NBS

NBSIR 84-2820 (FDA) Reference

TM"

ns Information on Polymeric Materials

Used In Orthopedic Devices

NAT'L INST. OF STAND & TECH

Alliob Q3HM71

Task 80-01, NBS-Bureau of Medical Devices Interagency Agreement

January 1984

-QC • 10G tPARTMENT OF COMMERCE, Malcolm Baldrige. Secretary

UOQ AL BUREAU OF STANDARDS, Ernest Ambler, Director

’ 84-2320

1934

NBSIR 84-2820 (FDA)

Information on Polymeric Materials

Used In Orthopedic Devices

J.M. Crissman

G.B. McKenna

Task 80-01, NBS-Bureau of Medical

Devices Interagency Agreement

January 1984

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

NATIONAL BUREAU OF STANDARDS, Ernest Ambler, Director

NATIONAL bureau

OF STANDARDS

LIBRARY
'
Abstract

This report provides information on the two polymeric

materials most commonly used in the fabrication of orthopedic

implants. The work was done as part of Task 80-01 NBS-FDA/BMD

Interagency Agreement. The two materials described are ultra

high molecular weight polyethylene UHMWPE and polymethylmethacrylate

(PMMA) bone cement. The report contains information on such

subjects as specifications (ASTM), raw materials characteri zati on

processing, morphology, mechanical properties, and wear.

Foreword

This report has been prepared for the Food and Drug Administration,

Bureau of Medical Devices, as part of Task 80-01 NBS-BMD Interagency

Agreement

The technical monitors for this effort were Mr. Daniel J. Chwirut

of the Center for Medical Device Analysis, and Dr. James G. Dillon of the

Division of Surgical and Rehabilitation Devices, Bureau of Medical Devices,

Department of Health and Human Services, FDA, Washington, D.C. The

principal investigator was Dr. John M. Crissman and the project leader

was Dr. Gregory B. McKenna.

The authors gratefully acknowledge the assistance of Dr. Fred A. Khoury

who provided the optical and electron micrographs concerned with ultra

high molecular weight polyethylene. In addition the authors wish to thank

the following publishers for permission to reporduce in this report certain

figures taken from journal articles: John Wiley & Sons, New York; Pergamon

Press, Inc., Elmsford, New York; Butterworth Scientific, Ltd., Surrey, England;

the American Society for Testing and Materials (ASTM) , Philadelphia, PA;

the Journal of Bone and Joint Surgery, Boston, MA. and United Trade Press

Limited, London.

1. Introduction

This report provides information on the two polymeric materials

most commonly used in the fabrication of components for orthopedic

implants. Specifications for both materials are listed in the American

Society for Testing and Materials. Each polymer is discussed in a

separate chapter with the headings and paragraph numbering system

designed to be compatible with those used in the chapters prepared

by Eagle Engineering, Inc., Houston, Texas describing the six

metal alloys used in orthopedic implants.

The two materials to be described are ultra high molecular

weight polyethylene (UHMWPE) and polymethylmethacrylate (PMMA) bone

cement. UHMWPE is used in the fabrication of implant components such

as acetabular cups (hip replacement) and tibial plateaux (knee

reconstruction). PMMA bone cement is used to secure the implant

components to bone as well as act as a space filler material in

repair of diseased bones.

Since in clinical use the function of the two polymeric materials

is quite different, the same information may not in all cases be

available for both. Moreover, specific properties or characteristics

which are important in determining performance may be different in the

two cases. The properties provided are typical and should not be

regarded as valid design numbers or design criteria.

A section on wear has been included in the chapter on UHMWPE.

Wear represents an important aspect of materials characterization and

performance in applications such as orthopedic devices, yet it remains

a most difficult phenomenon to quantify. As in the case of metals,

the results are often contradictory and in dispute. To a large extent

this is believed to be due to differences in test procedures, variations
in specimen geometry, and the methods in which wear is evaluated

and quantified. There remains a need for the development of more

standardized test procedures for the determination of wear and for

making quantitative comparisons of wear results.
TABLE OF CONTENTS FOR SECTION 2

Subject Page

2. Ultra High Molecular Weight Polyethylene 2-1

2.1 Trade Names and Specifications 2-1

2.1.1 Trade Names 2-1

2.1.2 Specifications 2-1

2.2 Description of Terms (ASTM Grade) 2-1

2.3 Virgin UHMWPE Powder Requirements (ASTM Grade) .... 2-2

2.3.1 Generic Properties 2-2

2.3.2 Nongeneric Properties 2-2

2.4 UHMWPE Fabricated Form Requirements (ASTM Grade) . . . 2-3

2.4.1 Compositional Requirements 2-3

2.4.2 Physical Requirements 2-3

2.4.3 Mechanical Requirements 2-3

2.5 Sampling (ASTM Grade) 2-4

2.6 Processing Information 2-4

2.6.1 Raw Polymer Characteri zation 2-4

2. 6. 1.1 Morphology 2-4

2. 6. 1.2 Density 2-6

2. 6. 1.3 Compaction .... 2-7

2. 6. 1.4 Melting Behavior 2-8

2.6.2 Compression Molding 2-8

2.6.3 Characterization of the Fabricated Form .... 2-10

2. 6. 3.1 Density 2-10

2. 6. 3. 2 Morphology 2-12

2. 6. 3. 3 Machini ng 2-15

2. 6. 3. 3.1 Sawing 2-15

2. 6. 3. 3. 2 Milling 2-16

2. 6. 3. 3. 3 Drilling 2-17

2.6.3. 3.4 Turning 2-17

2.7 Mechanical Properties 2-17

2.7.1 Constant Rate of Elongation 2-18

2.7.2 Impact Resistance 2-19

2. 7.2.1 Tensile Impact 2-19

2. 7. 2. 2 Izod Impact 2-19

2.7.3 Creep and Recovery 2-20

2. 7. 3.1 Short Time Creep and Recovery at

Small Deformations 2-20

2. 7.3.2 Long Time Creep at Small Deformations. . 2-28

2. 7. 3. 3 Creep at Large Deformations 2-33

2.7.4 Stress Relaxation 2-37

2.7.5 Fatigue 2-40

2-i
Page

2.8 Wear 2-44

2.8.1 Types of Wear 2-44

2.8.2 Methods for Wear Testing 2-45

2.8.3 Wear Rate Measurements 2-45

2.8.4 Wear Rate Data 2-48

2.9 Effects of High Energy Radiation 2-53

2.9.1 Density and Crystallinity 2-54

2.9.2 Melting Behavior 2-56

2.9.3 Tensile Yield Stress 2-58

2.9.4 Tensile Yield Point Elongation 2-60

2.9.5 Tensile Modulus 2-60

2.9.6 Impact Strength 2-62

2.9.7 Surface Hardness 2-62

2.9.8 Creep 2-64

2.9.9 Stress Relaxation 2-68

2.9.10 Fatigue 2-68

2.9.11 Effect of Radiation on Wear 2-69

2-ii
2. ULTRA HIGH MOLECULAR WEIGHT POLYETHYLENE (UHMWPE)

2 . 1 Trade Names and Specifications

2.1.1 Trade Names

UHMWPE used in the manufacture of surgical implants is

? marketed under the names and designations 1900 and Hostalen

RCH 1000 C

3

.

2.1.2 Specifications ^

The principal specifications for linear UHMWPE powder

intended for use in surgical implants, and fabricated

forms, are those contained in ASTM F-648-80, Standard Specification

for ULTRA-HIGH-MOLECULAR-WEIGHT POLYETHYLENE POWDER AND FABRICATED

FORM FOR SURGICAL IMPLANTS [ 1 ] 5

.

2.2 Description of Terms (ASTM Grade)

2.2.1 generic property- that property which is determined solely

by the chemical composition and structure of the virgin polymer.

2.2.2 virgin polymer powder- the form of UHMWPE as obtained from

the manufacturer and prior to fabrication into a bulk shape.

2.2.3 fabricated form. any bulk shape of UHMWPE, fabricated from

the virgin polymer powder, used during the process of fabricating

surgical implants prior to packaging and sterilization.

Certain commercial materials are identified in this report in order to

specify them adequately. In no case does such identification imply

recommendation or endorsement by the National Bureau of Standards.

2 Registered Trademark - Hercules Inc., Wilmington, Delaware.

3 Registered Trademark - American Hoechst Corporation, Somerville, New Jersey.

4 Reprinted, with permission from the Annual Book of ASTM Standards. Copyright,

ASTM, 1916 Race Street, Philadelphia, PA 19103.

^Numbers in brackets denote references to the literature found at the end of

this section.

2-1
2.3

Virgin UHMWPE Powder Requirements (ASTM Grade)

2.3.1 Generic Properties:

2. 3. 1.1 The virgin polymer shall be a linear homopolymer of

ethyl ene

.

2. 3. 1.2 The molecular weight of the polymer powder shall be

indicated by determining either the relative solution viscosity

or the flow value. The relative solution viscosity shall not be

less than 1.95. The prescribed method for the determination of

the relative solution viscosity is ASTM D1601 . The flow value

shall not be less than 0.2MPa when determined by ASTM F -648

Annex A2.

2.3.2 Nongeneric Propertie s

:

2. 3. 2.1 The polymer powder shall not contain extraneous matter

such as dirt, lint, silica and discoloring materials of diameter

greater than 300 urn. When viewing a sample prepared in accordance

with section 7.1.2 of ASTM F648-30, the concentration of visible

particles of extraneous matter shall not be greater than 10 per

2. 3. 2.

2

The concentration of trace elements in the polymer

powder shall not exceed:

Element ppm, max

A1

Ti

Ca

Cl

50

300

100

90

2. 3. 2.

3

The content of any other trace element shall not

exceed 10 ppm.

2. 3. 2. 4 Total trace elements shall not exceed 550 ppm.

2. 3. 2. 5 All powder shall pass a No. 16 (1.18mm) sieve.

2-2
2.4 UHMWPE Fabricated Form Requi remen ts - (ASTM Grade)

2.4.1 Compositional Requirements :

2. 4. 1.1 No stabilizers or processing aids are to be added to

the virgin polymer powder during manufacture of a fabricated form.

2. 4. 1.2 The surface of a fabricated form shall not contain

particles of residue of diameter greater than 300 y m. The

concentration of the visible particles shall not be greater than

2

10 particles per 400 cm .

2.4.2 Physical Requirements:

2. 4. 2.1 The fabricated form shall be fused to the extent that

light patches larger than 300 u m in diameter shall not be

observed

.

2. 4. 2. 2 The density of the fabricated form shall be within the

range from 0.930 to 0.944 g/cm^.

2.4.3 Mechanical Requirements:

2. 4. 3.1 UHMWPE in fabricated form from which surgical implants

are made shall meet the requirements listed in Table 2. 4. 3.1.

Table 2. 4. 3.1 UHMWPE Fabricated Form Requirements

Property Fabricated Form Type ASTM Test Method Requirement, min

Tensile strength at 23°C 0638 (Speed C)

1)1 timate

Yield

Elongation

U1 timate

Yield

Elongation

Ul timate

Yield

Elongation

I zod impact strength

Deformation under load

molded, machined

molded, machined

molded, machined

extruded, machined

extruded, machined

extruded , machined

compression molded

compression molded

compression molded

all types

all types 0 621 (A) (7 MPa

(1000 psi) for

24 h)

0 255(A)

34MPa (5000 psi)

21 MPa (3000 psi)

300°? 300S

21 MPa (3000 osi

)

200fj 27 MPa (4000 psi)

19 MPa (2800 osi)

350?' ncnbreak

2?1 deformation

after SO-min

recovery

Hardness all types D 2240(Shcre 0) 60

2-3
2.5 Samp 1 i nq (ASTM Grade )

2.5.1 Where applicable, the requirements for each lot of powder

and fabricated form by sampling sizes and procedures shall be done

according to ASTM Method D1898 Standard Recommended Practice for

SAMPLING OF PLASTICS.

2.6 Processing Information

2.6.1 Raw Polymer Characteri zation

2. 6. 1.1 Morphology

UHMWPE polymer, in its as polymerized (i.e., nascent) state

consists of fine particles. The particles are irregularly shaped

and vary in size from approximately 80 to 330 u m [2,3].

Scanning electron micrographs showing one example of the nascent

UHMWPE powder are presented in Figure 2. 6. 1.1. The surfaces

of the particles are highly convoluted and appear to be composed

of clusters of minute spheres. Whereas portions of the surface

of the particle are relatively smooth, others exhibit a highly

porous appearance. At higher magnification the gaps or voids

in the porous regions can be seen to be bridged by thin fibrils

which are 50 to 100 nm in diameter. The various origins of

fibrils in nascent polyolefins have been reviewed by Marchessault

et al [4]. The particle size distributions for the different

commercial powders have been found, from sieve analysis, to have

roughly log normal distributions [3].

2-4
Fig. 2. 6. 1.1 (a) Scanning electron micrograph of a particle

of raw UHMWPE. (b) The region to the right of the area y shown

at higher magnification. Note the fibrils bridging the gaps

or voids in the porous regions. (From reference [2]).

2-5

2.6.1 .2 Density

The density of nascent UHMWPE powder can be determined by one

or more of several different methods described in ASTM D792 [5] ,

ASTM B21 2 [6], ASTM B331 [6], and ASTM B527 [6]. Table 2. 6. 1.2

lists typical values of the density as determined by each method.

Table 2. 6. 1.2 Typical Values of Density for UHMWPE

Powders Used in the Fabrication of Surgical Implants

J

Measure of Density

3 Density g/cm Method of Determination

theoretical 0.945-0.964 ASTM D792

apparent 0.34 -0.40 ASTM B21

2

green^ 0.76 -0.82 ASTM 8331

tap 0.41 -0.47 ASTM 8527

''Data taken from reference [3]

2 For a compaction pressures greater than lOOMPa.

2-6

In processing, the apparent density is used to determine the

tooling size and storage volume. The apparent density is determined

to a large extent by particle size, size distribution, and morphology.

Tap density refers to the packing possible when the barriers to interparticle movement are overcome.

2. 5. 1.3 Compaction

The compressi bi 1 i ty of a polymer powder can be represented as a

compaction efficiency, or "densi fi cati on parameter" (DP), defined by:

np - green density - apparent densi ty

theoretical density - apparent density

The densif ication parameter increases with increased compaction pressure

and, in the case of UHMWPE powders, reaches a plateau value at pressures

above about 1 CO MPa [3,7]. Therefore, there aopears to be no particular

advantage in using compaction pressure greater than 100 to 200 MPa.

The maximum DP attainable for UHMWPE powders ranges from about 0.66 to

0.80 [3] .

2.6.1 .4 Melting Behavior

It is typical of crystal 1 izable polymers such as polyethylene

that their melting temperatures are dependent upon their previous

thermal and mechanical histories, and are lower than the equilibrium

melting point, i.e. the melting point of an ideal extended chain

crystal of the polymers (^lOK, rUI46°C)[9]. UHMWPE powders, as polymerized, exhibit a very high degree of crystal 1 i ni ty . Based upon

specific enthalpy of fusion measurements obtained from differential -

scanning calorimetry [2,3] , the percent crystallinity (x) is found to be

in the range x

= 66-78, and the melting point of the crystals is in

the range 413K-416K ( 1 40°C-1 43 °C ) . This feature of UHMWPE reflects

2-7
the favorable crystallization conditions present during polymerization.

Upon melting and subsequent recrystall i zation , the percent crystallinity

drops into the range X = 51-63, depending on the particular polymer

being used and the rate at which it is cooled from the melt. Upon

remelting the melting point drops into the range 408K-41 1 K ( 1 35°C-1 38°C)

depending on how rapidly the polymer is cooled following the first melting.

2.6.2 Compression Molding

Because of its high molecular weight and very high melt viscosity,

UHMWPE is not easily processed by conventional techniques such as

injection molding and extrusion. Processing techniques are further

restricted by the requirements in ASTM F648 that no stabilizers or

processing aids can be added to the virgin polymer powder. Fabricated

forms of UHMWPE polymer are most commonly produced by compression

molding of the powder. The finished prosthesis is either machined from

a block or sheet of previously compression molded material, or, in

some cases, is direct press molded from the powder. The latter process

is more difficult requiring uniform temperature control during both

heating and cooling, and if the shape of the piece is complex, the

pressure during cooling may not be evenly distributed over the whole of

the piece.

Each manufacturer of UHMWPE powder [9,10] has a recommended

5 procedure for compression molding of the powder. A list of several

of the more important procedures is provided below.

5 One manufacturer of UHMWPE supplies the polymer in sheet form only when it

is intended for use in surgical implants.

2-8
Procedures for Compression Molding UHMWPE

1. The mold cavity should have adequate depth to accomodate a

powder charge approximately 2.5 to 3 times the thickness of the

final piece.

2. The powder should be cold pressed at a pressure of from

10-20 MPa (1450-2900 psi) for a period of from 1-10 minutes.

This step is important in order to expel most of the entrapped

air from the powder. The cold pressing step can be done either

prior to heating of the press or during the early stages of

heating with the platens already at temperature.

3. After cold pressing the pressure is reduced into the range

2-5 MPa (290-715 psi) during heating. The press is heated to a

temperature between 200°C and 220°C and held at that temperature

for sufficient time to insure that all the material is in the

molten state.

4. A pressure in the range from 10-20 MPa (1450-2900 psi) is

again applied and maintained during cooling. Insufficient

pressure during cooling may result in cavities inside the piece

and sink marks on the surface.

5. The rate at which the mold is cooled will be determined

by the value of density desired for the molded form (see Section

2.6.3). Water cooling of the platens at a rate as high as 20°C

per minute cooling of the platens under ambient conditions will

3

generally yield a density in the range 0.930-0.940 g/cm .

2-9
2.6.3 Character! zation of the Fabricated Form

2. 6. 3.1 Density

The density of the fabricated form will depend upon

both the rate at which piece is cooled during molding and the

particular UHMWPE being used. Subsequent heat treatment may

increase or decrease the density. Examples of the density

variability which can be found in molded UHMWPE are given in

Table 2. 6. 3.1.

Table 2. 6. 3.1 Density (p) and Percent Crystallinity (x) of compression

molded UHMWPE Prepared under Varied Thermal Histories (From Reference [12])

UHMWPE

Polymer

Density3 g/cm

to t%

Crystal 1 i ni ty

1 (Quenched)

13 0.923 51

1 (Slowly Cooled) 0 0.935 59

1 (Annealed)

01 0.942 63

2 (Quenched)

13 0.925 52

2 (Slowly Cooled) 0 0.941 63

2 (Annealed)

01 0.953 71

a - Measured at 23°C in a water athanol density gradient column,

b - Obtained by reheating a sheet of the slowly cooled polymer to 200°C

between chromed photographic plates followed by submersion into ice

water.

c - Cooled in the press at a rate of ^1°C per minute,

d - Annealed under an Atmosphere at 130°C for 72 hours followed by

135°C for 48 hours.

2-10
The two UHMW polymers represent different lots of material

obtained from the same manufacturer. Polymers 1 and 2 had manufacturer

specified intrinsic viscosities* of £ 25 and £ 13 deciliters per gram,

which based on the manufacturers method for estimating molecular weight

from dilute solution viscosity measurements, correspond to molecular

6 6 weights of approximately 4x10 and 2x10 respectively . Both lots of

polymer satisfy the requirements contained in Section 4. 1.2.1 of

ASTM F-648

^ > 1.95) for UHMWPE used in surgical implants. Note,

however, that both quenched samples and one of the annealed samples

have densities which are not within the range specified in Section 5.2.2

of ASTM F-648. ( p =0. 930-0. 944 g/cm

3

)

*The determination of the intrinsic viscosity of UHMWPE from dilute

solution viscosity measurements is subject to a number of problems and

uncertainties which have been discussed in detail by Wagner and Dillon [13].

2-11
2. 6. 3. 2 Morphology

Unlike conventional low molecular weight polyethyl enes

,

compression molded UHMWPE retains a memory of the particulate

nature of the raw polymer powder. An example of this phenomenon is

shown in Figure 2. 6. 3. 2.1. This "grain memory" persists even in

samples for which the dwell temperature in the melt is as much as

90°C above the melting point. Since the mechanical properties of

the molded form will be influenced by the effectiveness of the

compaction and fusion of the raw polymer particles, it is important

that the dwell temperature in the melt be sufficiently high to

achieve the optimium fusion of the powder particles, but without

serious degradation of the polymer. For most molding operations a

dwell temperature in the range from 200-220°C is recommended.

Further examination of the bulk polymer reveals (Figure

2. 6. 3. 2. 2) that there are substantial variations in fine structure

in each cross-section as evidenced by the differences in the size of

the birefringent structural units from one region to another. In

some regions the birefringent units are barely visible and the

corresponding areas exhibit a fine grainy apoearance under the

polarizing microscope. Under normal conditions used to mold UHMWPE

large well developed spherulites are not observed. However, rotation

of the sample about the microscope axis while keeping the crossed

polarizer and analyzer fixed, indicates that some of the larger

birefringent units are incipiently spherulitic or axial etic in

character. The absence of well developed spherulites is likely a

consequence of a high nucleation density with the result that evolving

spherulites imoinge upon their nearest neighbors before attaining

‘mature 1 shapes.

2-12
Fig. 2. 6. 3. 2.1 Low magnification micrograph taken in a light microscope

(using transmission phase contrast optics) showing a cross-section

of a sheet of UHMWPE molded at 200°C. (Mw % 4x1 0^, cooling rate in

the press ^ 1 C per minute). Markers point to some of the boundaries

between the original raw polymer particles. (From reference 11 )

2-13
.
(b)

c ig. 2. 6. 3. 2. 2 Light micrographs of an area in a cross-section of a

sheet of UHMWPE (Mw % 4x10^) molded at 200°C. (a) Phase contrast

optics, (b) crossed polarizer and analyzer. Note the variation in

the sizes of the birefringent structures in (b). The fine variations

in contrast in (a) are associated with the birefringent structures

in the sample. (From reference 11 )

2-14

2. 6. 3. 3 Machi n 1 ng

Molded shapes of UHMWPE can readily be sawed, turned,

planed, milled, or drilled using equipment designed for fabricating

metals and wood [8,14]. To obtain high quality machined surfaces

sharp tools should always be used. In many applications normal

tool steel is adequate, but stellite or carbide tipped metal

cutting tools are recommended. The optimum cutting speed is

in the range from 250 to 1500 m/min. The lower cutting speeds

do not cause significant heat buildup and sticking, whereas water

cooling or dilute water soluble-oil lubrication is reauired when

cutting at the higher speeds.

2. 6. 3. 3.1 Sawing

UHMWPE shapes can be cut with standard power tools which

have sharp, wide-spaced teeth. Band saws are particularly

applicable since the blade carries off the heat. To prevent the

saw blade from becoming jammed, the teeth should have a slight

amount of set. When using circular saws, Tungsten carbide blades

are recommended. The higher the cutting speed the better the

surface finish will be, provided significant heat buildup

does not occur.

Table 2. 6. 3. 3.1 provides recommended settings and speeds for

sawing UHMWPE.

Table 2. 6. 3. 3.1 - Recommendations for Sawing UHMWPE Fabricated Forms

3

Band Saw Circular Saw Clearance

Angle, degrees 15 15 Angle, degrees 5-3 0-5

Feed, cm/rev. Hand Hand

Cutting speed, m/min 915-1840 915-2745

Tooth pitch, c:n .30-1 .0

a From reference 14

2-1 5
2. 6.3. 3. 2 Mi 11 ing

Semifinished forms of UHMWPE can be milled on any standard

or high speed top milling machine. To assure the most efficient

removal of chips, milling tools should be graduated to allow for

good tool clearance. Cutting speeds of from 180 to 540 m/min

are recommended, with a feed rate of approx imatel y 0.025 cm/rev.

2.6 .3. 3. 3 Dr i 1 1 i ntj

Drilling of UHMWPE can be done on a lathe, milling machine,

or drill press. With allowances for good chip removal , local

overheating can be largely avoided. If coolants, such as

compressed air, water, or dilute soluble oil are not used, the

drill should be removed from the hole frequently to clean out

the chips and prevent overheating. A set of speci fic ations

for drill i ng UHMWPE is given in Table 2.6 .3 .3 .3.

Table 2.6.3 .3.3 Specifications for Drilling UHMWPE Fabricated r a Forms

Hole Angle of Cl earance

,

Helix, Feed

,

Cutti ng Diameter, Poi nt, degrees degrees cm. /rev

.

speed

,

cm. degrees m/m i n

<2.03 60 15-20 0 .02 75

2.03-4.06 120 12-15 0 .02 75 >4.06 140 12-15 0 .038 100

a From reference 14

2-1 6
2. 6. 3. 3. 4 Turning

UHMWPE can be turned without difficulty on conventional

wood - or metal-working lathes. Cooling is generally only

required for large depth of cut. With feeds of from .0045-

.0127 cm/rev. surface speeds up to 180 m/min. are applicable.

2 . 7 Mechanical Properties

Table 2.7 lists the fabricated form requirements contained in

ASTM F-648 for the mechanical properties of UHMWPE used in surgical

implants. The minimum requirements listed are, for the most part,

based on relatively short time tests. However, the mechanical

properties of semi crystal 1 i ne polymers such as UHMWPE are highly

time dependent. The measured property may be quite sensitive to the

rate of straining or the environment. In such cases, data obtained

by these tests should not be considered reliable for applications in

which the load-time scale or environment is greatly different from that

of the test method. For engineering design purposes, tests should be

done over a broad load-time scale (including impact and creep) and

range of environment. In addition, the time dependent mechanical

properties of semi crystal! ine polymers are influenced by structure

or morphol ogy related features such as degree of crystallinity,

orientation, spherul ite size, and texture, all features which are

determined by both the polyner chemistry and processing.

Table 2.7 - UHKWPC Fabricated F-,r,n Regui regents (ASTM F-odS)

Property Fabricated Form Type ASTM Test Method Recju irenent

m i r.

T entile strength at 23 “C D 633 (Speed C)

U1 tnaate molded, machined 3 4 ( 50 CO us i

Yield molded, machined 21 M?a ' C CC C PSi)

El O'. get ion .'.elded, machine: 5 vO \ Ui cimate extruded, machined 34 A,?~ . 5 COO csi ' Yield extruded, machined 2! lira (j ) 83 ps ; Elongation extruded, machined ?co:

111 timate compression melded c /

'• r 3 v Vj J O 05 i , Yield compression molded 19 M?

(283 0 psi ) Ei ongation con press ton molded 35? * Izon impact strength all types D 255(A) ncncrea Deformation under load all types D 621(A) (7 MPa 2% defo •maticn aft

',1C 00 os i ) for 24 h) 33 -m in re :cver) Hardness all types 0 224 G (Shore D) 50

2-1 7
2.7.1 Constant Rate of Elongation

Uniaxial tensile stress-strain curves typical of UHMWPE are shown

in Figure 2.7.1. NMW LPE refers to normal molecular weight linear

polyethylene. UHMWPE materials in general have lower yield strengths

but higher ultimate tensile strengths than do NMW polyethylenes

In the post yield region UHMWPE does not exhibit the phenomenon of

macroscopic necking but does exhibit strai n-hardeni ng . The maximum

stretch ratio (A) attainable in UHMWPE is in the range 6-8, whereas

in NMW PE stretch ratios of 25 or greater can be attained.

5 - BA

Nominal strain (%)

Fig. 2.7.1 Uniaxial tensile nominal stress-strain curves for UHMWPE-A

([n] ^ 19.8 d t /g), UHMWPE-B ([ n ] £ 14.8 dc/g), and NMW PE (M £ 207,000).

w The data were obtained in accordance with ASTM D633. 1 psi = 6.895 x 10

3 Pa .

[i rom Bhateja, Polymer, 22 , 23 (1981). Reprinted by permission of Butterworoi

& c°. , Ltd. 0 ]. I 2-18
2.7.2 Impact Resistance

Impact resistance refers to the ability of a material to withstand flexural or tensile shock. The methods of measurement involve

the determination of the energy required to rupture specimens having

standardized geometries.

2.7.2. 1 Tensile Impact

Tensile impact can be determined using ASTM D1822,

Standard Method of Test for TENSILE-IMPACT ENERGY TO BREAK

PLASTICS AND ELECTRICAL INSULATING MATERIALS [5]. The tensile

impact strength of UHMWPE is reported to be in the range 2100-3000

kJ/m

2 [8,16].

2. 7.2.2 Izod Impact

Izod impact resistance can be determined using ASTM D256,

Standard Method of Test for IMPACT RESISTANCE OF PLASTICS AND

ELECTRICAL INSULATING MATERIALS [5]. Figure 2. 7. 2. 2 shows the

notched impact strength (V-notch) of UHMWPE as a function of

temperature [8,17].

200 |

IOl -200 -160 -120 -80 —40 0 40 80 120 °C

Fig. 2. 7. 2. 2 Notched Impact Strength of UHMWPE as a function of

temperature. (From reference 17)

2-19
2.7.3 Creep and Recovery

The ability of a material to resist creep and to recover

after deformation represents an important aspect of durability

and lifetime. In the case of semi crys tall ine polymers such as

UHMWPE the creep behavior can vary significantly depending upon

the time scale of the experiment, the magnitude of the applied

stress, and the temperature. The description of the creep and

recovery behavior is divided into the three following subsections:

(1) Short time creep and recovery at small deformations.

(2) Long time creep at small deformations.

(3) Creep at large deformations.

Short times refer to experiments in which the duration of the

creep is less than 10 seconds 1 day). Small deformations

refer to experiments in which the maximum creep strain is less

than about 3 percent.

2. 7. 3.1 Short Time Creep and Recovery at Small Deformations

Tensile creep and recovery curves for UHMWPE (M ^4x10^,

w p=0.935 g/cm^) are shown in Figure 2. 7. 3. 1.1 [21]. In this type

of experiment a constant tensile load is applied for a time t-j

at which time the load is removed and the strain during recovery

is determined as a function of time. The time t=0 for the

recovery corresponds to the time at which the load is removed.

The dashed lines represent the strain during recovery which is

predicted from viscoelasticity theory [20] assuming a superposition

principle of the form

e R (t) = £ c

( t+t ]) " £ c

( t ) . (1

)

2-20
where e^(t) is the strain during recovery at a time t after a

creep experiment of duration t-j , and e ( t+t-j ) and e (t) are

the strains at time t+t-j and t during a creep experiment.

Figure 2. 7. 3.1. 2 presents recovery data from a series of

experiments in which the applied stress is the same in each case,

but the duration of the creep, t^ , is varied. The dashed lines

have the same meaning as in the previous figure. While at the

very early times during recovery the predicted and experimental

values of strain are in good agreement, at the longer times

approaching t= t-j , significant deviations occur. The strain

during recovery does not continue to decrease toward zero as

predicted by equation (1), but tends to plateau to a finite

value. Even for very small deformations and short times under

load, UHMWPE exhibits "plasticity," i.e., the material retains

a permanent set after creep. Figure 2. 7. 3. 1.3 shows similar

creep and recovery data for UHMWPE at 38°C and an applied stress

of 4MPa. Creep curves at four different temperatures are shown

in Figure 2. 7. 3.1 .4.

2-21
Fig. 2. 7. 3. 1.1 Creep strain in uniaxial extension (open circles) and

strain during recovery (filled circles) versus time for UHMWPE

6 3

(M x4x!0 , p=0.935 g/cm ) at different levels of applied stress,

w

3

The duration of the creep step was 10 seconds and the temperature

23°C. The dashed lines represent the recovery predicted using equation (1).

(Data from reference 18)

: 2-22
Fig. 2. 7. 3. 1.2 Strain during recovery (after creep in uniaxial extension)

from experiments done at constant applied stress (8MPa) but varied creep

time, t, . The dashed lines represent the recovery predicted using equation (1).

JHMWPE (Mw

x4x 10^, p= 0.935 g/cm^) (Data from reference 18)

2-23
STRAIN

Figure 2. 7. 3. 1.3 Creep strain in uniaxial extension (filled circles)

and strain during recovery (open circles) versus tine for UHMWPE

(M x 4xl06

, p=0.935 g/cm3

) at constant applied stress (4MPa) but varied

w creep time, t-, . Temperature 33°C. (From reference 21)

2-24
STRAIN

Figure 2.7. 3.1 .4

p=0.935 g/crn^) at

O 47°C, and V

Creep curves in uniaxial extension for

four different temperatures . c^=4MPa.

62°C.

UHMWPE (Mw x4xl0 6

, A 23°C , 38°C

,

2-25
STRAIN

A comparison of the creep behavior of UHMWPE (M^ 4x1 0^) having

different thermal histories is given in Figure 2. 7. 3. 1.5. A significant

increase in the amount of creep results if the polymer is quenched

from the melt.

l

Fig. 2. 7. 3. 1.5 Comparison of creep behavior in uniaxial extension

for UHMWPE (M^4xl06

), 0 - quenched from the melt fc =0.923 g/cm

3 ), SC - w slowly cooled from the melt at £ 1°C per minute (o=0.935 g/cm"

5 ), and A -

annealed under nitrogen atmosphere at 130°C for 72 hours followed by

1 35°C for 48 hours. Applied stress 4MPa (Data from reference 19).

2-26
STRAIN

Figure 2. 7. 3. 1.6 provides a comparison of the creep behavior of

fi 3 UHMWPE (M £4x10 , p=0.935 g/cm ) in uniaxial extension and uniaxial

w compression. With the exception of the curves obtained at a stress of

15.6MPa no significant difference in behavior is observed between

extension and compression. At 15.6MPa the strain is as large as ten

3

percent at times approaching 10 seconds and the true stress is

significantly less in compression than in extension, even tough the

engineering stress is the same.

Fig. 2. 7. 3. 1.6 Comparison of creep behavior of UHMWPE in uniaxial

extension (open symbols) and uniaxial compression (filled symbols).

M ^4x10^, p= 0.935 g/cm . (Data from reference 21). w

2-27
2 . 7 . 3 . 2 Long time C ree p at Small Deformations

If the time under load in uniaxial extension is loneer

than about one day and the maximum creep strain remains small

(<5%), then the creep behavior deviates from that described

in Section 2. 7. 3.1. Figure 2. 7. 3. 2.1 illustrates this behavior

for UHMWPE in uniaxial extension. Similar data can be found

in reference [22] for both the extension and compression of

4

UHMWPE. A break occurs in the creep curve between 10 and

5 10 seconds beyond which the creep proceeds more slowly, and

on log-log coordinates the creep curve can be represented by

a straight line. The same behavior is observed in certain lower

molecular weight polyethyl enes and polypropyl ene [23,24].

Turner and Moore have reported [24] that the break point,

or “discontinuity" in the creep curve depends upon the processing

conditions in the following ways:

(1) The break point moves to longer times as the rate of cooling

from above the melting point decreases. This movement corresponds

to a horizontal shift only of the break time.

(2) The break point moves from one strain to a smaller one

in a vertical fashion as the storage time prior to testing

increases. The vertical shift is large for the quenched polymer

and undetectable for the fully annealed material.

(3) When plotted on log-log coordinates the creep curve

maintains the same shape for a range of material states and

therefore a range of coordinates of the break point.

2-28
(4) Beyond the break point the slope of the creep curve is

independent of the stress level and temperature for strains

below about 3 % in uniaxial extension and for temperatures

up to at least 60°C.

(5) The break point decreases in time with increased temperature

or applied stress.

From Figure 2. 7. 3. 2. 2 it can be seen that the same behavior occurs

in uniaxial compression as well. However, in the case of uniaxial

compression, the creep curves in the region from the break point up to

times of 600 hours or longer have the same slope for stresses at least

as large as 16 MPa and strains of 10 %.

Figure 2.7.32 .3 demonstrates that the break point shifts to shorten

times with increased temperature. An increase in temperature from23°C

to 64°C decreases the break point by approximately two decades in time.

2- 29
Nivais

0

m

c

a; +j

X

01

x

fO

c

cou

c

tn

ro

CTi oII

o

>)

s_

CL) >o

(_> a)

"Oc

rOCL

a>aj

s_

<_>

CMm

C\J

zn

2-30

at

23°C.

ck

=4MPa

and

t^

=

Duration

of

the

creep

step.

Creep

data

are

shown

only

for

t^

=

405

hours.

(From

reference

21).
CM'O

NlVdlS

r—

CM

a>uc c CD o i.

•r—

u

tO Ml/l CD

CD S_

s_Q.

(=

O

O s_ o u_ "—

r—

03

•r— • X 4-3 03 c • r- •r— sc o

CL c -X

•r— 03CD

--

—'• S_ co__ jD O CD '

JC

CTl 4-3 ID <4- CO O

CTi• C o o

II

•1

4-3 Q. 03u

CD o O r—

r—

X CD

4-> c5 03—

s: r— ^— Xo

LU s- Q_ Q_ 3 Q. S 03 3 0) -C

<+- +-> o

to CL CD CD 4—3 a> 03

CD CD •i— a

CCJ •-

CM I CO < • CD CM 0CO

CM

CD • i

4-> Ll_ 03

2-31
2-32
2. 7. 3. 3 Creep at Large Deformations

When the applied stress is larger than about 9 MPa

the creep behavior of UHMWPE in uniaxial extension changes in

character from that observed at smaller values of stress. Creep

curves typical of UHMWPE are shown in Figures 2. 7. 3. 3.1 and

2. 7. 3. 3. 2 for two different lots of polymer obtained from the

same resin producer. Characteri sti c of each curve is a region

in which the creep is very rapid followed by a plateau region

at large deformations where the creep proceeds very slowly. The

time range during which the rapid creep occurs depends upon the

magnitude of the applied stress. The maximum stretch ratio ( X)

attainable in these materials at 23 °C is from XM to W3

depending upon the magnitude of the applied stress and molecular

weight of the polymer. Similar creep curves for UHMWPE quenched

from the melt into ice water are shown in Figure 2. 7. 3. 3. 3.

The principal differences between these creep curves and the

creep curves for the slowly cooled polymer are that, for a given

applied stress, the quenched material experiences a much greater

deformation initially, and its maximum attainable stretch ratio

is less.

2- 33
STRAIN

Fig. 2. 7. 3. 3.1 Tensile creep of UHMWPE ( o=0 . 935 g/cm3

, M ^4xl0 6

) at

w various levels of applied stress at 23°C. The arrow indicates the time

of fracture for the specimen at a stress of 27.5MPa. (From reference 12).

2-34
10 2 10

3 10

4 10

5 10

8 10

7

TIME.s

Fig. 2. 7. 3. 3. 2 Tensile creep of UHMWPE (,-.=0.941 g/cirT, M 2x10^) at

w various levels of applied stress at 23°C. The arrow indicate the time

of fracture for the specimen at a stress of 27.5MPa. (From reference 11).

2- 35
Fig. 2. 7. 3. 3. 3 Tensile creep of quenched UHMWPE (°=0.923 g/cm^,

% 6

Mw ^4xl0 ) at various levels of applied stress. Temperature 23°C.

The arrow indicates the time of fracture for the specimen at a stress

of 27.5MPa. (From reference 12).

2- 36
2.7.4 stress_Rel axati on

A method commonly used to characterize the time

dependent mechanical behavior of polymeric materials is stress

relaxation. In this method the sample is subjected to a

constant deformation and the stress response is determined as

a function of time. By carrying out a series of tests at

different levels of strain, it is then possible to generate

a set of isochronal stress-strai n curves. Two such sets of

curves are shown in Figure 2. 7. 4.1 for different lots of UHMWPE.

In each case the sets of specimens were cut from compression

molded sheets which were cooled from the melt at approximatl ey

1°C per minute. Both materials exhibit nonlinear viscoelastic

behavior, even at the smallest values of strain for which data

are shown. At large strains (e> 0.10) the behavior of the UHMWPE

is not characteri Stic of most lower molecular weight polyethyl enes

,

which upon application of the step in strain undergo necking

and cold-drawing.

The effect of annealing on the isochronal stress-strai

n

behavior of UHMWPE is shown in Figure 2. 7. 4. 2. The isochrones

for the annealed polymer tend to flatten at the large strains

and during application of the largest step the annealed sample

necked. Increasing the crystal 1 i ni ty by only about 5

C

- can

significantly alter the isochronal stress-strai n behavior.

2-37
TRUE

STRESS

(MPa)

Figure 2. 7. 4.1 Isochrones of true stress versus strain for UHMWPE

(cooled from the melt at 1°C per minute) 0 - M 4x10^, •- M 2x10^.

w w

s - seconds (From reference 11).

2- 38
10

8

.001 0.1 0.1 1.0

STRAIN

Fig. 2. 7. 4. 2 Isochrones of true stress versus strain for UHMWPE ^ 6 o (M^xlO ). o - cooled from the melt at £ 1°C per minute ( O =0.934 g/cm 3 ),

• - annealed (p=0.942 g/cm ). (From reference 11).

2-39
2.7.5 Fa ti gue

Fatigue is generally studied under one of two conditions,

either at constant load (static fatigue) or with sinusoidal

loading. It is common practice to subject the test specimen to

a sinusoidal loading history, or strain history, of the form.

o(t) = P + Q sin wt,

where P is the mean level and Q sin wt is the oscillatory component

of the applied stress, or strain, w being 2 ?r times the test frequency.

Figures 2. 7. 5.1 and 2. 7. 5. 2 show fatigue data for UHMVJPE for

the case in which the specimens were subjected to tension only and

where P=Q. In Figure 2.7. 5.1 the logarithm of the time to failure

is plotted versus the peak stress (P+Q). For this type of loading

history the maximum strain in the specimen is not fixed, but rather

the specimen elongates during the course of the experiment and may

reach several times its original length before fracture occurs. Each

data point represents . the averaqe value of the loaarithms of the

failure times for a minimum of six specimens. The solid line represents

the results of tests carried out under a constant load P+0.' For the

given values of P and 0 and the range of freauency examined the lifetime under constant load defines a lower bound to the lifetime under

I

cyclic loading.

In Figure 2.7. 5.2 the logarithm of the time to failure is

plotted versus the test frequency for tests carried out at four

i different levels of peak stress (P+0). The dashed line represents

the behavior which would result if the lifetime were "cycle dependent".

By cycle dependent is meant that the failure occurs after a given

2-40
number of cycles independent of the frequency. The particular

case shown represents failure after 60 cycles. The line is not

unique, but will shift up or down depending upon the assumed

number of cycles. However the slope will remain independent of

the number of cycles assumed. For UHMWPE, the fatiuge lifetime

under conditions of sinusoidal loading is not cycle dependent.

2-41
Fig. 2. 7. 5.1 Log. time to failure versus log. engineering stress

for UHMWPE (Mw ~4xl0^, dens i ty=0 . 935 g/cm^). The solid line

corresponds to the behavior observed for specimens tested to

failure at a constant load equal to the peak engineering stress

under sinusoidal loading. (From reference 11).

2-42
106

"I 105

O

o

®

co u.oH

©

il 104

1031

10-4

45.0 MPa

10-3 10-2

Test Frequency, Hz

10-i

Fig. 2. 7. 5. 2 Log. time to failure versus log test frequency for

UHMWPE (Mw -4xl0^, density =0.935 g/cm^) tested at four levels of

peak engineering stress. The dashed line corresponds to the

behavior which would be observed if the lifetime were cycle dependent

where it has been assumed that failure occurred after 60 cycles.

(From reference 11).

2-43
2 .8 Wear

Wear represents an important aspect of materials character!' zation

and performance in applications such as orthopaedic devices, yet it

remains a most difficult phenomenon to quantify. In his review article

on wear (1982) entitled "WEAR-SCREENING AND JOINT SIMULATION STUDIES VS.

MATERIALS SELECTION AND PROSTHESIS DESIGN" Clarke [24] concludes:

It is disappointing that despite the voluminous literature in the areas

of materials wear- screening and joint wear-simulation studies 3 the data

are so fragmented and contradictory. Virtually every aspect concerning

the significance and effect of laboratory wear devices 3 test parameters 3

materials types3 and treatments on the resulting wear performance is

in dispute.

2.8.1 Types of Wear

Wear is generally classified into one of three types [25].

(1) abrasive wear - caused by hard asperities on the counterface or

by hard particles between the surfaces.

(2) fatigue wear - the detachment of material as a result of cyclic

stress variations on a localized scale.

(3) adhesive wear - transfer of material from one surface to the

other as a consequence of the forces of adhesion between them.

However, there can be considerable uncertainty in determining

the point at which abrasive wear, involving plastic deformation

and cutting of the polymer surface, becomes fatigue wear as the

topography of the different surfaces change with time.

2-44
2.8.2 Methods for Wear Testing

Laboratory wear testing of surgical implant UHMWPE is

done using either a wear-screeni ng machine or a joint simulator.

Such devices fall into one of three categories [25]. A category 1

wear-screeni ng device makes no special provisions to simulate

physiological conditions, and provides only a first-level comparison

of a materials' wear resistance.

Category 2 devices also make no attempt to reproduce a

prosthesis geometry nor provide simulation of the dynamic

joint-loading profiles. Parameters such as contact stress,

frequency, and stroke-amplitude are held to within a range

applicable to human joint performance. The three specimen

geometries most often used in wear-screening devices are the

disk-on-flat, annulus-on-flat, and pin-on-flat.

Category 3 devices are joint simulators which test. actual

prostheses with load and motion characteri s ti cs similar to the

natural joints behavior, and may use physiological fluids as a

1 ubri cant.

2.8.3 Wear Rate Measurements

Laboratory methods for determining wear in UHMWPE include

(1 ) the change in a particular specimen dimension, (2) specimen

weight loss, or (3) weight of collected wear debris [25] .

Most wear studies have been done using method (1). However, under

load UHMWPE experiences creep and upon unloading exhiDits both

a recoverable component and a plastic, or nonrecoverabl e

,

2-45
component of the creep. Creep complicates any test which

determines wear by a dimensional method since the methodology

must be able to separate the deformation due to wear from that

due to creep. The results of one study [33] (Table 2. 8. 3.1)

indicate that the non-recoverabl e creep component in UHMWPE

can be as large as the actual wear. At 6.4 MPa, 21 9 cm of

elastic (recoverabl e) creep and 17ym of plastic (non-recoverabl e)

creep were detected.

Table 2. 8. 3.1

COMPARISON OF CREEP AND WEAR

ESTIMATES FOR DISC-ON-FLAT WEAR

STUDIES OF POLYETHYLENE AT VARIOUS

Minimum

Contact Total creepk

STRESS LEVELS

0

Total wear Ra ti o stress (C ym) (W ym)W/C (%)

3.1 224 17 7.6

5.3 366 28 7.7

9.2 279 56 20

13.8 564 102 18

14.5 660 198 30

17.8 1219 279 23

a - From reference 33

b - Includes both the recoverable and nonrecoverabl e creep.

2- 46
Method (2) for assessing wear involves measuring the loss in

weignt of the specimen as wear progresses. This method can also

be subject to large errors unless the weight gain due to fluid

sorption is corrected for. In some cases the magnitude of the fluid

absorption weight gain has been found to be larger than the weight

loss from wear [34-37]. Wear data for UHMWPE are shown in

Figure 2.8.3. 1 where correction factors for bovine serum absorption

have been applied to pin-on-flat specimen geometries. In these

tests wear specimens and identical control specimens were presoaked

in serum for several weeks to minimize fluid sorption during the

wear test.

Method (3) in which the wear debris is collected and weighed is

used principally in studies involving cateaory 3 devices [joint simulators).

This method is subject to the limitations that (1) it cannot be used

with lubricants such as biological fluids because of contamination

by proteinaceous material, (2) rigorous washing and filtering of

debris to insure total particle entrapment is required, and (3)

artifacts may always be present due to fluid sorption by UHMWPE

particles and contamination by extraneous matter.

Figure 2.8.3. 1 Linear regression data for pin-on-flat wear-screening

tests. Data represented three polyethylene specimens run against 316

stainless-steel counterfaces with bovine serum lubrication. (From

references 38 and 39).

2- 47
2.8.4 Wear Rate Data

Quantitative comparison of polyethylene wear rates

is difficult because different types of polyethyl enes from

different manufacturers have been treated and evaluated under

widely different test conditions in a variety of wear test

machines. A representative sampling of different wear studies

reported in the literature is given in Table 2. 8. 4.1.

Table 2.8.4.

COMPARISON OF SOME OF THE COMPLEXITIES APPARENT IN TRYING TO COMPARE LA30RAT0RY WEAR PREDICTIONS FOR LHMu’PE

Polyethylene MWxlO Wear Stress range Study, [Ref.] type specific Wear machine technique Lubricant (MPa) 1

Seedhom et al . [27] a Pin-on-f lat D Synovial fluid 3.6

Gal ante and Restoker [23] Phillips 5095 Oisc-on-flat D Water 2. 3-6.

Hi FAX 1900 RCH 1000 Dumbleton et al . [29] HOP Hi FAX 0.1 Annulus-on-flat 0 PI asma 3.45

Hi FAX 1.25

Hi FAX 1000 3.0 Scales and Wright [3C] RCH 10C0

b Pin-on-f lat

c D Serum NS Dumbleton and Shen [31] Hi FAX 1900 3.0 Annulus-on-flat 0 Water 0.9-11.3

McKel lop et al . [32] Hi FAX 1900 Pin-on-flat W Serum 3.45-6.9

Note: 0, dimensional wear measurement; MS, not specified; W, weight loss technique of wear measurement.

a Various treatments of UHMW surgical grade high density polyethylene - all sterilized (2.5 Mrad)

b Two different suppliers,

c Pin of metal, flat of polyethylene.

(From reference 25.)

2- 48
Tables 2. 8. 4. 2 through 2. 8. 4. 6 present some results from

different wear studies concerned with the effects of minimum

contact stress, testing time, and molecular weight on the

wear of UHMWPE.

Table 2. 8. 4.

2

POLYETHYLENE WEAR DATA FROM DISC-ON-FLAT

WEAR TESTS

Minimum W vm

contact stress (wear/year) W(3.45 N/rnn

2

)

3.1 1.43 1.59

5.3 2.35 1 .53

9.2 4.71 1.77

13.8 8.57

14.5 16.64

17.8 23.44

Note: Test device = disc-on-flat; sliding distance = 127 km; years

effective use =11..9 years; contact stress = 3.1-17.8 M/mm2

; lubricant

water; temperature = 37°C; velocity = 1158 m.m/sec. (From reference 40

2-49
Table 2. 8. 4.

POLYETHYLENE WEAR DATA CALCULATED FROM

ANNULUS-ON- FLAT WEAR TESTS COMPARING

WEAR RATES AT 1 YEAR EFFECTIVE USE

Contact stress Wear

Test no. ( N/mnr ) (pm/year

1 1 .03 1.2

2 1.55 1 .2

3 2.24 3.3

4 2.76 4.9

5 3.79 7.3

6 4.48 7.7

7 5.93 10.1

8 7.07 9.1

9 9.17 10.0

10 11.20 CO

Note: Test device = annulus-on-flat; sliding distance = 10

years effective use = 1 year; contact stress = 1-11.2 N/mm”

water; temperature = room; velocity = 38.1 mm/sec. Also, W

2.19, r = 0.85. At P = 3.45, W = 5.1 um/year, at P = 6.9, W (From reference 31).

7 km;

lubricant =

= 0.845P +

= 8.0 m/year.

2 - 50
Table 2. 8. 4.

4

COMPARISON OF SH0RT-T0 LONG-TERM

POLYETHYLENE

TESTS OF

Contact Short-term Long-term Short- to Stress wear rate wear rate long-term

(N/mm^) (Hwi/year) (W^n/year) wear ratio

3.4 3.0 0.089 33.7

6.2 7.0 1 .98 3.5

Stress wear rate

ratio 2.3 22.2

Note: Test device = disc-on-flat; sliding distance = 25.4 km and greater;

2 years effective use = 2.4 and greater, contact stress = 3. 4-6. 2 N/mm ; lubricant = water; temperature = 37°C; velocity = not specified. (From

reference 40).

2-51
Table 2. 8. 4.

WEAR RATES OF POLYETHYLENE RELATED TO

MOLECULAR 'WEIGHT

Molecular weight

(X10

6

)

3.96

3.23

2.90

2.66

Variation = X 1 .5

Wear

(W um/year EU)

6.42

1.07

1.07

3.21

3.21

Variation = X 6

Note: Test device = disc-on-flat; sliding distance = 127 km; years

effective use = 11.87; contact stress = 3.4 N/mm2

; lubricant = water;

temperature = 37°C; velocity = not specified.

'Would not dissolve.

(From reference 40).

Table 2. 8. 4.

COMPARISON OF MACHINED POLYETHYLENE (A ,1 . 5X1

6 M0LWT)

AND MOULDED POLYETHYLENE (8, 3X1

6 MOL WT) FOR THREE

METAL ALLOYS

Polyethylene

(A:machined; No. Wear rate Wear

Counterface B -.moulded) specimens (Wum/year EU) ratio (%)

Ti -6-4 A 3 4.6 + 0.3 43

Ti-6-4 B 1 1.99

316 SS A 1 7.67

316 SS B 1 4.04 53 Co Cr A 1 7.67

Co Cr B 1 3.56 46

Note: Test device = pin-on-flat; sliding distance = not specified;

2 years effective use = not specified; contact stress = 3.4 N/mm ; lubricant = ringers; temperature = not specified; velocity = not

specified. (From reference 41).

2 - 52
The term "years effective use" is derived from the sliding

distance under the load phase of a walking cycle for a 22.2 mm

diameter hip. Based upon the results of several studies

[27, 32, 42] the sliding distance under load is arbitrarily set [25]

at about 10.7 km/year for a hip-flexion arc of approximately 55°. In

wear-screening studies, the wear rate is commonly reported in terms of

wear depth of the component "per year of equivalent use" with one

year representing 10.7 km sliding distance under load. Clarke [25]

has defined a "wear factor", or "wear index", given by

k =

V /LX,

where V is the wear volume, and L and X are the applied load and sliding

distance respectively. The measured volume of wear is normalized by

dividing by the sliding distance and load parameters. The unit micron/

year of effective use incorporates the 10.7 km sliding distance as

well as a nominal contact stress in the range 3.45 MPa.

2.9 Effects of High Energy Radiation

Sterilization of UHMWPE components in orthopedic prostheses

is done using high energy radiation, most comnonly gamma irradiation.

The typical dose is in the 2-4 Mrad range. In addition to sterilization,

high energy radiation also causes chemical crosslinking and/or

degradation of the polymer [43-45]. In the absence of oxygen, radiation

is said to produce a net crosslinking effect which occurs preferenti ally

in the amorphous regions [461. There is additional evidence that some

crosslinking occurs within the crystals, possibly limited to the fold

surfaces [ 47,48] . Crosslinking and/or degradation can alter the

mechanical behavior from that of the uni rr-?di ated polymer. Section 2.9

describes various aspects of the changes in physical and mechanical

behavior of UHMWPE brought about by high energy radiation

2-53
2.9.1 Density a nd Crystallinity

One effect of high energy radiation is to increase the

density, or percent crystal! ini ty , of the UHMWPE. Figures

2. 9. 1.1 and 2. 9. 1.2 indicate how the percent crystallinity and

density increase with increased irradiation dose.

Fig. 2. 9. 1.1 The degree of crystallinity, calculated from the density,

versus irradiation dose (gamma irradiation) for UHMWPE from two different

commercial sources. [From Roe, Grood, Shastri , Gosselin, and Noyes, J.

Biomed. Mat. Res., 1

, 209 (1981). Reprinted by permission of John Wiley

& Sons , Inc. 0 ]

2-54
0. 92- T

40

~i —r80 120 Dose (MRAD)

160 200

0. 91 -

Fig. 2. 9. 1.2 Density as a function of irradiation Dose (3

irradiation) for two different experimental samples of UHMWPE

(0,®) and normal molecular weight polyethylene (H) . [From

Bhateja, Andrews, and Young, J. Polymer Sci . , 2J_, Polymer Physics

Ed., 523 (1983). Reprinted by permission of John Wiley & Sons,

Inc. Q

]

2-55
Both sets of data indicate that the density increases

with dose for doses up to at least 10 Mrad. Two explanations

for the increase in density have been given; (1) highly

constrained tie molecules are preferential ly severed by the

irradiation, thus removing hinderance to further crystallization

[461; and (2) greater localized chain mobility occurs due to

energy absorption during irradiation, and the material tends to

reorganize in the direction of higher crystal 1 i ni ty [49 1. There

also is evidence that subsequent to the i rradi ati on , aging of

UHMWPE occurs, and the density continues to increase over a

period of many months [50 1.

2.9.2 Melting Behavior

The effect of high energy radiation on the melting behavior

of UHMWPE is less well defined than in the case of the density.

Bhateja et. al. [46 1 report that for an experimental sample of

UHMWPE, which was subjected to 6 radiation, an increase in the

temperature of the DSC melting peak of about 5-7°C occurred

over the first 16 Mrad of radiation. These results are shown in

Figure 2. 9. 2.1

2-56
Fig. 2. 9.2.1 Effect of 3 radiation dose on the DSC (differential

scanning calorimetry) peak melting temperature. e,« - UHMWPE, Qnormal molecular weight polyethylene. (From Bhateja, Andrews, and

Young, J. Polymer Sci . , 2J_, Polymer Physics Ed., 523 (1983).

Reprented by permission of John Wiley & Sons, Inc. £) .]

2-57
On the other hand, Roe et. al [49] compared the DSC

curves for two specimens of a commercial grade UHMWPE, which

were a irradiated to 0.0 and 10.0 Mrad respectively , and found

little or no change in position of the primary melting peak

centered around 137°C. A small secondary melting peak was

observed at around 120°C for the specimen irradiated to 10 Mrad.

2.9.3 Tensile Yield Stress

The tensile yield stress is defined as the peak engineering

stress occurring at tensile yield. Several studies have been

made of the effect of high energy radiation on the tensile yield

stress of UHMWPE. Data showing tensile yield stress as a function

of radiation dose, for doses up to 200 Mrad, are shown in Figure

2. 9. 3.1. The yield stress is found to increase with increased dose.

2-58
(a) y irradiated UHMWPE from two different commercial sources. Aging

time 6-24 months. [From Roe, Grood, Shastri, Gosselin, and Noyes, J.

Biomed. Mat. Res., _1_5, 2Q9 (1981). Reprinted by permission of John Wiley

& Sons, Inc. 0 ].

(b) b irradiated experimental UHMWPE. 9,* - UHMWPE, 9 - normal molecular

weight PE [ From Bhateja, Andrews, and Young, J. Polymer Sci., 21, Polymer

Physics Ed., 523 (1983). Reprinted by permission of John Wiley & Sons,

Inc. 0 ]

.

Fig. 2.9.3. 1 The effect of irradiation dose on the tensile yield stress

of UHMWPE

2-59
2.9.4 Tensile Yield Point Elongation

The effect of y irradiation on UHMWPE is to decrease

the elongation at yield [49]. Most of the reduction occurs during

the first 5 Mrad of irradiation (Figure 2. 9. 4.1).

2.9.5 Tensile Modulus

The effect of high energy irradiation on UHMWPE is to

increase the tensile modulus with increased dose. Data are

shown in Figure 2. 9. 5.1 for polymer subjected to fj - irradiation

[46] and y- irradiation [49].

Fig. 2. 9.4.1 Elongation at the yield point stress as a function of

irradiation dose for two different commercial UHMWPE polymers.

y - irradiation. [From Roe, Grood, Shastri , Gosselin, and Noyes, J. Biomed. Mat. Res., J_5, 209 (1981 ). Reprinted by permission of John

Wiley & Sons , Inc. Q

1

2- 60
Fig. 2

(a) 3 - irradiation, 0.5% secant modulus, 0,« - Two different

experimental samples of UHMWPE, 0 - normal molecular weight PE.

[From Bhateja, Andrews, and Young, J. Polymer Sci., 21_, Polymer

Physics Ed., 523 (1983). Represented by permission of John Wiley

& Sons, Inc. 0 ].

(b) y

- irradiation, commercial UHMWPE [From Roe, Grocd, Shastri

Gosselin, and Noyes, J. Biomed. Mat. Res., J_5, 209 (1981). Reprinted

by permission of John Wiley & Sons, Inc. 0].

9.5.1 Effect of radiation dose on the tensile modulus of UHMWPE.

2-61
2.9.6 Impact Strength

Du Plessis et al [53] report that the impact strength

of irradiated UHMWPE decreases significantly with an increase in

irradiation dose. Their data are shown in Figure 2.9.6. 1 for

specimens irradiated in the presence of three different environments.

2.9.7 Surface Hardness

Du Plessis et al [53] report that the surface hardness

(Type D-Shore) of UHMWPE increases when the polymer is irradiated

( Y rays) in the presence of crosslinking agents, but not so

irradiated in a nitrogen atmosphere. (Figure 2. 9. 7.1).

Fig. 2. 9. 6.1 Izod impact energy as a function- of dose for UHMWPE radiation

crosslinked (y rays) in the presence of crosslinking agents or nitrogen.

O-N^, X-CHrCH, and *-CH*CH + CTFE (chi orotri fl uoroethyl ene) . [From

du Plessis, Grobbelaar, and Marais, Radiat. Phys. Chem. , 9 , 647 (1977),

reprinted by permission of Pergamon Press Inc., c, ].

2-62
73

65 I 1 I l 0 20 40 60 80

Dose (Mrad)

Fig. 2.9.7. 1 Surface hardness of UHMWPE as a function of radiation

dose for UHMWPE radiation cross! inked (a rays) in the Dresence of

crosslinking agents or nitrogen, o-^, X-CH=CH, and » - CH=CH + CTFE.

[From du Plessis, Grobbelaar, and Marais, Radiat. Phys. Chem., 9_,

647 (1977), reprinted by permission of Pergamon Press Inc., cM.

2-63
2.9.8 Creep

Exposure to high energy radiation improves the resistance

of UHMWPE to creep, even at relatively low irradiation doses. An

example of the short time creep and recovery behavior at small

deformations in uniaxial extension is shown in Figure 2. 9.8.1 for

a commercial sample of UHMWPE (Mw-\,4xl0 and p=0.935 g/cm prior to

irradiation) subjected to relatively low doses of y

- irradiation.

3 The creep strain after 10 seconds decreases with increased dose.

A comparison of the room temperature creep behavior in uniaxial

extension is shown in Figure 2. 9. 8. 2 for a commercial UHMWPE polymer

both unirradiated and irradiated at three different doses. At the

longer times the creep strain is significantly less in more highly

irradiated samples.

Bhateja and Andrews [54] report a similar behavior for

samples of UHMWPE irradiated to doses of 3-irradiation as high as

64 Mrad. Their results are shown in Figure 2. 9.8. 3.

2-64
z

<

q:

H-f)

10

-3 10

10

-2

-3 10

10 10

TIME (s)

io

2 io

3

Fig. 2. 9.8.1 Effect of y irradiation on the creep and recovery behavior

in uniaxial extension of UHMWPE (Mw

~4xl0 6

, p=0.935 g/cm3

). Temperature

23°C, Engineering stress 2.5 and 4.0 MPa, length of creep step 10^ seconds.

0-0 Mrad, A-3.74 Mrad, and»*-6.28 Mrad dose. (From reference 21).

2-65
Fig. 2. 9. 8. 2. Room temperature uniaxial creep behavior of UHMWPE (M ^4x10 , w

3

p=0.935 g/cm ) irradiated with various does of Y

-i rradiation . Engineering

stress 4MPa. a - 0 Mrad, 0 - 1.18 Mrad, 7 - 3.74 Mrad, and - 6.28 Mrad.

( From reference 21 )

2-66
Fig. 2. 9. 8. 3 Creep strain (uniaxial extension) versus time for

UHMWPE at an engineering stress of 10.3MPa. A-OMrad, B-4Mrad,

C-16Mrad, and D-64Mrad. (p radiation) (From Bhateja and Andrews,

Polymer, 24, 160(1983). Reprinted by permission of Butterworth

Scientific, Ltd. 0 ].

2-67
2.9.9 Stress Relaxation

No data are available.

2.9.10 Fatigue

Only limited data are available on the effect of high

energy radiation on the fatigue behavior of UHMWPE. Table

2.9.10.1 shows some of the fatigue results of Nusbaum and Rose

[52]. The samples were cycled for the number of cycles indicated

and at the peak stress indicated. After 10

7 cycles at 1.7xl0

4

KPa,

several of the irradiated specimens showed a pronounced neck,

whereas unirradiated samples did not. They concluded that the

irradiation rendered the UHMWPE slightly weaker than the as-received

polymer, and where cyclic loading is concerned, irradiation may

aggravate creep under some conditions.

Table 2.9.10.1 Fatigue in Uniaxial Tension of

UHMWPE at 23 °C

2.5x10^ cycles in water at a 10

7 cycles in serum at a oeak

peak stress of 1 ,03xl0

4 kPa stress of 1 .72xl0

4 kPa Specimen As received Irradiated* As received Irradiated*

Yield Stress (XloVa)2.36 2.47

Strain 0.15 0.15 Lower Yield 2.22 2.1 2.19 2.05 Stress (XI 0

4 k?a

)

Strain 0.53 0.67 2.8 3.3

* 2.5 Mrad dose (S-radiation)

(From reference 52)

2-68
2.9.11 Effect of Radiation on Wear

A number of studies have examined the effect of radiation

on the wear properties of UHMWPE. Several of these are listed in

Table 2.9.11.1.

Table 2.9.11.1

WEAR-SCREENING STUDIES INVESTIGATING THE EFFECT OF RADIATION STERILIZATION

ON UHMW POLYETHYLENE 3

Study Polyethylene Sterilization Test device Lubrican

Dumbleton [29] Hi FAX 1900 20-1000 Mrad

(argon)

Annulus-on-flat Plasma

Scales [30] RCH 1000 2.5 Mrad

(ethylene oxide

autoclaved)

Pin-on-flat Serum

Ungethum [55] HOPE 0-15 Mrad

(air)

Pin-on-flat Serum

McKellop [56] Hi FAX 1900 0-7.5 Mrad

(air)

Pin-on-flat Serum

Grobbelaar [57] RCH 1000 0-30 Mrad

(acetylene)

Sand-slurry test

Jones [58] 0-5 Mrad Dry

(air)

a From reference 25

In Table 2.9.11.2 it can be seen that tne results of these

studies are quite varied. This may be a consequence of differences

in sterilization technique and wear test methodology. Tables

2.9.11.3-5 summarize the results of three different studies

concerned with the wear behavior of a - irradiated UHMWPE. Clarke

[25] concludes "It would appear that irradiation sterilization generally

increases polyethylene wear. At the levels generally used (2.5Mrad),

the likely increase is about 200 to 3000."

2-69
Table 2.9.11.2

SUMMARY OF IRRADIATION STERILIZATION EFFECTS ON POLYETHYLENE WEAR RATES

5

Steril ization Effect on we a r

Study range (Mrad) Decrease Increase

Dumbleton [29] 0-1000 a 3.6

Scales [301 0-2.5 2.1

Ungethum [55] 0-15 2-3

Jones [58] 0-5 None None

Grobbelaar [57] 0-80 0.3

McKellop [56] 0-7.5 8

a Wear increased 360% from 0-20 Mrad, returned to unsterilized wear rates

at 200 Mrad, and then increased 40% by 1000 Mrad.

5 From reference 25.

Table 2.9.11.3

SUMMARY OF WEAR DATA COMPARING WEAR RATES OF STERILE Hi FAX 1900

POLYETHYLENE

3

Wear/year EU Wear Overal

Treatment (W m) ratio vari ation

Unsterile 13.9 1.0 X2.5

20 Mrad 19.3 1.4

200 Mrad 5.4 0.4

1000 Mrad 7.5 0.5

Note: Test device = annulus-on-flat; sliding distance = 10.7 km; years

9 effective use = 1.0 year; contact stress = 3.4 N/mm"; lubricant =

serum; temperature = room; velocity =81.3 mm/sec.

a From reference 29.

2 - 70
Table 2.9.11.4

WEAR DATA FOR POLYETHYLENE DISCS (RCH 1000) SLIDING ON COBALT

CHROME (METAL PIN = 12.56 mm

2 CONTACT AREA

0

)

No. Wear rate Wear

Test Steril ization Treatment specimens (W pm/year) W(3. 45 N/mm) ratio

A Non sterile 7 130 + 20% 56 1.0

Nonsterile 3 4 142 + 3% 62 1.1

B 2.5 Mrad -- 274 t 9% 119 2.1

2.5 Mrad Boiled 14 days

b

2 189 82 1.4

2.5 Mrad Boiled 50 days 2 109 47 0.8

Note: Test device = pin-on-flat; sliding distance = not specified; years

effective use = not specified; contact stress = 7.96 N/mm

b

; lubricant = serum; temperature = 20°C; velocity = not specified.

a From different manufacturer.

b Boiled in saline solution.

°Froni reference 30

Table 2.9.11.5

Hi FAX 1900 WEAR DATA WHEN SLIDING ON STAINLESS STEEL

a

2 Stress N/mm

(c/min)

3.45(60)

6.9(100)

No. of

specimens

3

3

Wear

(W m/year)

0.36 + 15%

0.65 + 15%

Volumetric wear

(mm /year)

0.025

0.048

Note: Test device = pin-on-flat; sliding distance = 188 km; years

2 effective use = 17.6 years; contact stress = 3.45-6.9 N/mm ; lubricant =

serum; temperature = room; velocity = 60-100 c/min.

a From reference 38

2-71
References

[1] Annual Book of ASTM Standards published by the American Society For

Testing and Materials, 1916 Race St. Philadelphia, PA 19103, 1982

Edition, Part 46.

[2] G. B. McKenna, F. Khoury, and J. M. Crissman, "Relationship Between

Morphology and Mechanical Properties of Ultra High Molecular Weight

Polyethylene," First Annual Report to the FDA, Bureau of Medical

Devices, Task 80-01, NBS-BMD Interagency Agreement, October, 1980

(NBSIR 81 -2209FDA)

[31 G. W. Halldin and I. L. Kamel, Poly. Eng. & Sci . , 17_, No. 1 , 21 (1977).

[41 R. H. Marchessaul t, B. Fisa, and H. D. Chanzy, Crit. Rev. Macromol.

Sci. , 1, 315 (1972).

[51 Annual Book of ASTM Standards, Part 35.

[61 Annual Book of ASTM Standards, Part 9.

[7] K. C. Chen, E. J. Ellis, and A. Crugnola, SPE, 39th ANTEC, 270-272 (1981).

[8] For a discussion of the equilibrium melting point of polyethylene

see for example, J. D. Hoffman, G. T. Davis, and J. I. Lauritzen Jr.,

"The Rate of Crystal 1 izatian o:

? Linear Polymers with Chain Folding,"

Chapter 7 (see page 5o7), in Treatise on Solid State Chemistry Vo 1 . 3,

N. B. Kr.nrag Ed. « Plenum Press N.Y. (1 976).

[9] Luiletin, "Hostalen GUR UHMW Polymers", American Hoechst Corp.,

Somerville, New Jersey.

[10] Report No. DFG-18A, "Compression Molding Techniques," Hercules Inc.,

Wilmington, Delaware.

[11] J. M. Crissman, F. Khoury, and G. B. McKenna, "Relationship Between

Morphology and Mechanical Properties of Ultra High Molecular Weight

Polyethylene", Second Annual Report to the FDA Bureau of Medical Devices,

task 80-01, NBS-BMD Interagency Agreement, December, 1981 (NBSIR 82-2493 FDA).

[12] J. M. Crissman, G. B. McKenna, and F. Khoury, SPE, 40th ANTEC, 55-58 (1982).

[13] H. L. Wagner and J. G. Dillon, J. Biomed. Mater. Res., 1_3, 821 (1979).

[14] Report No. F-243, "Machining of 1900 UHMW Polymer Shapes," Hercules

Inc., Wilmington, Delaware.

[15] S. K. Bhateja, Polymer, 22_, 23 (1981 ).

[16] Brochure, 1900 UHMW Polymer, Hercules Inc., Wilmington, Delaware.

2-72
[17] J. Berzen, H. W. Birnkraut, and G. Braun, British Poly. Journal, 10,

281 (1978). ~

[18] L. J. Zapas and J. M. Crissman, Polymer, in Press.

[19] J. M. Crissman, L. J. Zapas, and F. Khoury, "Relationship Between

Morphology and Mechanical Properties of Ultra High Molecular Weight

Polyethylene", Third Annual Report to the FDA, Bureau of Medical

Devices, Task 80-01, NBS-BMD Interagency Agreement, March 1983 (NBSIR

83-2696 FDA).

[20] J. D. Ferry, Viscoelastic Properties of Polymers, John Wiley & Sons, Inc.,

New York, 1961.

[21] J. M. Crissman, "Mechanical Properties of Ultra High Molecular Weight

Polyethylene" , Fourth Annual Reoort to the FDA, Bureau of Medical

Devices, task 80-01, NBS-BMD Interagency Agreement, 1984 (NBSIR 84-?

FDA).

[22] J. Berzen and J. Theyssen, Kunststoffe, 68, 42 (1978).

[23] S. Turner, Poly. Eng. & Science, _6, 306 (1966).

[24] D. R. Moore and S. Turner, Plastics & Polymers, 42, 41 (1974).

[25] I. C. Clarke, Chemical Rubber Company (CRC) Critical Reviews in

Biomedical Engineering, 8, Issue 1, 29-91 (1982).

[26] J. K. Lancaster, Plast. Polym. , 44, No. 156, 297-306 (1973).

[27] B. Seedhom, D. Dowson, and V. Wright, Wear, 24, 35 (1973).

[28] J. 0. Galante and W. Rostoker, Acta Orthopaed. Scand., Suppl . 145

(1973).

[29] J. H. Dumbleton and C. Shen, Wear, 29, 163 (1974).

[30] J. T. Scales and K. W. J. Wright, Acta Orthop. Belg., 41_ (Suppl. 1),

160 (1975).

[31] J. H. Dumbleton and G. Shen, Wear, 37_, 279 (1976).

[32] H. Mekellop, I. C. Clarke, K. Markoff, and H. Amstutz, J. Biomed.

Mater. Res., ]_2, 895 (1978).

[33] W. Rostoker and J. 0. Galante, J. Biomed. Mat. Res., Y3_, 957 (1 979).

[34] I. Duff-Barclay and D. T. Spillman, Proc. Inst. Mech. Eng., 1 81 (3J),

90, London, (1966/1967).

2- 73
[35] H. C. Amstuiz, J. Biomed. Mat. Res., 3_, 547 (1968).

P 6 ] P. S. Walker, D. Dowson, M. Lonqfield, and V. Wright, Proc. Inst.

Mech. Eng., 1 81 (3J), 133-135, London, (1 967).

P 7 ] R. J. Pawluk and N. S. Eftekhar, Trans. Orthop. Res. Soc., 1,

56 (1976).

[3 8] H. McKellop, I. C. Clarke, K. Markolf, and H. Amstutz, J. Biomed.

Mat. Res. , 15, 619 (1981).

P 9 ] I. C. Clarke, P. J. McGuire, D. Hull, R. Okuda, and A. Sarmiento,

Trans. Orthop. Res. Soc.,'6_, 185 (1981 ).

[40] W. Rostoker and J. 0. Galante, J. Biomed. Mat. Res., ]_0, 303 (1976).

[4 1] D. Miller, R. Ainsworth, J. Dumbleton, E. H. Miller, and G. Shen,

Wear, 28, 207 (1974).

[42]

J. Charnley, Plastics in Medicine and Surgery , Plastics and Rubber

Institute, London, 3.T (1~975~H

[43] A. Charlesby, Atomic Radiation and Polymers, Pergamon Press, New York,

1960.

[44] F. A. Makhlis, Radiation Physics and Chemistry of Macromolecules , Wiley,

New York, 1975.

[45] M. Dole, Polym. Plast. Technol . Eng., 1_3, 41 (1979).

[46] S. K. Bhateja, E. H. Andrews, and R. J. Young, J. Polym. Sci . , 21. Polymer Physics Edition, 523-536 (1983).

[47] E. H. Andrews and I. G. Voi gt-Marti n , Proc. R. Soc. London Ser. A,

327 , 251 (1 972).

[48] R. Nagaeawa and K. Kobayashi , J. Appl . Phys., 41_, 4276 (1970).

[49] R. J. Roe, E. S. Grood, R. S. Shastri , C. A. Gosselin, and F. R. Noyes,

J. Biomed. Mat. Res., J_5, 209-230 (1 981 ).

[50l S. !(. Bhateja, Polymer, Commun., 23_, 654-655, ( 198- 2 ).

[51] C. T. Lue, E. J. Ellis, and A. Crugnola, 29th SPE ANTEC, Boston, MA,

246 (1981 ).

2-74
[52] N. J. Nusbaum, and R. M. Rose, J. Biomed. Mater. Res.,J_3, 557-576 (1979).

[53] T. A. du Plessis, C. J. Grobbelaar, and F. Marais, Radiat. Phys.

Chem. , 9, 647-652 (1977).

[5 4] S. i<. Bhateja and E. H. Andrews, Polymer, 24, 160-166 (1983).

[5 5] M. Ungethum, Artificial Hip and Knee Joint Technology , M. Schaldach

and D. Hohmann, Eds., Spri nger-Verl og , Basel, 491 (1 976).

[56] H. McKellop, G. Griffin, I. Clarke, and K. Markoff, Proc. 26th

Orthopaedic Research Society Meeting, Atlanta, GA, Feb. 5-7, 99 (1980).

[5 71 C. Grobbelaar, T. du Plessis, and F. Marais, J. Bone Jt. Surg.,

60-B, 370 (1978).

[581 W. Jones, W. Hady, and A Crugnola, NASA Technical Paper 1462,

May (1979).

2- 75

TABLE OF CONTENTS FOR SECTION 3

Subject

3. Poly(methyl methacrylate) Acrylic Bone Cement 3-1

3.1 General 3-1

3.2 Trade Names and Specifications 3-1

3.2.1 Trade Names 3-1

3.2.2 Specifications 3-3

3.2.3 Scope 3-3

3.2.4 Definitions 3-3

3.2.5 Physical Requirements 3-3

3. 2. 5.1 Liquid . . 3-3

3. 2. 5. 2 Powder 3-4

3. 2. 5. 3 Powder-Liquid Mixture 3-4

3. 2. 5. 4 Cured Polymer 3-4

3.2.6 Test Methods for Bone Cement Character!' zation

( ASTM Grade) 3-5

3.2.6. 1 Test Conditions 3-5

3. 2. 6. 2 Viscosity Change 3-5

3. 2. 6. 3 Doughing Time 3-5

3. 2.6.4 Maximum Temperature 3-6

3. 2. 6. 5 Setting Time 3-8

3. 2. 6. 6 Intrusion 3-8

3. 2. 6. 7 Compressive Strength 3-10

3. 2. 6. 8 Indentation and Recovery 3-11

3. 2. 6. 9 Water Sorption 3-11

3.2.6.10 Solubility 3-12

3-i
3.3 Polymer Properties 3-13

3.3.1 Raw Polymer 3-13

3. 3. 1 . 1 General 3-13

3. 3. 1.2 Liquid Monomer 3-13

3. 3. 1.3 Polymer Powder 3-14

3. 3. 1.4 Other Constituents 3-15

3.3.2 The Curing Cement 3-17

3.3.3 Cured Cement 3-19

3. 3. 3.1 General 3-19

3. 3. 3. 2 Physical and Chemical Properties 3-19

3. 3. 3. 3 Mechanical Properties of Cured Bone

Cement 3-27

3. 3. 3. 3.1 Static Properties 3-27

3. 3. 3. 3. 2 Fracture and Crack Growth

Characteri sti cs 3-33

3. 3. 3. 3. 3 Viscoelastic Behavior of

Acrylic Bone Cement 3-41

3.4 References 3-45

3-i i
3. Poly(methyl methacrylate) Acrylic Bone Cement

3.1 General

The poly(methyl methacrylate) (PMMA) used in orthopedics as a

bone cement consists essentially of a powder of methacrylate

polymer which is mixed with a liquid monomer. This mixture

forms a dough which reacts over a period of several minutes,

deDending on conditions, to form a hardened cement. The cement

thus formed is a sort of composite of a PMMA matrix material

(the polymerized liquid) filled with particles (the powder). The

resulting product, thus, does not necessarily exhibit the same

properties exhibited by commercially available poly(methyl

methacrylates) which are generally homogeneous, high molecular

weight polymers.

Certain of the materials labeled as PMMA bone cements are in

reality not oure methyl methacrylate homooolymers , but are copolymers.

A separate section will deal with this subject.

3.2 Trade Names and Specifications

3.2.1 Trade Names"'

Table 3.1 lists the trade names and major sources for

Acrylic bone Cements.

Certain commercial materials are identified in this reoort in order to

specify them adequately. In no case does such identification imply

recommendation or endorsement by the National Bureau of Standards.

3-1
Table 3.1

Trade Names and Manufacturers of PMMA

Bone Cements

Trade Name Manufacturer

CMW CMW Laboratories Limited

Bone Cement Division

Blackpool, FY4 40F

England

Palacos R Kulzer and Co.

GMBH

Bad Homburg

West Germany

Sul fiSulzer Brothers Limited

Medical Engineering Department

CH-8401

Winterthur

Swi tzerl and

Surgical Simplex P North Hill Plastics Limited

London N16

England

Distributed in U.S. by:

Howmedica, Inc.

Orthopedics Div.

359 Veterans Blvd.

Rutherford, N.J. 07070

Zimmer Bone Cement Zimmer USA

Warsaw, Indiana 46580

U.S. A.

Omni plastic Cintor Orthopedic

Division of Johnson & Johnson Products, Inc Braintree, MA. 02184

3-2
3.2.2 Specifications^

The principle specifications for acrylic bone cements

intended for use in surgical implant applications are

those contained in ASTM F - 45 1 - 7 6 Standard Soeci f i cati on for

ACRYLIC BONE CEMENTS [1 ]. The following are sections

excerpted from the ASTM specifications.

3.2.3 Scope

3.2. 3.1 The specification covers self curing resins

used primarily for the fixation of internal orthopedic

endoprostheses

.

3. 2. 3. 2 Units of premeasured powder and liquid are

suoplied in a form suitable for mixing to set "in vivo".

3. 2. 3. 3 While a variety of copolymers may be i ncorporated

,

the composition of the set cement shall contain

poly (methacryl ic acid esters) as the main ingredients.

3. 2. 3. 4 The specification covers compositional and

physical performance and packaging requirements but does

not cover toxicity or biocompatibility of the cement.

3.2.4 Definitions

3. 2. 4.1 Unit - one package or vial of preweighted

powder component or liquid component.

3.2.5 Physical Requirements (ASTM Grade)

3. 2. 5.1 Liquid:

Appearance - The liquid shall be clear and free

of deposits or sediment on visual examination in its

contai ner.

"* Reprinted, with permission from the Annual Book of ASTM

Standards. Copyright, ASTM, 1916 Race Street, Philadelphia,

PA 19103.

3-3
Stability - The liquid shall not increase in

viscosity to such an extent that its flow time is

increased more than 10?o when heated in the dark and

in its original container to 60±2°C for 48 hours.

Sterility - The liquid, as poured from its

container, shall pass the tests described in "Sterility

Tests - Liquids and Ointments" [ 2].

3 . 2 . 5 . 2 Powder :

Appearance - The powder shall be pourable and

free of extraneous materials such as dirt or lint.

Sterility - The powder, as poured from its

package, shall pass the tests described in "Sterility

Tests - Sol ids" [ 2 ]

3.2. 5. 3 Powder - Liquid Mixture: The material shall

exhibit the properties presented in Table 3.2.

Table 3.2

ASTM Specifications for Acrylic Bone Cement Powder-Liquid Mixture [1]

Maximum Dough Setting Time Maximum 0 Minimum

Time, minutes Range, minutes Exotherm, C Intrusion, mm

5.0 5 to 15 90 2.0

3. 2. 5. 4 Cured Polymer : The material, after setting,

shall have the properties presented in Table 3.3.

Table 3 .

Requirements for Cured Polymer After Settina [1]

Minimum Compressive Maximum Minimum Maximum Maximum Water 7 Strength, MPa Ide n tation ,mm Recovery ,% Water Sorption, S Sol ubi 1 i ty mg/cm

'

70 0.14 60 0.7 0.05

3-4
3.2.6 Test Methods for Bone Cement Characteri zati on ( ASTM Grade)

3. 2. 6.1 Maintain all equipment, mixing surfaces and

materials at 23±2°C at least 2 hours prior to testing and

conduct all tests at 23±1°C unless otherwise specified.

3. 2. 6. 2 Viscosity Change - Record the viscosity change

before and after the heating exposure by timing the flow

of the liquid level between the 0 and 5 ml marks of a

10 ml measuring pipet. Calculate the change as follows:

% change = x 100 (3.1)

where t

g

= flow time before heating, and tA

= flow

time after heating exposure for 48 hours in the dark

in a closed container at 60±2°C. An alternative method

for viscosity may be used.

3. 2.6.

3

Doughinq Time - Mix all the powder and liquid

of a single unit together as directed by the

manufacturer ' s instructions. Start a chronometer at

the onset of combining powder with the liquid and read

all subsequent times from this stop watch. Approximately

1.5 minutes after the onset of mixing, gently probe

the mixture with a surgically gloved finger. Take visual

notice as to the formation of fibers between the surface

of the mix and the finger as it leaves the surface.

Repeat this process of probing from that time at 15 second

3-5
intervals until the gloved finger separates cleanly.

Denote the time at which this is first observed as the

doughing time. Mix the mixture between determinations to

expose fresh material for each probing process.

Make two separate determinations of the doughing time.

The two values shall agree within 30 seconds of each other,

otherwise repeat the test on two additional units and

report the average of all runs.

3. 2. 6. 4 Maximum Temperature - Within 1 minute after

doughing time, gently pack approximately 25g of the dough

into the mold described in Fi gure 3. 1 . The mold shall be

made of polytetrafl uorethyl ene (PTFE) and be equipped with

a No. 24-gage wire thermocouple or similar device

positioned with its junction in the center of the mold

at a height of 3.0 mm in the internal cavity. Immediately

seat the plunger with a C-clamp or suitable press to

obtain the 6.0 mm specimen height. Upon producing plunger

seating remove the excess material and the C-clamp or

press for the remainder of the procedure. Continuously

record the temperature. Continue measuring the time from

the onset of mixing the powder with the liquid. Record

temperatures until cooling is observed. Report the maximum

temperature recorded to the nearest 1°C.

3-6
Fig. 3.1 Exothermic Heat Kold. (ASTM F451) [1]

3-7
The maximum temperature shall be the average of

two separate determinations reported to the nearest

1°C. If the difference between the two determinations

is greater than 5.0°C repeat the test on two additional

units and report the average of all four runs to the

nearest 1°C.

3.2. 6. 5 Setting Time - From the continuous time vs.

temperature recording from above, the setting time is

that time when the temperature of the polymerizing mass

is:

< Tmax

+ Tamb1ent )/? (3-2)

where Tmax

= maximum temperature

^ambient

= Ambient temperature of 23°+2°C

The setting time is reported to the nearest

15 seconds. Two separate determinations of setting

time are made and if the two values do not agree to

within 1 minute of each other, two additional runs are

to be made and the average of four determinations

reported as the setting time.

3. 2. 6. 6 Intrusion - The mold necessary for this test is

depicted in Figure 3.2. It will be made from PTFE. Mix

one unit of powder and liquid according to the

manufacturer ' s recommended procedure. When doughing is

achieved, gently pack the entire mixture flat into the

mold and insert the plunger. One minute after the

doughing time place a load of 49N (11 lb) on top of the

plunger for 1 minute. Then remove the weight and allow

the mixture to set.

3-8
100mm

Fig. 3.2 Intrusion Mold. (ASTM 451 ) [ 1

1

3-9
Following the set, remove the soecimen and measure

the average height of the intrusion into all four of the

1.0 mm diameter holes to the nearest 0.5 mm. Run this

test once. If the requirement is not met it must do so

in a repeat test.

3. 2. 6. 7 Compressive Strength - The test specimens shall

be cylinders 12 mm high and 6 mm in diameter. The ends

shall be flat and smooth and parallel to each other and

at right angles to the long axis of the cylinder.

The test cylinders are prepared five at a time by

placing five molds on a flat glass plate and slightly

overfilling by using one unit of mixed cement at the

commencement of dough time. Place a second flat glass

plate on top of the molds and hold the glass plates

and molds firmly together with a small C-clamp. The

ends of the specimens can be faced by polishing with 240

grit paper with water prior to removal from the mold.

The time lapse between the start of mixing and

the measurement of compressive strength testing shall

be 24±2 hours. The tests can be run on any universal

testing machine equipped to record load and deformation

time. Employ a crosshead speed of 20 or 25.4 mm/min.

Test the specimens without use of any type of pad

between the specimen and machine plattens. The failure

load shall be the load at the 2 % offset [ 2 % proof stress),

upper yield point, or at fracture, whichever occurs first.

The compressive strength is the failure load divided by

the original cross sectional area of the sample. It is

reported as the average for the five specimens to the

nearest 1 MPa.

3-10
3. 2. 6. 8 Indentation and Recovery - Determine indentation

using a Rockwell superficial hardness tester. Use the

disk from 3. 2.6.4. Remove any flash or other irregularity

which would prevent the disk from seating firmly on the

hardness tester. At 24±2 hours after the start of mixing,

subject the specimen in three separate areas to a minor

load of 29. 4N (6.7 lb.) employing a steel ball of 12.7mm

in diameter and adjust the dial gage on the hardness tester

to zero. Then bring the total load to 294 N (66.3 lb.),

for 10 minutes, at which time read the depth of the

indentation from the dial gage to the nearest 0.001 mm This is Reading A. Bring the load back to 29.4 N. After

10 minutes, read the dial gage and record the value as Reading B. Round the averages of Reading A and Reading B from the three areas to the nearest 0.01 mm Record

indentation as A (average). Calculate the percent

recovery as:

% Recovery = A(avg) - B(avg)

A(avg) ]

xl 00 (3.3)

3 - 2 '

6

- 9 Water Sorption - Perform the water sorption

test in duplicate on disks 50±1 mm diameter and 0. 5=0.1 mm

thick. The specimens may be made in a mold composed of a stainless steel ring 50 mm in diameter by 0.5 mm thick, mounted

between flat glass plates. Use sheets of cellophane or polyethylene between the mold and the glass plates. Place

the powder and liquid mixture into the mold 3 minutes after

the start of the mixing. Place a load of 98 N (22 lb.) on top of the glass plate for the duration of setting.

3-11
After setting remove the specimens from the mold and

lightly sand to dimensions using 240 and 400 grit

abrasive paper successively, backed by a hard, flat surface.

All surfaces of the disk will be smooth and top and

bottom will be flat.

Dry the disks in a dessicator containing thorouahly

dry anhydrous calcium sulfate (CaSO^) or silica gel

(freshly dried at 130°C) at 37° ± 2°C for 24 hours.

Remove to a similar dessicator at room temperature for

1 hour then weigh with a precision of 0.2 mg. Repeat

this cycle until the weight loss of each disk is not

more than 0.5 mg in any 24 hour period. This weight

shall be known as the conditioned mass, m^.

Immerse the conditioned disk in distilled water at

37±2°C for 7 days after which time remove the disks from

the water with tweezers, wipe with a clean, dry hand

towel until free from visible moisture, wave in the air

for 15 sec and weigh 1 minute after removal from the water.

The resulting weight shall be known as the mass after

immersion, m^. Calculate the water sorption as follows

for each disk:

v 2 Water Sorption = (mj-mQ)/40 cm (3.4)

Record the average of the determined values for two disks

2 to the nearest 0.01 mg/cm . Round the final value (average

2 of two determinations) to the nearest 0.1 mg/cm .

3.2.6.10 Solubility - After the final weighing in (3. 2. 6. 9),

recondition the disks to constant weight in the dessicator

3-12
at 37±2°C as done in 3. 2. 6. 9 above. This shall be known,

as the reconditioned mass, m

R

. Calculate the value for

solubility as follows:

mr m D Solubility = (3.5)

40 crri Round the final value (average of two determinations) to

2

the nearest 0.01 mg/cm .

3.3 Polymer Properties

3.3.1

Raw Polymer

3. 3. 1.1 General - The characteri zation of the raw

polymer must be considered in light of the means of

employment, i.e., as a medical material which will

be implanted into the body while it is undergoing

a chemical ( polymeri zati on ) reaction. The properties

of each of the two phases of this material prior to

handling by the surgeon will affect both its ultimate

properties as a bone cement and its handling

characteristics in the operating theater. In the

following DaragraDhs are discussed the major constituents

of bone cements. In the next section the mechanical

properties of the cement will be discussed.

3. 3. 1.2 Liquid Monomer - Bone cements are obtained

from various suppliers and as such the details of

composition can vary. However, the major constituents

of self-curing and non-cross! i nkabl e bone cements

consist of a liquid monomer and a polymeric oowder

[3-7] . The monomer is a clear, volatile liquid which

can polymarize spontaneously even at room temperature,

3-13
for which reason an inhibitor, usually hydroaui none , is

added for purposes of storage [3,4]. This tendency

to polymerize explains the need for the ASTM requirement

to measure viscosity changes upon heating the liquid

(see section 3.2). Also added to the monomer is an

initiator, often a tertiary amine such as

dimethyl paratol uidi ne , which has no effect until

it reacts with the activator in the powder.

3. 3. 1.3 Polymer Powder - The powder phase of the

cement is a prepolymeri zed polymethylmethacrylate

to which is added an activator, usually benzoyl

peroxide. Upon mixing monomer (liquid) with the

powder the cement cures to a solid mass which has

many of the properties of commercial PMMA, i.e. the

cement at body temperature is an amorphous, glassy

polymer.

The properties of the powder phase itself may

affect the way in which the cement handles, as well

as the ultimate properties of the cement. Particle

size, shape and distribution seems to vary with manufacturer's empirical determination of what combination

of handling characteri sti cs , cure character!' sti cs ,

Dorosity of the final cement, etc. represent the best

compromises for the final product. In Table 3.4 are

shown some typical properties of the bone cement powders.

Figure 3 shows a typical micrograDh of a bone cement

3-14
powder. In some bone cements the powder is a copolymer

of styrene and methylmethacrylate (See Table 3.5).

3. 3. 1.4 Other Constituents - There are two other

ingredients often used in the bone cement. Barium

sulfate is used as a radio-opaque additive to enable

the cement to be observed on medical x-ray radiographs.

Table 3.4

Some Properties of Powders Used in Commercial Bone Cements

Bone Cement Description of Powder

Powder Molecular

Wei qht Reference

CMW Spherical and comminuted powders

ranging in size from 5-75 ym

l.Ol x 10

5 [7]

Palacos R Spherical beads ranging in size

from 5 to lOO^m with clusters

of 10-15um size particles

4 . 5 9x 1

0

5 (a) [7]

Sul fix-6 Similar to Palacos R but with

fewer clusters.

1.19x1

0

5 (a) [7]

Surqical Simplex

P Large fraction of polymer

"flour" mixed with spheres of

similar size to the others

8.90xl0

4 (a) [7]

Surgical

Simplex P

. 3.2xl0

4 (b) [8 ]

Surgical Simplex P Spheres of varying sizes and 4.4xl04

- (b) [4]

irregular clusters of size 1.95x10;? (a) [4 ] 1000 to 100,000um. The most 1.98x10^ (b) [4]

frequent sphere s^ze was 25um

and the largest 40um.

(a) . Viscosity Average Molecular Weights

(b) . Number Average Molecular Weight, From GPC, polystyrene Calibration

3-15

Fig. 3.3 Scanning electron micrograph of powder component

of the cement. No difference was observed between

radiopaque (barium sulfate-containing) and

radiolucent (barium-free) powder. [4]

3-16

The barium sulfate containina powder is

indistinguishable from the pure one in the

microscope [4]. In addition, various antibiotics

can be added to the cement to help fight deep

infection [9,10]- Table 3.5 shows the ingredients

of typical bone cement powder and liquid components.

Table 3.5

Compositions of Powder and Liquid Phases of

Some Commercial PMMA Bone Cements L 7 j

Material

Powder

Powder Component Liauid Component wt %

CMW 40g + (2.5-5.0)g BaS04 20g 58-69

Palacos R 40g (may include 0.5g

gentamycin) with ZrC>2

20cc • 67-68

Sul fi x-6 40g PMMA with ZrC^ Ibg 71

Surgical

Simp! ex-P

40g '

5g D oly(methyl methacrylate)

)30g methyl methacryl ate-

| styrene copolymer

4g BaS04 20cc i

63

1 a 7 Methyl

Methacryl ate

J 2 . 6'o N, N-cimethyl -p-

| toluidine

75r 15ppm hydroqui none

3.3.2 The Curing Cement

The mixing of the powder and liquid commences a

chemical reaction of polymeri zati on of the liauid monomer.

As the cement cures it passes through a "dough" stage during

which time it can be handled by the surgeon and implanted.

The handling of the dough in the surgery environment can

introduce body fluids and air into the final cured cen’ent.

3-17
Thus, while most of the laboratory data for bone cement is

obtained on samples which have been carefully handled to

insure reproducible and controllable results, in order to

allow comparisons of behavior, the actual state of the cement

"in vivo" may be modified significantly from that obtained

in the laboratory.

During polymerization of the bone cement, a considerable

amount of heat is generated [5,13,14 ]. The exact geometry

of the cement (surface to volume ratio), the speed of

reaction, the surface boundary conditions (ambient temperature,

heat transfer coefficient at the surface), etc. will obviously

affect the peak temperature obtained. As seen in Table 3.5

these temperatures can be quite high and efforts have been

made to reduce the curing exotherm [13,14]. However,

the extent of the problem is controversi al and the cement

is used quite successfully in spite of it.

Table 3.6

Values of Peak Temperatures Obtained on Curing PHMA 3one Cement Hasses

Material Powder/Liquid Ratio Peak Temperaturea

,°C Reference

Surgical Simplex P 1/1

" 2/1

" 3/1

" 2/1

" 2/1

" 3/1

48° [13]

39° [13]

33 [ 13]

83-110° [14]

75-87° b

[ 5] 73 [ 5]

a The experimental procedures used in [13], [14] and [5 ] were auite

different. This is one of the problems in determining the clinical

importance of cure exotherm of PMMA bone cements.

^Deoended upon variables such as relative humidity, storaqe conditions,

presence of radio-opaque material, etc.

3-18
3.3.3 Cured Cement

3.3.3. 1 General - The cured cement is a two phase

system. It consists of the now polymerized monomer

filled with the particles of powder. Although the cement

is a glassy poly(methyl methacrylate) polymer, its

properties can differ appreciably from those of

commercially available acrylics. One reason is its

two phase nature, simply the existence of a particle/

matrix interface can affect the properties of the

material. In addition, the processing of the material

at ambient temperature and pressure and as a "dough"

results in considerable porosity and resultant change

in properties relative to those of the commerical

polymer. Finally, the bone cements tend to have

lower molecular weights than do the commercial materials.

This also affects the material properties.

3. 3. 3. 2 Physical and Chemical Properties - After cure,

the bone cement is primarily poly (methyl methacrylate)

with small residues from the initiatior, inhibitor and

activator materials. In addition, the agents added for

radio-opacity and antibiotic activity remain. In a

major study of one bone cement, for example, Brauer [4 ],

et al . looked at residual peroxide, residual monomer,

and residual hydroquinone after cure. They found

that the amount of residual substance changed with

time after cure, and "storage" environment.

3-19
In Table 3.7 are shown the results of storage

time on the residual peroxide in a cement (Simplex P)

which originally contained 2 % benzoyl peroxide in the

powder phase. As can be seen, length of time and storage

temperature affect the amount of peroxide remaining.

The peroxide appears to continue to react after "cure"

is complete. As the author's note [4], however,

the amount of peroxide remaining in the cement is

small and since the peroxides (either benzoyl peroxide

or polymeric peroxide) are insoluble in water, it is

unlikely that any appreciable amount of peroxide will

leach out of the cured cement under clinical conditions.

Of, perhaps, greater interest is the observation

that residual monomer remains in the cured cement. The

amount undoubtedly depends on the exact conditions of

cure, but Brauer, et al . [4 ] cite various literature

sources as finding monomer concentrations up to 21%

and that tissue concentrations adjacent to the cement

showed up to 1.2% cement. Brauer, et al [4 ] found

residual monomer contents of up to 33% during mixing

of the dough. This had decreased to only 3.3% after

one hour. Their results are summarized in Figure 3.4 and

Table 3.8.

3-20
Table 3.7

Residual Peroxide in Cured Bone Cement [4 ]

Peroxide/Resi

n

Length of Storage (g/100 g)

1 hr in air, room temp. 0.81

14 months in air, room temp. 0.76

21 days in water at 37°C 0.33

21 days in water at 37°C a 0.21

15 months in water at 37°C 0.14

a Radi opaque material containing barium sulfate.

Table 3.8

Gas-Chromatographic Analyses for Residual

Monomer Content of Cured PMMA Bone Cement [ 4

]

Monomer

Storage Time Concentration {%)

30 min 3.25

60 mi n 3.36

60 mina 2.83(3.04) e 4 hr 2.87

20 hr 2.70

13 days 2.30

45 days 2.48

49 days 2.21

57 days 3 2. 12(2.27) e 174 days b 2.60

207 days b 2.48

215 days 2.42

215 days b 2.26

136 days 0 1.44

3 months 3 1.08

a Bone cement with radiopaque filler.

^Samples mixed in 3:1 Dowder-l iquid ratio.

c In water at 37°C, air-dried for 30 min.

d In water at 37°C, dried 1.5 months under 10mm vaccum at 37°C.

e Monomer concentration based only on resin content of bone cement.

3-21
Fig. 3.4 Change of residual monomer in cured cement

with storage time in air at room temperature.

[4]

3-22
Of additional interest is the fact, which can

be seen from Figure 3.4 and Table 3.8 that the monomer

content of the cured cement decreases more if the

material is stored in water. The monomer is apparently

leached out in the water.

The amount of residual N, N-Dimethyl-p-toluidine

(DMPT) in the cured cement is of the order of 0.4%

[4] and there is no evidence that it leaches out

into .the environment after storage in water. Table 3.9

shows the amounts of hvdroquinone in the monomer,

powder and cured bone cement. The amount of residual

hydroquinine can be seen to be quite small.

The cured cement has a molecular weight similar

to, but slightly higher than that of the powder. In

Table 3.10 are summarized some molecular weight data for

the powder and cured cements. Also presented are data for

a commercial polymer. Figure 3.5 shows the molecular

weight distribution for the powder and for the cured

cement. It is not surprising, given their higher

molecular weights> that the industrial P MM

A

1 s have better

mechanical properties than do the bone cements. However,

for the rapid diffusion of monomer and swelling of powder

consistent with appropriate doughing times, a

comparatively low molecular weight poly(methyl

methacrylate) powder is, perhaps, a desirable ingredient

for bone cements [4]

.

3-23
Table 3.9

Results of Analysis for Hydroquinone in the

found

Monomer, Powder, ana Cured Cement [ 4 ].

Sample Analyzed HQ

a mg 0/

,0

20 ml methyl methacrylate + 1.50mg

H0 1.50 -

19.5 ml methyl methacrylate + 0.08ml

DMPTb + 1.00 mg HO 1.06 -

19.5 ml methyl methacrylate + 0.5ml

DMPT + 1.50 mg HQ 1.59 -

20 ml methyl methacrylate + 1.50mg HQ

+ 1.50 mg Q

c 3.10 -

Cured bone cement ( 41 . 20g

)

-d. Cured bone cement ( 35 . 61 g

)

0.53 13 x 10

4

0.17 5 x 10

4 Bone cement powder (40g), average of

A 3 runs 50.05 *51 X 10

4

a HQ = hydroquinone.

b DMPT = dimethyl -p-tol ui di ne.

C Q

= benzoquinone.

Table 3.10

Molecular Weights of Powder and Cured Materials

Material State ;_1 n

Simplex P Powder 44,000

Cured 51 ,000

CMW Powder - Cured

Palacos R Powder - Cured -

Sul fi x-6 Powder

Cured

Simplex P Powder - Cured -

Simplex P Powder 32,100

Cured 45,300

Cast PMMA Tube 546,000

(Industrial Grade)

Mv M1 'w Reference

135,000 198,000

[4] 203,000 242,000

101 ,000

150,000 - C7]

459,000 _ [7] 435,000 §

119.000

127.000 [7]

39.000

97.000

- [7]

-f

'

l

1

00 i

—

- 639,000 [12]

3-24
MOLECULAR WEIGHT M * 10

3

Fig. 3.5 Molecular weight distribution of (-X-) the

cement powder and (-0-) cured product. From

gel-permeation chromatoaraphi c measurements

of a 0.25^ solution of the polymer. Mobile

phase: tetrahydrofuran ; stationary phase:

Styragel . [4]

3-25
One of the major differences between bone cement

and commercial PMMA is the porosity of the bone cement.

The mixing of a very viscous "dough" introduces air

into the cement and the final product can have considerabl

porosity. The amount of porosity will depend

upon the exact mixing procedures as well as the powder/

liquid ratio, powder particle size, etc. In Table 3.11

are shown some typical cement porosities obtained by

Kusy [ 7] for several bone cements. Obviously there

is considerable variability in the data. The precise

significance of a relatively high pore volume is

difficult to judge since other factors also play a

role in the performance of a cement.

Table 3.11

Porosity of Two Phase Bone Cements [7 ]

Material Percent Porosity

CMW 1 .7±1 .2

Palacos R 8.0±3.0

Sul fix 6 1 . 6 ±1 . 5

Surgical Simplex P 5.0±4.1

• 3-26
3. 3. 3. 3 Mechanical Properties of Cured Bone Cement

3.3. 3. 3.1 Static Properties - The most

commonly measured properties of bone cement are

the quasi-static tensile and compressive strengths

and modulus of elasticity. The tests used to

produce values for these parameters are generally

run following accepted standardized tests for

rigid plastics such as ASTM D638 [15] ASTM

D695 [1 6 ] and ASTM D45[ 1].

As noted in the previous sections, the

physical and chemical properties of the cured

cement can depend upon the conditions of cure.

In addition, the cement changes spontaneously

after cure due to either continued curing [3 ]

or loss of monomer [5 ]. Also, in-vivo o>"

in simulated in-vivo environments the PMMA

bone cement takes up water which can also change

the properties [4,5].

The data in Table 3.12 show some typical properties of bone cement after cure and at ambient

conditions. Because the chemical and physical

properties of the bone cement are evolving with

*Since PMMA bone cement is a viscoelastic material, this is an apparent modulus value at some arbitrary strain rate.

3-27
Material

CMW

PALACOS R PALACOS RG' Sul fix-6

Surgical Simplex

Surgical Simplex

Surgical Simplex

Palacos R Surgical Simplex

Sul fi x-6

Palacos R Commercial PMMA

Fransverse Tensile Stn

time after cure, it can be expected that the

mechanical properties will also evolve. Haas,

et al. [5 ] performed indentation and recovery

tests (see section 3.2) on cured bone cement and

observed changes which varied with temperatures

of molding and of ageing (storage after cure)

Their results are shown in Figure 3.6. These

results correlated well with monomer loss in

the cement after curing [4 ].

Table 3.12

Typical Mechanical Properties for PMMA. Bone Cements at Room Temperature

Modulus of Tensile Compressive

Elasticity (GPa) Strength (MPa) Strength (MPA) Reference

2.76 42.7 - [7]

2.62 46.2 - [7]

2.55 46.2 -

[7]

2.41 48.3 - [7]

2.55 33.8 - [7]

2.30 56.

a 93.1 [5]

- - 80.7 [18]

- , •- 82.0 [9]

2.4 - 90 [19]

2.6 - 91 [19]

2.4 - 87 [19]

2.83 73.8 [7]

American Dental Association Specification No. 12 [17]

3-28
RECOVERY

- %

INDENTATION

-

^zm

Fig. 3. 6 Average curves for indentation and recovery of bone

cement. Curve 1 is an average of individual curves

for three specimens, curve 2 for two specimens, and curve 3 for four specimens, using the indicated mold

and aging temperature. [5]

3-29
After curing, the bone cement absorbs water

if it is placed in an aqueous environment [7 ].

This water abosrption can adversely affect the

mechanical properties of the bone cement.

For example, Rostoker, et al , [20] have

tested the flexural strength of bone cements

after up to 26 months implantation in rabbits.

Their results are shown in Table 3.13. Although

there was much scatter in the data (large

standard deviations) Rostoker, et al . concluded

that the degradation after 26 months was

significant.

Table 3.13

Effect of Implantation Time on Bending

Strength of PMMA Bone Cement [ 2QT

Time After Molding Environment Failure Stress *

1 day

2 months

12 months

26 months

ai r "in vivo

"in vivo

"in vivo

48.3:3.8

43.8±4.9

50.9:5.7

46.6:6.0

* Based on Maximum "Elastic" Stress in Beam in 3 ooint bending.

Comparison of test results for PMMA bone

cements made at varying strain rates and

temperatures must be made with caution since

PMMA is a viscoelastic material. Figure 3.7

shows the effect of strain rate on some

properties and Figure 3.8 shows the effect of

temperature.

3-30
70

Fig. 3.7 Effect of Strain Rate on the Mechanical Prooerties

of PMMA Bone Cement. [19]

, 3-31
Fig. 3.8 Effect of Temperature on the Mechanical Properties

of PMMA Bone Cement. a o =0.1% Proof Stress, oq = Compressive Strength. E=Young's Modulus. [19]

3-32
3. 3. 3. 3. 2 Fracture and Crack Growth Character!

- sties

Failure of bone cements has been treated using

the concept of Fracture Mechanics*. In this, now

widely accepted model, it is assumed that

failure in the material initiates at an "inherent"

flaw when the stress reaches some critical

value, a . The inherent flaw is thought of as

a crack of length and can be treated

stati sti cal ly

.

In static tests, there are two parameters

which are often used to describe the resistance

of the material to crack propagation: (a) The

fracture energy, y, required to create new

surface or (b) the fracture toughness, Kjq.

Under plane strain conditions the relationship

between failure stress, a-p, and y is [21]:

a f

= [

2 E(l-v2)C, -r (3.6)

where E is the Young's modulus, C

Q

is the

inherent flaw size and v is Poisson's

ratio (~0.35 for PMMA). Thus, the higher y

the greater the resistance to crack propagation.

* The reader is referred to reference [21] for a detailed discussion

of fracture mechanics and its application to polymers.

'3-33
Similarly the failure stress, Of, in terms of

the fracture toughness, Kjq, (in plane strain

Mode I conditions) is [2 1 ]

_

(

K IC X °f"(-p ) Q

2 c 0

(3.7)

where Q is a parameter depending on test

geometry. In an elastic material , K

2

_ IC t i 2\

Y £ (1 -V )

. (3.8)

As shown in Table 3.14, bone cement when

tested in air at ambient temperatures has a

lower fracture toughness, K

j

q

,

than does

commercial PMMA. However, as Freitag and Cannon

[22] have shown in their report on fracture

toughness of bone cements, there is considerable

variability in the reported values for Kjq

in the literature. Thus, the literature values

cited by them range from 0.6 to 1.73 times the

value they measured on a commercial PMMA. Kusy

[ 7] measured fracture energies for bone cements

and commercial PMMA and found that the values

for the base cement fell within the range of

values he measured on two different samples of

industrial material. His results are shown in

Table 3.15. The reasons, then, that the bone

cements have lower tensile strengths than do

commercial PMMA lie in the values of the "inherent"

3-34
flaw size, also shown in Table 3.15. Because of

either its two phase nature or the porosity of

the material the bone cements have inherent flaws

which are 5 to 10 times the size of those of

commercial PMMA*.

Table 3.14

Fracture Toughness, Kjq, of Bene Cement

Tested in Air at Ambient lemperature

Material <IC> WN m

' 3/2

Surgical Simplex P 1.04

Zimmer Cement 1.24

Commercial PMMA 1.26

Surgical Simplex P 1.0

Commercial PMMA 1 .

7

Surgical Simplex P 1.6

Table 3.15

Fracture Parameters for PMMA Cements at

Zl °C in air [ 7 IT

Reference

[ 22 ] [ 22 ] [ 22 ] [23]

[23]

[24]

_r o Mean Inherent Flaw

Material Fracture Energy, yxlO~ 5 (erg/cni ) Sixe, C n xlCr (mm)

CMW 2.4 + 0.6 25

Palacos R 2.9 + 0.8 25

Sulfix 6 3.4 X 1.0 25

Surgical Simplex P 2.6 ± 0.6 37

Commercial PMMA 1.3

3.9

X

X 0.3

1 .8 4.8

We note here that based on the values in Table 3.15, the fracture

mechanics analysis does not agree exactly with the inherent flaw

concept and fracture mechanics. However the idea is useful for

analysi s

.

3-35
Other variables can affect changes in the

bone cement fracture behavior. One widely

studied one being the pressure of fabrication.

Thus, Freitag and Cannon [22], Beaumont and

Young [23] and Si h and Berman [24] have found

improvements in Kj

c with increasing fabrication

pressure. However, given the limitations of the

operating theater, it is unlikely that increasing

fabrication pressures to decrease porosity is

likely to be a practical method of improving

bone cements.

An important consideration in the fracture

of bone cements is the effect of environment

on the fracture toughness. Simply testing bone

cement in water results in an increase in Kjq

[ 23]. Similar results are found for testing

in Bovine Serum [22]. Both these increases in

toughness may be due to plasticization at the

crack tip by the fluid. However, the major

interest is the effects of environment on the

material after long times of exposure to the

envi ronment.

Kusy [7] has measured the fracture energy

and inherent flaw size of bone cements after 10

months in distilled water. Results are presented

3-36
in Table 3.16. In comparino these results with

those of Table 3.15 it must be noted that the two

sets of tests were carried out at different

temperatures . Literature values reported by

Kusy [7], however, show little effect of

temperature changes from ambient to 37°C on

either y or C

Q

and it does, therefore, appear

that the ageing of the cement in aqueous

environment leads to little change in fracture

toughness, but large increases in C 0 and,

therefore, decreases in tensile strength.

Table 3.16

Fracture Parameters for Pf-'MA Cements

after Immersion in Distilled Water [71*

Material KjQ(MNm ~^) yx10~ ^ (ergs/cm^) C 0 xlO^(mri)

CMW - 34

Palacos R 83

Sulfix 6 98

Surgical Simplex P - 160

"'After 10 months immersion at 37°C. Tests carried out at 37°C in air.

3-37
The fatigue behavior of bone cement has

been little characterized. Figure 3.9 shows

some typical S-N behavior for a bone cement

after work by Freitag and Cannon [25].

Similar results were obtained by Pilliar, et al

[26]. Table 3.17 indicates that the lifetime in

fatigue is greater in Bovine serum than in air.

The conclusions drawn on differing bone cements

in reference [25] are disputed in reference [27].

Table 3.17

Comparison of Fatigue Lifetime of PMriA Bone Cement in Bovine Serum and in air

Tt""T7'"C T'25]V

~ - * Test Environment Fatigue Lifetime, Cycles

Air 2.7 x 10

5 Bovine Serum 3.8 x 10^

*Rotating Beam Fatigue. Stress level not reported.

A problem in characteri zi ng fatigue behavior

of polymers is their viscoelastic nature. Thus,

at high frequencies thermal buildup may cause

premature failure. In addition, total fatigue

lifetime is frequency sensitive in such a

fashion that prediction of failure time or cycles

to failure at 1 hz from data obtained at 35 hz,

for example, can lead to an incorrect estimate of

3-38
STRESS,

psi

lifetime. Figure 3.10 shows how test frequency

affects the fatigue lifetime for a commercial

PMMA [28]. Discussions of this problem can be

found in references [29] and [30].

Fig. 3.9 Life comparison of Simplex-P hone cement fabricated at pressures of

(A) 5, () 25, and (O) 50 psi and tested in air at 22°C. [25]

3-39
TEST FREQUENCY, Kz

Fig. 3.10 Effect of Test Frequency on the Lifetime of

Commerical PMMA Subjected to Zero-Tension

Sinusoidal Fatigue Loading. [28]

3-40
3. 3. 3. 3. 3. Viscoelastic Behavior of Acrylic

Bone Cements - PMMA is a material which shows

non-linear viscoelastic behavior [12]. Therefore, it should be expected that PMMA bone

cements exhibit similar behavior. The effects

of various fabrication and cement variables on

the viscoelastic properties of the cured cements

have been little studied. The most complete

work to date is that of Chwirut [31], although

other studies have been made [32,33]. Figures

3.11 and 3.12 show compressive creep data for

various bone cements at 37°C and in a normal

saline environment. It is of interest to note

that (a) in Figure 3.11, the creep rate is

approximately the same for all stress levels,

except at the highest stress. Chwirut [31]

found that the data fit well to an exponential

creep function, e=f(c)t

n

; (b) in Figure 3.12, the

amount of creep for different cements varies

quite significantly from .006 to .011 for the

unreinforced materials. This could be significant

and may be due to differences in porosity,

molecular weights or particle sizes of the cements.

Certainly more work is required to fully

characterize the behavior of the cements.

-41
An interesting finding in reference [31]

is that, at long times and stresses which are

in the range to be expected for implanted bone

cements, mi crocracki ng occurs in the cements.

This is important and may be related to "in

vivo" failure of cements.

3-42
CREEP

STRAIN

TIME, hours

Fig. 3.11 Creep of Zimmer Regular Bone Cement at 37°C in Normal

Saline Solution. [31]

3-43
TIME, hours

Fig. 3.12 Compressive Creep of Bone Cement at 37°C in

Normal Saline Solution. Stress =12.1 MPa.

A. Zimmer Regular Cement. B. Zimmer Low Viscosity

Cement. C. Zimmer Carbon Fiber Reinforced Cement.

D. Omniplastic Cement. E. Surqical Simolex P. [31]

3-44
1

• Annual Book of ASTM STANDARDS published by the American Society

for Testing and Materials, 1916 Race Street, Philadelphia, PA.

19103, 1982 edition, Part 46.

2. U.S. Pharmacopeia, Revision 19, Mack Publishing Co., Easton, Pa.,

1975, pp 593-599.

3. D. F. Williams and R. Roaf, Implants in Surgery , W. B. Saunders

Company, Ltd, Philadelphia (1973).

4. G. M. Brauer, D. J. Termini and G. Dickson, "Analysis of the

Ingredients and Determination of the Residual Components of

Acrylic Bone Cements", J. Biomed. Mater. Res., ]]_, 577-607 (1977).

5. S. S. Haas, G. M. Brauer and G. Dickson, "A Characteri zati on of

Polymethylmethacryl ate Bone Cement", J. Bone Joint Sura., 57-A,

380-391 (1975).

6. S. S. Haas, G. Dickson and G. M. Brauer, "A Proposed Specification

for Acrylic Bone Cement", J. Biomed. Mater. Res. Symposium, _6,

105-117 (1975).

7. R. P. Kusy "Characterization of Self-Curing Acrylic Bone Cements",

J. Biomed. Mater. Res., J_2, 271 -305 ( 1978).

8. S. C. Bayne, E. P. Lautenschl ager , C. L. Compere and R. Wildes,

"Degree of Polymerization of Acrylic Bone Cement", J. Biomed. Mater.

Res'. , 9, 27-34 (1975).

9. E. P. Lautenschlager, G. W. Marshall, K. E. Marks, J. Schwartz and

C. L. Nelson, "Mechanical Strength of Acrylic Bone Cements

ImDregnated with Antibiotics", J. Biomed. Mater. Res., J_0 , 837-845

(1976).

10. R. C. Nelson, R. 0. Hoffman and T. A. Burton, "The Effect of Antibiotic

Additions on the Mechanical Properties of Acrylic Cement", J. Biomed.

Mater. Res., 1_2, 473-490 (1978).

11. A. B. Welch, "Antibiotics in Acrylic Bone Cement. In Vivo Studies",

J. Biomed. Mater. Res., VZ, 843-855 (1978).

12. G. B. McKenna and L. J. Zapas "Non-linear Viscoelastic Behavior of

Poly(methyl methacrylate) in Torsion", J. Rheology, _23, 1 51 -1 66 (1979).

13. R. C. Turner, P. E. Atkins, M. A. Ackley and J. B. Park, "Molecular

and Macroscopic Properties of PMMA Bone Cement: Free Radical

Generation and Temperature Change Versus Mixing Ratio", J. Biomed.

Mater. Research, 1_5 , 425-432 (1981 ).

14. H. B. Lee and D. T. Turner, "Temperature Control of a Bone Cement by

Addition of a Crystalline Monomer", J. Biomed. Mater. Res., TJ_,

671-676 (1977).

3-45
15 . ASTM 0638-76 "Standard Test Method for Tensile Properties of

Plastics", Annual Book of ASTM Standards, published by the

American Society for Testing and Materials, Philadelphia, (1977)

part 35.

16. ASTM D695-69 "Compressive Properties of Rigid Plastics" Annual

Book of ASTM Standards, published by the American Society for

Testing and Materials, Philadelphia, (1977), part 35.

17. American Dental Association, Guide to Dental Materials and Devices , Ed. 6, p. 201, American' Dental Assn., Chicago, 1972.

18. E. P. Lautenschlager, J. J. Jacobs, G. W. Marshall and P. R. Meyer,

Jr. "Mechanical Prooerties of Bone Cements Containing Large Doses of

Antibiotic Powders", J. Biomed. Mater. Res., ]_0, 929-938 (1976).

19. A.J.C. Lee, R. S. M. Ling and S. S. Vangala, "The Mechanical

Properties of Bone Cements", J. Med. Eng. Techno!., 2_, 137-140,

(1977).

20. W. Rostoker, P. Lereim and J. 0. Galante, "Effect of 'In Vivo'

Environment on the Strength of Bone Cement", J. Biomed. Mater.

Res. , 1_3 , 365-370 (1979).

21. A. J. Kinloch and R. J. Young, Fracture Behavior of Polymers , Applied Science, New York, 1983.

22. T. A. Freitag and S. L. Cannon, "Fracture Characteri sties of

Acrylic Bone Cements", J. Biomed Mater. Res., J_0, 805-828 (1976).

23. P.W.R. Beaumont and R. J. Young "Slow Crack Growth in Acrylic Bone

Cement", J. Biomed Mater. Res., 9_, 423-439 (1975).

24. G. C. Sih and A. T. Berman, "Fracture Toughness Concept Applied to

methyl methacrylate", J. Biomed. Mate. Res., _1_4, 31 1 -324 (1980).

25. T. A. Freitag and S. L. Cannon "Fracture Characteri sti cs of Acrylic

Bone Cements. II. Fatigue", J. Biomed. Mater. Res., JJ_, 609-624

(1977).

26. R. M. Pill i a r , R. Blackwell, I. MacNab and H.U. Cameron, "Carbon

Fiber Reinforced Bone Cement in Orthopedic Surgery", J. Biomed.

Mater. Res., ]_0, 893-906 (1976).

27. C. F. Stark, "Fracture and Fatigue Characteri sti cs of Bone Cements",

J. Biomed. Mater. Res., 1_3 , 339-342 (1979).

28. G. B. McKenna and R. W. Penn, "Time Dependent Failure in Poly(methyl

methacrylate) and Polyethylene", Polymer, 2]_ , 213-220 (1 980).

29. G. B. McKenna and R. W. Penn, "Time Dependent Failure of a Polyolefin

Rubber Candidate Material for Blood Pump Applications", J. Biomed.

Mater. Res., 14, 689-703 (1980).

3-46
.30. G. B. McKenna, J. M. Crissman and F. A. Khoury, "The Relationship

Between Morphology and Mechanical Properties of Ultra High Molecular

Weight Polyethylene", Second Annual Report to the Bureau of Medical

Devices, Food and Drug Administration, NBS-BMD Interagency

Agreement, Task No. 80.01, October, 1981.

31. D. J. Chwirut, "Long-Term Compressive Creep Deformation and Damage

in Acrylic Bone Cements", J. Biomed. Mater. Res., In Press (1983).

32. R. W. Treharne and N. Brown, "Factors Influencing the Creep Behavior

of Poly(methyl methacrylate) Cements", J. Biomed. Mater. Res.

Symposium, 6 , 81-88 (1975).

33. S. Pal and S. Saha, "Stress Relaxation and Creep Behavior of Normal

and Carbon Fibre Reinforced Acrylic Bone Cement", Biomaterials, 3_, 93-96 (1982).

3-47

NBS-114A I RE V. 2-8C ) U.S. DEPT. OF COMM.

BIBLIOGRAPHIC DATA

SHEET (See instructions)

4. TITLE AND SUBTITLE

1. PUBLICATION OR

REPORT NO.

NB5IR 84-282U

2. Performing Organ. Report No.l 3. Publication Date January 1 y84

Information on Polymeric Materials Used in Orthopedic Devices

5. AUTHOR(S)

J. M. Crissman and G. B. McKenna

6. PERFORMING ORGANIZATION (If joint or other than NBS. see in struction s) 7. Contract/Grant No. NATIONAL BUREAU OF STANDARDS

DEPARTMENT OF COMMERCE 8. Type of Report & Period Covered

WASHINGTON, D.C. 20234

9. SPONSORING ORGANIZATION NAME AND COMPLETE ADDRESS (Street. City. State. ZIP)

Food and Drug Administration

Bureau of Medical Devices

8257 Georgia Ave., Silver Spring, MD 20910

10. SUPPLEMENTARY NOTES

E2J 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

bi bl iography or literature survey, mention it here)

This report provides information on the two polymeric mate’rials most commonly

used in the fabrication of orthopedic implants. The work was done as part of

Task 80-01 NBS-FDA/BMD Interagency Agreement'. The two materials described are ultra high molecular weight polyethylene UHMWPE and polymethylmethacrylate

(PMMA) bone cement. The report contains information on such subjects as specifications (ASTM), raw materials characteri zation , processing, morphology,

mechanical properties, and wear.

12. KEY WORDS (Six to twelve entries; alphabetical order; capitalize only proper names; and separate key words by semi colon s) Acrylic; bone cement; mechanical properties; morphology; orthopedic implant;

poly(methyl methacrylate); processing; ultra high molecular weight polyethylene.

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