NBSIR 73-420

Survey on Metallic Implant Materials

J. R. Parsons and A. W. Ruff

Metallurgy Division

Institute for Materials Research

National Bureau of Standards

Washington, D. C. 20234

December 1973

Final Report

Prepared for National Institute of Arthritis, Metabolism

and Digestive Diseases

Westwood Building

5333 Westbard

Washington, D. C. 20016

NBSIR 73-420

SURVEY ON METALLIC IMPLANT MATERIALS

J. R. Parsons and A. W. Ruff

Metallurgy Division

Institute for Materials Research

National Bureau cf Standards

Wastiington, D. C. 20234

December 1973

Final Report

Prepared for National Institute of Arthritis, Metabolism

and Digestive Diseases

Westwood Building

5333 Westbard

Washington, D. C. 20016

U. S. DEPARTMENT OF COMMERCE, Frederick B. Dent, Secretary

NATIONAL BUREAU OF STANDARDS, Richard W. Roberts, Director

ABSTRACT

The application of metallic materials as orthopedic implants in the

human body is reviewed, concentrating on materials presently in clinical

use and undergoing laboratory evaluation for possible future use. The

criteria considered explicitly are tissue compatibility, mechanical

properties, corrosion resistance, and toxicity. The three principal

metallic implant materials, stainless steel, cobalt alloys, and

titanium, are discussed in detail. Wherever possible, comparisons are

made between the materials in terms of the intended application.
I:

V : ,
INTRODUCTION

This report concerns metallic biomaterials used in orthopedic surgery.

Orthopedics can be defined as the medical specialty concerned with the preservai i t| tion and restoration of the form and function of the musculoskeletal system and

supporting structures. The focus of this report will be on orthopedic implants,

both protheses and fixation devices. It will not consider external prostheses,

e.g., artificial limbs; it further excludes non-orthopedic implants, e.g.,

artificial heart valves. Artificial biomaterials is itself a large category,

including metals, ceramics, plastics, composite materials, glasses, etc. The

focus of this report will be on metallic materials, however, this is not to

suggest that equally important problems and solutions do not involve ceramic

implants, for example. Other material types would require complete studies in

themselves. The aim of this report is to survey the available literature on the

subject, summarize the principal findings, and suggest items of future importance.

Other review articles on this and on allied subjects are available, including

1 2 3 4 5-7

those by Cohen , Weisman , Katz , Hulbert et aj_. , and others , and may be

consulted for additional information, beveral symposia on this subject are

referenced later in specific connections. A recent National Academy of Science

g workshop report is available.

Metals have been used as internal fixation devices since the days of Hippo-

' crates, ca. 400 B. C. At first, pure metals alone were available, including

gold, silver and iron. Around 1900, plated steel bone plates and screws were

applied. Cobalt-base alloy devices were introduced in about 1911, and in 1926

stainless steel was first utilized as an implant. At present time, three

, metallic materials dominate the field,
2

stainless steel, cobalt (Vitallium type*) alloys, and titanium. Each alloy was

developed initially for other applications but each exhibits particular properties

that are crucial for successful implant materials.

;,v Implant Material Requirements

The concept of an ideal implant material is of little importance since the

demands associated with different applications are found to vary a great deal.

2

However, the general requirements can be set down as follows, after Weisman .

1. Tissue compatabi 1 i ty

2. Suitable mechanical properties (including wear resistance)

3. Corrosion resistance

4. Non-toxic, non-carcinogenic

5. Appropriate cost

Each requirement will have a different effective weighting factor, depending on the

application. For example, short-time fixation devices do not need the high degree of

tissue compatabi 1 i ty that an implanted artificial joint might require. However, every

implant material must have documented performance in the above areas.

Tissue compatabi 1 ity test protocol has been discussed by Cohen^ and others; a

uniform testing method would be desirable and is not presently available. Long-time

compatabi 1 i ty data is essential as surgeons begin to implant devices in younger

patients. Certain mechanical properties values can be accurately determined by laboratory

testing in air or in vitro , for example, tensile strength and durability. Other

properties such as fatigue life, require careful environmental control during measurement. _ . Identification of proprietary materials by trade name is made solely for the purpose

of adequate description and connection with reported data.
j 3

' We note that strength beyond the required level may not be desirable, since loss of

ductility frequently accompanies increased strength. A certain amount of ductility

, is required in implants to prevent fracture on sudden impact loading, and to permit

, tailoring the device to fit to the patient.

The required corrosion resistance of an implant material also depends on many

factors; expected implant time, body reaction to corrosion products, effect of corrosion

I on mechanical strength, and others. Design factors can be important; crevices should

! be avoided where low pH and oxygen concentrations can destroy the passive, protective

film on the metal implant. Installation problems, such as the coupling of dissimilar

, metals, must be avoided. The inertness of the implant and its products, whether

I corrosion-or wear-produced, is obviously essential. Attention has recently been

focussed on the migration of particles from artificial implants through the body to

jj remote locations. Clearly, it is of great concern to obtain as complete an understand! nc

I as possible of such processes. Infections do occur at the implant site and it is

important to clarify the role of tne implant metal and its products in that matter.

Properties of Bone

It is essential to consider the response of bone to the attachment of implants

that will transmit mechanical loads for various periods of time. Accurate knowledge

of the mechanical properties of bone in vivo are required. Bone consists of oriented

fibers of collagen, hydroxyapati te crystals, amorphous calcium phosphate and

mucopolysaccharides, combined to form a dynamic, viscoelastic, anisotropic material.

The compressive strength of human bone is about 20,000 psi^* The strength varies

significantly with bone type, age, and physiological status. The elastic modulus

I of wet human femur is about 2 x 10 psi , longitudinally, and about one-third of that

value, radially. Bone is highly strain-rate sensitive. Since damage is repairable

**English units are employed in this report to be consistent with convention in the

field. Conversion to-SI units is shown in sone cases.
4

in living bone, the resistance to fatigue failure is very high. One important

mechanical factor in connection with artificial implants concerns the elastic

modulus mismatch with bone. Ideally, a material substituting for bone should have

a comparable modulus value, thus avoiding local elastic stress concentrations at the

joint interface. Given the viscoelastic and anisotropic properties of bone, an

ideal match is nearly impossible. Recent attention to porous implant materials

(or coatings on solid implants) where tissue ingrowth can take place, is of interest

here in terms of the intermediate modulus values offered in such materials in the

region of ingrowth. Further information on the properties of bone is available in

1 4,9,10 several sources

" Format of this Report

An extensive discussion will follow on the three metallic materials principally

used in clinical orthopedic implant activities; stainless steel, cobalt alloys, and

titanium. The physical properties and corrosion behavior of each metal will be

discussed, followed by a presentation of past and current information on implant

applications. A concluding section will discuss other metals applied in this

connection. Where possible we attempt to compare materials for use in a given

application, however, conflicting requirements associated with the application and

the lack of accurate data makes this a difficult task and leads to very qualitative

conclusions.
II. STAINLESS STEEL

Physical Properties

Stainless steels, as the name implies, are more corrosion resistant than

plain carbon steels or low alloy steels. This superior resistance to corrosion

is produced principally by the addition of chromium. Small amounts of other

elements such as copper, aluminum, silicon, nickel and molybdenum can further reduce

corrosion. Combinations of chromium and nickel are particularly useful.

The American Iron and Steel Institute (AISI) designate 12% chromium as the

cutoff point between "alloy" steel and "stainless" steel. Between the limits of

12% and about 30% chromium, with additions of other elements, are a spectrum of

types of stainless steels. First, there is the group of stainless steels which are

ferri tic and not hardenable by heat treatment (for example, AISI types 405, 430, 430F

and 446). Typical applications of these steels are in automotive and architectural

trim, manufacture handling of nitric acid, and medium temperature unstressed machine

parts. Second J there are the martensitic stainless steels (designated AISI type

403, 410, 414, 501 and 502 for example). Unlike ferritic stainless steels, these

steels are hardenable by heat treatment, have good wear characteristics and

have found applications as tools and cutlery. Neither of these two types of stainless steels have application in the implant field due to poor corrosion resistance.

The martensitic type 440C is used extensively in surgical instruments.
Third, and most important from an implant materials standpoint are the

austenitic stainless steels. Some AISI types are 301, 304, 304L, 316, 316L,

317 and others. Virtually all stainless steels implants are manufactured from

steels in this group^^. The austenitic stainless steels are not hardenable by heat

treatment. However, a wide range of tensile, hardness and related properties can

be obtained through cold working , provided the shape of the product lends itself

to such treatment, i.e., rods, bars, sheets, and forgable shapes. In the fully annealed

5 state, austenitic stainless steels have an average tensile strength of about 1.0 x 10 p 10 2 (2.0 X 10 N/m ). In a heavily cold-worked state the tensile strength may increase

5

to 2.0 X 10 psi or more. Ductility, on the other hand, is greatly reduced by coldwork, going from about 50% (elongation to rupture) in the annealed state to about 4%

in a fully work-hardened condition. Similarly, impact strength or toughness (ability

to withstand shock loading) is reduced by cold work.

The modulus of elasticity of the austenitic stainless steels is approximately

6 12

28 X 10 psi. This modulus is at least twice as high as cortical bone and an

13 order of magnitude higher than that of trabecular bone . The importance of this

modulus difference will become evident later.

_

Deformation of a metal at a temperature below the recrystal 1 ization temperature

to produce a permanent strain.
7

The average fatigue limit, the stress level below which failure due to cyclic

4

stresses does not occur, is approximately 4.0 x 10 psi for the austenitic steels.

Fatigue strength is somewhat dependent on work hardening, increasing with tensile

strength values. However, it is more affected by such variables as orientation

texture, temperature, surface finish, microstructure and environment.

In summary, the austenitic stainless steels are strong, ductile, impact

resistant and have adequate fatigue properties for many design situations. Additional

technical information concerning austenitic stainless steels is readily available

14 15

from several sources

'

Corrosion Behavior

The corrosion resistance of austenitic stainless steels is high, compared to

the ferritic and martensitic steels. The combination of chromium and nickel additions

produces high corrosion resistance, even in some active chemical environments.

Like many other corrosion resistant metals and alloys, the austenitic stainless

steels develop this resistance by forming an adherent, inert, passive metal-oxide

film^^'^^. This film, when formed, inhibits further oxidation or corrosion. The

presence of nickel with small amounts of molybdenum can significantly improve the

1

8

stability of the passivating film . The nature and mechanism of passivation are

19 discussed in various corrosion texts

There are circumstances where a passive film breaks down and corrosion will

occur. Environments which are oxidizing usually strengthen the passive film, whereas,

a reducing environment may break down the film, causing the metal to corrode. For

example, in a high- temperature oxidizing atmosphere, i.e., heat exchangers, jet

engines, etc., there may be negligible corrosion. However, in sulphurous gases
3

(such as hydrogen sulphide) rapid attack can take place. There are also circumstances

jnder which stainless steels will undergo local corrosion. A localized disruption

3f the passivating oxide layer can take place followed by continuing attack. Possible

reasons include wear or fretting, galvanic coupling, intergranular corrosion (at grain

boundaries), corrosion at crevices and inclusions, and mechanical surface damage.

Jnder certain circumstances, austenitic stainless is subject to any of these corrosion

)henomena.

Galvanic corrosion and fretting corrosion are perhaps the most basic and well

jnderstood of the corrosion effects mentioned. These problems can be largely overcome

through adequate design and knowledgeable application of corrosion principles. Dissimilar

letals having different electrochemical potentials should not be placed in contact in

J corrosive electrolytic media. Likewise, continuous mechanical disruption (fretting)

)f the passive film of a metal in a corrosive environment should be avoided. Interjranular corrosion, once a major problem of austenitic stainless used in high

:emperature service, has been largely overcome. This type of corrosion was caused by

:he precipitation of chromium-rich carbides at grain boundaries following high

:emperature exposure. This left small areas directly adjacent to these boundaries

leficient in chromium and thus less corrosion resistant. Through the efforts of

!ain et a_l_.^^ and others in the 1930's this problem was isolated and remedies developed.

)uring World War II, the low carbon type austenitic stainless steels (304L and 316L)

/ere developed. The low carbon content of these steels effectively prohibits carbide

)recipitation and subsequent pitting attack.
I!

Not yet as well understood are pitting corrosion, crevice corrosion, corrosion

fatigue, and stress corrosion of austenitic stainless steels. In marine environments,

i 14

I it is well documented that austenitic stainless steels show crevice and pitting

corrosion. As stated earlier, however, the addition of molybdenum significantly

improves the stability of the passive film and can retard, but not eliminate pitting

and crevice type corrosion. Similarly, stress corrosion and corrosion fatigue processes

21 are accelerated in chloride and caustic solutions. For example, data on corrosion

fatigue indicates up to a 50% increase in corrosion fatigue damage of austenitic

22

stainless tested in salt water as opposed to fresh water. Other sources report

substantial loss of tensile strength (up to 90% in 42% boiling MgCy of stressed

austenitic stainless steels in chloride solutions due to stress corrosion cracking.

Additions of molybdenum have little effect on these types of corrosion. Damage can

be minimized, however, by lowering stresses and/or by altering the environment.

23

For additional information the reader is referred to the work by Fontana and Green

which provides discussion and examples of all the types of corrosion mentioned above.

In summary, the corrosion characteristics of the austenitic stainless steels,

like the mechanical properties, are generally very good. These steels are not

inherently inert, however, but become resistant due to the formation of a protective

oxide film. The metal itself, with no such film has poor corrosion resistance.

This will be a significant problem in low oxygen concentration environments.

Stainless Steel as an Implant Material

In the 1930 's, orthopedic surgeons had become disenchanted with the open reduction

of fractures using bone plates. This disenchantment was due to a rather high rate of
10

failure caused by adverse tissue reactions or cracking of the carbon or vanadium

24

steel plates . When the austenitic stainless steels became available to the medical

Icommunity, they were naturally applied as fixation devices.

The basic type 302 stainless steel was the first to be used. The success rate

1Df open fracture reduction immediately improved. However, corrosion-related

failures of the fixation devices and adverse tissue reactions remained as significant

Droblems. Later, the more corrosion resisiant types 304, 316 and 317 stainless steels

' 24 25 vere employed ' . With the advent of World War II and the associated injuries

the demand for and use of implant materials increased greatly.. It was during these

/ears that the reputation of stainless steel, mainly type 316 was established. It

vas also during this period that cobalt alloys (see III), developed originally for

dental use, gained wide acceptance as an orthopedic implant material. Results from

the use of 316, 316L and 317 material as implants have, in general, been good despite

Droblems that are aetailed below. They are regarded as some of the best orthopedic

implant materials available. Some property and composition values are given in Tables i

and II.

Strength, Corrosion and Bio-^Acceptance

Most surgeons and engineers agree that the mechanical properties of stainless

steels type 316, 316L, and 317 are entirely adequate for many implant situations.

Barring design or manufacturing flaws, improper application or insertion of a device,

or critiCdl time-dependedt degradation (corrosion, fatigue, dtc), stainless steel

should function well as an implant material. Unfortunately, the above mentioned

Droblems are very much a reality and some failures of stainless steel implants

occur.
n

Fatigue and corrosion-related fatigue have been isolated as a source of failure

of stainless orthopedic implants. This is not totally unexpected. Orthopedic

devices such as bone plates, hip nails and joint prostheses are subject to cyclic

26 27

loads of a magnitude up to four times body weight ' . Much higher localized

stresses can result from improper insertion, design or manufacturing flaws, and

flaws induced by localized corrosion. The presence of notches, such as screw threads,

28 can drastically reduce fatigue strength . To date, relatively little data is

28 1 available concerning in vivo fatigue processes. Grover and Cohen point out the

difficulties in acquiring complete information on actual, failed implants. However,

29 some information is available. Cohen examined many implants removed for various

reasons and reported definite evidence of fatigue failures. Likewise, investigations

30 31-34

by Cahoon and Paxton and others ~ indicate fatigue initiated by corrosion to be

35 33

a major cause of failure. Wheeler and James and Colangelo report accelerated

fatigue failure of type 316 stainless steel in simulated body fluids as compared

to that in air.

Somewhat less information is available concerning the stress corrosion

failure of stainless steel implants. As mentioned earl ier, it is well known that

austenitic stainless steel is subject to stress corrosion in chloride or caustic

environments. It is then a very real possibility that a highly stressed stainless

orthopedic device in physiological solutions could fail as a result of stress

corrosion. Indeed, some investigators have indicated stress corrosion to be a

31 36 37

problem. Hughes and Jordon , Williams and earlier, Zapffe report cases of

stress corrosion in stainless steel implants. In vitro experiments by Fink and Smathe

indicate that stressing an implant increases corrosion. Other investigators allude

36
12

to stress corrosion as a possible explanation of observed failure of stainless

39 orthopedic devices. On the other hand, recent experiments reported by Galante

showed no evidence of stress corrosion in 316 stainless steel.

Intergranul ar corrosion should not be a problem in stainless steel implants

of the 316 and 317 variety, provided the materials have undergone the proper

heat treatments. Several cases have been reported"^^ ,36,40

^^j^q^^^ faulty heat

treatment (particularly in regard to welded devices) or failure to meet compositional

specifications (including mislabeling) has been isolated as the problem. The best

practice is to use low carbon type (316L, 0.03% carbon max.) stainless steel for

impl ants'^^

.

Crevice corrosion can be a problem in stainless steel implants. Susceptible

locations include, for example, screw hole regions of bone plates or other devices

fixed by screws. Work by Colangelo and Green^^ indicated that 45% (24 out of 53)

of the stainless steel orthopedic devices studied after removal were corroded. The

predominant form of corrosion in the devices was crevice attack. In multi-component

31 42 devices, 91% showed significant corrosion. Other surveys ' of stainless steel

implants removed for various reasons support the above findings. Several in vitro

43 44 studies ' also indicate crevice corrosion and/or pitting to be the predominant

type of corrosion in stainless steel implants. For these reasons stainless steel can

45 46

only be considered marginally satisfactory ' in terms of resistance to corrosion

in the body. The importance of proper passivation and pre-treatment of implants,

47

particular stainless steel, has been noted .

In mul ti -component devices, mechanical disruption of the protective oxide film

46 47

may lead to corrosion attack. It has been suggested ' that crevice corrosion
13

i is initiated by "fretting" of screws and plates and maintained by virtue of the low

oxygen environment of the screw hole. In this respect, stainless steel is probably

' inferior to both titanium and chromium-cobalt alloys in that its protective film is

not as resistant to damage nor so easily repaired once breached. Recent work by

48 49

j Homsy et al_. and Williams has been directed toward this problem of crevice

i corrosion in stainless steel multi -component devices.

i Galvanic coupling of stainless steels of different types or of entirely different

metals is no longer the problem that it was. However, even now mislabeling,

compositional discrepancies, inhomogeneities, or lack of consideration of the phenomena

42

does occasionally result in the failure of an implant. Scales et a]_. reported

instances of incorrect labeling as to type of materials used in implants, notably

steel implants. In several cases, the result was severe corrosion. Other cases of

40 50 incorrect labeling and compositional discrepancies in steel implants ' have been

reported. Note that stainless steel, being less noble, would undergo galvanic

corrosion if used together with either titanium or cobalt alloy components.

Absence of a quantitative definition of biological acceptance or biocompatabi 1 ity

of an implant material, combined with the' lack of any standardized compatabil ity test,

has led researchers down widely varied paths in determining the "degree of compatabi 1 ity

of various materials. Stainless steel is certainly no exception. Having been used

rather successfully for thirty or more years, the data base amassed on the compatabil ity

of stainless steel is extensive; however, much of it is unconnected.

In either bulk or particulate form, intact, passivated stainless steel appears

to be well tolerated by both calcified and soft tissue. That is, barring the

hazard of bacterial infection or purely mechanical irritation of tissue, no major

detrimental effect is ordinarily found. The normal tissue reaction is to isolate the

implant. In the case of solid metal implants, isolation is accomplished by
14

sequestration of the implant by a fibrous tissue membrane ' . This reaction has been

detailed by numerous researchers and is reproducible in laboratory animals. Histologic

study of the encapsulating tissue, particularly in regard to thickness and cellularity,

provides some measure of the compatabil ity of the material. In comparing stainless

51 steels with other prosthetic materials, Laing et al_. found types 316 and 316L

to be as well tolerated as the cobalt alloys and many ceramics. Titanium exhibited

an even greater degree of tolerance. This study also correlated the degree of tissue

reaction with the measured amount of corrosion products observed in adjacent tissues.

The relative toxicity of released ions was considered an important factor. Previously,

50 40 . Cohen and Faultz and Cohen and Hammond had pointed out the importance of corrosion

products, following studies of tissue surrounding corroded stainless steel fixation

42 53 devices. Other reports ' corroborate this picture all of which indicates that

stainless steel may have insufficient in vivo corrosion resistance compared to othen

43 54 implant materials. Various in vitro corrosion studies ' have indicated that

type 316 stainless steel is inferior in corrosion resistance to titanium and cobalt

alloys.

The tissue reaction to stainless steel in particulate form is much the same as

that to the bulk material. If large enough, the particules may become encapsulated.

Otherwise, they are phagocytised ( intercellularly isolated) by histiocytes or in some

instances, remain extracellular and migrate, thus reducing the metallic particle

concentration. Long distance migration of particles or corrosion products is a 39

potential hazard that requires more study. In recent experiments, Galante evaluated

fibroblastic and histocytic reactions in the tissues surrounding various particulate
15

materials. Stainless steel was well tolerated, ranking near a cobalt alloy studied.

No severe or unusual reaction was noted.

55 56 Oppenheimer et_ aj_. and others have reported the possibility of stainless steel

57

being a carcinogen. An earlier study also linked nickel with cancer, bone necrosis

45 52

and inflamation. These reports have been to some extent ignored ' since cancerous

growths have not been positively correlated in practice with the implanation of

45 53 stainless steel or other currently used prosthetic metals. It has been suggested '

that metal ions or corrosion products, in combination with complex host chemistry could

produce antigens and thereby evoke an imnunological rejection of a metallic implant.

Like the potential carcinogenic properties just mentioned, this idea has not been

confirmed. It should be noted, however, that recently Barranco and Soloman^^ reported

a case of allergic skin reaction (eczematous dermatitis) from exposure to nickel from

a stainless steel screw in the patella.

Another source of possible irritation is incompatability between the physical

or mechanical properties of the implant and the host bone or tissue. For instance,

the difference in coefficient of friction between an implant and cartilage is a

possible source of irritation^ ^ . Another is the incompatability resulting from

13

differences in elastic modulus of the implant material and adjacent bone . Such

differences can result in high stress concentrations at the bone-implant interface.

Bone reaction to these stresses is poorly documented, however, instances of bone

resorption and/or loosening of implants and even bone necrosis are known. In the

case of type 316 stainless steel the modulus mismatch with bone is apppoximately

an order of magnitude. Various approaches have been taken which reduce this problem

in cases where modulus mis-match is obviously a disadvantage. The use of self-curing
16

methacrylate as a filler provides a buffer-zone of compatable modulus between the

stiff implant (a hip prothesis for example) and the more elastic host bone. Other

possibilities for reducing the modulus mismatch include the use of porous metal or

60 sintered fiber metal materials . These ideas and materials will be discussed in

subsequent parts of this report.

To conclude, it is clear from some thirty years of successful implantation

that stainless steel is well tolerated in most instances. However, considerable

evidence indicates that a build-up of corrosion products in tissues adjacent to the

implant can occur and evoke acute reactions. With this in mind, the long time

compatabi 1 i ty of stainless steel must still be questioned. In cases where implants

are expected to remain in situ for many years, it may be appropriate to use materials

deemed more corrosion resistant, provided their mechanical properties are adequate.

Further, the mismatch in elastic moduli of bone and stainless steel can have serious

biologic implications.

I Summary

At present, stainless steels (316, 316L and 317) are used for many prothestic

6

1

implants. Results are usually quite good, however, there are failures reported.

Such failures can usually be traced to one of four broad problem areas; surgical

insertion, biologic-pathologic problems, materials failure, or a combination of any

of the above. Of primary interest in this report are materials failures or combinations

of problems involving materials failures, i.e., corrosion and subsequent biologic

reaction to corrosion products or released ions. With this in mind, some comments

regarding stainless steel as an implant material follow.
17

1. Type - AISI types 316, 316L and 317 are used almost exclusively today.

In the United Kingdom, 316, 316L and 317 may be covered under one stanuard, EN58J,

or more recently appear as 316SI2, 316SI6 and 317ST6. Most other countries use the

standard AISI codes.

2. Physical Properties - Stainless steel is a strong and a ductile material.

The physical properties are more than adequate for most orthopedic implant situations.

Corrosion can degrade the physical strength of a device with time. The modulus of

elasticity of stainless steel is much higher than that of bone. In temporary, short

time rigid fixation (bone plating) this may be an advantage. In the case of a

permanent prosthesis, however, adverse reaction to localized hiqh stresses induced

by modulus mis-match can occur.

3. Fatigue - Stainless steel can and does fail in situations of cyclic stress,

provided the endurance limit of the metal is exceeded. Corrosion can accelerate the fatigue process. Fatigue failures occur in orthopedic implants of stainless steel. This is largely a design problem, and can be overcome in many cases.

However, long-time implant situations involving cylical loading constitute a difficult area.

4. Corrosion Resistance and Bio-Compatability - In the human body, stainless

steel is subject to localized corrosion by various mechanisms. Pitting and crevice

corrosion are the greatest problems. Evidence also supports the existence of stresscorrosion-cracking and corrosion-fatigue as failure mechanisms. Intergranular corrosion does not appear to be a problem since the introduction of 316, 316L and 317 as

implant materials. However, in view of the corrosion problems found, performance of

stainless steel implants roust be considered only marginally satisfactory. Stainless

steel is well tolerated by adjacent tissues, provided corrosion does not Occur to any

great extent. There are some reports of severe tissue reaction, however. Additional

research and development is required to improve these alloys for long term and critical

function applications.
18

'j. Manufacture - Stainless steel is the least expensive and most formable

implant metal of those in current use. Recent vacuum melting techniques have

improved the metal quality. Controls on composition accuracy are important, since

mul ti -component devices are manufactured from different production lots.
19

i III. COBALT ALLOYS

I Physical Properties

Cobalt-based alloys have found many engineering uses, including magnetic

j materials, high temperature creep-resistant materials, electrical resistance alloys,

cutting tool materials, and in the dental and medical fields. The elements and

proportions added to the cobalt-base will be determined by the requirements of the

application. Elements usually added are chromium, nickel, molybdenum, titanium,

tungsten, and iron. Cobalt alloys may exist in either cubic or hexagonal crystal

structure, or a mixed phase of those two, depending on the alloy composition and

processing treatments. While the alloys of interest generally have large amounts

of other element additions, many of the characteristics of pure cobalt are also

found in the alloys. These include low coefficient of friction, good wear

resistance, good resistance to oxidation and the effects of sulfurous environment,

and excellent high temperature properties. There are many heat-resistant alloys,

for example, HS31, HS36, HS95 and others. They were developed since 1941 for hot

gas turbine applications, and will not be discussed further here. However, one of

the wrought alloys, HS25, is used for implant applications.

The cobalt alloys vary considerably in strength and ductility. Both cast

14

alloys and wrought alloys are available . Yield strengths vary from 65 to 175,000 psi,

depending on composition and structure. The elastic modulus ranges from 29 to 36 x 10^

psi depending on the alloy composition. The wrought alloys are hardened by precipitation, usually chromium-carbide precipitates, by solid-solution hardening,

usually tungsten and chromium solute additions, and by plastic deformation hardening

during forming operations. The cast alloys are hardened by the first two mentioned
20

mechanisms. Recent production advances include the use of high purity raw materials,

vacuum melting (particularly useful to reduce porosity) and improved de-oxidizing

methods. These advances have led to improved corrosion resistance, ductility and

strength

Corrosion Behavior

The cobalt alloys generally exhibit a high resistance to corrosion, even in

aqaressive environments^^. They show a good resistance to sulfurous corrosion, much

superior to stainless steel in that regard. The high chromium content contributes

significantly to this property. In most environments a passive, adherent, oxide

1

9

film develops on the alloy, consisting in part of chromium oxide layers . Cobalt

alloys are significantly more resistant to nitric acid (10 to 40% concentration)

14 sodium hydroxide, and concentrated acetic acid than is austemtic stainless steel

Nickel additions and molybdenum additions (about 10%) to the wrought alloy improve

corrosion resistance. The alloys will differ in resistance to local pitting or

crevice corrosion, depending on composition and other factors. However, in most

applications involving aqueous, saline environments, local corrosion has not been

noted as a problem. Galvanic corrosion can occur when the cobalt alloy is coupled

with a less noble metal.

Cobalt Alloys as an Implant Material

The alloy Vitallium was first used as an implant material in 1937. Originally,

the alloy had been developed and applied as a dental inlay material. Improvements

in casting methods, inspection, and machining techniques have led to the production

of a wide variety of orthopedic devices from this alloy. Other applications include

artificial valves for the human heart, using the cast HS21 alloy^^. One example of
21

production problems that can develop through poor casting is shown by Williams .

One difficulty that remains today concerns the casting of small objects (i.e., screws)

where the large grain size may result in having only a few grains across the section

45

of the object, a potentially weak situation . The initial low ductility of the

cast alloy can be significantly improved by solution heat- treatment procedures that

lead to a more homogeneous microstructure. New cast alloys have been developed in

recent years, principally for high temperature applications, however, they have low

53

ductility (order of 1%) and this would limit most implant applications.

Wrought cobalt alloys that were suitable for implant applications became

available in 1952. Several advantageous properties are offered by alloys such as

HS25 (also referred to as wrought Vitallium, an unfortunate term since the composition

and properties differ considerably from Vitallium). These include improved ductility,

higher strength, greater fatigue strength and impact strength. The composition is

given in Table III together with other pertinent specifications. The chromium

concentration is reduced from about 30% to 20%, while about 10% nickel and about

15% tungsten are added, relative to cast Vitallium. There are important property

differences between the wrought and the cast alloys that will be discussed below

(see also Table I for mechanical properties values).

Implants constructed from the cobalt-base alloys now enjoy a wide acceptance

throughout the medical world, A great deal of data has been accumulated from both

research and clinical studies that justifies much of the confidence in applying these

alloys. Simple fixation devices from both wrought and cast alloys find constant

application; complex devices such as the total hip prosthesis in cobalt alloy form
22

have been implanted in many individuals. Principal findings and conclusions from

both research and clinical studies are summarized in the following sections.

Implant Strength, Corrosion and Bio-Acceptance ..

:. Mechanical properties of both the cast and wrought form of cobalt-based alloys

are summarized in Table I, together with the current ASTM specification values.

The range of strengths and ductilities for the wrought alloys arise from different

thermo-mechanical processing methods. Increased forming deformation leads to higher

tensile strengths and to low ductility values. It is seen that considerably improved

strength values and ductility values can be obtained in wrought alloys compared to

cast alloys. Ductility is important whenever shape changes must be imposed during

installation of a device to obtain a proper fit. Certain devices lend themselves to

production by casting; others due to size or shape require mechanical forming. The

hardness of cast Vitallium is greater than the wrought material, and this is of

importance in joint prostheses, such as the femoral head of an artificial hip. Note,

however, that the impact strength of the cast alloy is only about 10% that of wrought

HS25. The elastic modulus, hence the stiffness, of these alloys is slightly higher

32

than stainless steel. This could be advantageous where rigid fixation (for example,

of a bone fracture) is required. The high surface hardness of these alloys has the

further advantage of minimizing damage and scratching during installation of an implant

Since cobalt-base alloys are in use in articulating prosthetic designs (total hip

or knee devices) it is important to consider the wear and friction properties of the

63

material. Moral discusses some data obtained from vacuum studies that show the

importance of surface crystal lographic orientation and structure. He points out that

relatively low wear rates are found in most applications but that little data was then
23

available pertaining to implants. Recently, several studies have been reported

concerned with wear in simulated prosthetic joint situations. Duff -Barclay and

Stillman^'^ used a simulator machine under dry, saline, and blood plasma environments.

They found cobalt alloy metal -to-metal prostheses to give the lowest wear rate.

Titanium and stainless steel materials exhibited high wear rates and were judged

68 impractical. Scales et al_. also concluded that a cobalt-chromium-molybdenum

alloy bearing on itself was the most likely long-term satisfactory combination.

They emphasized the importance of surface finish on the frictional values. They

found that the alloy bearing in a cup of RCHIOOO polymer (polyethylene) gave even

lower frictional torques and wear rates (although long-time use was not as favorable).

69

Recently, Galante and Rostoker have also reported that Vitallium bearing on

RCHIOOO showed the lowest wear rates of several metals and ceramics tested. It

appears that cobalt-base alloys do offer good wear resistance, hardness, and acceptable

friction values for joint prostheses applications.

There is ample evidence in the literature of the generally high corrosion

42

resistance of cobalt alloy implants. In a 1959 study. Scales et al_. reported on 15

cobalt alloy implants removed from patients, some of which showed evidence of

corrosion. The other implants removed in this study, a total of 549 mostly ironchromium-nickel alloys, showed evidence of corrosion in 228 instances, about 41%.

Scales et al_.^^ reported no sign of corrosion in two cobalt alloy hip nails removed

43

from patients after 5 and 28 months. Hoar and Mears concluded that vitallium alloys

could tolerate long exposures to chloride solutions, as a result of a comparative

44

study of different materials. Subsequently, Bultitude and Morris reached the same

conclusion after studying three vitallium implants in vitro . On the basis of this
24

and other evidence, it appears that general corrosion of the cast cobalt alloy

implants is minimal. However, there is evidence of local, pitting corrosion problems

and of a variation of corrosion amount with composition of the cobalt alloy.

Rose et aj_.^^ report a study of a removed nail plate, broken in situ , where

subsequent analysis showed the nail to be Vitallium and the plate HS25, wrought

alloy. The failure was identified as a fatigue-induced fracture, influenced by

corrosive pitting attack associated with the different corrosion resistance of these

72

two alloys. In vitro studies by Arndt et al_. have also shown that HS25 is more

susceptible to crevice corrosion than is cast Vitallium. Coupling of these two

different cobalt alloys should be avoided. The accumulated experience with cast

Vitallium implants should not be thought to apply directly to wrought alloys of

other compositions.

The surface preparation of implants may have a significant effect on the

73

subsequent behavior. Revie and Greene studied a wrought cobalt alloy in vitro

and found a significant, beneficial effect of pre-treatment in establishing a passive

film (initial polarization current densities were decreased by a factor of over 100).

They also found an improved resistance of electropol ished surfaces relative to

74

rougher, finished surfaces. Recently, Homsy et al_. have considered the problem

of tissue and bone adhesion to surfaces of cobalt alloy prostheses. While these

cast surfaces show excellent corrosion resistance, the oxide film present together

with the as-molded surface roughness produces an interface that permits some ingrowth

or adhesion. Such a characteristic can make removal of the implant a difficult

procedure. These authors found that filling the surface pores with a polymer by

a burnishing technique can reduce this adhesion problem.
25

The biological acceptance of cobalt alloys has been demonstrated by over

30 years of implantation of various devices. An even longer period of successful

dental applications (referring to cast Vital 1 ium material ) can be cited. However,

specific long-term data, greater than 10 years continuous implant time, is difficult

to locate. In a 6 month study of tissue reactions to implantation of both cast and

51

wrought cobalt alloy specimens in rabbits, Laing et aj_. found satisfactory tissue

acceptance of both alloy groups. While the concentration of released elements was

lower for the cobalt alloys than for titanium, the tissue reaction was thicker for

the cobalt alloys. A significant release of elements into the surrounding tissue was

found, even in this short period of time. An excellent discussion of methods and

the results of determining tissue tolerance has been given by Cohen^. He notes that

characteristically, cobalt alloy implants show thin scar encapsulation with no continuing inflamation. However, the implant is never strictly inert; some interaction

is always noted. An indication of possible long term problems of great concern is

given in the report of a carcinogenic response of rat muscle to wear particles from

75

a cobalt-alloy prosthesis. Heath e^ a]_. injected wear debris from a machine-operated

prosthesis into muscle tissue. Subsequently, 14 tumors were detected in periods of

time up to 15 months; 7 of the tumors were analyzed. The authors point out that

finely divided cobalt alloy wear debris may pose a serious long term problem in humans

and that little if any relevant data is on hand.

Summary

Cobalt alloy prostheses and fixation devices have been implanted with generally

successful results for over 30 years, following even longer experience with application

in the dental area. The good corrosion resistance, biological acceptance, and mechanica

properties of these alloys indicates continuing use. This material was one of the
26

first used for artificial heart valves, chosen because of the good experience to

date. Some problems that appear worthy of current attention include corrosion in

mul ti -component devices using different cobalt alloys, local corrosive attack on

wrought alloy surfaces, and the body tolerance of wear particles from articulating

devices. A summary of the properties of these alloys is as follows:

1. Type - Both cast cobalt-chromium-molybdenum and wrought cobalt-chromiumnickel -tungsten alloys are available as orthopedic devices. The longest experience

is with the cast alloy (Vital lium type).

2. Physical Properties - The alloys show a high hardness, high elastic modulus,

and moderately high strength and ductility values. The porosity and impurity content

of the cast alloy has been markedly reduced in recent years by production controls.

Greater ductility and strength are found in the wrought alloys. Wear rates (between

identical surfaces) are relatively low, and excellent surface finishes can be obtained

and maintained.

3. Fatigue - The fatigue limit is typically about 30 to 40% of the tensile strengt

a similar ratio as that of austenitic stainless steel. Corrosion induced fatigue

failure has been observed in an implant.

, 4. Corrosion Resistance and Bio-Compatabil ity - The cobalt alloys of interest are

not subject to significant general corrosion in the human body. Examples of local,

pitting or crevice attack have been noted on the wrought alloys. Biologic acceptance of

the cast alloy is excellent; the usual tissue reaction is minimal. There is some

concern over body response to wear particles of this alloy, particularly over long

time periods.

5. Manufacture - The investment-casting of Vital lium and related alloys is an

established procedure, applicable to many devices excepting small sizes and unusual

shapes. The wrought alloy types are possible to heat-treat and mechanically form

into a variety of shapes. Machining and polishing methods are established.
27

IV. TITANIUM

Physical Properties

Titanium is a relatively new engineering material. Despite the fact that

it was first discovered as early as 1791, it was not available on a commercial

76

basis until about 1950 . Prior to corranercial production, titanium was in many

cases a chemical nuisance - an impurity. Today, titanium is surpassed in commercial,

66

structural uses only by aluminum, iron and magnesium .

Currently, titanium is manufactured by the Kroll process. Discovered by

W. J. Kroll in 1938, this process made it economically feasible to produce

titanium on a commercial basis. In 1947, before implementation of the Kroll

process, the cost of a pound of titanium was approximately $3000. By the

end of the 1950's, titanium was priced at about $3.00 a pound^''^. The process

takes titanium from the ores rutile (TiO^) or ilmerrite (Fe0-Ti02) and combines

it with chlorine to form ti tanium-tetrachloride (TiCl^), a heavy colorless liquid.

The titanium-tetrachloride is then reduced in an inert atmosphere with molten

magnesium or sodium to form magnesium chloride (MgCy or sodium chloride (NaCl)

and pure titanium in the form of a porous solid (sponge). The porous solid is

melted (again under an inert atmosphere) in an electric furnace and cast into

an ingot. •.

The metal resulting from the above process is light, strong and ductile.

The density is about 60% that of steel. The strength-to-density ratio is greater

for titanium than annealed austenitic stainless but less than cold worked stainless^^.

The mechanical properties of unalloyed titanium are strongly dependent on the impurity

content of the finished product. Titanium has a strong affinity for the gases
i 28

1 hydrogen, nitrogen and oxygen. Hydrogen can be absorbed at elevated temperatures

i (from water vapor, acids, oils, or other hydrocarbons. Absorbed hydrogen beyond

I \0.02% can cause severe embri ttlement. Fortunately, most absorbed hydrogen can be

M removed by vacuum annealing. Oxygen and/or nitrogen can also embrittle titanium.

; (However, within specified limits, contamination by these elements can have a

I strengthening effect. There are four grades (ASTM) of conmercially pure titanium

available in the U.S. Grades 1 and Z are the most pure being at least 99,2% titanium,

' with tensile strengths from 40,000 to 60,000 psi. The tensile strengths of grades

I 3 and 4 (99.0% pure) are from about 70,000 psi to 95,000 psi.

1 Grades 1 and 2 are slightly more ductile than the higher grades. The higher

' grades are more sensitive to stress concentrations (notch sensitivity) and have

reduced toughness. Fatigue strength is very good, about equivalent to stainless

I steel, and increases slightly with grade. The modulus of elasticity is 15.0 x 10^ psi,

slightly more than half that of stainless steel. Unalloyed titanium is not hardenable

by heat treatment. Some hardening by plastic deformation is possible but not nearly

to the extent of stainless steel.

At room temperature, titanium exists in a hexagonal structure, termed the

alpha phase. At 1625°F a transition to a cubic structure, the beta phase, occurs

: which remains stable to the melting point of approximately 3140°F. By selective

I addition of alloying elements, advantage can be made of this two-phase behavior and

I a wide range of properties can be obtained. There are three basic alloying effects.

: First, additions of aluminum will raise the alpha phase transition temperature.

I These alpha alloys have several attributes: weldability and retention of strength

I and oxidation resistance at elevated temperatures. Second, some elements when added

to titanium do not affect the transition temperature but act simply as a solid

I
29

solution strengthener. Tin and zirconium are often used to achieve this effect.

Third and most important is the alloying effect of chromium, columbium, copper,

iron, manganese, tantalum and vanadium. Additions of these elements reduce the

alpha-beta transition temperature so that a stable two-phase alpha - beta alloy is

obtainable at room temperature. These two-phase alloys can be significantly

strengthened by heat treatment. One example is titanium-6% aluminum, 4% vanadium

(Ti-6A1-4V). Over 50% of the titanium produced is used in making this alloy.

67

It shows considerable promise as a prothestic metal

In the fully annealed state, Ti-6A1-4V has a tensile strength of 135,000 psi

and a yield strength of 120,000 psi. Heat treatment can increase the tensile

strength to 170,000 psi and the yield strength to 150,000 psi. The strength of this

alloy matches that of all but the most heavily work-hardened stainless steel. The

modulus of elasticity, 16.5 x 10^ psi, is near that of the commercially pure titanium.

Ti-6A1-4V is less ductile (8% elongation to failure) and harder than unalloyed

titanium. It has a higher fatigue strength than pure titanium, being over 50% of

the tensile strength. This ratio is better than most stainless steels. However,

tests on notched Ti-6A1-4V show a decrease in fatigue/tensile strength ratio to

about 20%.

In summary, titanium and some titanium alloys (Ti-6A1-4V) are strong, light

materials. The fatigue strength of titanium and its alloys are generally good.

Impact resistance is likewise good. Additional information can be obtained from

14

the ASM Metals Handbook and manufacturing sources.
30

Corrosion Behavior

18 23

The corrosion resistance of titanium is very good ' . At ordinary

temperatures, commercially pure titanium is nearly unattacked by strongly oxidizing

acids, aqueous chloride solution, moist chlorine gas, sodium hypochlorite, sea

water, brine solutions and many other equally corrosive mediums. Most titanium

alloys are nearly as resistant. Titanium and its alloys are protected by an

inert, tightly adherent, passive oxide film. When passivated in this manner, titanium

acts as the noble metal when coupled galvanically with all structural alloys. Consideration must be given to this relation if titanium is used in conjunction with

another metal in a multi -component device in a corrosive medium.

Titanium and most of its alloys are rather immune to localized corrosion, such

as pitting and crevice corrosion. However, in locations of restricted geometry in

an acidic and a low oxygen environment, crevice corrosion can take place in titanium.

78

McMaster reports that the fatigue behavior of titanium is nearly unaffected by

corrosive media, although a slight decrease in fatigue strength was noted in brine

14 solutions. Other sources corroborate this observation. For practical purposes,

stress corrosion cracking is not a problem in titanium or its alloys.

In certain applications, wear is a significant problem in titanium and titanium

alloy devices. In cases of titanium bearing on titanium or on another metal, wear

may be excessive despite lubrication. Galling is one significant type of wear^^.

Surface treatments and/or alloying may provide an answer to this problem.

Titanium and Titanium Alloys as an Implant Material

As early as 1940, research was being conducted to determine if titanium could
31

be used as an implant material. In that year, after conducting animal experiments,

79

Bothe et aj^. reported that "titanium was fully as well tolerated as stainless steel

80

and Vital lium." In 1951, after experimental work on rabbits and rats, Leventhal

81 came to the same conclusion. In vitro tests by Clark and Hickman further confirmed

the inertness of titanium in physiological solutions. About this time, titanium

was beginning to be produced commercially. The price of the material dropped

82

and quality improved. By 1957, Leventhal had reported on fifteen cases in which

titanium prostheses were inserted in humans. By the mid-1 960 's nearly every type

of implant previously made with stainless steel or cobalt alloys had been successfully

constructed from titanium^^.

Acceptance of titanium as an implant material, particularly for orthopedic use,

has been slow. This is especially true in the United States. Reasons for this

reluctant acceptance are many. First, for approximately thirty years, stainless

steel and cobalt-chromium alloys have proven generally successful, although certainly

less than perfect. Hence, there is a hesitance to adopt a new implant material that

would require modified procedures, additional experience, etc. Second, until

recently, the cost of raw titanium and titanium products was high relative to those

two alloys. Thirdly, there still remains some doubt among manufacturers and

physicians whether titanium and titanium alloys are actually superior to other

currently available prosthetic metals. Titanium and titanium alloy devices are in

clinical use at this time. Widely varied research studies are still being conducted.

The following sections summarize the major conclusions thus far.
;32

I Implant Strength, Corrosion and Bio-Acceptance

I As discussed earlier, the yield strength of grades 1 and 2 titanium is

1from 40,000 to 60,000 psi. Rose^ and others have noted that this level of

: (Strength may be insufficient for bone plate and other stressed applications.

I V However, clinical evidence^^ '^^ available indicates mechanical failures of titanium

: care not a problem. Naturally, other factors enter in here, such as degree of load

I "bearing and duration of implantation. In applications where strength is a factor,

I (the stronger Ti-6A1-4V alloy should prove adequate. Table I lists some mechanical

I ; properties values for the two alloys.

, < In a previous section the modulus of elasticity of implant materials was

discussed in terms of its effect on adjacent tissue. There are two schools of

I > thought here. First, there are researchers and clinicians who feel that a stiff,

high modulus material such as stainless steel is a necessity if completely rigid

45 64

I fixation of fractures is to be accomplished ' . In this view the lower modulus

of titanium is a disadvantage. However, adverse stress-induced biologic reactions,

implant loosening, and reduced damping or energy absorbing ability have led to a

3 77

second school of thought. Katz and others have indicated that a modulus of elasticity

more comparable to that of bone would be desirable, especially in cases where

removal of the implant is not anticipated. In this view the stresses are transferred

between the implant and the bone more uniformly. Here the lower modulus of titanium

and titanium alloys is an advantage.

The fatigue strength of titanium is approximately the same as that of the

austenitic stainless steels. The fatigue properties of the Ti-6A1-4V alloy appear to
33

be better than stainless steels currently in use for surgery . Fatigue failures

have not been reported clinically.

The corrosion resistance of titanium and Ti-6A1-4V in saline solutions has been

43 shown to be high. Hoar and Mears studied titanium in vitro and found a greater

resistance to corrosion that for stainless steel and the cobalt alloys. Mueller

54 44 84

and Greener and others also report a high degree of passivity. Fraker et al

have reported electrochemical measurements and electron microscope surface studies

of titanium and several alloys. They find very good resistance to attack, but

point out the need for careful initial surface preparation. Available information-.

indicates that titanium may be the most corrosion resistant implant material among

the principal three materials.

The tissue reaction to titanium is that found for most compatible materials,

isolation in a fibrous tissue envelope. Small particles of titanium may be

51

phagocytised or remain extracellular. Early work in rabbits by Laing et al_. indicated

that titanium was equally as well tolerated as stainless steel or cobalt-based alloy

materials. More recent experiments confirm this, and in fact, it appears that titanium

and Ti-6A1-4V may actually be better tolerated than stainless steel or cobalt alloys.

39

Galante reports good tissue acceptance of bulk titanium and titanium particles.

85 86 Likewise, Predecki et aj_. ' indicate excellent tolerance of titanium in tissue

adhesion studies in animals.
34

3

In connection with the question of bio-compatibility, it is appropriate to carefully

consider the growing interest and work on porous implants. In such materials the

surface/volume ratio is greatly increased relative to a solid, bulk implant. The

f purpose is to promote adhesion by tissue ingrowth. The large tissue-implant interface

5 places particular demands on the compatibility and corrosion resistance of the implant

1^ material

.

87

g. After initial success in fabricating a porous cobalt alloy implant , several

88

l3 groups became interested in titanium materials. Hahn and Palich applied a titanium

^ oxide coating on a wrought titanium implant. They reported no compatibility problems

60

p and a large increase in fixation strength. Galante et al_. fabricated and studied

ingrowth on sintered fiber titanium aggregates, having bulk densities between 45%

and 65%. Implantation in rabbit and dog femora demonstrated ingrowth and good shear

I strength, in the range of that of bone. Following a previous study of porous cobalt

89

(.alloy materials, Hirschhorn et al_. prepared and studied porous titanium and porous

^Ti-6A1-4V materials. They also reported good ingrowth and tissue acceptance, noting

( however, the need for careful corrosion studies on such materials. Recently, Nilles

90

. et al^. have studied porous Ti-6A1-4V materials, having controlled pore size and

91

, density, with applications to knee prostheses attachment. Young has reported a

f current study of porous titanium dental implants.

, Wear is a sig-nificant problem in applications of titanium to articulating

I surfaces. In view of the recent interest in total joint replacement, serious

consideration has been given to this matter'^^'^^ Recent in vitro experiments

69

by Galante et al_. and others concerning wear in artificial joints have indicated

the difficulties. On the other hand, clinical results of the orthoplasty series

84

of Emenus et al_. did not indicate a wear problem. With the aim to produce joint

protheses with projected lifetimes considerably longer than at present, wear phenomena
35

must be considered in greater detail. Investigation into specific surface treatments

or perhaps new alloying combinations is essential if titanium materials are to be

used in this application.

Most of the available information concerning the clinical use of titanium as an go

orthopedic implant material comes from abroad. Emneus and Berg reported on a series

of 110 Moore arthroplasties done between 1961 and 1965. Of the prostheses used,

55 were titanium and 55 were a cobalt alloy. No differences were reported between

the groups. In cases where titanium appliances were removed, there was no apparent

irritation of tissue even though in some instances where mul ti -component (plates and

screws) devices were used, a discoloration appeared. Part of this observed pigment

52 was phagocytised and part remained extra-cellular. Previously, Emneus and Stenram

reported on favorable tissue acceptance of titanium in animal experiments. Although

compatibility was excellent, in some cases a dark pigment (Turnbull negative) was noted

in tissue adjacent to the implant. Subsequent x-ray spectroscopy and quantitative

analysis revealed titanium to be the principle constituent of the discoloring pigment.

Long range implications of this finding have not been reported.

41

More recently, Williams and Meachin reported examination of implants and

adjacent tissue in 190 cases. Of these 190 implants, 49 were of titanium. The

titanium implants were well -tolerated and did not corrode to any observable extent.

Local tissue, however, showed titanium contents (dark particles) in concentrations

that varied from patient to patient. The maximum found was about 0.2%. There was

no definite adverse reaction to released titanium observed. However, the authors

suggest a more detailed study of the possibility of systemic release of titanium
36

and also distant hypersensitivity effects, if any. Additional information concerning

the compatibility of titanium comes from the field of oral surgery. At present, titanium

92

oral endosteal implants are being used routinely with clinical success . These blade

implants are inserted directly into the alvellar bone ridge. Stabilization is dependent

92

upon bone growth around and into vents in the implant. Weiss et al_ report that

titanium is the metal of choice in this application, and that 90% of the endosteal

implants inserted in the free world are titanium.

Summary

Virtually all implant devices constructed of stainless steel or cobalt-based

alloys have been also constructed from titanium. The use of titanium and its alloys

is more widespread abroad than in the U.S. Available clinical reports are generally

favorable, particularly in regard to corrosion resistance and compatibility. There

is some concern among researchers and clinicians that unalloyed titanium does not

have the mechanical strength to perform adequately in highly stressed situations.

As of yet, however, clinical results do not reveal this problem. Experimentally,

wear appears to be a problem in articulating prostheses. These findings and

others are summarized as follows:

Type - Several grades of unalloyed titanium and a Titanium-6 Aluminum-4

Vanadium alloy have been used as implant metals.

2. Physical Properties - Unalloyed titanium is not as strong as stainless steel.

The stronger Ti-6A1-4V alloy can be used where high strength is required. Elastic

modulus mismatch with bone is less with titanium and titanium alloys than with

stainless or cobalt-based alloys.
37

3. Fatigue - The fatigue properties of unalloyed titanium are as good as most

stainless steels. The fatigue properties of Ti-6A1-4V are better than stainless steel.

4. Corrosion Resistance and Bio-Compatibility - In the human body, titanium

and its alloys do not appear subject to significant corrosion, local or general.

Titanium alloys are very well tolerated by adjacent tissues. Some titanium is

apparently released resulting in dark local pigmentation of the tissue.

5. Manufacture - Titanium and its alloys are nearly as formable and machineable

as stainless steel, if proper equipment is available and procedures followed. The

machineability of Ti-6A1-4V is as good as stainless steel. At present, costs

of finished products should not be more than 20% higher than stainless steel and

about the same as cobalt-based alloy devices.
38

j

V. OTHER METALS

Early history of the use of metallic surgical implants includes many metals

that are no longer thought suitable. These include iron, silver, bronze, aluminum,

24

brass, magnesium and others . However, there are other metals and alloys that '

are in current use or thought of as being worthy of further study and evaluation,

as prosthetic implants. These will be discussed briefly below.

Tantalum

The metal is used as wire sutures and gauze for abdominal wall repair. It is

] quite inert and compatible, however, its low strength prevents application to structural

repair problems. Considerable tissue reaction in terms of the membrane thickness

51 surrounding the tantalum implant has been reported .

Zirconium

Some results are available on the implantation of zircalloy (1.5% tin, 0.15% iron,

93

0.1% chromium, balance zirconium) in animals. Galante and Rostoker implanted specimens

j in the back muscles of rabbits for twelve months. Evidence of pitting corrosion was

found in two of the four specimens. The authors concluded, however, that the alloy 1

was suitable for consideration as an implant material, without appearing to offer I

51

any advantages over presently available materials. Laing et al_. report that measurements

on the membrane thickness surrounding zirconium implants are quite variable; a strong

tissue response is found.

Ni tinol

This alloy is a nearly equi-atomic binary alloy of nickel and titanium, first developed

at the Naval Ordinance Laboratory, Washington, D. C. By suitable processing, items can

be fabricated from this alloy that will undergo a specific, controllable shape change

at some later time, as a result of an impressed thermal change. Implant applications
39

include bone plates, fasteners, and artificial muscle fibers. Corrosion data and

94 95

bio-compatibility data are becoming available and early results ' appear encouraging.

Cobalt-Nickel-Chromium-Molybdenum Multi -Phase Alloy

96

This wrought, quarternary alloy is both age-hardenable and work-hardenable.

It is quite different in metallurgical properties from the cobalt alloys discussed

earlier. Plastic deformation as a result of forming operations produces a mixture

of cubic and hexagonal phase. Heat treatment can bring about additional strengthening

through precipitation of a Co^Mo phase. Yield strength values from 60,000 to over

96

200,000 psi are attainable , with elongations from 70% to 10%, respectively.

The corrosion resistance in saline solutions appears good. Preliminary implant

studies in animals have been conducted. Tissue compatibility in rabbits appears as

good as Vital lium.
40

CONCLUSIONS

The lack of an ideal implant material is made clear after one becomes familiar

with the results to date using presently available materials. In many cases, compromises

are made, for example, in the tensile strength of a material, in order to obtain

improved corrosion resistance or acceptable tissue response. There is clearly room

for further development of new materials for implant purposes. It is the authors'

view that this probably involves only minor modifications or deviations from the

presently available materials, rather than the introduction of completely new

materials. The period of time required for a thorough and accurate evaluation

of a new implant material is so long and the entire testing protocol so extensive,

that completely new thrusts become a major effort. Further, it appears to us that

a minor tailoring of a set of a few basic materials to suit specific implant

problems may produce the "arsenal" of implants that the surgeon requires. As an

example, an extensive choice of alloys with their associated properties is available

in the cobalt-base system. Very high strength wrought alloys exist and could be

further improved for specific purposes, such as fracture fixation devices. Other

I cobalt alloys show excellent wear resistance for consideration as artificial joint

prostheses, particularly if further tissue tolerance data are obtained in connection

with the wear debris that is produced. The problem of variation in the susceptibility

to local corrosive attack among the cobalt alloys of different compositions can surely

be attacked and solved by minor alloy modifications. It is hoped that this report

will highlight some of these specific areas that need attention.

In attempting to compare the results of studies by different investigators on

the same implant material, the need for a more uniform method of testing and evaluation

became apparent. Laboratory corrosion testing, for example, is today a relatively

well-defined area. Procedures are established and validated. Further studies of

implant materials would be well served by establishing a specific procedure for
41

determining the corrosion resistance and tissue compatibility of implants in vivo .

Many excellent test methods already exist and need only be formulated together into

one program. Joint activity among the American Academy of Orthopedic Surgeons,

the American Society for Testing and Materials, and other appropriate groups is

envisaged. Guidance could be obtained from the materials screening and evaluation

activities associated with the artificial heart program supported by the Federal

government through the NIH.

One implant area worthy of special attention concerns artificial joint

prostheses. It appears that a significant increase in the use of this surgical

procedure should be expected in the next decade. The success in total hip

implantations in recent years is certainly outstanding, and this effort will be

expanded to other joints of the body. Serious materials problems exist here today,

primarily involving wear, adequate fixation, and body tolerance to wear and

corrosion products. A programatic approach to activities in this implant area

would be of benefit to all concerned.

ACKNOWLEDGEMENTS

Many individuals have contributed information and ideas during the period of

time that this report was being assembled; their assistance is gratefully acknowledged

One of the authors (RLP) visited many of the laboratory groups whose work is mentioned

here. The information obtained in that way was an important element of this

study and their cooperation is most appreciated. Special mention should be made of

the assistance received from Anna Fraker and Wilford Nusser throughout this activity

One of the authors (RLP) was a Presidential Intern while conducting this study and

acknowledges the support obtained from that program.
i
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I

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Table II. Con^osition and Specifications for Stainless Steel (amounts in

weight percent) for Surgical Implants

.

Element Material

316L ASIM

F56-71

Specification

(Type B)

Chromium

Nickel

Molybdenum

Manganese

Carbon

Phosphorous

Sulfur

Silicon

Iron

16-18

10-14

2-3

2

0.03 max,

0.045

0.03

1

bal.

17-20

10-14

2-4

2 max.

0.03 max.

0.03 max.

0.03 max.

0.75 max.

bal.

Note: British specifications are nearly the same (BS 3531; 1968).
Table III. Composition and Specifications for Cobalt Alloys (amounts in weight percent)

for Surgical Implants.

Element Material

Vitallium, cast

(HS21 similar)

Wrought alloy

(HS25)

ASTM

F75-67

Specification

Cast

ASTM

F90-68

Specification

Wrought

Chromium 30 20 27-30 19-21

Nickel 10 2 . 5 max. 9-11

Molybdenum 5 0.1 5-7

Manganese 0.5 1.4 1 max. 2 max.

Tungsten 14 14-16

Carbon 0.4 0.01 0.35 0.05-0.1,

Silicon 0.3 0.5 1 max. 1 max.

Iron 0.7 0.2 0.75 max. 3 max.

Cobalt bal. bal. bal. bal.

Note: British specifications are nearly the same (BS 3531; 1968).
Of;

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BIBLIOGRAPHIC DATA

SHEET

irpirni lCATION OK KlfPORT NO.

.„..W61UL.7J3-42Q

Tiri.r'; and suhmti.e

Survey on Metallic Implant Materials

Al!TJIOR(.S.) ._iL.-.R^.Piiiis0ns__and-A,_.lii-_Ruff-.

I PFHFOHMING (M(G ANiZATlON NAME AND ADDKESS

i NATIONAL I31IKEAU OF STANDAIUXS

I . DliPARTMENT OF COMMERCE

I WASHINGTON, D.C. 20234

2, Gov't Accessioil \ 3. tlvci pit nt'.s Act:'js:. ion N

No. I

i 1?. ^f.--.)H.soting Organization Name atul Address;

I

I National Institute of Arthritis, Metabolism & Digestive

I Diseases

j Westwood Bl-dg., 5333 Westbard, Washington, D. C. 20016

5. J-'i.jhli= ;\rion Dare

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14. Spon r-.onnj; Agr.ncy

I 15. SUPPLEMENTARY NOTES

i 16. A!5ST[<A(..T (A 200-word or lt>?s factual summary o{ most significant information. If document include.', a f,i;;nifir ant

bibliojit aphy or literature siitvcy, mention it here.)

The application of metallic materials as orthopedic implants in the human body is

reviewed, concentrating on materials presently in clinical use and undergoing

laboratory evaluation for possible future use. The criteria considered explicitly

are tissue compatibility, mechanical properties, corrosion resistance, and

toxicity. The three principal metallic implant materials, stainless steel, cobalt

alloys, and titanium, are ..discussed in detail. Wherever possible, comparisons are made between the materials in terms of the intended application.

' K{ \ WOIiPS '.•Mpli,il>ctii .il orck't, .sepiiratcd by .s cnucoiini s)

bioiiiaterials, corrosion, implant materials, mechanical properties, metals

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UNCI, ASSil-TED

20. SECURITV CLASS

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UNCLASSHiEO

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53 pp

22.

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