the technique gave results

similar to those observed for prostheses embedded in actual femurs.

They studied loads acting in the medial -lateral plane at 0° , 10°, and

20° from the vertical. Here also, high stresses were observed for

varus positioning. Most importantly they observed high stresses in

the acrylic which they felt were high enough to cause acrylic fatigue

fracture and interface failure. They recommend that future designs of

prostheses should attempt to reduce bone cement stresses and increase

interfacial strength.

Results of Beam Theory Applied to Total Hip Replacements

Beam theory seems to have been more applied to calculating stress

distributions in the natural femur itself than to analysis of total

hip replacements. Although tangential to the emphasis of this report

a few of these studies will be mentioned briefly. The first of these

was the work of Koch (1917) who used beam theory to verify Wolff's

ideas (1892) of bone architecture. Toridis (1969) introduced the

mathematical refinement of using the theory of curved beams to analyze

these problems. Finally, Rybicki, Simonen, and Weis (1972) compared

17
beam theory to the finite element method. Their findings indicated

that beam theory was appropriate for the shaft of the femur, but that

the more complicated continuum elasticity theory was required in the

areas near muscle attachments and geometrical irregularity such as the

femur head and greater trochanter. This would be expected due to

Saint Venant's principle (Papov, 1976).

Two investigations have used beam theory for hip implant problems.

Bartel (1977) calculated stresses in bone, bone cement, and the prosthesis as they are affected by diameter of the prosthesis stem and

medullary cavity. Calculations of stresses along the length for both

bone and the prosthesis are presented. Barberi et al . (1978) used

beam theory applied to the actual geometry based upon Legendre polynomials, measured an interfacial stiffness experimentally, and computed

the stress state which minimized the strain energy. Torsional loads

were neglected. They computed the distributed compressive force and

flexural moment in the bone and the maximum flexural stress in the

stem for nine combinations of stem stiffness and length. They conclude

that longer stems do not necessarily lower all the stresses in the

cement and that, due to the nature of the flexible support of the

femur, stiffer stems do not have lower stresses.

Apparently beam theory studies, while of some use in understanding stresses in implants, have been superceded by the more technologically sophisticated methods of strain gauge instrumentation and

finite element method computer analysis. It should be pointed out,

however, that in at least one investigation, beam theory calculations

corresponded more closely to the results of rigorous three dimensional

finite element analysis than did a simplified two dimensional finite

element analysis (Valliappan et al., 1977).

18
Approaches Used for Finite Element Analysis of Total Hip Replacements

A number of different investigators have applied the finite

element method to various problems of interest with regard to total

hip replacements. With some exceptions, e.g. Yettram and Wright (1979)

and Svensson, Valliappan, and Wood (1977), they have not been concerned

with the stresses at the interface.

Table III-l summarizes the modelling approaches used by some of

these investigations. Specific details of the results are described

in the next sections. Some aspects of these studies are quite similar,

but there are also significant differences. For example there is no

consensus as to the best way to model the boundary condition between

the femoral stem and the bone cement. Yettram and Wright (1979) use a

displacement continuity condition. This is equivalent to perfect

adhesion at the interface with no slip between the two components.

Svensson et al . (1977) employ a no-slip condition on the medial and

lateral surfaces and a no-adhesion condition on the anterior and

posterior surfaces. Also, in this regard it should be pointed out

that another study of the biomechanical behavior of the hip implant

has considered the boundary condition to be that of a perfect grouting

material, i.e., frictional forces at the interface (Berme and Paul,

1979). Some investigators have used two dimensional finite element

models and others employ three dimensional models. In principle, the

three dimensional approach is more mathematically rigorous (and much

more complex to use) and should give a better approximation to reality.

There also does not seem to be consensus as to what joint loads should

be used for finite element analysis. A vertical load or the forces

present in a one-legged stance seem to be the most common choices.

19
TABLE III-l

VARIOUS APPROACHES EMPLOYED FOR FINITE ELEMENT ANALYSIS OF

TOTAL HIP REPLACEMENTS

A. Type of Model

Investigators

Yettram and Wright (1980)

Yettram and Wright (1979)

Cook, Klawitter, and Weinstein

(1980)

Scholten, Rohrle, and Soil bach

(1978)

Hampton, Andriacchi, and Gal ante

(1980)

Svensson, Valliappan, and Wood”

(1977)

Svensson, Valliappan, and McMahon

(1980)

B. Boundary Conditions

Investigators

Yettram and Wright (1980)

Yettram and Wright (1979)

Cook, Klawitter, and Weinstein

(1980)

Model

2 dimensional, plane stress, 751 isoparametric quadrilateral elements,

834 nodes, orthotropic elastic properties used for bone

2 dimensional, plane stress, 500 con- stant strain triangular and quadrilateral elements, 524 nodes, isotropic

elastic properties used for bone

3 dimensional, tetrahedron volume

elements with square displacement

function at the edges, isotropic

elastic properties for bone with

spatial nonuniformity in properties

taken into account

3 dimensional, 31 isoparametric

hexahedron elements, 200 nodes,

isotropic elastic properties for bone

2 dimensional, 118 quadrilateral

elements, isotropic elastic properties for bone

3 dimensional, 167 isoparametric

elements with 20 nodes each

Interface Assumptions

displacement continuity, i.e. no slip,

at the metal/bone cement and the

bone/bone cement interfaces

not explicitly given; elements between

the metal prosthesis and bone itself were given different values of elastic

properties to simulate different

fixation mechanisms such as bone cement

and porous coatings

Same as above

not explicitly given

20
Table III-l B. (continued)

Scholten, Rdhrle, and Sollbach

(1978)

this was a variable investigated in

the study; both displacement con- tinuity and tangential displacement

of prosthesis relative to bone were

examined

Hampton, Andriacchi, and Gal ante

(1980

displacement continuity at both

i nterfaces

Svensson, Valliappan, and Wood

(1977)

displacement continuity at the medial

and lateral bone cement/metal interfaces;

non-adhesive, i.e. slip, conditions for

the anterior and posterior interfaces

Svennson, Valliappan, and McMahon

(1980)

Goodman joint elements which correspond

to slip and no normal stress transmission

through the element for tensile contact

stresses

C. Type of Loading

Investigators Joint Forces

Yettram and Wright (1980) 10^ lb vertical load

Yettram and Wright (1979) same as above

Cook, Klawitter, and Weinstein

(1980)

1250 N acting at 20.5° to vertical

at the femur head, 778 N acting

upward at 22° to vertical at greater

trochanter (abductor muscle traction);

i.e. the one-legged stance

Scholten, Rdhrle, and Sollbach

(1978)

the one legged stance of Pauwels (1965);

femur axis inclined 9° to vertical, joint

reaction force 175 kgf at 16° to vertical,

130 kgf abductor muscle traction at 21° to

vertical

Hampton, Andriacchi, and Gal ante

(1980)

stresses normalized to a 1 N load in

each of the 3 coordinate directions

(vertical, anterior-posterior, and

medial-lateral) were calculated

Svensson, Valliappan, and Wood

(1977)

the one legged stance of McLeish and

Charnley (1970); 1620 N downward at

24° to vertical as the joint reaction

force, abductor traction of 1062 N

upward at 29° to vertical, a second

muscle traction of 265 N vertically

downward at the trochanter

Svensson, Valliappan, and McMahon

(1980)

3 kN downward vertical load acting

at 7° to axis of prosthesis

21
Muscle tractions are sometimes neglected, and only static situations

are considered.

Results of Finite Element Method Studies of Total Hip Implants

A number of studies have been conducted with the finite element

method which focus on various aspects of implant design. Cook et al

.

(1980) studied the factors which promote a satisfactory stress distribution in the calcar region. Noting the current trend toward larger

stem sections to prevent fatigue failures of the stem, they point out

the possibility of stress shielding and therefore bone resorption at

the calcar region. The resorption can cause implant loosening proximally and therefore higher stresses distally which still could lead to

fatigue failure. As would be expected for a material of lower stiffness which would be less prone to stress shielding, their results

indicated that hip implants made of carbon would produce a higher,

more physiologically normal stress pattern in the calcar region. It

also was shown that shortening the length of the stem did produce

higher average stresses but the loading pattern was not physiological.

Scholten et al . (1978) studied stresses in the surrounding bone

for eight different fixation combinations involving support or nonsupport at the prosthesis shoulder area (i.e., with or without a

collar); direct stem to bone contact or stem to bone contact through

an intervening layer of bone cement; and sliding or displacement

continuity boundary conditions. Most of the calculations involved

stress in the bone although some study of the stresses in the prosthesis itself was carried out. A prosthesis of conventional design was

compared to a Ritter-Grunert prosthesis which has a tension bar through

the trochanteric region which is connected to a fixation plate. They

22
found high compressive radial stresses, high tensile tangential

stresses, and axial compressive stresses in the bone on the medial

side for sliding boundary conditions (as would be expected from a

prosthesis acting as a wedge). They also calculated the stress distributions in the natural bone and found that the Ritter-Grunert prosthesis produced a stress distribution much closer to the physiological

state than did the conventional prosthesis.

Additional studies have been concerned with the stresses in the

metal stem component itself. Hampton et al. (1980) used a three

dimensional model and determined the longitudinal stress per unit load

in the stem of hip implants for loads directed, respectively, along

the X, y, and z coordinate axes (see Figure III-l). The stresses are

highest in the region of the implant about 1/3 to 1/2 of . the way along

the implant from the distal end. Their conclusion was that the vertical load was responsible for most clinical failures. However, the

lateral and posterior loads were significant enough to possibly

endanger survival of the device. The posterior load (which can be

high in activities such as stair climbing, rising from a chair, or in

certain phases of the gait cycle, see Figure IV-1) when combined with

the vertical load produces significant stress in the anteriolateral

corner of the stem, a site at which cracks often originate in failed

devices. This failure site has also been described by Harris and Tarr

(1979).

The hypothesis advanced in the next chapter is that the posterior

force is important to understand the failure of the metal /bone cement

interface. Unfortunately no information on this subject is available

from Hampton's study because it was optimized to determine stresses in

the metal rather than in the acrylic.

23
Yettram and Wright (1980) used a two dimensional finite element

model (which was found suitable for parametric studies) to calculate

the tensile stresses along the lateral surface of the stem of the

femoral component of the total hip replacement. The variables examined

were stem taper, stiffness of the cement, stiffness of the prosthesis,

and the presence or absence of a shoulder on the prosthesis. It was

found that cement stiffness and implant shape had a minor effect. The

Young's modulus of the stem material was found to be the most important

factor. Decreasing stress levels were observed in less stiff implants

(as would be expected from a simple consideration of stress shielding

effects).

Yettram and Wright (1979) in another study reported their findings

with regard to the stresses in the bone cement. As the stem material

becomes more flexible, the stress distribution in the cement becomes

more uniform and the stresses tend to increase. The normal stresses

in the cement in the proximal portion of the stem both laterally and

medially is compressive. This study contained calculations of the

stresses at the metal /bone cement interface and is discussed in more

detail in the next section.

Svensson et al. (1977, 1980) have compared the results of finite

element analyses with strain gauge measurements. The conclusion from

the first paper was that agreement with the experimental results

required a slip boundary condition between the prosthesis and cement.

The second paper concluded that the three most critical stress values

in the system are the tensile stress on the lateral side of the prosthesis, the tensile stress across the bone/acrylic interface, and the

shear stress across the bone/acrylic interface. The authors concluded

24
that a meaningful extension of their work would be studies of the

effect of different metal/acrylic interface conditions. This work,

Svensson et al. (1980) is discussed in more detail in the next section

in which calculations of interfacial stresses are considered. In

contrast to other studies described above, this work determined that a

heavier prosthesis would be beneficial and that no significant improvements can be obtained by varying the mechanical properties of the bone

cement or metal prosthesis.

Calculations of Stresses at the Metal/Bone Cement Interface

From the viewpoint of stress analysis the term interfacial stress

is an imprecise one. What actually can be calculated is the stress in

the metal at the interface and the stress in the acrylic at the interface. Both of these factors are important in regard to the structural

integrity of the total hip replacement. The stress in the metal will

determine the number of load cycles possible before fatigue failure of

the stem will occur. High levels of stress in the acrylic at the

interface will lead to failure of the interface due to fatigue,

overload, or as has recently been suggested, by creep of the acrylic

(Gibbons and Buran, 1980).

Two studies which have been concerned with use of the finite

element method to calculate the stresses in the bone cement at the

interface are those of Svensson et al . (1977) and Yettram and Wright

(1979). The results of these calculations for the variation along the

at the interface is shown in Figure III-2. The direction of the

normal stress is perpendicular to the implant surface. The stress

distributions are qualitatively quite similar even though differences

25
I 0 -I 0-2

T cr,

SVENSSON

FIG. m-2

YETTRAM AND WRIGHT

1062 N 1620 N

et al ( "no slip", —"no tension") one legged stance

Normal Stress CTp and Shear Stress T in Bone

Cement at the Cement/Stem Interface

MPA

26
exist between the two studies with respect to prosthesis shape, assumed

boundary conditions, and the forces applied by the body weight and

muscle tractions.

If the assumption is made that the interface will fail where the

acrylic is subjected to the highest shear stresses or the highest

tensile stresses, these studies would indicate that interface failure

would more likely occur (a) along the medial surface than the lateral

and (b) at the distal end rather than the proximal.

Interestingly, as described in Chapter IV, the clinical observations are exactly the opposite. A possible explanation for the discrepancy is that the finite element analyses do not account for the

anterior-posterior transverse forces in the joint (see Chapter IV).

These transverse forces produce torsional loading of the interface

which probably should also be considered in calculations of the stress

state of the acrylic at the interface.

27
Chapter IV

CONCEPTS FOR IMPROVEMENT OF STRESS ANALYSES OF

TOTAL HIP IMPLANTS

This chapter briefly reviews the literature pertaining to failures

of hip implants which are observed by clinicians. Although there are

a great many causes of failures of total hip replacements, one important concern is the integrity of the metal /bone cement interface. As

discussed by Markolf and Amstutz (1976) lack of fixation of the stem

and the cement at the interface can cause a stress state in the stem

which could ultimately lead to fracture of the device. The lack of

correspondence between the analyses of stress at the interface given

in Chapter III and the actual positions of prevalent interface failure

is examined. The probable cause for the discrepancy is seen to be the

neglect of anterior-posterior forces in the analysis. This would

indicate that the behavior of metal /bone cement interfaces under

conditions of torsional loading is an area which should be further

investigated in the future.

Clinical Observations of Failure Modes of Hip Implants

A number of studies have described observations of fractures of

femoral stem components (Carlsson et al . , 1977), (Galante, 1980),

(Collis, 1977). However, the majority of total hip replacement procedures which are performed are successful and, of those which are not,

only a portion of them are failures due to biomaterials or biomechanics

related causes. Nonetheless failures of the latter type are a significant problem, and it is important to understand why they occur.

Statistics in regard to failure frequencies of implants are far

from complete, but a recent study by Nolan et al . (1975) puts the

28
matter in perspective. In a study of 3,204 total hip replacements,

3.9 percent required reoperation after 2 to 5 years. The complications

which necessitated the revision were (from most to least frequent):

infection, dislocations, trochanteric problems, ectopic bone, and

loosening of the femoral prosthesis.

Loosening of the femoral stem component seems to be often associated with biomaterials or biomechanics related hip implant failures.

Other factors also often implicated are lack of support at the calcar

region and varus orientation (Galante, 1980). Loosening can occur at

either the metal /bone cement interface or the bone/bone cement interface, although the latter problem has received more attention in the

literature and can cause considerable pain to patients.

In a recent study by Gruen and Amstutz (1977) in which failure

modes of total hip replacements were investigated, the incidence of

loosening (as determined by the presence of radiolucent lines in

radiographs) of the two interfaces was about equal: 10.3 percent at

the metal /bone cement interface and 11.1 percent at the bone/bone

cement interface. Failure as defined in their study represented a

failure in the mechanical integrity of the total hip replacement and

not clinical failure which is characterized by pain and restrictions

in function, motion and muscle power.

In the work of Gruen and Amstutz it was shown that loosening at

the metal /bone cement interface does not occur with equal probability

at all points on the surface of the implant. One hundred percent of

the failures occurred at the proximal end and greater than 90 percent

of these occurred at the proximal lateral surface. As was previously

pointed out, these clinical observations are in sharp contrast to the

29
predictions of the finite element studies described in the last section

of Chapter III. Those investigations considered the total hip replacement to be loaded only by forces whose lines of action lie in the

medial-lateral plane. Transverse forces acting in an anteriorposterior direction (which would create torsion) were neglected.

Actually, as is considered in the next section, these forces can have

significant magnitudes and improvements in theoretical and experimental

stress analyses could certainly be made if they were not neglected.

Joint Loads

The hip is a very highly loaded joint. Estimates of the peak

force which is present in the hip over the course of the walking cycle

range up to six times body weight (Williams and Svensson, 1968).

Measurements with instrumented femoral prostheses (Rydell, 1965),

(English and Kilvington, 1979) have shown 2.5 times body weight to be

present.

Of particular interest to the subject of this report are the

magnitudes of the vector components of this force. This topic has

been considered by Berme and Paul (1979). Their results show the

components of force to vary over the walking cycle as shown in Figure IV-1. The X, y, and z axis system used is the same as that which

appears in Figure III-l.

The stress analyses of Chapter III assume that only force components Fy and F^ are present, yet, as shown in Figure IV-1, the peak F^

force is about as large as the others. This would mean that a significant twisting moment would be produced about axis 66 (which is taken

to be a principal axis of inertia of the stem in the y z plane, see

Figure III-l) due to application of an anterior-posterior force at the

30
31

FIG.

Bf-1

Variation

of

Joint

Force

Components

with

Time

During

the

Walking

Cycle
ball of the femoral component. This would mean that the bone cement/

metal interface experiences a significant torsional load during a

considerable portion of the walking cycle.

It appears that insufficient attention has been paid to the

biomechanical significance of this transverse force with respect to

its effect on the hip joint either in regard to trauma of the natural

joint or the behavior of artificial implants. The biomechanics literature is apparently quite sparse in this respect. The search which was

pursued here found only one reference to torsional loading; Mizrahi et

al . (1979) showed that the shearing stresses in the metal implant

corresponding to torsional loading are greatest at the proximal end of

prosthesis (i.e., in the area most often observed clinically to be

prone to interface failure). Also, they showed that these stresses

increase with the degree of flexion and abduction (i.e., at anatomical

positions different from the upright stance and for which the anteriorposterior force would be larger).

The importance of considering torsion for better understanding of

the hip joint (whether natural or totally replaced) can be inferred

from a few clinical studies as well. Wroblewski (1980) has suggested

that several types of hip injuries and diseases are related to transverse loading of the joint. In another study of fractured femoral

stems (Wroblewski, 1979), the findings indicated that the fractures

were due to the combination of torsion and bending stresses. Carlsson

et al . (1977) in examinations of a small number of fractured femoral

stems also reports evidence of torsional deformation.

The information considered above implies that torsional loading

of metal/bone cement interfaces could exert a significant influence on

32
their failure. Since little, if any, research has been done on the

torsional characteristics of the metal/bone cement interface, experimental studies in this area were initiated. These are described in

Chapter V.

33
Chapter V

CHARACTERIZATION AND TESTING OF METAL/BONE CEMENT INTERFACES

Initially it was thought that ASTM standard tests could furnish

guidance as to appropriate testing methods for determining the strength

of metal/bone cement interfaces. As is discussed below, this turned

out not to be the case. The literature pertaining to mechanical

testing of bone cement and bone cement/metal interfaces was then

reviewed. This research is described in this chapter. Since there

have been no previous studies of the behavior of the interface in

response to torsional loading (a failure mode which, as previously

discussed, could be important for lack of success of total hip replacements in some cases), experimental pilot studies to devise a torsion

test for the interface were initiated. The results are described here

in Chapter V.

Interfacial Bond Strength Tests

A number of different testing methods to determine the strengths

of metal to adhesive bonds have been sufficiently well characterized

that they have become ASTM Standard Tests. Among these are ASTM

D1002-72 (1978) Strength Properties of Adhesives in Shear by Tension

Loading, D1 062-78 Cleavage Strength of Metal -to-Metal Adhesive Bonds,

D2293-69 Creep Properties of Adhesives in Shear by Compression Loading,

D3433-75 Fracture Strength in Cleavage of Adhesives in Bonded Joints,

D3658-78 Determining the Torque Strength of Ultraviolet Light-Cured

Glass/Metal Adhesive Joints, D2095-72 Tensile Strength of Adhesives by

Means of Bar and Rod Specimens, D2182 Strength Properties of Metal-toMetal Adhesives by Compression Loading, and E229-70 Shear Strength and

Shear Modulus of Structural Adhesives. The modes of loading involved

in some of these tests are shown in Figure V-1.

34
TENSION

CLEAVAGE

^

F

///////777.

SHEAR BY TENSILE LOADING

SHEAR BY COMPRESSION LOADING

SHEAR

FIG. V-1 Loading Modes Employed for ASTM Standard

Tests of Strength Properties of Adhesives

35
The applicability of tests suited for testing adhesives to evaluation of the bone cement/metal interface is uncertain. First of all

bone cement does not function like, or only like, an adhesive. It is

true that bone cement does bond to a metal surface via a mechanical

interlocking mechanism and (like an adhesive) does chemically bond to

the metal. The latter effect has been demonstrated by Keller et al.

(1980) and others who found that different materials with the same

surface finish or the same material given different chemical treatments

of the surface have different strengths. Bone cement also functions

as a grouting material, however. A grouting material is a space

filling substance (such as mortar used to fill chinks or cracks).

Thus unlike an adhesive which is a planar layer between two massive

materials, the bone cement has significant thickness itself. Although

a bone cement layer could be prepared to conform to the geometry of

the adhesive tests, it is quite possible that the polymerization

process could be altered such that the mechanical properties of the

cement (and therefore the strength of the interface) would not be

representative of those of bone cement when prepared in bulk. Another

limitation of using the ASTM tests is that there is no standard for

torsion testing of interfaces in which one component is enclosed by

the other, which is probably the most meaningful test with regard to

similarity to clinically observed failure.

Studies of the Mechanical Behavior of Bone Cement

and Bone Cement/Metal Interfaces

The polymethylmethacrylate bone cements which are used for the

total hip replacement have been extensively investigated biomaterials.

The literature which has been generated will be briefly reviewed here;

36
only the papers dealing with interfacial strengths will be examined in

more detail.

Studies by Kusy (1978), Haas et al . (1975), and Edwards and Thomasz

(1981) have used various chemical, physical, and microscopic techniques

to characterize and evaluate the chemical composition, degree of

polymerization, porosity, and other microscopic features of the cement.

The latter two investigations were also concerned with the curing

characteristics of the material. Other studies have also been concerned with the curing characteristics of the cement iji situ . Sih,

Connelly, and Berman (1980) studied the temperature- time history of in

vitro samples of setting bone cement. Ahmed et al. (1979) calculated

the thermal stresses produced as bone cement solidifies.

Other research has dealt with effects (both positive and negative)

on the mechanical properties of PMMA bone cement which results when

different components are added to the basic composition for various

purposes. Pi 1 liar et al . (1976) and Saha and Kraay (1979) examined

the influence of fiber reinforcing materials. Combs and Greenwald

(1979) studied the effects of barium sulfate. Bayne et al. (1975) and

Rijke et al . (1977) investigated the dependence of properties on

porosity in the bone cement. Lautenschlager , Marshall, et al . (1976);

Lautenschlager, Jacobs et al. (1976); Nelson et al . (1978), and Moran

et al . (1979) have determined the influence of antibiotic additions on

the mechanical behavior.

Many other studies of the mechanical behavior of bone cement have

been conducted. Holm (1977) measured flexural strength and Young's

modulus. Lee et al . (1977) performed bending, compression, and tensile

tests. Freitag and Cannon (1977) studied the fatigue behavior of bone

37
cement. A number of investigations have been pursued in which a

fracture mechanics approach has been used to study bone cement. These

include (Beaumont, 1977), (Beaumont, 1979), (Beaumont and Young,

1975a), (Beaumont and Young, 1975b), (Owen and Beaumont, 1979), (Sih

and Berman, 1980), (Robinson et al., 1981), and (Kusy, 1978).

Two studies have focused on the effect of the i_n vivo environment

on the mechanical behavior of bone cement. Rostoker et al . (1979)

tested samples of PMMA bone cement which had been implanted in rabbits

and found a statistically significant decrease in fracture stress

after 26 months. Gibbons and Buran (1980) examined bone cement samples

which had been retrieved from total hip replacements. Their findings

were plastic deformation resulting from creep, abrasive wear of the

cement, and cracks radiating from the corners of the medial edge.

Several investigations have been conducted which tested the

strength of the bone cement/metal interfacial bond. These will be

discussed in detail. Beaumont and Young (1977) performed push-through

tests on stainless steel cylinders embedded in bone cement. Failure

occurred either by shear fracture of the cement or disruption of the

interface. The load-displacement curves exhibited a linear portion up

to a maximum load P^. and then a slowly decreasing load with further

displacement in a saw-tooth pattern corresponding to stick-slip failure. The interface shear strength was defined as:

x -3 = P./ndL^ (1)

where d is the rod diameter and is the length embedded in acrylic.

Various types of surface conditions of the rod were examined,

for a 1 mm deep thread was 45 MPa (which according to the investigators corresponds to the shear strength of the cement), was

38
observed to be insensitive to surface finish below center-line averages

(determined with a Talysurf measuring device) of 8 pm. For these surfaces

x.g was seen to be 9.2 ± 3 MPa. The presence of water at the interface

had a pronounced adverse effect on t.^ and decreased it from 9.2 to 4.7 MPa.

In a study which used a somewhat similar approach, Greenwald and

Wilde (1974), measured to be 0.207 MPa for bone cement/high density

polyethylene interfaces, between 2.12 and 3.35 MPa for cancellous

bone/bone cement interfaces, and 1.35 MPa for stainless steel/bone

cement interfaces. The investigators speculated that knurling or

roughening the surface of the metal might enhance the bond strength.

There was quite a descrepancy between the investigations of

Beaumont and Young; and Greenwald and Wilde with regard to the shear

strength of the bone cement metal interface. Perhaps the difference

lies in the length of the samples used. The influence of sample

length on test results was considered in a more refined treatment of

the problem by Beaumont and Plumpton (1977). They performed pull out

tests on bone cement/metal interfaces and analyzed their results in a

more complete manner than previously using shear-lag theory [see

Greszczuk (1969) for a description of shear-lag analysis]. The experimental results indicated that the average shear stress x^^ (identical

to X. above) increases as L decreases.

The reason for this effect is that the stress at the interface in

a pull-out test can be shown by a more complete analysis to vary along

the length. The stress is not constant as implied by equation 1 but

rather is a function of distance x along the length away from the point

of application of the force. The functional dependence is given below

(Beaumont and Plumpton, 1977):

39
T (X) =

P .a ~

2^ (sinh ax - coth aL^ cosh ax) (2)

where

a

2Gi

1/2

(3)

In the above equation r is the radius of the rod, is its Young's

modulus, b^. is the interface thickness (which can be estimated from

the surface roughness), and G. is the interface shear modulus, z ,

the maximum interfacial stress, occurs at x = 0 and is given by:

X max = , 4.U , at coth aL

I e e av

As L ^ 0, X ^ X . Thus, measuring x vs L and extrapolating to

e ’ av max

’ ^ av e

k »

L = 0 determines x . For any L , x /x can then be determined

e max e ’ max av from a pull out test. With this information a can be obtained, from

which G. can be extracted.

1 For dry interfaces and interfaces in contact with blood, interface

weakening was also observed in their study due to the fluid contact.

X was found to be 18 and 10 MPa respectively. G. was observed to max 1 3 2

be 10 and 10 MPa respectively. Beaumont and Plumpton cite the

finite element studies of McNeice and Amstutz (1975) which find

stresses at the interface of 16 ± 2 MPa and concluded that interface

failure could occur early in the life of hip prostheses.

Shear-lag analysis seems to offer promise to assist in the data

reduction for the experimental torsion studies of the type for which

40
preliminary results are described in the next section. It may also

offer insight into the torsional stresses present in hip implants due

to the action of the anterior-posterior force which is present during

the walking cycle. These possibilities should be futher investigated.

Another study of the bond strength between bone cement and implant

material has been conducted by Raab et al . (1981). They measured

interfacial shear strengths by push through tests for bone cement/metal

interfaces and by shear through compression loading tests for UHMWPE/

bone cement interfaces. The metals investigated were Co-Cr-Mo,

Ti-6A1-4V, and 316 LVM. Tests were performed at a strain rate of 1.3 x -3 -1

10 s since it was observed that interface strength and elastic

modulus increased with strain rate above a plateau region which ended

at about 10 ^ s Tests were performed in both Ringers solution at

37 °C and in a dry environment. Interface stresses were calculated

using finite element techniques. For pullout tests their calculations

indicated that the stress is a function of the length, a result similar

to the work of Beaumont and Plumpton cited above. Along a radial

direction from the cylinder the stresses are highest at the metal

surface and decline with increasing distance. A fracture mechanics

approach was used for tests with other specimens to find the crack

extension force. In still other studies Raab et al . measured the

fatigue behavior of interfaces. S-N curves which appear to exhibit

endurance limit type behavior were found.

Most interesting from the standpoint of this report were their

studies which measured the interface strengths for the different

metal/bone cement surfaces after various times of immersion in saline

solution. The results are shown in Table V-1. The different metals

41
TABLE V-1

SHEAR STRENGTHS OF BONE CEMENT/MATERIAL INTERFACES FOR VARIOUS

METALS AND UHMWPE^, T.g (MPa)

Interface Dry 30 days

(sal 1 ne)

60 days

(sal i ne)

Dry

(conditioned)

Co-Cr-Mo/B.C.^ 6.9 ± 0.7 5.3 ± 0.8 5.3 ± 1.0 6.0 ± 0.6

T1-6A1-4V/B.C. 12.5 ± 1.8 6.3 ± 0.6 6.7 ± 1.3 12.2 ± 1.6

316L VM/B.C. 11.2 ± 1.9 6.2 ± 0.4 6.4 ± 1.1 8.2 ± 1.4

UHMWPE/B.C. 2.3 ± 0.4 0.5 ± 0.3 0.4 ± 0.2 1.6 ± 0.2

^ after Raab et al (1981)

^ B.C. - bone cement

42
showed a considerable range of interfacial shear strengths when tested

dry. Co-Cr-Mo/bone cement interfaces were less than 60 percent as

strong as Ti-6Al-4V/bone cement interfaces. The surface finishes on

these materials were 3.0 pm or better. This value lies within the

range where

^ was observed by Beaumont and Young (1977) to not

depend upon surface finish. The metals were prepared according to

ASTM standard F86, Standard Recommended Practice for Surface Preparation

and Marking of Metallic Surgical Implants. They were grit blasted and

passivated in HNO^ 24 h before testing and autoclaved 2 h before testing.

Conditioning surfaces by exposure to laboratory air for two weeks prior

to testing diminished interfacial strength in 316L and UHMWPE. In all

cases the metal/bone cement bonds were much stronger that the UHMWPE/bone

cement bonds. Differences in x^. ^ between metals diminished with time of

saline exposure as 31 6L and Ti -6-4/bone cement bonds deteriorated in

strength. The authors concluded that the strength deterioration is an

interfacial effect rather than a bulk effect in the bone cement (since

other tests on bone cement itself did not show degradation of its

properties with comparable saline exposures). They attributed the

degradation of surfaces due to conditioning by exposure to laboratory

air to formation of surface oxide.

It can be presumed that chemical interaction or bonding between

the two materials of the interface could occur since different materials with similar surface finishes have different x.^. Similar findings

were made by Keller et al . (1980).

They also studied the variables which affect the strength properties of bone cement/metal interfaces. The method used in their investigation was to measure the interfacial bond strength in tension a^.

.

This test has the disadvantage that the failure which is produced

43
occurs both at the interface and also within the acrylic itself so

that some parts of the surface are covered with cement. Keller et al

.

studied the influence on a. of interface formation time (i.f.t.),

i.e., the time after the onset of mixing that contact was made between

the metal and bone cement, and curing time (the time elapsed between

sample preparation and mechanical testing). The materials they investigated were 316L, Ti-6A1-4V, and Co-Cr-Mo. In one series of tests

the samples used were polished to 0.1 pm and subsequently electropolished and prepassivated. In another series the samples had a 15 pm

fi nish.

The results of this investigation showed that differences in a^.

resulted for different cure times and interface formation times. The

various metals behaved differently, however, so that no general trends

can be observed in the data, except that late i.f.t. 's lead to inferior

properties. The results indicated that, in contrast to the observation

of Raab et al. (see Table V-1), 316L/bone cement interfaces were

stronger than the other interfaces (6.5-11.7MPa for 316L depending

upon cure time and i.f.t. as compared to 5.1-9.7MPa for Co-Cr-Mo and

1.9 to 8.4 MPa for Ti-6A1-4V). This effect was more pronounced at 1 h

than after 1 week. Thus the work of Keller et al . also indicates more

similar behavior between metals as time goes on. Somewhat surprisingly,

the rougher surfaces had lower values of a^. than the prepassivated

surfaces. This could mean that roughness is not a factor below 15 pm

but that passivation treatments promote chemical interactions and

bonding between the metal and acrylic. Also, as suggested by these

investigators, roughness might lead to air bubble entrapment at the

surface.

44
For 316L, tests in another series of samples were run in 37 °C

Ringer's solution. Like Raab et al . , reduction in was noted in this

study as well. The effect was most pronounced for late i.f.t.'s when

the doughy mass of bone cement has high viscosity and would be expected

to have considerable porosity at the interface leading to water contact

with the interface by sorption processes.

Experimental Studies

Since the literature search previously described led to the

conclusion that the response of the bone cement/metal interface to

torsional loading appears to be a question which should be investigated to obtain further insight into failures of total hip replacements, a series of experiments was pursued in this area. The specimen

used for the mechanical tests is shown at the top of Figure V-2. A

cylindrical metal rod of 0.635 cm diameter was embedded in a rectangular prism of bone cement such that their axes were coincident. Torque

was applied to the free end of the rod which tended to twist the metal

surface relative to the acrylic surface at the interface. This type of

sample geometry was selected over that used in ASTM Standard E229-70

Shear Strength and Shear Modulus of Structural Adhesives (see the bottom

of Figure V-1). It was thought that because the metal is surrounded by

the acrylic this sample represents a closer approximation to the geometry

of the total hip replacement.

Ten samples of this type were prepared. The manufacturer's

instructions to the surgeon were followed in the preparation of the

bone cement (Howmedica Surgical Simplex P). The powder (consisting of

PMMA and methyl methacrylate-styrene copolymer with or without barium

sulfate added for radi opacity) and methyl methacrylate monomer were

mixed in the recommended 2 g to 1 ml ratio. The contents were stirred

45
T/L

(N-m/m)

FIG. !2-2 Length -normalized Torque T/L^q vs. Twist

Angle X Interfaces 7, 8 and 9.

46
for two minutes and the doughy mass was kneaded approximately two

minutes longer, after which the bone cement was brought into contact

with the metal rod. Thus the interface formation times ranged from

4 to 4-1/2 minutes.

The metal samples were 31 6L stainless steel polished to a 15 pm

finish. Before the interface was formed the rods were washed in a

detergent solution, rinsed, and ul trasonical ly cleaned in ethanol. The

interface was formed through use of a Teflon mold. The metal sample

was aligned parallel to the axis of the mold. The assembly was equipped with a Teflon plunger to compress the bone cement as it was hardening to promote better contact with the metal and reduce porosity, thus

producing higher quality interface formation. Some interfaces were

formed using a small weight for pressure on the plunger, as was done

by Keller et al . (1980) to simulate a currently used technique in

total hip arthroplasty. Other interfaces were formed by applying as

much pressure as possible by hand to the plunger for 5-1/2 minutes as

the cement hardened.

It was observed that the interfaces made with the low pressure

were poorly formed and fragile compared to those made with the high

pressure. The decision was made to form the interfaces which were to

be mechanically tested by the higher pressure technique. It was felt

that, although the higher pressure method less closely simulated the

clinical situation, this would not be a significant disadvantage

because so many other variables also affect interface formation i

n

vivo in the operating theater. The higher pressure technique seemed

to offer the advantage of producing a more reproducible, higher quality

interface with a greater true surface area of contact between the

47
metal and bone cement. Reproducibility is important for studies where

different surface treatments of the metal (e.g., polished, passivated,

roughened, scribed, etc.) are examined with respect to their influence

on the strength of the metal /bone cement interface.

Interfaces which were formed with the high pressure method were

tested in torsion. The testing apparatus used was a modification of a

torsion machine which has been used in corrosion fatigue studies as

described by Imam et al . (1979). In its modified form the machine

comprised a dead weight loading system. A cable, pulley, loading

platform, and weights were used to apply a torque to a shaft which

could be calculated from the weight W and geometrical configuration.

At the other end of the shaft was a grip attached to the free end of

the 31 6L rod. The end of the rod embedded in acrylic was held in a

grip which fixed the acrylic block rigidly in position.

The twist angle y was measured by determining the change in

position of a laser beam which was reflected from a mirror whose axis

was coincident with that of the shaft and which was attached at the

end where the torque was applied. The twist angle is produced predominantly by distortion of the acrylic block as torque is applied to the

rod embedded within it or by relative movement of the two components

at the interface. The metal rod, grips, and shaft which applies the

torque are presumably virtually rigid in comparison.

Preliminary experiments were performed in which the W vs y

response was measured for samples in which the length of the rod

embedded in the acrylic L ranged from about 1.25 cm to 5 cm. These

dC

experiments demonstrated that there was a linear relationship between

W and y which was reproducible upon loading, unloading, and subsequent

48
reloading. They also indicated that smaller values of L would have

to be used in order to stress the interface beyond the region of

linear elastic response without exceeding the maximum loading capacity

of the dead weight system.

Results from three tests in which torque T vs twist angle y was

determined for samples in which L was under 0.635 cm are shown in

Figure V-2. The torque is normalized by dividing by L to afford a

fairer comparison of the different tests. Several conclusions may be

drawn from this data. The curves obtained experimentally validated

the concept of using a sample of the geometry shown at the top of

Figure V-2. That is, "failure" of the interface itself was observed

rather than fracture of the acrylic which can occur sometimes with

other types of geometries. "Yielding" at the interface seems to occur

when the T vs y curve departs from linearity. The twist which results

when the T vs y curve becomes nonlinear is permanent deformation.

This could be easily seen at high twist angles when the sample had

been visibly rotated in the acrylic.

Quantitatively the elastic portion of the T vs y curve was fairly

reproducible for the three tests. Table V-2 gives the values observed

for the proportionality constant G" between T and y for the linear

portion of the curve governed by the equation T = G"y. As yet, an

analytical treatment of how to calculate G" from the sample geometry

and the shear moduli of the metal and bone cement has not been carried

out. However, undoubtedly the effective stiffness is increased as the

length of metal embedded in acrylic L increases. As shown in the

dC

table, when G" is normalized to L the standard deviation was within

3C

20 percent of the mean.

49
TABLE V-2

EXPERIMENTAL VALUES OF PROPORTIONALITY CONSTANT G"

BETWEEN TORQUE AND TWIST ANGLE

Interface G"

(N-m/degree)

Lac (cm)

G"/L

ac (N/degree)

7 0.691 0.47 147.0

8 1.134 0.56 202.5

9 0.837 0.58 144.3

Mean 164.6 ± 32.9

50
The value of torque required to produce yielding at the interface

Ty, could easily be quantitatively determined from the T vs y curve.

Although this point does not correspond to complete fracture of the

interface in which the two materials are separated (since the torque

does not go to zero) it probably represents a stress level beyond

which the interface should not be stressed in order to avoid subsequent

damage and can be taken to represent a practical criterion for failure

of the interface.

Time dependent effects were observed to be associated with

yielding. When a weight increment was applied in the region beyond the

linear region, it would take several seconds for significant twist to

occur but the twist would gradually increase with time at a decreasing

rate until it ceased. This behavior seems similar to a creep phenomenon.

The Ty observed in the tests shown in Figure V-2 varied widely.

Ty for interface 7 was about 0.23 N-m, for interface 9 was about

2 N-m, and for interface 8 was over 2.8 N-m and could not be observed

within the loading capacity of the torsion testing machine. Part of

this variation is to be expected. Keller et al . (1980) in extensive

mechanical studies of metal/bone cement junctions did observe standard

deviations which exceeded 50 percent of the mean interface strength.

Interfaces 8 and 9 could have a Ty range compatible with this figure.

However, interface 7 failed at a much lower value. Presumably the

reason lies with differences in the preparation of the samples. The

cure time for interface 7 was 9 days compared to 7 days for interfaces

8 and 9. The total amount of bone cement prepared (from which the

mass used to form the acrylic block was taken) for interface 7 was

51
half of that used for 8 and 9. The properties and quality of the bone

cement does depend somewhat upon the quantity which is polymerized.

Also, interfaces 8 and 9 were formed from the radiopaque bone cement

formulation whereas interface 7 was formed from the plain cement

composition. The material used for interface 7 had been manufactured

about 6 years prior to that used for interfaces 8 and 9.

In future tests it is probable that the variability could be

reduced by more uniform preparation of the interface, through use of

hydraulically applied pressure to the mold, and mixing of identical

amounts of material. Also, equal cure times should be employed for

all samples. Testing with more sophisticated mechanical testing

machines (than that used here) which can measure the time dependent

behavior of the interface could also be of assistance in this regard.

52
Chapter VI

SUMMARY AND RECOMMENDATIONS

The main conclusions to be drawn from this literature search can

be summarized as follows:

1. Until biomaterials exist which match bone in its elastic properties and are also otherwise suitable for implantation, stress shielding

of bone in the vicinity of implants will occur. Bone resorption is a

possible consequence. The area of the total hip replacement most

likely to be adversely affected is the calcar femoral e. Calcar resorption can create high stress situations in the implant itself leading

to its eventual failure. Implant designs which minimize stress shielding are to be preferred.

2. Most of the experimental and theoretical techniques which are

used by stress analysts have been applied to bone and many problems of

importance to hip implants have been examined. The unique conditions

present within the body create uncertainties with regard to these

analyses. The theoretical analyses are particularly sensitive to

uncertain boundary conditions, irregular geometry, and nonuniformity

in elastic properties.

3. Few studies have been conducted of the stresses present iji vivo

at the metal /bone cement interface. The theoretical studies which

have been done do not show maximal stress patterns in the areas where

clinical failures are most likely seen.

4. The reason that there is a discrepancy between theoretical calculations and clinically observed failures is probably because of the

neglect in the computations of the anterior-posterior joint forces

which are present during the walking cycle.

53
5. Since failures of the bond between metal and bone cement occur in

practice and can lead to failures of the device itself, development of

testing methods to determine the bond strength are important. ASTM

tests which are used for measurement of properties of adhesives are

not suitable.

6. Many studies of the physical and mechanical properties of bone

cement have been conducted. Studies of the strengths of the bone

cement/metal interface in tension and shear have been made. The

strength is dependent upon many variables including metal surface

preparation and roughness, cure time, interface formation time, composition of the metal, air or Ringers solution environment, time between

preparation of the metal surface and interface formation, etc.

7. Since no studies of the bone cement/metal interface have been

done when torsional forces act at the interface and the literature

search indicated that torsion could be most important for interface

failure, preliminary experiments were performed. The results indicated

that a sample in which a cylindrical rod was embedded in an acrylic

block could allow an interfacial yield strength to be measured.

Based upon the results of this research the following recommendations can be made:

1. Further investigation of the importance of torsional loading to

the stresses developed at bone cement/metal interfaces should be

pursued to gain increased insight into service failures of total hip

replacements. Methods such as shear-lag analysis or three dimensional

finite element analysis should be employed for this purpose.

54
2. Further study should be pursued of the optimum test sample geometry for torsion specimens. Refinements should be made in the manner

in which interfacial strength and the proportionality constant between

torque and twist are determined from the experimental data.

3. Since time dependent effects were observed in this study and

creep has been observed in failed bone cement, further mechanical

tests to be pursued should be performed with a testing machine in

which strain rate can be controlled. Also since both a torsional

force and a downward force exist J_n vivo during the walking cycle,

some tests with a machine with biaxial capabilities should be performed.

4. Systematic studies of ways to improve the strengths of bone

cement/metal interfaces should be investigated. The common implant

metals and various types of surface finish (such as polished, roughened, porous, passivated, knurled, etc.) should be studied. Preparation techniques for the interface should be carefully controlled.

5. Scanning electron microscope studies of the interface between

bone cement and implant metals should be pursued to determine the true

contact areas between the two materials. This research should be

correlated with mechanical testing data to determine the intrinsic

strength of the interface based upon actual contact area. This would

allow results of laboratory tests of interfacial strength to be extrapolated to the clinical situation if SEM studies of retrieved bone

cement samples were done to determine the contact area present ijn vivo

at the metal /bone cement interface.

55
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63

Part III

SURFACE PREPARATION AND CORROSION BEHAVIOR OF TITANIUM ALLOYS

FOR SURGICAL IMPLANTS

A. C. Fraker, A. W. Ruff, P. Sung, A. C, Van Orden

and K. M. Speck
SURFACE PREPARATION AND CORROSION BEHAVIOR OF TITANIUM ALLOYS

FOR SURGICAL IMPLANTS

A. C. Fraker*, A. W. Ruff’^, P. Sung=^**, A. C. Van Orden*

and K. M. Speck**

*National Bureau of Standards, Washington, DC 20234

**Uniformed Services University Medical School,

Bethesda, MD 20016

***Food and Drug Administration, Silver Spring, MD 20857

ABSTRACT

Surface preparation and corrosion behavior of titanium alloys were

studied; both topics deal with the formation of surface oxide films.

When metals are prepared for surgical implant use, an effort is made to

produce an optimum surface. Effects of the surface treatments on

surface morphology, surface film composition, and structure are shown in

transmission electron micrographs and electron diffraction patterns.

Roughened surfaces were produced and surface films of TiO and/or Ti02

occurred on some specimens. Na0 ‘xTi 02 occurred on the specimens washed

in NaOH solution. Open circuit potential vs time curves have been

measured to show some electrochemical effects of various surface treatments. Results show that titanium alloys immersed in Hanks' physiological

solution reach the same final open circuit potential after approximately

two weeks' exposure regardless of prior surface treatment. Measurements

of the anodic polarization behavior of titanium alloys and other surgical

implant alloys show the effects of alloy composition and testing solution on the passive region and breakdown potentials of these materials.

In general, for titanium alloys, the differences are not great but the

presence of nickel results in a significant lowering of the breakdown

64
potential. The materials studied were titanium, T1-6A1-4V, Ti-Ni (memory

alloy), Ti-13Cu-4Ni and Ti-4.5Al-5Mo-l.5Cr. In all cases titanium

materials are more corrosion resistant in Hanks' solution than the

Co-Cr-Mo alloys, 316L stainless steel or the Co-Ni-Cr alloy. Repassivation measurements show the rapid formation and the high degree of

stability of the protective film on the titanium and titanium alloys.

65
SURFACE PREPARATION AND CORROSION BEHAVIOR OF TITANIUM ALLOYS

FOR SURGICAL IMPLANTS

Introduction

The purpose of this paper is to present the effects of surface

preparation on electrochemical behavior of titanium alloys in a simulated

physiological environment and to show the general corrosion behavior

and repassivation characteristics of titanium alloys in physiological

solutions. While these topics are related and both deal with the formation

of surface oxide films, the specifics of each are presented separately.

Titanium alloys have a high corrosion resistance in saline solutions

which are typical of physiological environments. This has been the

finding of Hoar and Mears [1] and a number of other research investigators [2-8]. Solar [9], in a thorough review of the corrosion resistance

of titanium surgical implant alloys discussed the nature of the passive

film, ^ vitro corrosion measurements, iji vivo biocompatibility and

clinical observations. This review concluded that the titanium alloys

probably are the most corrosion resistant and biocompatible metals in

use today.

Several investigations involving titanium alloys are reported in

this paper, and all support the superior corrosion resistance of titanium

alloys over other implant metals when exposed to saline solutions.

These studies have dealt primarily with anodic polarization, open

circuit potential and repassivation measurements. Anodic polarization

curves show the current density in the passive region, the span in

potential of the passive region and the breakdown potential. Repassivation studies provide a means for determining kinetics of surface film

66
formation and subsequent stability, and these data have shown that

titanium alloys repassivate readily and have stable surface oxide films

up to a potential of 2 volts. Open circuit potential measurements and

transmission electron microscopy observations of surface preparation

effects on titanium alloys indicate that acid or alkaline treatments

cause surface roughness and affect initial surface films. Washing in

boiling water or boiling saline water would be better, assuming that the

material is sufficiently clean.

Part 1. The Effects of Surface Preparation

Electrochemical Measurements . Earlier work of Revie and Greene [3]

showed that effects of passivating on the corrosion behavior of titanium

materials in saline solution were negligible while such treatments

significantly improved the corrosion behavior of 316L stainless steel.

The prescribed surface treatments of ASTM F-86 for metals to be placed in

the body will result in different effects for the various currently used

surgical implant alloys. This can be explained partly by considering

the data in table 1, taken from the "Handbook of Physics and Chemistry"

[10] to show oxide solubility.

Fraker and Ruff [11] showed that titanium alloys reach the same

rest potential in Hanks' physiological solution regardless of prior

treatment. Some examples of this are shown in figures 1 a-e. All voltages

are given in reference to the hydrogen electrode. The solid curve in

figure la shows that the initial open circuit potential for a mechanically

polished specimen immersed in Hanks' physiological solution at 4 °C is

-0.10 volts (e|^) which rapidly changes to +0.14 volts. Results from a

mechanically polished and steam sterilized specimen at room temperature.

67
Table 1 . Solubility of Selected Metal Oxides [ 10 ]

Oxide Solubility in acid (a), alcohol (al), alkaline (alk)*

Brookite

TiOp Anatase

Rutile

s H2S0^,alk; i a

TiO s dil. H2SO4; i HNO^

Ti^Os s dil. H2S0^; i HNO3, HCl

FeO s a; i al , al

k

s HCl, si s HNO3

s Cone, a; i al , eth

Fe (OH)^ s a, NH^Cl; i alk

Cr02 s HNO3

CrO i dil HNO3

i a, alk, al

Cr202*xH20 s a, alk; si s NH^(OH)

Co (0H>2 s a, NH^ salts; i alk

CoO s a; i al, NH^OH

CO2O3 s a; i al

C03O4 V si s a; i aq. reg.

* s, soluble; v, very; si, slightly; a, acid; alk, alkaline;

al , alcohol; i, insoluble; eth, ether.

68
23 °C, are shown in the dashed curve and have a rest potential of

+0.22 volts. All specimens for figures 1 b-e were kept at 23 °C.

Figure lb shows that a similar specimen, washed in boiling sodium

chloride prior to sterilizing has an initial potential of +0.14 volts

after sterilizing and a rest potential of +0.24 volts after 14 days.

Figure Ic shows the open circuit potential measurements for a titanium

specimen washed in 30% HNO^ at 60 °C, steam sterilized and immersed in

Hanks' solution. Here the specimen has a positive open circuit potential

of +0.10 volts which levels off to +0.25 volts after 15 days. Washing

titanium and titanium alloy specimens in alkaline solutions such as

NaOH results in the dissolving of the surface oxide film [10]. The

titanium specimen of figure Id had this treatment and the initial open

circuit potential was -0.13 volts but after 5 minutes reached 0 volts

and quickly rose into the positive voltage range. This specimen eventually reached a rest potential of +0.20 volts. Figure le shows open

circuit electrode potential measurements for titanium which was passivated

in 30% HNO^ at 60 °C and was unsterilized. Note the initial high

positive potential of +0.68 which changes with time to a rest potential

of +0.22 volts.

It is seen in figures la-e that all specimens exhibit open circuit

potentials which rise and after two to four weeks fall to rest potentials

ranging from +0.20 volts to +0.24 volts. The similarity in the rest

potentials of the different specimens indicates that in this simulated

physiological solution, the surface film on titanium materials will

tend toward an equilibrium for this given environment regardless of

prior treatments such as alkaline washes, acid washes or sterilization.

Thicker films which would give a significantly more positive initial

69
open circuit electrode potential could be a source of ion release as

the surface film changes to a film more stable in the given environment.

Surface oxide films as a possible source for titanium ion release have

been discussed previously [12]. The oxide films merit further study

since titanium is highly corrosion resistant to body fluids, and it is

difficult to explain titanium metal ion release into tissue surrounding

an implant on any other basis.

Characterization of Surfaces and Films . Oxide film morphology and

composition on titanium alloys also are affected by the type of surface

treatment applied. Titanium is a metal which undergoes a series type

of oxidation sequence. The surface film changes from a lower to a

higher oxide as oxidation progresses and as temperatures are increased.

The following phases form; Ti + 0 Ti(0) TigO -> Ti^O ^ Ti20 TiO -»

Ti^Os ^ ^ Ti02- This has been reported for gaseous oxidation of

titanium by Kornilov [13] and for aqueous oxidation from TiO to Ti02 by

Fraker and Ruff [14]. If titanium has not been exposed in aqueous

environments in excess of 100 °C, the surface film will be of a lower

oxide type such as TiO.

The titanium oxide films are insoluble in most acids and remain

protective even after an acid washing treatment. This was indicated by

the positive electrode potential measurements of figure le after the HNO^

passivating treatment. The titanium oxides are not stable in alkaline

solutions, and the surface film resulting after treating in these

solutions is a mixture of titanium oxide and the alkali metal oxide.

Titanium thin foils were prepared by electropolishing [14] in a

solution of 200 ml of methanol, 115 ml of 1-butanol and 20 ml of 70 percent

perchloric acid at a temperature of -40 °C to -50 °C and at a voltage

70
of 13 to 15 volts. The thin foils were given designated treatments and

were studied using transmission electron microscopy techniques. The

transmission electron micrographs of figures 2 through 6, respectively,

show the pure untreated titanium, titanium after immersing in boiling

water for 30 minutes, titanium after immersing in boiling NaCl solution

for 30 minutes, titanium after immersing in NaOH solution for 15 minutes,

and titanium after passivating in 30% HNO^ at 60 °C for 30 minutes.

All specimens except the untreated titanium were steam sterilized

final ly.

Electron diffraction patterns are inserted in figures 3 through 6.

The presence of the oxide, TiO, after the plain water or saline water

treatment is indicated by the diffuse rings. The diffraction spots are

from the titanium metal. After steam sterilization, some patches of

higher titanium oxides can occasionally be found but these are not

predominant [7]. Immersion of titanium in alkaline solution such as

0.5N NaOH results in a combined form of surface oxide which is not as

protective as any titanium oxide form. The film covering the surface

in figure 5 is a form of sodium titanate, Na0*xTi02. After immersing

in hot HNO^ and steam sterilizing as shown in figure 6, the resulting

surface film is Ti02» anatase. Earlier studies [15] have shown that

lowering the solution pH enhances the nucleation and growth of the

titanium surface film, Ti02» anatase. It is apparent from these

electron micrographs that the titanium surfaces are smoother and less

affected after treatments in either plain or saline water than after

acid or alkaline treatments.

71
Part 2. Anodic Polarization Behavior and Repassivation Studies

Anodic Polarization . Anodic polarization measurements were used

to determine the passive range and the breakdown potential for a number

of metals which are used currently as surgical implants. Experimental

details of the procedure and composition of the materials have been

given in a paper [16] describing corrosion behavior of the memory

alloy, Ti-Ni and Ti-6A1-4V. All materials were implant grade as specified

by ASTM designated compositions. Specimens were prepared by mechanically

polishing through 0.05 pm Al20^ and then were anodically polarized. The

potential was increased at a stepping rate of 0.006 V/min. in Hanks'

physiological solution. All potentials were measured in reference to a

saturated calomel electrode. Hanks' solution was prepared [16,17] and

tests were conducted at 37 °C and a pH of 7.4. Results showed that the

titanium alloys tested had superior corrosion resistance over the other

metals tested by a large margin. The breakdown potentials for these

materials are given in table 2.

Table 2. Breakdown potentials for implant metals in Hanks' solution

METAL BREAKDOWN POTENTIAL. (VOLTS vs S.C.E.)

316L Stainless Steel +0.2 to +0.3*

Co-Cr-Mo +0.42

Co-Ni-Cr +0.42

Ti-Ni +1.14

Ti-6A1-4V +2.0

Tantal urn +2.25

Titanium (pure) +2.4

Ti-4.5Al-5Mo-l.5Cr +2.4

^Breakdown varies and is in this range.

72
Anodic polarization curves for titanium and the three titanium alloys

listed here are shown in figure 7 along with polarization curves for the

other metals. Note the wide passive region. The effects of alloying

elements are not great except in the case of the memory alloy which

contains 50 percent nickel and the Ti-13Cu-4Ni alloy. The breakdown

potential for pure nickel in Hanks' solution is less than +0.1 volt [16].

Even the nickel bearing titanium alloys were far more passive than the

cobalt-chromium alloys or 316L stainless steel. Note that tantalum also

has a high breakdown potential.

Repassivation Studies . The surface oxide film of a metal in

service may be removed or disturbed by fretting, abrasion, cracking or

scratching. The repassivation kinetics and the passive film stability

become important if the material is to retain its corrosion resistance

and integrity. Repassivation studies were conducted in Hanks' solution

at 37 °C. The specimen, the Ti-6A1-4V alloy, was abraded with a SiC paoer

and had its passive film removed. The anodic current transient was

recorded on an oscilloscope immediately following the passive film

removal. Figure 8 shows the current transient at applied potentials of

zero, +600 mV, +1200 mV and +1500 mV. All voltages in the repassivation

studies are in reference to the saturated calomel electrode. Within

this voltage range, there is little change in the repassivation rate as

indicated by the peak height and decay. There also is little change in

the film stability which is indicated by the leveling off of the current

transient curve. At higher potentials, the (I /I ) increases and the

IT)3X passive film becomes unstable. This is shown in figure 9 for potentials

of +3.0 V, +4.0 V, +5.0 V, +6.0 V, and +7.0 V, respectively. Note that

at +6.0 V, the steady state current is higher and the film becomes

73
unstable. The degree of passivation which occurs is expressed by I

is the steady state current and I is the maximum current as indicated

in figure 9. When compared with the other metals of table 2, titanium

alloys have, by far, the most stable surface film. For example films on

316L stainless steel become unstable at 0.2 V and those on Co-Cr-Mo become

unstable at 0.42 V [18] while films on pure titanium are stable to 2.4 V.

CONCLUSIONS

The excellent corrosion resistance of titanium and titanium alloys

in saline and physiological environments has been reaffirmed. Some

specific conclusions relating to surface preparation and corrosion

behavior of titanium are as follows:

1. Open circuit electrode potential measurements in Hanks'

physiological solution indicate that titanium alloys will

exhibit the same rest potential after a period of 2 to

3 weeks regardless of prior treatment. The surface film will

change to approach an equilibrium state for the given solution.

2. Transmission electron micrographs show that the morphology of

the titanium surface is affected by prior treatment. Washing

in acid or alkaline solutions results in a roughened surface,

an effect which is not observed at neutral pH.

3. Electron diffraction patterns along with the transmission

electron micrographs show that the nature and composition of

the titanium surface film change with variation of treatment.

Alkaline washes produce a surface film consisting of a mixed

oxide containing both titanium oxide and the alkali oxide.

74
Hot acid washes produce the surface oxide, Ti02 anatase, and

hot water or hot saline water washes result in a surface

oxide which is TiO or a lower form of titanium oxide. Steam

sterilization can produce TiO, lower oxides and occasional

areas of Ti 02 anatase.

4. Anodic polarization measurements in Hanks' solution of numerous

titanium alloys and other surgical implant metals show the

titanium materials to be the most corrosion resistant with

more positive breakdown potentials. This breakdown potential

is dependent on alloy composition, with nickel acting to

lower it.

5. Repassivation studies show that titanium and other surgical

implant metals repassivate readily. The stability of the

surface oxide film on titanium persists to +2.4 volts, which

is +1.98 volts beyond the next most stable oxide on the

Co-Cr-Mo alloy.

75
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pp. 465-470.

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4. Mueller, H. J. and Greener, E. H. Journal of Biomedical Materials

Research, Vol. 4, 1970, pp. 29-41.

5. McMaster, James A., "Titanium for Prosthetic Devices", IMD, DentalMedical Committee, AIME, Cleveland, Ohio, Oct. 21, 1970.

6. Hille, Geunther H. , Journal of Materials, Vol. 1, June 1966,

pp. 373-383.

7. Fraker, A. C. , Ruff, A. W. and Yeager, M. P. , Proceedings, Second

International Titanium Conference, Plenum, New York, 1973, pp.

2447-2457.

8. Solar, R. J. , "An InVitro Investigation of the Corrosion and

Capacitative Behavior of Titanium Surgical Implant Alloys," Ph. D. Thesis, Univ. of Pennsylvania, Philadelphia, 1977. Solar, R. J.

,

Pallack, S. R. and Korostoff, E. , Journal of Biomedical Materials

Research, Vol. 13, No. 2, 1979, p. 217.

9. Solar, R. J. , "Corrosion Resistance of Titanium Surgical Implant

Alloys: A Review," Corrosion and Degradation of Implant Materials,

ASTM STP 684, B. C. Syrett and A. Acharya, Eds. , American Society

for Testing and Materials, 1979, pp. 259-273.

10. Handbook of Chemistry and Physics, R. C. Weast, Ed., The Chemical

Rubber Publishing Co., Cleveland, OH, 1971.

11. Fraker, A. C. and Ruff, A. W. , "Studies on the Effects of Surface

Preparation of Titanium Alloys", Second Annual Meeting of the Soc.

for Biomaterials, Philadelphia, PA, April 9-13, 1976.

12. Solar, R. J. , Pollack, S. R. and Korostoff, E. , "Titanium Release

from Implants: A Proposed Mechanism," Corrosion and Degradation of

Implant Materials, ASTM STP 684, B. C. Syrett and A. Acharya, Eds.,

American Society for Testing and Materials, 1979, pp. 161-172.

13. Kornilov, I. I., Doklady Akad. Nauk. USSR, Vol. 183 (5), 1968, p. 1087.

14. Fraker, A. C. and Ruff, A. W. , Corrosion Science, Vol. 11, 1971,

pp. 763-765.

76
15. Fraker, A. C. and Ruff, A. W. . "The Effect of Solution pH on the

Saline Water Corrosion of Titanium Alloys", Proceedings, Second

Internation Titanium Conference, Plenum, New York, 1973, pp. 2655-

2663.

16. Speck, K. M. and Fraker, A. C. , Journal of Dental Research, Vol.

59(10), 1980, pp. 1590-1595.

17. Hanks, J. H. and Wallace, R. E. , "Relation of Oxygen and Temperatures in the Preservation of Tissues by Refrigeration," Proceedings

of the Society of Experimental Biology, Washington, D.C., Vol. 71,

1949, pp. 196-200.

18. Sung, P. , and Fraker, A. C. , "Repassivation Kinetics of Ti-6A1-4V,

316L Stainless Steel, Co-Cr (Cast) and Co-Ni-Cr (MP35N), Transactions, Seventh Annual Meeting, Soc. for Biomaterials, Rensselaer

Polytechnic Institute, Troy, NY, May 28-31, Vol. 4, p. 28, 1981.

77
>

X

Q>

O

CL

Fig. la. Open circuit potential vs. time curve for

titanium, mechanically polished, steam sterilized

and immersed in Hanks' solution.

TIME (days)

Fig. 1b. Open circuit potential vs. time curve for

titanium, mechanically polished, washed in

boiling 3.5% NaCI for 30 min., steam sterilized

and immersed in Hanks' solution.

78
Fig. 1c. Open circuit potential vs. time

titanium, mechanically polished, washed for

15 min. in 30% HNO3 at 60°C, steam sterilized

and immersed in Hanks' solution.

TIME (days)

Fig. Id. Open circuit potential vs. time curne for

titanium, mechanically polished, washed for

2 min. in hot 5N NaOH,

immersed in Hanks' solution.

79
POT.,

6h

(V)

TIME (days)

Fig. 1e. Open circuit potential vs. time curve for

titanium, mechanically polished, washed 15 min.

in 30% HNO3 at 60°C, unsterilized and

immersed in Hanks' solution.

80
Fig. 2.

Titanium after electropolishing.

Fig. 3

Titanium after electropolishing, immersing

in boiling water for 30 minutes and steam

sterilizing. Diffuse ring indicates a surface film, probably a lower titanium oxi e

Fig. 4. Fig. 5.

Titanium after immersing in boiling 3 . 5 % NaCl Titanium after immersing in hot .5N NaOH for

I for 30 min. and steam sterilizing. Diffuse 15 min. and steam sterilizing. Surface film

! rings indicate TiO. is a form of Na0'XTi02.

81
Fig. 6.

Titanium after immersing in hot 30% vol . HNO^

for 30 min., and steam sterilizing. Surface*^

film is Ti 02 anatase.

Fig. 7 Anodic polarization curves for titanium, titanium alloys and

other implant metals. Potentiostatic measurements v/ere made

in Hanks' solution.

82
Fig. 8. Current (i) vs. Time (t) curves showing

repassivation of Ti-6AI-4V at 0 (bottom curve),

+600mV, +1200mV and +1500mV, respectively.

Fig. 9. Current (i) vs. Time (t) curves showing

repassivation of Ti-6AI-4V at +3.0V (bottom

curve), +4.0V, +5.0V, +6.0V and +7.0V,

respectively.

83
Acknowleqments

The helpful discussions with Sidney Weisnan, Howmedica, Inc. and

his assistance in providing bone cement and metal specimens are gratefully

acknowledged. Also appreciated are materials received from Zimmer-USA and

Carpenter Steel, discussions with Dr. Gerald Brauer, National Bureau of

Standards and the support of Messrs. Daniel J. Chwirut and Donald M. Marlowe,

Bureau of Medical Devices, the Food and Drug Administration.

84
IBS-114A (REV. 2 «ec) U.S. DEPT. OF COMM.

BIBLIOGRAPHIC DATA

SHEET CSee instructions)

1. PUBLICATION OR

REPORT NO.

NBSIR 92-2563

2. Performing Organ. Report No. 3. Publication Date

October, 1982

4. TITLE AND SUBTITLE

Metallurgical Studies of Interface Bonding on Implant Alloys

«. AUTHOR(S) and Speck, K. M.

Fraker, A. C., Ruff, A. W., Bundy, K. J., DeMontigny, S. A., Sung, P., Van Qrden, ATT 7/

C. PERFORMING ORGANIZATION (If joint or other than NBS, see instructions)

NATIONAL BUREAU OF STANDAROS

DEPARTMENT OF COMMERCE

WASHINGTON, D.C. 20234

7. Contract/Grant No. 22 4795023

8. Type of Report & Period Covered

Oct. 1981 - July, 1982

9. SPONSORING ORGANIZATION NAME AND COMPLETE ADDRESS (Street. City. State. ZIP)

Bureau of Medical Devices

Food and Drug Administration

HFK-220

8Z57 fieorgia Ave. , Silv.er Spring, MD 20910

:10. SUPPLEMENTARY NOTES

I I Document describes a computer program; SF-185, FIPS Software Summary, is attached.

11. ABSTRACT (A 200-word or less factual summary of most significant information. If document includes a significant

bibliography or literature survey, mention it here)

A literature review covering articles on stress analyses in total hip replacement

pros theses and the metal / bone cement interface is presented. The literature indi called

the need for a test which utilized loading in torsion, and such a test was developed

and is described. The test can be used to determine the influence of various parameters

including surface roughness and passivation and sterilization treatment on the strength

of the metal /bone cement bond. Some preliminary tests were conducted and results are

given. Future work using the mechanical test developed is discussed. Additional

studies were conducted on the surface preparation of titanium, and data are presented

to show that changes in initial electrochemical behavior and varying degrees of surface

roughness occur depending on whether the metal receives a neutral, alkaline or acidic

washing treatment.

12. KEY WORDS (Six to twelve entries; alphabetical order; capitalize only proper names; and separate key words by semicolons)

bone cement, hip prosthesis, stress analysis, surface preparation, surgical implant metjals

lest method . titanium ! 13. AVAILABILITY

Unlimited

03] For Official Distribution. Do Not Release to NTIS

Qn Order From Superintendent of Documents, U.S. Government Printing Office, Washington, D.C.

20402.

Order From National Technical Information Service (NTIS), Springfield, VA. 22161

14. NO. OF

PRINTED PAGES

95

15. Price

$ 10.50

USCOMM-DC 6043-P80
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