Kutz M. Handbook of materials selection

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2 ORTHOPEDIC BIOMATERIALS: TOTAL HIP ARTHROPLASTY 1169

to develop testing methodologies for this purpose, and many have been compiled

by the American Society for Testing and Materials (ASTM), the International

Standards Organization (ISO), and other bodies interested in standardization. For

instance ISO-7206 (1989) discusses many aspects of testing for partial and total

hip replacements. Thus, individuals who are working to design new materials

for biomedical implants do not need to reinvent the wheel when it comes to

evaluation tests.

Finally, when the mechanical strength of a potential implant material is eval-

uated, the failure behavior of the material must be taken into consideration. The

ability of the material to absorb energy during potential abnormal loading events,

such as a fall or a jump, will affect its overall, long-term performance. Metals

are signiﬁcantly less brittle than ceramics, and as a result are less likely to fail

during a high-energy or high-rate loading event. Ceramics are also more sus-

ceptible to failure in bending than metals, due to their relatively low tensile

strength and low resistance to crack propagation. These factors do not eliminate

ceramics from consideration for orthopedic implants; however, an understanding

of a material’s behavior under a range of loading modes and rates is needed

before a ﬁnal determination of its suitability for a particular application can be

determined.

Stress Shielding. In bony prosthetics, exceeding the minimum strength re-

quirements of the application is not the only mechanical behavior of concern.

Bone is a living tissue and not merely a structural material. As a result, bone

changes in response to its loading environment. In 1892, Wolff ﬁrst noted that

the pattern of the trabecular bone in the head and neck of the femur was similar

to the stress trajectories of a Culmann crane, an engineering structure with a

similar loading pattern (Wolff, 1892). Based on this observation, Wolff hypoth-

esized that bone develops and remodels in response to the load that it experi-

ences. Over 100 years later, Wolff’s law is still the governing principle behind

our understanding of how bone behaves in response to stress. In general terms,

bone requires a minimal, time-averaged stress in order to maintain its mass. If

the stress falls below this threshold level, bone will be lost. If the stress increases

beyond this threshold, bone will be added until the stress experienced returns to

its ‘‘desired’’ level. This is evidenced by the loss of bone seen in individuals

after sustained bed rest or in astronauts after missions spent in low-gravity en-

vironments.

As a result of this characteristic property, bone is susceptible to a phenomenon

termed stress shielding. This occurs when the stress in bone is reduced below

its maintenance threshold as a result of the mechanical role of an implanted

structure. Figure 3a shows the radiograph of a femur in which signiﬁcant bone

loss has occurred around the proximal end of the implant, with stress shielding

being one of the underlying causes. Bone loss such as this is one of the leading

reasons for implant failure and revision. The formation of a callus at the distal

end of the implant, as also shown in Fig. 3b, is additional evidence of the

response of bone to its stress environment. In many hip implant designs, while

the stress in the bone in the proximal region of the implant is reduced, an

increase in stress is seen in the distal region that results in the deposition of

additional bone.

1170 SELECTION OF MATERIALS FOR BIOMEDICAL APPLICATIONS

Fig. 3 X-rays of hip implants showing changes in bone mass resulting from stress shielding

around an implant. (A) Circled areas indicate regions of signiﬁcant bone loss in the proximal

region of the femur. The arrow indicates the area where the collar of the prosthesis is no longer

in contact with bone on the medial side. (B) Proximal bone loss (indicated in the circled regions

at the edge of the prosthesis) is accompanied by increased bone deposition (indicated by the

rectangle) in the distal region of the implant. The latter phenomenon is caused by increased

stress transfer in this region.

In any loaded structure consisting of two or more materials, the distribution

of stress and strain between the component materials will depend on their ge-

ometric arrangement and relative material properties. For a structure in which

the materials are in parallel with the loading axis and where the materials are

sufﬁciently well bonded—such as a well-ﬁxed femoral implant—it can be as-

sumed that the materials deform to the same extent and therefore experience the

same strain. In this isostrain condition, the stress in one of the components of a

two-phase composite can be calculated from the equation:

␴

⫽ (1)

EA ⫹ EA

11 2 2

where P is the total load on the structure, and E and A are the Young’s modulus

and cross-sectional area of each of the components. Thus, the fraction of the

load carried by each material, and the resulting stress, is related to its Young’s

modulus and cross-sectional area in relation to those of the other components

of the composite structure. The stiffer materials in the composite will carry a

greater proportion of the load per unit cross-sectional area.

As stated above, the important parameter for maintenance of bone is the stress

that it experiences compared to a threshold value. If bone in its natural state is

compared to bone with an implant, the effect of this intervention on the stress

2 ORTHOPEDIC BIOMATERIALS: TOTAL HIP ARTHROPLASTY 1171

in the bone, and therefore its remodeling response, can be estimated. For a

femoral implant, or a bone plate as another example, the implant is in parallel

with the bone tissue with respect to the main loading axis. Therefore, Eq. 1 can

be used to estimate the effect of the implant on the stress in the bone. The

applied load can be assumed to be the same pre- and postimplant, which yields

the following equations for the stress in the bone in the two conﬁgurations:

Preimplant (E

implant

⫽ 0; A

implant

⫽ 0)

EP P

bone

␴

⫽⫽ (2a)

bone

EA A

bone bone bone

Postimplant

bone

␴

⫽ (2b)

bone

EA ⫹ EA

bone bone implant implant

Thus, the stress in the bone is reduced by the inclusion of the implant, with the

amount of the reduction dependent on the modulus and area of the implant.

Implants with a higher modulus and a larger cross-sectional area will shield the

bone from a greater proportion of its normal, physiological stress, resulting in

bone loss according to Wolff’s law.

Joint Motion

Friction. Frictional forces between the articulating surfaces of a joint have

two primary effects: (1) to increase the muscle force required to overcome the

internal friction and allow motion to occur and (2) to increase the torque ex-

perienced by the implant and/or bone, such as at the location of the femoral

neck. Large internal bending moments due to high frictional forces may lead to

failure of the implant, and therefore should be avoided.

The natural joint, with its cartilage bearing surfaces and synovial ﬂuid lubri-

cation, possesses a remarkably low coefﬁcient of friction that minimizes the

tangential and bending forces at the joint. It has been recognized that, in an

artiﬁcial hip, the coefﬁcient of friction between the femoral head and the ace-

tabular cup must be minimized in order to most closely approximate the normal

physiology. While no material combinations currently provide a coefﬁcient of

friction of the level seen naturally, a number of options have been identiﬁed that

provide sufﬁciently low friction forces in the joint. Table 2 provides the coefﬁ-

cients of friction of some commonly paired materials used in hip replacement

in comparison to the natural state. Once implanted, the joint will be lubricated

with physiologic or synovial ﬂuid, although it will no longer be contained within

the original joint capsule. Therefore, the friction values provided were measured

using physiological ﬂuid or bovine albumin as a lubricating material. Alumina

has a relatively high surface tension, allowing it to develop a good lubrication

ﬁlm that minimizes friction in vivo (Ravaglioli and Krajewsk, 1992).

As friction forces are proportional to the coefﬁcient of friction and the normal

force between the contact surfaces, the loading conﬁguration of an implant will

also determine the forces that inﬂuence joint motion and applied bending mo-

1172 SELECTION OF MATERIALS FOR BIOMEDICAL APPLICATIONS

Table 2 Coefﬁcient of Friction for Sample

Material Combinations Used in Total Hip

Replacement

Material Combination Coefﬁcient of Friction

Cartilage/ cartilage 0.002

CoCr/ UHMWPE 0.094

Zirconia/ UHMWPE 0.09–0.11

Alumina/ UHMWPE 0.08–0.12

CoCr/ CoCr 0.12

Alumina/ alumina 0.05–0.1

UHMWPE, ultra-high-molecular-weight polyeth-

ylene; CoCr, cobalt–chromium alloy.

Source: Data from Park and Lakes (1992), Strei-

cher et al. (1992).

ments. Therefore, materials that will provide a sufﬁciently reduced friction en-

vironment for a relatively low force location, such as the wrist, can result in a

friction force in the hip that may cause implant failure to occur or cannot be

adequately overcome by normal muscle forces to allow unhindered motion.

Thus, once again, the speciﬁc application of an implant must be considered when

selecting materials—generalized assumptions and choices do not necessarily

produce the optimum selection.

Wear. Whenever contact surfaces and motion are combined, material wear

must be taken into consideration. Wear is the process whereby one object,

through motion, removes material from the surface of the contacting object.

Generally, the harder material will cause wear to occur on the softer material.

Three basic types of wear can occur: abrasive wear, adhesive wear, and third-

body wear. Abrasive wear exists when a hard material, such as a metal, moves

cyclically against a soft material, such as a polymer. Adhesive wear involves the

sliding motion of two similar materials, where molecular bonds can be formed

at the interface of the structures. In rough materials, the surfaces appear as a

series of peaks and valleys. The two articulating surfaces typically come into

contact at the peaks of the surface roughness, concentrating the contact load

over a much smaller area and increasing the contact stress. As the molecular

bonds between the objects are broken through motion, they also break off par-

ticles of the underlying material. Third-body wear includes the effect of particles

between the articulating surfaces that tend to accelerate wear.

The amount of wear that occurs between two surfaces will depend on several

factors: (1) the hardness of the two materials (p), (2) the normal force at the

surface (F

), (3) the Archard coefﬁcient for the pair of materials (k), (4) the area

of contact between the surfaces during the cyclic motion, and (5) the number of

cycles expected. Archard’s coefﬁcient is similar in concept to the coefﬁcient of

friction and describes the degree to which a normal force at the surface is trans-

lated into a wear-producing force. A harder material will always sustain less

wear when in motion against a given surface than a softer material. As two

materials sweep out a larger surface of contact for each cycle of motion, there

will be a greater fraction of the surface area experiencing wear, thus resulting

2 ORTHOPEDIC BIOMATERIALS: TOTAL HIP ARTHROPLASTY 1173

Fig. 4 Illustration of the development of a new socket within an

acetabular cup due to excessive wear.

in an increase in the overall rate of wear. And, with little explanation required,

if a given amount of wear occurs per cycle, then an increased number of cycles

will result in a greater total amount of wear.

The wear rate, or volume of wear particles produced (V), can be approximated

for adhesive wear by the equation:

kF x

V ⫽ (3)

where x is the total sliding distance between the surfaces (Black, 1999). The

total sliding distance can be determined for a single cycle, a deﬁned period of

time (e.g., one hour), or estimated for the entire life of the implant, thus allowing

for the calculation of a rate for wear particle production.

Wear can be exacerbated through corrosion or fatigue processes, indicating

that in vivo wear rates may differ from those measured in laboratory. The oc-

currence of wear in an implant may affect both its mechanical function and the

response of the body to the implant. The former will be discussed here, while

the effect of wear debris will be described in the discussion of biocompatibility.

As wear occurs on one or both of the opposing surfaces of an articulating

joint, the shape of the implant may change as material is removed. This is

especially evident when signiﬁcant wear occurs in a polymeric acetabular cup

within a total hip replacement. Figure 4 shows how signiﬁcant wear can result

in the creation of a new socket, typically with a reduced diameter. This change

in the conﬁguration of the acetabulum can affect the range of motion of the hip

as well as whether any joint impingement occurs. As the socket deepens, the

femur will shift upward, which will increase the laxity of some of the ligaments

and tendons, thus affecting overall joint performance.

1174 SELECTION OF MATERIALS FOR BIOMEDICAL APPLICATIONS

Table 3 Ideal and Practical Electrochemical Series for Common Metals

Ideal Series

Half-Cell Potential (mV) of Ideal Series with

Respect to H / H⫹ Half-cell Practical Series

Cathodic

Gold

⫺1.50 Platinum

Platinum

⫺0.86 Gold

Silver

⫺0.80 Passivated stainless steel

Copper

⫺0.47 Titanium

Lead 0.12 Silver

Tin 0.14 Unpassivated stainless steel

Nickel 0.23 Copper

Iron 0.44 Tin

Chromium 0.56 Lead

Aluminum 1.70 Wrought iron

Titanium 2.00 Aluminum

Magnesium 2.40 Magnesium

Anodic

The half-cell potentials listed are for the ideal series and are with respect to a H / H

⫹

half-cell.

Source: Data from Black (1999), Park and Lakes (1992).

2.2 Biocompatibility

Once a material is selected for an implant application based on the functional

requirements, it must be evaluated in terms of material–body interactions.

An implant material will react chemically with the local environment, with

the type of reaction dependent on the class of material. Metals are susceptible

to corrosion, polymers experience leaching and absorption, while ceramics are

generally considered to be chemically inert—unless designed to be bioactive.

The effects of chemical degradation may affect both the tissue and the material

itself, especially its mechanical properties, and so both aspects must be consid-

ered. In addition, degradation products can affect the physiology locally, at a

remote location, or systemically.

Corrosion

Metallic materials are susceptible to corrosion, particularly in the ionic ﬂuid

environment of the body. To assess the corrosion potential of a metal, it is

necessary to examine the half-cell potential of that metal—which will act as an

anode when it releases electrons—with respect to the material acting as the

cathode. This cathode may be another metal or the ionic environment itself. An

electrochemical series lists the half-cell potentials of metals in order from the

most noble (or cathodic) to the most anodic. When two materials are in contact

with each other directly or through an ionic solution, the metal listed ﬁrst in the

list will act as the cathode while the other will behave as the anode. Practical

electrochemical series typically relate half-cell potentials as measured in an ap-

plication-speciﬁc environment and may include alloys. This contrasts with ideal

series, which list only pure metals as measured with respect to a hydrogen half

cell reaction. The ideal series approximates the behavior of metals in pure water.

Table 3 shows the half-cell potentials of common metals as measured in an ideal

2 ORTHOPEDIC BIOMATERIALS: TOTAL HIP ARTHROPLASTY 1175

Fig. 5 Pourbaix diagram for chromium in aqueous solution of 1 N Cl

⫺

. The diagram indicates

the electrochemical environment of various regions of the body. [From John H. Dumbleton and

Jonathan Black, An Introduction to Orthopaedic Materials, Thomas Publishing,

Springﬁeld, IL, 1975, with permission].

electrochemical series, as well as the qualitative series of metals in a saltwater

solution. This latter series reasonably approximates what would be found in the

body, where sodium chloride is a major constituent of the ionic soup.

The low half-cell potentials of gold and platinum are a result of their essential

inertness. For other materials, their resistance to corrosion may be due to the

formation of a protective oxide layer on the object’s surface. A second way to

determine how a metallic material is likely to behave in vivo from a chemical

standpoint is to examine Pourbaix diagrams that have been developed based on

theoretical chemical relationships. These curves, such as the one shown in Fig.

5, describe the expected corrosive behavior of materials as a function of pH and

the surrounding electrical potential. By determining the expected environment

at an implant site, it is possible to predict which materials are likely to meet the

outlined requirements for implantation. In regions of buffered pH, which are

common within the human body, the partial pressures of O

and H

can be used

to estimate the electrical potential at the site (Black, 1999).

Three general behaviors can exist for a metallic material in an ionic environ-

ment: corrosion, passivation, or immunity. Corrosion is the chemical reaction in

which a metal is oxidized, producing metallic ions within the ﬂuid environment.

As a metal corrodes, it can effect both the overall implant properties—resulting

in premature failure—and the body. In the latter case, the ions produced through

the oxidation reaction can interfere with the normal physiologic process of the

body at either a local or a systemic level. It cannot be assumed that the ions

produced will have no adverse reaction in the body if they are normally present

in trace amounts. For instance, iron, a mineral required for normal red blood

cell production, will prove to be toxic at elevated concentrations in the body,

causing liver and pancreatic failure (Smith, 1983). Therefore, it is important to

select materials that will exist within their passivation or immunity regions in

the in vivo environment. Passivation is the creation of a protective, oxide or

1176 SELECTION OF MATERIALS FOR BIOMEDICAL APPLICATIONS

hydroxide layer on the surface of the metal. Immunity is the region in which

minimum oxidation is going to occur as the equilibrium concentration of metallic

ions is low. (The equilibrium concentration refers to the levels of ions in the

immediate, surrounding environment at which the oxidation reaction reaches

equilbirium, thus producing no further net release of ions.) Both passivation and

immunity require that less than 10

⫺

M of ions be produced (Park and Lakes,

1992). Thus, in neither of these cases is oxidation eliminated for a metallic

material; however, the effect on the implant’s properties and the ion concentra-

tion produced will be negligible. As a ﬁnal check of biocompatibility, it is im-

portant to ascertain that the trace levels of ions produced by passivated or

immune metals will not produce toxic effects at even those low concentrations.

Pourbaix diagrams can only provide information on the ideal behavior of a

metal within its immunity or passivation regions. As the limitation of corrosion

in these environments is due to the establishment of a chemical equilibrium, the

ionic environment of the implant surface and the surrounding physiologic ﬂuid

must remain undisturbed. Therefore, any damage to the passivation layer, or any

ﬂuid ﬂow that removes the equilibrium ions from the region immediately adja-

cent to the implant, may cause renewed corrosion. Cyclically loaded implants

could be expected to experience both surface damage, due to wear or contact

with other tissues, and movement of the ﬂuid surrounding the implant. In ad-

dition, the local pH at a site may be affected by common physiologic insults,

such as injury and infection. Finally, most Pourbaix diagrams are determined

for pure water environments, and therefore do not include the role of additional

ionic components in the physiologic ﬂuid. This would also be expected to inﬂu-

ence the corrosion behavior of a material. Therefore, the behavior predicted by

the Pourbaix diagrams may be taken as a base line expectation, but must be

conﬁrmed through both animal and clinical testing.

Some additional rules can be applied to minimize the occurrence of corrosion

for an implant with metallic components. These are:

1. Minimize the number of different metals used in the implant. Two metals

that differ in electrochemical potential, due to differences in elemental

composition or processing, will create a galvanic potential when con-

nected either directly or through an ionic ﬂuid. This can include parts

manufactured of the same material but by different companies or different

techniques, as the processing schedule may differ enough to create dif-

ferences in the electrical potentials of the components.

2. Minimize the possibility of surface damage. Any scratches to the surface

of a metallic implant can serve as foci for corrosion once it is implanted.

Thus, it is important to protect the implant from the stage of ﬁnal proc-

essing through implantation to prevent damage.

3. Reduce the number of crevices or regions where oxygen depletion is pos-

sible. A region with a low-oxygen concentration, such as the interfaces

between modular components of a total joint implant or the threads of a

bone screw, is likely to become anodic with respect to the surrounding

metal—even with respect to the remaining bulk of the metal. Corrosion

will therefore occur with the small, anodic region providing electrons to

2 ORTHOPEDIC BIOMATERIALS: TOTAL HIP ARTHROPLASTY 1177

the entire cathode. If there is a discrepancy in the size of the regions, as

would normally be expected, then corrosion within the crevice can be

accelerated, leading to potential mechanical failure.

Leaching and Absorption

Polymers placed in a ﬂuid environment can experience two opposite phenomena.

In leaching, unreacted monomer molecules, ﬁllers, or small chains of polymers

can diffuse from the bulk of the polymer to the surrounding ﬂuid. As in corrosion

products, these released molecules may have a negative effect on the local phys-

iology or, if transported through the bloodstream or lymphatic system, on sys-

temic or remote processes. In addition, signiﬁcant leaching may reduce the

density of the polymer and consequently have an adverse effect on the properties

of the structure. Absorption occurs when water molecules, proteins, or lipids

diffuse from the ﬂuid into the mass of the polymer. The absorbed molecules

become distributed between the molecules of the polymer, reducing the me-

chanical strength of the structure and increasing its susceptibility to wear.

Both absorption and leaching occur as a result of diffusion processes across

the surface of the implant. Fick’s ﬁrst law of diffusion describes the rate of

solute transfer across a permeable barrier:

⭸C

F ⫽⫺D (4)

冉冊

⭸x

where F is the rate of solute transfer per unit cross-sectional area of the surface,

D is the diffusion coefﬁcient, and

⭸C/⭸x is the concentration gradient. The dif-

fusion coefﬁcient is dependent on the solute (the molecule that is being absorbed

or released), the matrix through which it is moving (the remainder of the bulk

polymer), and the type of diffusion that is occurring. Thus larger molecules

moving through more tightly bonded matrices are likely to leach out at a lower

rate than small molecules diffusing through a more open, amorphous structure.

All materials, including metals and ceramics, can absorb molecules—

particularly water—from the surrounding environment. However, this occurs

much more readily in the relatively loosely bonded polymers. Absorption in

polymers can also result in swelling, due to their low elastic modulus, which

may cause geometric changes that interfere with the performance of an implant.

The strain that a polymeric object experiences due to swelling may induce cracks

and may also reduce the ultimate strength of the object. This latter phenomenon

occurs because, due to the new baseline strain in the material, less stress is

needed to reach the material’s ultimate strain. If the absorbed molecules are

small, such as water, they will act as plasticizers and weaken the bonds between

the polymer chains, thus reducing the Young modulus of the material. If a pol-

ymer is hydrophobic in nature, it is less likely to absorb water. However, ab-

sorption of nonpolar molecules such as lipids may still occur.

Leaching generally has a smaller effect on the mechanical properties of a

polymeric structure than absorption. However, as was the case in corrosion, local

changes in properties may occur as material is lost. In fact, if sufﬁcient leaching

occurs to create adjacent or expanded voids within the chemical structure of the

polymer, these regions may act as stress risers. A high amount of leaching will

1178 SELECTION OF MATERIALS FOR BIOMEDICAL APPLICATIONS

act to increase the porosity of the polymer. In both of these cases, the elastic

modulus and the mechanical strength of the polymer would be affected.

In general, low-molecular-weight or highly amorphous polymers are more

susceptible to leaching or absorption processes than high-molecular-weight,

highly crystalline, or highly cross-linked polymers. As diffusion coefﬁcients in-

crease for smaller molecule sizes, in general, it can be expected that additives

and free monomers have a higher rate of leaching than the large polymer chains

themselves, while the relatively small water molecule will be absorbed by hy-

drophilic polymeric structures at a relatively high rate. The behavior of free

monomers and additives within polymers is of particular interest, as they may

have a different effect on the surrounding tissue than tests based on the bulk

polymer suggest. In fact, free monomer should be expected to have a different

physiologic effect than molecules that have reacted to form the bulk polymer,

due in part to the difference in chemical reactivity of the structures.

Wear Debris

The material produced through the wear process is typically in the form of

particulate debris. The size of the particles is dependent on the material involved

and may range from submicon dimensions to millimeter-sized pieces. In poly-

ethylene, particles typically range from 0.5 to 50

␮

m in their largest dimension,

while polymethylmethacrylate (PMMA, bone cement) particles tend to be sig-

niﬁcantly larger (Willert and Semlitsch, 1996). These particles can have several

adverse effects on implant performance beyond the geometric changes discussed

above. First, if the debris becomes trapped between the articulating surfaces of

the joint, it will act as a collection of third-body particles to accelerate the wear

process. Second, the presence of particulate debris in the tissue surrounding the

bone triggers an immune response that can result in signiﬁcant bone loss through

the process of osteolysis. When this occurs without an associated infection, it is

termed aseptic loosening. The presence of histiocytes, macrophages, and foreign

body giant cells in locations with both mild and severe osteolysis indicate that

the process is inﬂammatory or immunological in nature. It is currently hypoth-

esized that the presence of particulate debris results in macrophage and giant-

cell recruitment (Jacobs, et al., 1994). The larger particles are engulfed by giant

cells through phagocytosis and form granulation tissue, while macrophages react

with the smaller particles. Normally, foreign debris on a small scale will be

removed through the lymphatic system. However, if the volume of debris pro-

duced overloads the lymphatic system, then macrophages at the implant site may

release cellular mediators that trigger the bone resorption observed.

2.3 Current Material Selection

The traditional materials that have been used for total hip arthroplasties are ultra-

high molecular-weight polyethylene (UHMWPE) for the articulating surface of

the acetabular cup and a metal, today typically an alloy of titanium or of cobalt–

chromium, for the femoral stem, head, and backing of the acetabluar cup. In a

smaller number of designs, ceramics such as alumina (Al

) and zirconia

(ZrO

) have been used for a modular femoral head component, both to reduce

friction within the acetabulum and minimize the number of metallic components

that may exacerbate corrosion. Table 4 provides a summary of some of the