NBSIR 85-3166

Studies of Porous Metal Coated

Surgical Implants

A. C. Fraker1

, A. W. Ruff1

, A. C. Van Orden1

, H. Hahn2

, A. J. Bailey 1 and C. D. Olson 1

’U.S. DEPARTMENT OF COMMERCE

National Bureau of Standards

Gaithersburg, MD 20899

June 1985

Prepared for: FDA Contract 22 4795023

QC

100

. U56

85-3166

1985

NATIONAL bureau

OF STANDARDS

LIBRARY

NBSIR 85-3166

STUDIES OF POROUS METAL COATED

SURGICAL IMPLANTS

A. C. Fraker 1

, A. W. Ruff1

, A. C. Van Orden1

, H. Hahn2

, A. J. Bailey1 and C. D. Olson 1

'U.S. DEPARTMENT OF COMMERCE

National Bureau of Standards

Gaithersburg, MD 20899

2 Artech Corporation

Falls Church, VA 22042

June 1985

Prepared for: FDA Contract 22 4795023

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

NATIONAL BUREAU OF STANDARDS. Ernest Ambler. Director

Contents

I. Brief Literature Review of Metal Porous Coatings

Anna C. Fraker

II. Corrosion Behavior of Sintered Porous Coated Co-Cr-Mo

Surgical Implant Metal

Ann C. Van Orden and Anna C. Fraker

III. Corrosion-Fatigue Behavior of Arc Plasma Sprayed Ti or Ti-6A1-4V Coatings on Ti-6A1-4V Substrates

Anna C. Fraker, Henry Hahn, Alison J. Bailey and Craig

D. Olson

Page

1

21

38

IV. Acknowledgements 50

Note

:

The use of materials and equipment in these studies does not in

endorsement by NBS.
'

'' " ' • *' t :•• . • ' * . , •> • ' ' „ ./ „ *
Part I

Brief Literature Review of Metal Porous Coatings

Porous Metal Surgical Implants

Introduction

This report is on the topic of porous metal coated metal surgical implants

which are in use or clinical study. This is a rapidly developing technology, and

the introduction of additional types of porous coated metals or metal combinations

can be expected. Porous coated implants may be rough surfaces, a single layer of

spheres, fibers, etc., or porous coatings may consist of several layers. The pore size

probably should not be less than lOOum, and the porosity should be interconnecting.

The most used porous coatings at present include those which are sintered on to the

substrate^, those sintered together and mechanically held to the substrate 2 and coatings

applied by the arc plasma spraying process^. These coatings can be subject to phase

changes due to sintering, oxidation and contamination. Application of the porous

coating should not affect the mechanical strength of the prosthesis. Descriptions

of these porous materials should include chemical compositions of substrate and

porous coating, mechanical properties and all processing parameters and temperatures

involved in the production of the prosthesis. This review will present some background

information relating to the need for improved orthopedic materials, some comments on

internal fixation of devices, and an overview of porous metals used in orthopedic

devices.

1
Background Information Relating to Need

Two major factors dictate the need for more reliable implants: the

average life span has increased and the number of younger people receiving

bone implants has increased. The necessity of the continued development of

surgical implant materials can be illustrated with the following statistics. In

1976, approximately 120,000 total joints were implanted in the United States.

This included 80,000 hips, 30,000 knees, 8,000 fingers, and 2,000 other joints.

4

An Arthur D. Little report 5 in 1982 gave annual figures for total hips of 100,000,

total knees of 50,000 and other joints of 12,000.

The need for increased studies of various aspects of total hip replacement

was highlighted by the National Institutes of Health Consensus Development

Conference which was held March 1-3, 1982 at the Masur Auditorium on the NIH

campus. 6 At this meeting it was stated that approximately 100,000 total hip

replacements are made in the United States each year. Each total hip repacement

costs around $9,500.^ Any complications or revisions can result in multiples

of this figure. Earlier estimates of the patient cost for failed implants were

100 million dollars per year. 6 Since many of the patients receiving total joints

are over 65, much of this cost is borne directly by the Federal Government through

Medicare. The increased pain and problems for the patient are beyond quantification

Nevertheless, the prognosis for total hip surgery is good, and a conservative figure

for the success rate is 90 percent. The longer term success rate going beyond 10

years is less well known.

Studies of groups of patients have given figures of 12 percent for prosthesis

loosening over a period of 8 years. 8 These data were reported from The Hip Society

by Richard J. Johnston, M.D., who emphasized the success of the current mode

of total hip surgery and predicted a 75 percent success rate at the end of

20 years. He cautioned against abandoning this successful procedure in favor of

2
other procedures. Incidences of loosening after 10 years as indicated radiographically have been reported to be as high as 11.3 percent for the acetabular

component and 29.9 percent for the femoral component. 9 This radiographic

evidence did not indicate prosthesis failure or the need for revision but it did

show that the prosthesis fixation was somewhat weakened. Indications for

success in the total hip surgery are at such a high level that it seems feasible

to consider that many more of the current failures can be eliminated. Failures

associated with prosthesis loosening originate with the bone/bone-cement

interface and with the metal/bone cement interface or a combination of both.

Osteroarthritis is the major cause of joint replacement and mainly affects

older people. Treatment of joint disease by replacement with artifical materials

has become a highly successful mode of treatment for thousands of people suffering

from disabling joint disease. Total joint replacement is a highly successful

procedure, giving relief of pain and improved function to many people. In the next

twenty years, the increased number of people coupled with the increasing mean age of

the population could produce a potential patient population well above 180,000 persons.

The cost, patient needs and the large number of joint replacements justify biomaterials development.

3
Some Methods of Prosthesis Fixation

Surgical prosthesis fixation in the bone has been a subject of concern and study

for many years. Prosthesis fixation has been accomplished by impaction of the prosthesis

into the medullary cavity of the bone and by mechanical means such as using screws

10,11,12. In the late 1950 ' s , Charnley13 used poly (methyl methacryl ate) cement, and

this method of fixation is credited for the high success rate of joint prostheses

since that time. However, each of these methods has problems and disadvantages which

lead to the failure of the arthoplasty.

Devices impacted into the medullary cavity are/were held in place by compressive

stress action. Stress concentrations can occur which result in overloading the implant

material, and failure occurs. A small degree of motion can adversely affect bone growth

A decrease in bone growth leads to loss of attachment and loosening of the device.

Mechanical internal fixation has been used as one of the methods of attachment of

the acetabular portion of the toal hip prosthesis. Fixation with screws should

not be used if indications are that loosening would be expected. This method has

been used successfully^ but less so than the cemented prosthesis.

Fixation with cement suffers from problems related to the nature of the

cement^. Poly (methyl methacrylate) polymerizes in situ and secures the prosthesis

in place. Heat develops during polymerization and may damage surrounding bone

tissue. A toxic monomer may be released from the cement and this raises doubts

about the use of cement from a biological viewpoint3

. Some evidence indicates

that delayed failures in cemented total hip prostheses may result from poor stability

resulting from relative motion between the acrylic cement and the bone^Jf).

Other disadvantages include cement cracking due to voids, brittleness and loss

of creep resistance. Improvements in fixation are being sought through means of

boney ingrowth or cement attachment to a porous coated prosthesis.

4
Overview of Porous Metals for Bony Ingrowth

Pore Size and Bone Growth , A comprehensive review of porous metal implants

has been given by Spector^who discusses production methods, histology, mechanical

properties and clinical trials. Spector noted that no adverse tissue growth into

porous titanium, Co-Cr-Mo or 316L stainless steel had been observed and tissue

ingrowth occurred in the pore size range of lOOum to 400um. A Porocoat (Trade

name of a DePuy product) brochure^ 8 which contains seventy five references

pertaining to porous metal implants, states that the best pore size range is lOOum

to 300um. Pilliar^, in a review dealing with porous ingrowth and processing of

Co-Cr-Mo orthopedic implants, reports that ingrowth occurs in animals over a

pore size range of 50um to 400um. Earlier work by Hulbert, et al.^O indicated

tissue ingrowth into ceramic implants with pore sizes ranging from 45um to 150um

and with more rapid growth in the larger pore size. A recent study^l of the effect

of pore size on peel strength of tissue used pore sizes ranging from 50um to

200um. It was recommended on the basis of this study that the average pore size

should not drop below 90um and that pore sizes greater than 200um might be better for

ingrowth. Porous sintered Co-Cr-Mo substrates with porosity less than lOOum

have exhibited loosening^ 8

.

Hungerford, et al.,22»23 indicated that the optimum pore size for implantation in humans would be 400um to 500um and reported that cementless knee prostheses

with pore sizes of 400um were successful. Nine knees were implanted in osteoarthritis patients and seven knees were implanted in rheumatoid arthritis

patients. Scores showing increased joint usage of the osteoarthritis group ranged fron

42 pre-operati ve to 92 post-operative and scores for the rheumatoid arthritis grvj;

ranged from 39 before to 86 after surgery.

5
Healthy bone ingrowth for fixation of the prosthesis is needed as soon as

possible following surgery. One way to increase this ingrowth is by electrical

stimulation 2 ^ Prosthesis stabilization so that minimal or no movement occurs between

the tissue and implant is required for bony ingrowth. Electrical stimulation is more

effective in the early stages of bone growth and should be studied as a meansof assisting with initial stabilization. Published proceedings from a symposium on Bioelectric

Effects on Bone and Cartilage25 and from a symposium on Electrically Induced Osteogenesis 25 provide more details of research which has been conducted on this subject.

Another technique which has been used to stabilize prostheses and to assist in

bony ingrowth is the use of bioceramics in the form of calcium phosphates. Hydroxyapatite, Ca^PO^OH, has been used as a filler material and as a coating to stimulate

bone growth. Hydroxyapatite increases the amount of bone growth during the first four

weeks after surgery but after that the bone grows at the same rate with or without the

hydroxyapatite. Whitlockite or beta tricalcium phosphate (TCP), Ca3 (P04)2, is another

material which is considered for treatment of bone lesions and as filler material^5 *^

.

TCP is biogradable and in particulate form can be incorporated into the bone but some

adverse reactions at residual TCP sites were found in animal studies. TCP needs

further testing and evaluation before it is put into use. Brown and Chow30 have

i have developed a dicalcium phosphate-tetracalcium phosphate mixture, which when com-

||

bined with water, sets by a process involving formation of hydroxyapati te. Its

setting properties and its apparent non-resorbabil i ty should give it advantages over

beta-TCP for some applications. More information on the fundamental chemistry of

calcium phosphate and numerous references to their use may be found in a book entitled

Bioceramics of Calcium Phosphates" edited by Klaas de Groot3 ^

.

It is unclear from reported studies^ 5 whether smaller or larger pores result

in a stronger bone/porous-surface interface bond. Most studies indicate increased

6
strength with increasing pore size within a specified size range. Push-out test

data show shear strength values ranging from 14.5 MPa to 25.5 MPa^2. Predecki,

et al.33 conducted studies of bone growth into screw type interfaces. Harris, et a 1 34

used three screws to rigidly place porous coated (Co-Cr-Mo sintered spheres) acetabular

components in canine hips. Bone growth covered 53 percent of the implant and no

resorption was observed at the screws.

Some types of porous or irregular coatings are described in the following paragraphs, and many of these are in clinical trials today. The Judet^ 5 ^

hip

prosthesis has an irregular cast surface with circular cavities for bone attachment,

but it is nonporous. This method appears to be useful for fixation, but some stem

fractures have occurred and more data on this system are needed^?. Over 2000 cases

cases have been reported with this type of fixation.

Another type of porous prosthesis is that developed by Lord 3 ? in which tiny spheres

about 1mm in diameter are cast on the surface about 0.5mm apart. Bone attachment occurs

with this nonporous surface. This coating is known as the madreporic porous coating.

Porous titanium implants have been successfully produced by a power metal

sintering process to produce a void-metal composite3 ^. Tests of this material

in animals have shown healthy bone ingrowth with high shear strength at the bone/

metal interface. The mechanical properties and elastic modulus in compression

for the void-metal composite are closer to those of living bone. One consideration

with this type of porous implant is that voids reduce the strength of the prosthesis.

No reference to clinical use of this material was found.

7
Wire mesh and fiber metal titanium coating on a Ti -6 A1 -4 V substrate have been

studied 39 . These coatings are made of 50um to lOOum diameter wire. Pressure sintering of the wire mesh permits the use of sintering temperatures of 800° to 900°C

which are below the beta titanium transformation temperature of 975°C for the T i -6A1-4V

used in the study. There was some loss of strength in these coatings after sintering.

Coatings of 316L stainless steel were made and bone tissue grew equally well into

both materials. Lining the inside of the pores of the Ti-6A1-4V fiber metal coating

with a 20um layer of hydroxyapatite^? enhanced bone ingrowth initially by a factor

of four, but this was a short term effect lasting only a few weeks. Interfacial

decohesion between bone and implant was increased by a factor of two. Porous coatings

on Ti - 6 A1 -4 V have been made by sintering spheres of Ti or Ti-6A1-4V on to the

Ti-6A1-4V substrate 40,41

> but sintering above the beta titanium transformation

temperature results in an elongated alpha titanium microstructure and a reduction

in fatigue strength. This loss in fatigue strength also was attributed to stress

raising effects in the porous surface and at the particle/substrate interface.

One of the first studies of porous metal surgical implants was made by

Hirschhorn and Reynolds^ w ho compacted Co-Cr alloy powders of alloy HS-21

(the castable alloy) and HS-25 (the wrought material.) Hirshhorn and Reynolds

demonstrated the feasibility of using porous implants by obtaining bone growth into the

implant and by attaining a lower elastic modulus for the porous material. The

recommended porosity was 30 percent for an alloy such as HS-25.

Today, a widely used sintered porous Co-Cr-Mo coated Co-Cr-Mo implant is made

by high temperature sintering of the spheres onto the surface^. Metal powders or

spheres made from material of the same composition as the substrate are produced

by the gas atomization process or the spun electrode process. Resulting irregular

('

i

I

8
shapes and/or spheres have a fine cast structure prior to sintering. These

particles may be held to the substrate by means of an inorganic binder. The

binder could be a source of contamination, but this is not necessarily the case,

depending on the nature of the binder. Sintering at temperatures in the

1200°C to 1300°C range results in interparticle or particle/substrate connections

or necks which form probably by a diffusion process. The microstructure within the

spheres changes to large grains similar to that of the substrate with second phase

regions present in both coating and substrate. The shear strength of this porous

coating exceeds 21 MPa43. Kilner44 discusses melting of the interdendritic phase

(a eutectic consisting of carbide, sigma and cobalt rich phases) and the resulting

detrimental effects on the mechanical properties, and suggests that very slow cooling

from the sintering temperature to 1235°C helps to prevent long continuous carbide phase

formations. This slow cooling may not be practical in actual production. However,

due to the presence of these phases and the possible detrimental effects on mechanical

properties, the processing temperatures, cooling procedures and composition

are important aspects of porous coated Co-Cr-Mo implants. It is important for

the porous coating to have the same composition as the substrate and if not, to

identify phase compositions and other factors which would adversely affect the strength

or corrosion behavior of the porous coated implant.

Fiber metal implants 2,45 are ma(j e 0 f p Ure j-j fibers, 0.19mm in diameter and 6.35mm

long and are kinked and pressed into the needed shape prior to sintering. Fibers

also are cut in lengths ranging from 1.3 to 3.8cm. The piece is sintered in

vacuum at 1070°C to 1240°C for one to four hours. This piece can be pressed a

certain amount after sintering to correct for any shape changes. Machining is not

recommended as it tends to close surface porosity. The fiber metal part shows

good ingrowth and high bone-implant interfacial strength. This fiber metal part

9
may be used as a bone segment or the fiber metal pressed shape can be fitted mechanically into a shaped area of a hip stem or other prosthesis to provide for attachment

to a solid core. This fitting may be followed by sintering. Sintering at temperatures

above the beta transus should be avoided and care should be taken at any sintering

temperature not to introduce oxygen into the material since detrimental effects on

the fatigue life and other mechanical properties of the prosthesis could result

This type of porous coating has been used in clinical trials and data should

be available soon on the result or actual use of this material. Co-Cr-Mo and

316L stainless steel wires^ have been studied for use as fiber metal parts, but

titanium has been the material of choice for extensive investigation.

The arc plasma sprayed porous metal coating 3,46 on a metal substrate is being

used clinically for dental implants, for backs of acetabular cups, and on various

areas of femoral stems. This coating usually is Ti or T i -6A1 -4V applied to a

Ti-6A1-4V substrate. Other materials such as Co-Cr-Mo also have been applied to

metal substrate using this process. The coating is applied by spraying metal

powders in size ranges of 200um to lOOum or other specified size using a carrier

gas such as argon. The plasma flame may have a temperature of 15,000°C and the

gas speed is 3000 m/sec. Prior to the application of the coating, the substrate

is grit blasted with A1 2 O 3 to produce a roughened surface. This grit blasting

is a possible source of contamination since some A1 2 O3 particles or other grit

blasting material may become embedded in the surface and not be removed in the

cleaning process which follows. Porous surfaces applied in this manner form strong

bonds to the substrate and may be applied to produce a wide range of porosities.

Mechanical properties, including fatigue life, are not adversely affected by the

application of this porous coating. The Ti-6A1-4V material must not be sintered

above the beta titanium transformation temperature if desirable mechanical properties

of the T i -6A1 -4V substrate are to be maintained. Indications are that this coating

10
does not need to be sintered.

Corrosion

The biocompatibil ities of the materials discussed in this paper have been studied

by Laing 47 * 48 and others, and the referenced investigations of the porous materials

indicate that no extensive unusual cellular behavior in the bony ingrowths was observed.

Questions remain about metal ion release and short and long term effects.

Initial studies of this topic have been made by Black4 ^, Woodman88 ’ 8 ^ and others

to show that ions are present and are carried through the body. Metal ions are

released by the process of corrosion. Corrosion studies of porous Ti -6A1 -4

V

fiber metal implants by Lautenschl ager , et al. 88

, and of sintered Co-Cr-Mo spheres

by Cameron et al 88 indicated no changes in the electrochemical behavior except

for increased corrosion currents which resulted from the increased surface area.

The rate of reaction per unit of surface area was the same for the porous and the solid

implants. Other corrosion studies84 of sintered Co-Cr-Mo spheres support these

earlier findings. There still are considerations regarding creviced areas

and variations in the composition and location of second phases in the implant

which could affect the corrosion behavior. Preliminary studies 88 on some other

porous materials indicate that contamination can occur during the production of

the implants which affects the corrosion resistance of the materials.

Clinical Results

Some results of clinical studies are being reported. Engh 88 reports on 26

patients who were monitored over a five year period. These patients received the

fully porous coated Moore prosthesis. Engh reports the short term success of

this prosthesis and technique to be about the same as that for the cemented hip

prostheses. He noted that bone remodelling in these cases had been a continuous

process and no stress shielding or bone resorption had occurred. The important <•

11
of the porous surface for increased bony ingrowth and three dimensional interlocking

was emphasized as well as the design of the Moore prosthesis which fits tightly in

the bone. No infection occurred and no prostheses were removed from patients in

this study group. All reports on stabilization of the implanted device by bony

ingrowth stress the importance of avoiding micromotion of the device. The design

of the device, the fit of the device and the rate of bony ingrowth are important

factors which contribute to the degree of success of the prosthesis.

Lord and Bancel 5 ^ gave a seven year report on 1509 cementless total hips with

madreporic prostheses. The prosthesis stem includes a single layer of spheres cast

together with the prosthesis. This coating is not porous but has an irregular

surface for bony attacnment. The acetabular component is a threaded metal back

which screws into place, and this component is lined with ultra high molecular

weight polyethyl ene. In this number of cases, problems such as infection (26 cases),

dislocations (68 cases), femoral shaft fractures (57 cases), hematomas (35 cases),

transient sciatic palsy (4 cases), greater trochanter avulsions (57 cases), ectopic

bone formation (131 cases) and stem fractures (3 cases) were observed. The numbers

are broken down into various groups relating to time of installation or problem.

Overall results for 235 cases for a five year period are 139 rated as excellent, 57

as good, 24 as fair and 15 as poor. The authors state that the madreporic prosthesis

is a safe, reliable procedure for individuals with reasonably good bone stock.

RingSS reports on 471 uncemented metal on plastic hips installed in 1980 - 1982.

The results obtained were better than those with uncemented metal on metal prostheses

and were comparable with results from cemented metal on plastic prostheses. The

acetabular component of the Ring prosthesis is a conical shape and must be axially

located. Infection is less of a problem than with the cemented prosthesis and

revision surgery can be accomplished with ease and without having to remove all components of the prosthesis.

12
Kuo, et al.59 report on reconstruction of long bones following tumor removal

accomplished by implanting segments of solid core T i -6A1 -4 V alloy coated with pure

titanium fiber metal. Results were excellent in 11 cases, satisfactory in five and

there was one failure. The one failure resulted from infection and the patient's

request not to undergo further surgery. Function was possible for this patient with

the use of a brace. Additional surgery and bone grafting were required for the satisfactory cases. Case summaries are given for the 17 patients.

Clinical studies of total knees were summarized by Hungerford and Kenna 23 .

Other research studies 66 on porous polysulfone coated Co-Cr-Mo stems placed in dogs

show that this coating supports bony ingrowth. Clinical studies of cementless

fixation using a polyacetal resin coated metal prosthesis are in progress 61

. The

polyacetal resin alone proved to be too elastic and the metal was added for strength.

The bone remodelling continues over many months. Studies in dogs by Chen 62

,

et al showed progressive ingrowth in three and six month observations. These

studies showed some osteoporosis which did not appear to be progressive. Progressive

active bone remodelling was observed through the experimental period. This study

also stressed the importance of initial firm fixation and close contact of the

prosthesis with the bone as well as the design of the device. This study also

mentioned the potential possibility of prosthesis removal which necessitates leaving

as much bone as possible in the initial installation.

The difference in bone growth rates in animals and older humans is striking 63 .

Histology studies were made of bone tissue taken from an eighty seven year old woman

after 40 days of implantation and of an eighty two year old man after a 6.5 month

implantation period. Bone ingrowth and adaptive bone modelling were observed but

the degree of ingrowth and fixation were less than that observed in animal models.

This difference was attributed partially to the patient's ages and inactivity.

The overall histological appearance of tissue adjacent to the prosthesis was that t

v m13
tissue was benign and free of foreign cell reaction. No macrophage activity and no

inflammatory responses were observed. Both patients suffered from some osteoporosis

and bone disease prior to implantation of the devices.

Galante^ summarized the current status of cementless fixation briefly in introductory statements to several articles referenced here. He noted that cementless fix

ation is more widely used in Europe and is not always based on bony ingrowth but also

on bony attachment to irregular surfaces. Two important prerequisites for bony ingrowth are close contact between prosthesis and bone and rigid immobilization. More

clinical data are needed on porous coated and other cementless prostheses. Reports

by clinicians and researchers will be useful in evaluating and developing these

devices.

14
References

1. Welsh, R. P., Pill iar, R. M. and MacNab, I., Surgical Implants: The Role of

Surface Porosity in Fixation to Bone and Acrylic, J. Bone and Joint Surgery,

53A , (5), 963-967 (1971).

2. Galante, J., Rostoker, W. , Lueck, R. and Ray, R. D., Sintered Fiber Metal

Composites as a Basis for Attachment of Implants to Bone, J. Bone and Joint

Surgery, 53-A (1), 101-114, 1971.

3. Hahn, H., and Palich, W. , J. Biomed. Mater, Res., Vol . 4, p. 571 (1970).

Preliminary Evaluation of Porous Metal Surfaced Titanium for Orthopedic

Impl ants.

4. Proceedings of the Workshop on Internal Joint Replacement, Northwestern University,

Chicago, Illinois, March 3-5, 1977.

5. Arthur D. Little Report, 1982.

6. NIH Consensus Development Conference, Total Hip Replacement, March 1-3, 1982,

Gordon, Stephen L., coordinator. National Institute of Arthritis, Diabetes

and Digestive and Kidney Diseases.

7. Kelsey, Jennifer L., Yale University School of Medicine, Epidemiology and

Impact." ibid.

8. Johnston, Richard C., The Hip Society, "Long Term Outcomes - Hip Society

Collective U.S. Experience," ibid.

9. Galante. Jorge 0., Rush Presbyterian-St. Lukes Medical Center, "Total

Hip Joint Replacements," ibid.

10. Klawitter, J. J. and Weinstein, A. M., The Status of Porous Materials to

Obtain Direct Skeletal Attachment by Tissue Ingrowth, Acta Orthopaed. Belgica,

Tome 40, Fasc. 5-6, 755-765, 1974.

11. Ring, P. A. J Bone and Joint Surg., 50B, 720-731, 1968.

12.

Charnley, J., J. Bone and Joint Surg., 47B , 354-363, 1975.

13.

Charnley, J., J. Bone and Joint Surgery, 42B , 28 (1960).

14. Ducheyne, P., Martens, M. , Aernoudt, E. Mulier, J. and De Meester, P.,

Skeletal Fixation by Metal Fiber Coating of the Implant, Acta Orthopaed.

Belgica, Tome 40, Fasc. 5-6, 1974.

15. Hornsey, C. A., Tullos, H. S., Anderson, M. S., Diferrante, N. M. and King, J. W.

Clinical Orthoped. and Related Research, No. 83, 317, 1972.

16. Amstutz, H. C., Lurie, L. and Bullough, P., Clinical Orthoped. and Related

Research, No. 84, 163, 1972.

15
17. Spector, Myron, Bone Ingrowth into Porous Metals, Biocompatibil ity of Orthopedic

Impl ants ,Vol . 1 1 , CRC Series in Biocompatibility, Boca Raton, FL, Ed. David F. Williams, Chapter 5, 89-128, 1982.

18. Porocoat, DePuy, Div. of Boehringer Mannheim Corp. , P.0. Box 988, Warsaw, IN

46580 Approx. 1982.

19. Pilliar, R. M., Powder Metal-Made Orthopedic Implants with Porous Surface

for Fixation by Tissue Ingrowth, Clin. Orthoped. and Related Res.,_l (6), 42-51,

1983.

20. Hulbert, S. F., Morrision, S. J. and Klawitter, Tissue Reaction to Three

Ceramics of Porous and Non-porous Structures, J. J., J. Biomed Mater.

Res. _6 , p. 347-374 (1972).

21. Bobyn, J. D., Wilson, G. J., MacGregor, D. C., Pilliar, R. M. and Weatherly,

G. C., Effect of Pore Size on the Peel Strength of Attachment of Fibrous

Tissue to Porous-Surfaced Implants, J. Biomed. Mater. Res., 16., 571-584, 1982.

22. Hungerford, D. S., Kenna, R. V. and Krackaw, K. A., The Porous Coated Anatomic

Knee, Orthop. Clin. North Am: 103-122, 1982.

23. Hungerford, D. S. and Kenna, R. V., Preliminary Experience With Porous Coating

Used Without Cement, Clin. Orthop. 176: 95-107, 1983.

24. Salman, N. N., The Effect of Direct Electrical Current Stimulation on the Bone

Growth into Porous Polymeric, Ceramic and Metallic Implants, PhD. Dissertation,

Clemson University, Clemson, SC, 251 pages, 1980.

25. Brighton, Carl T. , Guest Ed., Urist, Marshall R. , Ed., Symposium on Bioelectric

Effects on Bone and Cartilage, Clinical Orthopaedics and Related Research,

No. 124, May, 1977.

26. Brighton, Carl T. , Guest Ed., Symposium on Electrically Induced Osteogenesis, The

Orthopedic Clinics of North America, W. B. Saunders Co., Phil., PA, J_5 (1 ), 1984.

27. Ducheyne, P., Hench, L.L., Kagan, A., Martens, M. , Burssens, A. and Mulier, J. C.,

Effect of Hydroxyapati te Impregnation on Skeletal Bonding of Porous Coated

Implants, J. Biomed. Mater Res.,J_4, 225-237, 1980.

28. Lemons, J. E. , Tricalcium Phosphate Biodegradable Ceramic, ASTM Spec. Tech. Pub.

859, Symposium on Corrosion and Degradation of Implant Materials, Eds. A. C.

Fraker and C. D. Griffin, to be published, 1985.

29. Goldstrohm, G. L., Roberts, J. M. and Mears, D. C., Replacement of Canine Segmental

Bone Defects with Tricalcium Phosphate Ceramic Implants., Trans. Orthop. Res. Soc.,

8, 235, 1983.

30. Brown, Walter E. and Chow, Laurence C., Patent Pending.

31. De Groot, Klaas, Ed., Bioceramics of Calcium Phosphate, CRC Press, Inc. Boca Raton,

FL, 1983.

16
32. Nilles, J. H., Coletti, J. M. , Jr. and Wilson, C., Biomedical Evaluation of BonePorous Material Interfaces, J. Biomed. Mater. Res., 7_, 231 , 1973.

33. Predicki, P. Stephan, J. E., Auslaender, B. A., Mooney, V. L. and Kirkland, K.

,

Kinetics of Bone Growth into Cylindrical Channels in Aluminum Oxide and Titanium,

J. Biomed. Mater. Res.,_6 , 375 , 1972.

34. Harris, William H., White, Richard, E., Jr., McCarthy, Joseph C., Walker, Peter S.

and Weinberg, Edith H., Bony Ingrowth Fixation of the Acetabular Component in

Canine Hip Joint Arthroplasty , Clin. Orthopaed. and Rel . Res.,_l_ (6) 12, 1983.

35. Judet, R. , Sigiuer, M. and Judet, T. , An Uncemented Total Hip Prosthesis.

J. Bone and Joint Surg: _57_ (B), 396, 1975.

36. Judet, R. , Sigiuer, M. and Judet, T. , A Non-Cemented Total Hip Prosthesis,

Clin. Orthop. , 137 , 76-84, 1978.

37.

Lord, G. A., Hardy, J. R. and Kummer, F. J., An Uncemented Total Hip

Replacement. Experimental Study and Review of 300 Madreporique

Orthoplasties, Clin. Orthop. 141 , 2, 1979.

38. Wheeler, K. R. , Karagianes, M. T. and Sump, K. R. , Porous Titanium Alloy

for Prosthesis Attachment, Titanium Alloys in Surgical Implants,

Eds. H. Luckey and F. Kubli, ASTM STP 796, 241-254, 1983.

39. Ducheyne, P., Martens, M. , DeMeester, P. and Mulier, J. C., Titanium

Implants with Porous Structures for Bone Ingrowth: A General Approach,

ibid., 265-279.

40. Clemov, A. J. T. , Weinstei, A. M. , Klawitter, J. J., Koeneman, J. and

Anderson, J., Interface Mechanics of Porous Titanium Implant, J. Biomed.

Mater. Res.,_15, 73-82 , 1981.

41. Yue, S., Pill i a r , R. M. and Weatherly, G. C., The Effect of a Porous

Coating on the Fatigue Resistance of Ti-6A1-4V Alloy, Trans. 10th Ann.

Meet, of the Soc. for Biomater., Wash. DC, Vol . VII, 332, 1984.

42. Hirschhorn, J. S. and Reynolds, J. T. , Powder Metallurgy Fabrication of

Cobalt Alloy Surgical Implant Materials, Research in Dental and Medical

Materials. Plenum Press, New York, 137-150 (1969).

43. Pilliar, R. M., Cameron. H. U. Macnab, I., Porous Surface Layered Prosthetic

Device, J. Biomed. Eng. _10, 126 , 1975.

44. Kilner, T., The Relationship of Microstructure to the Mechanical Properties of

a Co-Cr-Mo Alloy Used for Prosthetic Devices, Ph.D. THesis, University of

Toronto, Toronto, Ontario, Canada, 1984.

45. Galante, J. and Rostoker, W. , Fiber Metal Composites in the Fixation of

Skeletal Prosthesis, J. Biomed. Mater. Res. Sym. John Wiley & Sons, Inc.,

NY, NY (4) 43-61, 1973.

17
46. Hahn, H., Rowe, R. , Lare, P. J., Fraker, A. C. and Ordway, F. , Symposium

on Corrosion and Degradation of Implant Materials, May 1983, ASTM STP 859,

To be published, 1985.

47. Laing, P., Compatibility of Biomaterials, Orth. Clinics of N. Am., £ (2),

249-275, 1973.

48. Laing, P., Clinical Experience with Prosthetic Materials: Historical Perspectives, Current Problems and Future Directions, Corrosion and Degradation of

Implant Materials, Eds. B. Syrett and A. Acharys, ASTM STP 684, 199-211, 1979.

49. Black, J., Systemic Effect of Biomaterials, Biomaterials, 5, 11-18, 1984.

50. Woodman, J. L., Black, J. and Nunamaker, D. N., Release of Cobalt and Nickel

from a New Total Finger Joint Prosthesis Made of Vitallium, J. Biomed.

Mater. Res.,JL7, 655-668, 1983.

51. Woodman, J. R. , Jacobs, J. J., Galante, J. 0. and Urban, R. M., Titanium

Release from Fiber Metal Composites in Baboons - A Long Term Study, Trans.

Orthop. Res. Soc.,_7, 166 , 1982.

52. Lautenschlager, E. P., Sarker, N. K. , Acharya, A., Galante, J. 0. and

Rostoker, W. , Anodic Polarization of Porous Fiber Metals, J. Biomed, Mater.

Res. , 8, 189, 1974.

53. Cameron. H. V., Pilliar, R. M. and MacNab, I., Porous Vitallium in Implant

Surgery, J. Biomed. Mater. Res., 8., 283 , 1974.

54. Van Orden, A. C. and Fraker, A. C., Effects of Porous Coatings on the

Corrosion Behavior of Co-Cr-Mo Material, Trans. 10th Ann. Meet. Soc. for

Biomater., Wash, DC, VII, 123, 1984.

55. Fraker, A. C., unpublished research.

56. Engh, Charles, A., Hip Arthoplasty with a Moore Prosthesis with Porous Coating:

A Five-year Study, Clin. Orthopaed. and Rel . Res., (6) 52, 1983.

57. Lord, G. and Bancel , P., The Madreporic Cementless Total Hip Arthoplasty:

New Experimental Data and a Seven-year Clinical Follow-up Study, ibid. 67.

58. Ring, P.A., Ring UPM Total Hip Arthropl asty , ibid. 115.

59. Kuo, Ken K.; Gitelis, Steven; Sim, Franklin H.; Pritchard, Douglas; Chao, Edmund;

Rostoker, William; Galante, Jorge 0. and McDonald, Patrick, Segmental Replacement

of Long Bones Using Titanium Fiber Metal Composite Following Tumor Resection,

ibid. 108.

60. Spector, M. , Davis, M. D. Lunceford, E. M. and Harmon, S. L., Porous Polysulfone

Coatings for Fixation of Femoral Stems by Bony Ingrowth, ibid. 34.

18
61. Morscher, E. W. and Dick, W. , Cementless Fixation of "Isoelastic" Hip Endoprostheses Manufactured from Plastic Materials, ibid. 77.

62. Chen, Po-Quang, Turner, Thomas M. , Ronnigen, Helge, Galante, Jorge, Urban, Robert,

and Rostoker, William, A Canine Cementless Total Hip Prosthesis Model, ibid. 24.

63. Bobyn, J. Dennis and Engh, Charles, Human Histology of the Bone-Porous Metal

Implant Interface, Orthopedics , _7 , (9), 1410, 1984.

64. Galante, Jorge 0., Guest Editor Editorial Comment, Clin. Orthopaed. and Rel . Res.,

1 , (6) 2, 1983.

19

Corrosion Behavior of Sintered Porous Coated Co-Cr-Mo Surgical Implant Metal

Corrosion Behavior of Sintered Porous-Coated Co-Cr-Mo Surgical Implant Metal

Introduction

Porous coatings are being applied to metal prostheses in an effort to improve

fixation of the device by bony ingrowth or in some cases, by providing stronger

mechanical attachment for the poly(methyl methacrylate) cement. The Co-Cr-Mo

material in the cast condition has been used successfully for permanent implants

since the 1930s 1 . In recent years, improvements in the mechanical properties

and especially the fatigue strength of the Co-Cr-Mo material have been made by

forging 2 and hot isostatic pressing2

. These processes result in the reduction of

the grain size from that of the cast material by many orders of magnitude. The

Zimmer, USA company uses isostatically pressed material and the forged high

strength material is available from Howmedica, Inc., Rutherford, NJ. These materials

probably are available from other sources as well.

The sintered porous coatings produce another modification of the Co-Cr-Mo

material which is not for the purpose of increasing strength but for providing

a porous surface which bone may grow into and attach itself^. The Co-Cr-Mo

porous coated material is prepared by binding Co-Cr-Mo spheres to the Co-Cr-Mo

substrate and sintering in a temperature range of 1150°C to 1300°C. The binder

should burn off and leave no residue. Details of the preparation of the porous

coating such as sintering temperature, cooling rate, binder, etc. are proprietary.

There always is concern regarding metal ion release or the release of any foreign

substance into the body. Earlier corrosion studies were made by Cameron, et al. 5 who

indicated no adverse effects on corrosion resistance due to the porous surface.

The purpose of the present investigation was to study further the corrosion behavior

of the porous coated Co-Cr-Mo material and to compare test data from solid specimens

21
with data from porous specimens. Compositional changes resulting from phase

changes and variations in the surface structure may alter the corrosion behavior.

Results of the present study show that the corrosion per unit area is the same

for the porous surface as for the smooth surface. Still, this gives an increase

in metal ion release since the surface area of the porous material is increased

by two or three times. Pitting and crevice corrosion were not observed with the

materials studied but variations in porosity or composition could change this.

The Co-Cr-Mo prosthetic devices, when placed in the body, will be in the passive

condition most of the time. This condition could be altered by scratching or rubbing

of the device or changes in the local pH in the area surrounding the implant. The

passive condition would minimize ion release in both the solid and porous forms.

The passive condition results from the formation of a surface oxide film, and studies

were conducted to analyze for the composition of this film and to determine the oxidation

state of the Cr ion which is predominant in the film. Details and results of the

investigation are presented in this report.

22
Materials and Methods

Materials used in this study were of a composition which fell within the

limits of the ASTM F-75 specification for Co-Cr-Mo material. The porous coated

materials were supplied by Howmedica, Inc., Rutherford, NJ and by DePuy, Inc.,

Warsaw, IN. The material from Howmedica was in the form of knee prostheses with

two or three layers of 850 to 1000 urn diameter spheres. Materials were provided

in the unpassivated condition and in the passivated condition. The later condition

involved passivation in 20 to 30 percent HNO3 as described in ASTM F-86. The

material from DePuy was provided in the form of coupons. The spheres were smaller

and ranged in diameter from 210 to 250 urn. The porous coating consisted of

three to four layers of spheres.

Gas adsorption techniques were used to determine the surface area of the porous

materials. A monolayer of gas is adsorbed on a porous specimen and then is desorbed

and measured. This is compared with the measured exposed area of the substrate base

which has no spheres. The surface area for the layers of larger spheres was four

times greater than that of the solid surface, and the surface area for the multilayered smaller spheres was six times greater than that of the solid surface.

Representative specimens were cut in cross section, mounted, polished and

etched for microscopic observation. Surface structure and surface films on the

porous and solid specimens were analyzed using scanning electron microscopy and

energy dispersive x-ray analysis (EDX). These techniques permit the determination

of the composition of the surface films and the underlying substrate as well as

the morphology of the surface film over different porous surfaces or second phases.

Flat specimens were used for study with the electron spectroscopy for chemical

analysi s( ESCA) equipment. These measurements provided data on the chemical : -

23
sition of the material and on the oxidation state of the elements present. It is

possible with this method to do a profile across the oxide film down to the metal

substrate. Emphasis in this study was on the valence of the chromium ion in the

oxide layer, and standards containing Cr+6 and Cr+3 were used as references.

Corrosion tests included anodic polarization measurements and pitting and

crevice corrosion testing. The porous surfaced specimens were cleaned in water

and rinsed in ethanol prior to corrosion testing. Anodic polarization measurements were made on specimens exposed to Hanks' physiological saline solution.

The solution was held at a temperature of 37°C and had a pH of 7.2. Specimens

were polarized potentiostatical ly at a rate of 0.006 V/ min. until breakdown

occurred. Voltage was measured versus a standard calomel el ectrode(S.C.E. ) and all

voltages mentioned in this report are with respect to this reference. These measurements involve the application of a voltage and a measurement of the current which

follows. The corrosion equipment is connected to a computer which is programmed to

run the test and to take the data. Three or more anodic polarization tests were

made on the cast material, the material with the layers of large spheres and the

material with multilayered small spheres.

Pitting and crevice corrosion tests were conducted using procedures described

in the ASTM F- 746 Recommended Test for Pitting and Crevice Corrosion. This is a

test in which a stimulation voltage is applied to the specimen to determine if

pitting is initiated. The stimulation potential for the tests with the Co-Cr-Mo

material was raised from 0.8V to 1.0V. These tests were conducted in a 0.9

percent NaCl solution.

24
Results and Discussion

Microstructures

The microstructure of the cast material is shown in the light micrograph of Figure 1. The grain size is large and there is a second phase area which may comprise a

mixture of the sigma phase, a cobalt rich phase and a carbide phase. The cross section

of the spheres which were sintered onto the surface is shown in Figure 2. The structure

dendritic due to rapid cooling in the production of the spheres. After sintering

the spheres on to the solid substrate at elevated temperatures, the fine structure

disappears and large grains form which contain second phase areas at grain boundaries

and other locations. This microstructure is shown in Figure 3 and is of the material

with the larger spheres. The second phase segregates in the neck areas between the

spheres and at locations where the spheres attach to the substrate. The amount of

this phase area and its constituents are related to the thermal treatment, cooling

procedures and composition of the material. Other views of the spheres and the

neck areas after being exposed to corroding conditions are shown in the scanning

electron micrographs of Figures 4 and 5. Some voids (see arrow) are present in

the neck areas.

Compositions of the secondary phases at the neck areas and of the matrix material

are shown in Figure 6. These results are for the material with the multilayers of

210 to 250um sized spheres. The secondary phase areas are enriched in Cr and Mo, and

this enrichment would correspond to the presence of the carbide phase(s) and

the sigma phase, respectively.

25
Figure 1. Light micrograph of cast Co-Cr-Mo.

26
Figure 3. Light Micrograph showing microstructure of Co-Cr-Mo spheres

sintered to a Co-Cr-Mo substrate.

27
Figure 4. Scanning electron micrograph of sintered Co-Cr-Mo

spheres on a Co-Cr-Mo substrate after corrosion

testing.

28
Figure 5. Scanning electron micrograph of neck region

connecting sintered Co-Cr-Mo spheres after

corrosion testing.

29
100

80

60

40

20

0

MATERIAL B

MATRIX AND SECONDARY

PHASES

MATRIX

SECONDARY

Fi Si Mo Cr Co

ELEMENT

Figure 6. Bar graph of EDS data showing smounts of selected elements i phases of Co-Cr-Mo after sintering.

different

30
ESCA Results

Electron spectroscopy for chemical analysis (ESCA) is a method used to examine

the outer 15 to 30 A of the surface. Specimens were analysed to determine the

valence of the chromium which was in the surface oxide film. Spectra were taken of the

films formed in Hanks' physiological saline solution. Spectra were taken initially

after exposure to the solution and again after ion sputtering for time intervals ranging

from 20 to 40 minutes. The selected data in Figure 7 show two chromium peaks.

The comparison of these peaks from the metal surface , curve C, with that of the

standard peaks for Cr+3

, curve A, shows a good match. The comparison of the

peaks from the oxide film, curve D, shows that this curve matches curve A for the

Cr+3 better than curve B which represents the +6 oxidation state. The distance between

the two peaks also was taken in to account in making this comparison. There is

some shift in the peaks for the oxide film, but the indication is that the chromium

is in the +3 oxidation state.

Electrochemical Measurements

The results of the anodic polarization measurements are shown in Figures 8 and 9.

This type of curve of current vs. potential shows the magnitude of the current, the

passive region and the breakdown potential for a given material. The breakdown

potential is denoted by a sharp rise in the current. Figure 8 shows curves for

the solid specimen (Cast), the specimen with multilayers of 210 to 250um sized spheres

(B) and the specimen with multilayers of 850 to lOOOum sized spheres (C). These

curves show higher current readings for the porous materials. The surface area of

of the material that yielded curve C is four times that of a solid specimen. The

surface area of the sample with the porous surface, represented in Curve B, is six

times that of a solid specimen. It is seen in Figure 8 that the current (and release

of metal ions) is greater for a porous surface of a given flat substrate area than it is

for a smooth surface of the same geometrical area.

31
Counts/Second

x

10

Figure 7. ESCA spectra for (A) O2O

3

standard for Cr+3

, (B) KpC^Oy standard for Cr+ 6

, (C) Cr metal surface,

and (D) oxide film.

32
LOG

CURRENT

(uR)

3

0

-1

-2

Figure 8

i i i i i i i i i j 2 1 0 .1 .2 .3 .4 .5 .6 .7 .8

APPLIED POTENTIAL CV vs SCE)

. Anodic polarization curves for Co-Cr-Mo in the (cast), (B) sintered

spheres of 210um to 250um diameter and (C) sintered spheres of 850u~

lOOOum diameter in Hanks' physiological saline solution at 37°C and

a solution pH of 7.2.

33
LOG

CURRENT

DENSITY

(

Ru

/

c

m

2

RPPLIED POTENT I RL CV vs SCE

)

Figure 9. Anodic polarization curves for Co-Cr-Mo in the (cast), (B) sintered

spheres of 210um to 250um diameter and (C) sintered spheres of 850um

to lOOOum diameter in Hanks' physiological saline solution at 37°C and

a solution pH of 7.2.

34
There appeared to be no basic differences in corrosion behavior between the

porous and the smooth materials studied in this investigation. The current density

versus the applied potential for these same specimens is shown in Figure 9. The current

per unit area for the total surface is used in this figure, and this brings all of the

curves closer together. The small differences which remain could be due to error

in the gas porosimetry measurements of the area of the porous surface.

The breakdown potential for all specimens were very similar and approximately

0.47 volts versus S.C.E. Passivation of these Co-Cr-Mo specimens resulted in an

increased breakdown potential which occurred at approximately 0.87 volts versus S.C.E.

Pitting and crevice stimulation tests were conducted on these materials, and in all

cases the test was continued until the voltage reached the breakdown potential

without indications of an onset of pitting or crevice corrosion. This type of local

corrosion could depend on depth and size or pores as well as the composition of second

phases, and the results of this study are specific to the type of materials used in

this investigation.

Summary

This work was an investigation of the corrosion behavior of sintered porous

Co-Cr-Mo in terms of anodic polarization behavior, pitting and crevice corrosion

resistance and surface film analysis. The results of this study may be summarized

as follows:

1. Microstructural analysis of the sintered materials showed the presence

of a second phase which often was concentrated at the neck region where

the spheres sintered together and at the interface region where the spheres

joined to the substrate. The second phase was found at the grain boundaries

and at other locations throughout the specimen. Three different batches of

35
material were examined, and the amount of this second phase was less in

two of the batches,

2. Energy dispersive x-ray analysis showed the composition of the second phases

in the neck regions to be enriched in Cr and Mo.

3. Anodic polarization measurements showed corrosion currents to be higher for

the porous materials, but the current density for the porous and solid

materials was essentially the same. The breakdown potential for both the

porous and solid materials in the Hanks' solution was 0.47 volts vs. S.C.E.

4. Electron spectroscopy for chemical analysis was used to study the composition

of the oxide films which were found to be a mixed oxide, predominantly

chromium oxide. Results indicated that the oxidation state of the chromium in

the surface film is +3 and not +6. This determination was made using standard

salts containing chromium of known valence.

36
References

1. Venable, C. S. and Stuck, W. G., The Internal Fixation of Fractures, C. C.

Thomas, Springfield, 1947.

2. Devine, T. M. and Wulff, J., Cast vs. Wrought Cobalt-Chromium Surgical Implant

Alloys, J. Biomed. Mater, Res.,_9 (2) 151-167, 1975.

3. Luckey, H. A. and Barnard, L. J., Improved Properties of Co-Cr-Mo Alloy by

Hot Isostatic Pressing of Powder, Mechanical Properties of Biomaterials and

Bioceramics Symposium, University of Keele, England, G. W. Hastings, Chmn., 1978.

4. Welsh, R. P., Pi 11 i ar , R. M. and MacNab, J., Surgical Implants: The Role of

Surface Porosity in Fixation to Bone and Acrylic, J. Bone and Joint Surgery,

53A , (5), 963-967, 1971.

5. Cameron, H. V., Pilliar, R. M. and MacNab, J., Porous Vitallium in Implant

Surgery, J. Biomed. Mater. Res., 8, 283, 1974.

37

Part III

Corrosion-Fatigue Behavior of Arc Plasma Sprayed Ti or Ti-6A1-4V Coatings on

Ti-6A1-4V Substrates
Corrosion-Fatigue Behavior of Arc Plasma Sprayed Ti and Ti-6A1-4V Coatings on

Ti-6A1-4V Substrates

Introduction

Commercially pure titanium and Ti-6A1-4V alloy coatings were applied to

T1-6A1-4V alloy by arc plasma spraying^ *2 >3. The coatings were applied in a

manner which produced a porous coating. The coatings also can be applied

in a more dense form and of other materials and then surface finished for

intended uses other than surgical implants. Application of coatings in

this manner results in a coating which is strongly bonded to the substrate.

Test results showed that fracture occurred within the coating and not at

the coating/substrate metal interface. Corrosion-fatigue behavior is not

adversely affected by application of this coating, and it is subject to further

improvements. The interface bond strength, fatigue strength and other mechanical

and chemical properties of this coating are subject to change depending on

the variation of conditions used in the production of the coating. Coatings

used in this investigation were prepared under controlled conditions, and

microstructures and other properties were analysed. A number of variables

could influence the integrity of this coating including; the purity of the

metal powder which is being sprayed, the material used for the grit blasting,

the cleaning process after grit blasting, the purity of the carrier gas and other

aspects of the application of the coating relating to the control of the gun,

spray nozzle, etc.

38
Materials and Methods

Cylindrical specimens of 2.86 cm diameter were coated on the ends and epoxyed

together for tension testing. Flat specimens were cut from Ti-6A1-4V to a size

of 2.54 cm x 0.64cm x 7.50 cm, and an arc plasma sprayed coating was applied to

one side. These specimens were used in the lap shear tests. Specimens for corrosionfatigue testing were made from 0.250" diameter rod of extra low interstitial (ELI)

grade Ti-6A1-4V which was obtained from Dynamet, Inc., Clearwater, Florida.

These specimens were cut into 7.6 cm lengths, and the middle 12.7 mm was machined

to a decreased diameter of 5mm. This center section was arc plasma sprayed to

apply the coating, and then was exposed to the corrosive solution during torsion

fatigue testing.

The arc plasma sprayed coating was applied in the following manner. Prior to

coating, the designated area of the specimen was grit blasted with aluminum oxide

abrasive to produce a uniform surface texture. The grit blasted area was then

coated with a base layer of dense material, followed by a porous layer of material

to produce a coating with graded porosity. The density and/or porosity can be controlled by varying the spraying parameters. Materials used in the coatings were

-30 +325 mesh Ti-6A1-4V powder from Consolidated Astronautics Division, United International Research, Inc., Concord, MA and ground sponge titanium from Reactive Metals, Inc

Niles, OH. A schematic of the gun for the arc plasma spraying is shown in Figure l 4

.

Tests to determine the bond strength of the coating to the substrate were

conducted using a lap shear method. Porous coated surfaces were filled and bonded

together with To rr Seal* adhesive resin. The adhesive resin was obtained from

Farian Associated, Palo Alto, CA. The two specimens were properly aligned and

clamped together. The resin was cured at 60°C (140°F) for 1.5 hours, and th.

resin was sanded off. Tension tests of the bond strength were conducted by coating the

39
Plasma Flame

(temp., 15000°C) Plasma Gas Suppl;

(Argon, Hydrogen

Metal Powder- earn*' argon gas

Gas Speed

(3000 m/sec.)

Metal particles (grain size: 0.05-0.1 mm) are fused. Particle velocity, about 600 m/sec.

\\

Water Cooling

Figure 1. Schematic drawing of the arc plasma spray gun

Figure 2. Liqht micrograph or APS Ti-6A1-4V coating on Ti-6A1-4V after sintering

at' 982°C (1800°F).

£fi
flat ends of cylindrical rods and filling and bonding together with the resin.

The lap shear and tension specimens were tested with a Model TTC Instron Testing

Machine with a 45 KN (10,000 lb) capacity and a cross head speed of 0.25 mm (0.01 in.).

Corrosion-fatigue tests of coated cylindrical specimens were conducted in fully

reversible torsion in a saline solution (Hanks' solution) environment. The applied

shear strain amplitude was +_ 0.010 to 0.022 and the cyclic twist was at a frequency

of 1 Hz. The solution was held at a temperature of 37°C and had a pH of 7.4.

The open circuit potential in volts was measured versus a standard calomel electrode

(S.C.E.) and recorded throughout the test. This potential was at a more positive

value when the specimen was at rest and after the specimen broke. This provided

another method, in addition to the cycle counts, for determining the number of

cycles to failure. The open circuit potential is more negative during cyclic twist

due to breaking and straining of the surface oxide film.

Results and Discussion

The cross sections of specimens showing the microstructures of the substrate and

coating after heat treating at 982°C (1800°F) and at 899°C (1650°F) are shown in

Figures 2 and 3, respectively. The porosity probably is interconnecting although it

may not appear to be so in this cross section. The specimen in Figure 2 was sintered

above the beta transformation temperature 980°C (1796°F) as is evident from the ordered

"needle or platlet" structure. The material in Figure 3 was sintered below the beta

transus, and the needle or lathe structure is not present. Sintering above the beta

transus has deliterious effects on the mechanical properties, and sintering at 982°C

(1800°F) was discontinued.

The surface structure and porosity are shown in Figures 4 and 5. The size of the

pores and the percent of porosity can be controlled by controlling the size of the

powder which is sprayed and other parameters of the process. Figure 4 is a ph t.emicrograph of a Ti sponge coating sprayed on a T i - 6 A1 - 4 V substrate. Figure 5

41
a Ti-6A1-4V coating sprayed on a Ti -6 A1 -4 V substrate. There was no great difference

in the physical appearance of the coating.

The results 2 of the lap shear tests of the Ti sponge top coating gave a strength

of 8.36 MPa (1213psi ) in the unsintered condition and 7.47 MPa ( 1 084ps i ) after sintering.

The strength of the T i -6 A1 -4 V top coat after sintering was 10.66 MPa( 1 546 ps i ) . The

strength of the Ti -6 A1 -4V alloy is higher than that of pure titanium in either solid or

porous form. The tensile strength of the Ti sponge coating was 8.54 MPa ( 1 238psi ) in

the unsintered condition and 19.36 MPa(2808psi) in the sintered condition. The tensile

strength of the Ti-6A1-4V sintered coating was 28.45 MPa (41 27 ps i )

.

More studies are needed regarding the effects of sintering on these coatings.

There is the possibility that the oxygen which is inadvertently introduced during the

production of the coating diffuses into the titanium. This would provide a strengthening effect in tension but the fatigue strength, especially torsion fatigue would be

reduced.

Figure 6 shows the substrate interface and porous coating. Microhardness

measurements were made across this type of section for both the unsintered and the

sintered material. Some of these data are summarized in Table 1. Comparison of

data in row A with that in row B for the Ti coating and the Ti-6A1-4V coating shows

that after sintering, a trend exists for an increase in hardness in the material.

Hardness measurements made in areas of the coating other than the spherical particles,

gave hardness values exceeding 900. These measurements were limited, and these data

are not shown in the table. The confidence level of the values listed in the table

could be raised by increasing the number of measurements.

Data on the corrosion-fatigue life are given in Table 2. The data

presented here are from tests using an applied shear strain amplitude of

_+ 0.010. Ten other corrosion fatigue tests were conducted in the shear strain

range of _+ 0.004 to + 0.022. The data shown in Table 2 can be used to

42
Figure 3. Light Micrograph of an APS T i - 6A 1 - 4V coating on T i - 6A 1 - 4V after sintering at 900°C (1650°F).

Figure 4. Light Micrograph showing surface morphology of an APS Ti sponge coating on a Ti - 6A1 -4V substrate.

43
Figure 5. Light micrograph showing surface morphology of an APS Ti -6A1-4V coating on a Ti-6A1-4V substrate.

Figure 6. Light micrograph of APS T i -6A1 -4V coating on Ti-6A1-4V

after sintering at 870°C (1600°F) and making hardness

measurements (see arrows).

44
KNOOP HARDNESS NUMBERS, HK 10 q. FOR Ti AND Ti-6A1-4V POROUS COATINGS

Type of

Coating on Ti-6A1-4V

and Treatment

Spherical Particles Interface

within coating

Substrate .004" Substrate

from interface

HK ioo

a (S.D.

)

HK 100 (S.D.) HK 100 (S.D.) HK 100 (S.D.)

(A) -20+40mesh

T1-6A1-4V, 626 (7) 569 (87) 445 (56) 380 (22)

sintered,

870°C/16hrs

(B) -20+40mesh

T i -6 Al-4 V , unsintered

431 (134) 530 (116) 336 (12) 348 (8)

(A) Ti sponge,

sintered, 599 (72) 617 (65) 501 (194) 349 (24)

870°C/16hrs

(B) Ti sponge,

unsintered 350 (33) 484 (149) 347 (22) 330 (8)

a Mean HK^qq value, 4 to 8 measurements

b (S.D.), Standard Deviation

45
Table 2. Corrosion-Fatigue Test Results at an Applied Shear Strain Amplitude of + 0.010

Specimen* and Treatment Number of Cycles to Failure

M - Machined and polished, unsintered >10 6 **

F - Grit blasted, unsintered > 10 ^**

G - Grit blasted, unsintered >10^**

L - Ti-6A1-4V coating, unsintered >10^**

J - Ti-6A1-4V coating, sintered, 900°C (1650°F) 9.3 x 10^

K - Ti-6A1-4V coating, sintered, 900°C (1650°F) 9.4 x 10^

D - Ti sponge top coat, unsintered >10^**

C - Ti sponge top coat, sintered, 870°C (1600°F) 2.2 x 10^

B - Ti-6A1~4V top coat, unsintered 4.8 x 10^

A - Ti-6A1-4V top coat, sintered 870°C (1600°F) 7.5 x 10^

E - Ti-6A1-4V top coat, sintered 982°C (1800°F) 9.8 x 10^

H - Ti-6A1-4V, uncoated, grit bl asted ,si ntered ,982°C (1800°F) 4.3 x 10^

I - T i - 6 A 1 -4V , uncoated ,gri t bl asted , sintered ,982°C (1800°F) 4.4 x 10^

Specimen substrate material was ELI Ti-6A1-4V in an annealed condition.

** Tests were stopped after exceeding 10^ cycles, but these specimens did

not fail.

46
compare effects of surface treatment, sintering and sintering temperature.

It is evident from this table that sintering at a temperature above the

beta transus temperature decreases the corrosion- fatigue life either with

the coating or in the grit blasted condition. It can be expected that grit

blasting results in a compressive stress on the surface which has a beneficial

effect on the fatigue strength.

The specimens are designated by letters A through M in Table 2. Specimens

sintered to 982°C (1800°F) had a low fatigue life, data from Specimens C, A, J

indicate that sintering at the lower temperatures also adversely affected

the corrosion-fatigue life but to a lesser degree than sintering above the beta

transus. The loss in fatigue strength for these specimens was not due to the

the microstructural transformations but probably to the diffusion of oxygen into

the material. The control material, specimen M, compares equally with the grit

blasted material with no coating and with the coated material which is unsintered.

These test data show that sintering adversely affects the corrosion- fatigue life

of the material

.

47
Summary

Porous coatings of Ti and H-6A1-4V were applied to Ti-6A1-4V substrates

using the technique of arc plasma spraying. Microstructural analyses and

microhardness measurements were made on the coated specimens after sintering at

high temperatures and after mechanical testing. Mechanical tests were conducted

to measure shear strength, tensile strength and corrosion fatigue life. The

following conclusions may be drawn from this investigation.

1. Arc plasma spraying can be used to apply porous coatings of either

titanium sponge or Ti-6A1-4V to Ti-6A1-4V substrates.

2. Grit blasting can result in some imbedded particles which do not wash out,

and this whould be minimized. Grit blasting appears to have no detrimental

effect on the fatigue strength.

3. The strength of the interface bond formed between the coating and the

substrate is high. Failure in lap shear tests always occurred within the

coati ng.

4. Parameters of the coating process can be varied to give a coating with

either a large or a small pore size or to give a coating having a graduated pore size.

5. Control of plasma spraying techniques and subsequent heat treating are

important to prevent oxidation and also to prevent microstructural changes

in the substrate metal.

6. Microhardness measurements increase within the material after sintering

which is indicative of oxygen diffusion into the material.

7. Application of the porous coating has not been shown to reduce the

corrosion-fatigue life of the substrate if the coated specimen is not

sintered. Sintering can lower the fatigue resistance.

48
References

1. Hahn, H, and Palich, W., Preliminary Evaluation of Porous Metal Surfaced

Titanium for Orthopedic Implants, J. Biomed. Mater. Res.,4_, 571-577 , 1970.

2. Hahn, H., Bone Implant, U. S. Patent 3 605 123, 1971.

3. Hahn, H., Rowe, R., Lare, P. J., Fraker, A. C. and Ordway, F., ASTM Spec.

Tech. Pub. 859, Symposium on Corrosion and Degradation of Implant Materials,

Eds. A. C. Fraker and C. D. Griffin, to be published, 1985.

4. Sutter, F. , Schroeder, A. and Straumann, F. , Engineering and Design Aspects

of the I.T.I. Hollow Basket Implants, J. Oral Impl antol ogy , Vol. X, No. 4, 1983.

49
Acknowledgements

The receipt of materials and specimens from the following individuals

and companies Is gratefully acknowledged; Larry Gustavson, Howmedica, Inc.,

Rutherford, NJ,; Todd Smith, De Puy Mfg. Co., Inc., Warsaw, IN 46580; Henry

Hahn, Artech Corp., Falls Church, VA.

50
NBS-114A (rev. 2 -ec) U.S. DEPT. OF COMM.

BIBLIOGRAPHIC DATA

SHEET (See instructions)

1. PUBLICATION OR

REPORT NO.

NBSIR -85-3166

2. Performing Organ. Report No. 3. Publication Date

June 1985

4. TITLE AND SUBTITLE

Studies of Porous Metal Coated Surgical Implants

5. author(S) A> c> Fraker; A. W. Ruff; A. C. C. D. Olson

Van Orden; H. Hahn; A. J. Bailey and

6. PERFORMING ORGANIZATION (If joint or other than NBS, see instructions)

national bureau of standards

DEPARTMENT OF COMMERCE

WASHINGTON, D.C. 20234

7. Contract/Grant No.

8 . Type of Report & Period Covered

9. SPONSORING ORGANIZATION NAME AND COMPLETE ADDRESS (Street. City. State. ZIP)

10.

SUPPLEMENTARY NOTES

| | 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)

The material in this report deals with the subject of metal porous coated surgical

implants which are used primarily for orthopedic applications. The report is

presented in three parts. The first part gives a brief history of the development of

various types of metal porous coated implants and discusses the need for improved

fixation of orthopedic devices which led to the investigation of porous coatings for

bony ingrowth attachment. The second part of the report contains experimental data on the corrosion behavior of sintered spheres of Co-Cr-Mo and analysis of surface films

after exposure to saline solution. The third part reports on a corrosion-fatigue study

of arc plasma sprayed porous Ti and Ti-6A1-4V coatings. The porous coatings on the

materials studied did not adversely affect the corrosion or mechanical properties of

the material. Changes in processing, contamination, etc. could have detrimental effect^

on the chemical and mechanical behavior of the materials.

12.

KEY WORDS (Six to twelve entries; alphabetical order; capitalize only proper names; and separate key words t

y

cobalt-chromium-molybdenum alloy; corrosion; fatigue; metal porous coating;

surgical implants; titanium-6 aluminum-4 vanadium alloy

13.

AVAILABILITY

|~y~] Unlimited

| 1 For Official Distribution. Do Not Release to NTIS

Order From Superintendent of Documents, U.S. Government Printing Office, Washington. D.C.

20402. O^er From National Technical Information Service (NTIS), Springfield, VA. 22161

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