Selection of Materials Biomedical Application(1)

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CHAPTER 37SELECTION OF MATERIALS FORBIOMEDICAL APPLICATIONS

Michele J. GrimmBioengineering CenterWayne State UniversityDetroit, Michigan

1 INTRODUCTION 1165

2 ORTHOPEDIC MATERIALS:TOTAL HIP ARTHROPLASTY 11662.1 Function 11672.2 Biocompatibility 11742.3 Current Material Selection 1178

3 BLOOD-CONTACTINGBIOMATERIALS: VASCULARPROSTHESES 11803.1 Function 1181

3.2 Biocompatibility 11833.3 Current Material Selection 1186

4 SPACE-FILLING BIOMATERIALS:BREAST IMPLANTS 11884.1 Function 11884.2 Biocompatibility 11884.3 Current Material Selection 1191

5 SUMMARY 1191

BIBLIOGRAPHY 1192

1 INTRODUCTION

Materials have been used for medical implant applications for centuries. Startingeven before George Washingtons famous wooden teeth, before the use of co-conut shells in the 1800s by South Seas natives to replace missing portions of the skull (Sanan and Haines, 1997), humans have attempted to use materialsfrom biological and inorganic sources to replace diseased or damaged tissues.In fact, some examples of biomaterial implants, particularly in the form of gold

and silver, date back to prehistoric times (Sanan and Haines, 1997). These ma-terial selections of the past, based more on the availability of materials andanecdotal evidence than scientic method, resulted in varying degrees of success.The development of aseptic surgical technique in the midnineteenth century byLister provided the basic medical tool needed for the more widespread andsuccessful use of biomedical implants (Park and Lakes, 1992). Since the twen-tieth century, as knowledge of the biological mechanisms behind the interactionsof implanted materials and tissues has increased, the selection of materials formedical implants has been based on progressive improvements and experimental

evidence.When selecting a material for use in a medical implant application, two gen-eral considerations need to be taken into account: the functional requirements

Handbook of Materials Selection, Edited by Myer KutzISBN 0-471-35924-6 2002 John Wiley & Sons, Inc., New York

Copyright 2002 John Wiley & Sons Retrieved from: www.knovel.com


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1166 SELECTION OF MATERIALS FOR BIOMEDICAL APPLICATIONS

Fig. 1 Anatomy of the hip.

of the implant and how it will interact with the body. The function of the implantincludes the physiologic role that it will replace, as well as the length of timethat it is designed to fulll that role. The interaction of the body and the im-planted material must be examined from two perspectivesthe effect of the

biological environment on the material properties and the effect of the material,and any degradation that may occur, on the local and systemic physiology of the body.

It is not possible within the scope of this chapter to discuss the optimumchoice of materials and the supporting rationale for every medical implant ap-plication. This task is made even more difcult by the fact that the developmentand selection of materials is an ever-changing eld. In light of this fact, thefollowing pages will use three implant examples to serve as the background fora discussion of the considerations in the selection and evaluation of materials

for medical implants. Each section will rst be organized around the functionalrequirements of the implant, with the resulting concerns regarding materialtissue interaction being discussed for the materials that would rst meet thefunctional demands. Geometrical considerations and overall implant design willalso play a major role in the success of an implant; however, this chapter willfocus on the selection of biomaterials as a discrete step in the design process.

2 ORTHOPEDIC BIOMATERIALS: TOTAL HIP ARTHROPLASTY

The hip (Fig. 1) is one of the most commonly replaced joints, due to the highincidence of both osteoarthritis and osteoporotic hip fracture within the popu-lation. As with the other joints in the skeletal system, the hip has two mainfunctions: (1) to transfer load from one bone to another and (2) to allow formotion between the bones. The loads on the hip have been found to vary from0.5 to 8 times body weight during activities of daily living, including brisk walking (Paul, 1999). The loads can be expected to reach even higher levelsduring events such as a stumble. The hip is essentially a ball-and-socket joint



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2 ORTHOPEDIC BIOMATERIALS: TOTAL HIP ARTHROPLASTY 1167

Fig. 2 Example of total hip implant, including femoral and acetabular components.

with 6 degrees of freedom constrained by the bony and ligamentous structuresof the joint. These functional constraints have generally dened the goals fordesign of hip implants over the past half-century.

The total joint replacement for the hip is divided into three components: the

femoral stem and femoral head, which may be integrated or modular, and theacetabular cup (Fig. 2). The combination of the femoral head and the acetabularcup provide the bearing surface for the joint. The femoral stem transfers theload to the femur and provides resistance to the bending moment caused by theanatomy of the joint. Total hip replacement began in the 1930s, with stainlesssteel used for both the acetabular cup and the femoral head, which was boltedonto the natural femoral neck. In the 1950s, the general design of the implantwas expanded by McKee to include the stemmed femoral component familiartoday. In 1959, Charnley introduced a plastic acetabular component, which ar-

ticulated with the metallic femoral component. This idea for low-friction arthro-plasty remains the basis for the predominant implant designs used today(Swanson, 1977).

2.1 Function

Load Support

Because of its primary role in mechanical support, the stem of a femoral pros-thesis can realistically be manufactured from a metal, a ceramic, or a compositematerial. For this discussion, composite materials will not be included, as they

have not yet been utilized in commercially available total joint replacements(they are available for bone plate applications). However, as development of



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Table 1 Representative Properties of CorticalBone

Compressive strength (MPa) 131224 longitudinal106133 transverse

Tensile strength (MPa) 80172 longitudinal5156 transverse

Shear strength (MPa) 5370

Source: Data from Cowin (1989).

composite materials for this application continues, they must meet the samerequirements to be discussed for ceramic and metallic materials.

Strength. The rst property to be considered for a load-bearing implant isits mechanical strength. The loading of the femur is dynamic and, while esti-mated to range up to 8 times body weight, is difcult to determine precisely.Therefore, as the implant will be loaded in essentially the same way as thenatural bone, it is reasonable to assume that a material that will provide the sameor greater load-bearing capacity as bone will meet the necessary mechanicalrequirements. The mechanical properties of cortical bone (Table 1) form thelower limit for properties of materials to be selected for the femoral stem. If theneck of the femoral stem is designed to be longer than the normal range of femoral neck lengths [approximately 9 cm, from the edge of the greater tro-chanter to the apex of the femoral head (Center et al., 1998)], then the increasedbending moment in this region should also be considered. The primary mode ormodes of loading that will be seen for a particular implant application will alsobe important factors in the determination of whether a material meets the con-straints regarding mechanical strength. The hip will be loaded primarily in bend-ing and compression, due to its unique geometry. Tensile and shear strength of any replacement material therefore become of great importance.

In addition to the yield or ultimate strength of an implant material, the fatiguestrength is also important for structures that will be cyclically loaded over anextended period of time. At the current time, hip replacements are designed to

last for approximately 20 years. As an individual typically loads and unloadsthe joint thousands of times per day during normal daily activities, fatigue-induced failure at this location is of paramount concern. In contrast to the ma-terials that might be chosen to replace the original tissue, healthy bone has thecapacity to repair any microfractures that may occur through continuous cycling.This normal remodeling function helps to eliminate the occurrence of fatiguefractures in normal bone.

Based on the cyclic loading that a material is likely to undergo when im-planted in the body, it is reasonable when evaluating selections to compare the

endurance limit of the materials under consideration to the experimentally de-termined strength values for bone. One further complication in this process,however, is that materials will fatigue differently in a typical, air-based labora-tory environment than they will in the ionic soup of the human body. Becauseof this, it is important to measure the fatigue performance of materials in anenvironment that closely mimics that of the implant locationincluding theionic composition, temperature, and pH. Substantial work has been conducted



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to develop testing methodologies for this purpose, and many have been compiledby the American Society for Testing and Materials (ASTM), the InternationalStandards Organization (ISO), and other bodies interested in standardization. Forinstance ISO-7206 (1989) discusses many aspects of testing for partial and total

hip replacements. Thus, individuals who are working to design new materialsfor biomedical implants do not need to reinvent the wheel when it comes toevaluation 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. Theability 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. Metalsare signicantly less brittle than ceramics, and as a result are less likely to failduring 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 tensilestrength and low resistance to crack propagation. These factors do not eliminateceramics from consideration for orthopedic implants; however, an understandingof a materials behavior under a range of loading modes and rates is neededbefore a nal determination of its suitability for a particular application can bedetermined.

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, bonechanges in response to its loading environment. In 1892, Wolff rst noted thatthe pattern of the trabecular bone in the head and neck of the femur was similarto the stress trajectories of a Culmann crane, an engineering structure with asimilar 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, Wolffs law is still the governing principle behindour 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 increasesbeyond this threshold, bone will be added until the stress experienced returns toits desired level. This is evidenced by the loss of bone seen in individualsafter 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 phenomenontermed stress shielding. This occurs when the stress in bone is reduced belowits maintenance threshold as a result of the mechanical role of an implantedstructure. Figure 3 a shows the radiograph of a femur in which signicant boneloss 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 leadingreasons for implant failure and revision. The formation of a callus at the distalend of the implant, as also shown in Fig. 3 b, is additional evidence of theresponse of bone to its stress environment. In many hip implant designs, whilethe stress in the bone in the proximal region of the implant is reduced, anincrease in stress is seen in the distal region that results in the deposition of additional bone.



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Fig. 3 X-rays of hip implants showing changes in bone mass resulting from stress shieldingaround an implant. (A) Circled areas indicate regions of signicant bone loss in the proximal

region of the femur. The arrow indicates the area where the collar of the prosthesis is no longerin contact with bone on the medial side. (B) Proximal bone loss (indicated in the circled regionsat the edge of the prosthesis) is accompanied by increased bone deposition (indicated by therectangle) 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 distributionof stress and strain between the component materials will depend on their ge-ometric arrangement and relative material properties. For a structure in whichthe materials are in parallel with the loading axis and where the materials aresufciently well bondedsuch as a well-xed femoral implantit can be as-

sumed that the materials deform to the same extent and therefore experience thesame strain. In this isostrain condition, the stress in one of the components of atwo-phase composite can be calculated from the equation:

E P1 1 (1)1 E A E A1 1 2 2

where P is the total load on the structure, and E and A are the Youngs modulusand 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 Youngsmodulus and cross-sectional area in relation to those of the other componentsof the composite structure. The stiffer materials in the composite will carry agreater proportion of the load per unit cross-sectional area.

As stated above, the important parameter for maintenance of bone is the stressthat it experiences compared to a threshold value. If bone in its natural state iscompared to bone with an implant, the effect of this intervention on the stress



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in the bone, and therefore its remodeling response, can be estimated. For afemoral implant, or a bone plate as another example, the implant is in parallelwith the bone tissue with respect to the main loading axis. Therefore, Eq. 1 canbe 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 yieldsthe following equations for the stress in the bone in the two congurations:

Preimplant ( E implant 0; Aimplant 0)

E P Pbone (2a)bone E A Abone bone bone

Postimplant

E Pbone (2b)bone E A E Abone bone implant implant

Thus, the stress in the bone is reduced by the inclusion of the implant, with theamount 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 thebone from a greater proportion of its normal, physiological stress, resulting inbone loss according to Wolffs 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 theinternal 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 femoralneck. Large internal bending moments due to high frictional forces may lead tofailure of the implant, and therefore should be avoided.

The natural joint, with its cartilage bearing surfaces and synovial uid lubri-cation, possesses a remarkably low coefcient of friction that minimizes thetangential and bending forces at the joint. It has been recognized that, in anarticial hip, the coefcient of friction between the femoral head and the ace-tabular cup must be minimized in order to most closely approximate the normalphysiology. While no material combinations currently provide a coefcient of friction of the level seen naturally, a number of options have been identied thatprovide sufciently low friction forces in the joint. Table 2 provides the coef-cients of friction of some commonly paired materials used in hip replacementin comparison to the natural state. Once implanted, the joint will be lubricatedwith physiologic or synovial uid, although it will no longer be contained within

the original joint capsule. Therefore, the friction values provided were measuredusing physiological uid or bovine albumin as a lubricating material. Aluminahas a relatively high surface tension, allowing it to develop a good lubricationlm that minimizes friction in vivo (Ravaglioli and Krajewsk, 1992).

As friction forces are proportional to the coefcient of friction and the normalforce between the contact surfaces, the loading conguration of an implant willalso determine the forces that inuence joint motion and applied bending mo-



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Table 2 Coefcient of Friction for SampleMaterial Combinations Used in Total HipReplacement a

Material Combination Coefcient of Friction

Cartilage /cartilage 0.002CoCr / UHMWPE 0.094Zirconia / UHMWPE 0.090.11Alumina / UHMWPE 0.080.12CoCr / CoCr 0.12Alumina / alumina 0.050.1a UHMWPE, ultra-high-molecular-weight polyeth-ylene; CoCr, cobaltchromium alloy.Source: Data from Park and Lakes (1992), Strei-cher et al. (1992).

ments. Therefore, materials that will provide a sufciently reduced friction en-vironment for a relatively low force location, such as the wrist, can result in afriction force in the hip that may cause implant failure to occur or cannot beadequately overcome by normal muscle forces to allow unhindered motion.Thus, once again, the specic application of an implant must be considered whenselecting materialsgeneralized assumptions and choices do not necessarilyproduce the optimum selection.

Wear. Whenever contact surfaces and motion are combined, material wearmust 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, movescyclically against a soft material, such as a polymer. Adhesive wear involves thesliding motion of two similar materials, where molecular bonds can be formedat the interface of the structures. In rough materials, the surfaces appear as aseries of peaks and valleys. The two articulating surfaces typically come intocontact at the peaks of the surface roughness, concentrating the contact loadover a much smaller area and increasing the contact stress. As the molecularbonds between the objects are broken through motion, they also break off par-ticles of the underlying material. Third-body wear includes the effect of particlesbetween the articulating surfaces that tend to accelerate wear.

The amount of wear that occurs between two surfaces will depend on severalfactors: (1) the hardness of the two materials ( p), (2) the normal force at thesurface ( F n ), (3) the Archard coefcient 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. Archards coefcient is similar in concept to the coefcient 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 lesswear when in motion against a given surface than a softer material. As twomaterials sweep out a larger surface of contact for each cycle of motion, therewill be a greater fraction of the surface area experiencing wear, thus resulting



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Fig. 4 Illustration of the development of a new socket within anacetabular 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 cycleswill 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 nV (3)3 p

where x is the total sliding distance between the surfaces (Black, 1999). Thetotal sliding distance can be determined for a single cycle, a dened 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, indicatingthat 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 theresponse of the body to the implant. The former will be discussed here, whilethe 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 isespecially evident when signicant wear occurs in a polymeric acetabular cupwithin a total hip replacement. Figure 4 shows how signicant wear can resultin the creation of a new socket, typically with a reduced diameter. This changein the conguration of the acetabulum can affect the range of motion of the hipas well as whether any joint impingement occurs. As the socket deepens, thefemur will shift upward, which will increase the laxity of some of the ligamentsand tendons, thus affecting overall joint performance.



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Table 3 Ideal and Practical Electrochemical Series for Common Metals a

Ideal SeriesHalf-Cell Potential (mV) of Ideal Series with

Respect to H/H Half-cell Practical Series

Cathodic

Gold 1.50 PlatinumPlatinum 0.86 GoldSilver 0.80 Passivated stainless steelCopper 0.47 TitaniumLead 0.12 SilverTin 0.14 Unpassivated stainless steelNickel 0.23 CopperIron 0.44 TinChromium 0.56 LeadAluminum 1.70 Wrought ironTitanium 2.00 AluminumMagnesium 2.40 Magnesium

Anodic

a 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 functionalrequirements, it must be evaluated in terms of materialbody interactions.

An implant material will react chemically with the local environment, withthe type of reaction dependent on the class of material. Metals are susceptibleto corrosion, polymers experience leaching and absorption, while ceramics aregenerally considered to be chemically inertunless designed to be bioactive.The effects of chemical degradation may affect both the tissue and the materialitself, especially its mechanical properties, and so both aspects must be consid-ered. In addition, degradation products can affect the physiology locally, at aremote location, or systemically.

Corrosion

Metallic materials are susceptible to corrosion, particularly in the ionic uidenvironment of the body. To assess the corrosion potential of a metal, it isnecessary to examine the half-cell potential of that metalwhich will act as ananode when it releases electronswith respect to the material acting as thecathode. This cathode may be another metal or the ionic environment itself. Anelectrochemical series lists the half-cell potentials of metals in order from themost noble (or cathodic) to the most anodic. When two materials are in contactwith each other directly or through an ionic solution, the metal listed rst in thelist 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-specic environment and may include alloys. This contrasts with idealseries, 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



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Fig. 5 Pourbaix diagram for chromium in aqueous solution of 1 N Cl . The diagram indicatesthe electrochemical environment of various regions of the body. [From John H. Dumbleton and

Jonathan Black, An Introduction to Orthopaedic Materials, Thomas Publishing,Springeld, IL, 1975, with permission].

electrochemical series, as well as the qualitative series of metals in a saltwatersolution. This latter series reasonably approximates what would be found in thebody, 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 theformation of a protective oxide layer on the objects surface. A second way todetermine how a metallic material is likely to behave in vivo from a chemicalstandpoint is to examine Pourbaix diagrams that have been developed based ontheoretical chemical relationships. These curves, such as the one shown in Fig.5, describe the expected corrosive behavior of materials as a function of pH andthe surrounding electrical potential. By determining the expected environmentat an implant site, it is possible to predict which materials are likely to meet theoutlined requirements for implantation. In regions of buffered pH, which are

common within the human body, the partial pressures of O 2 and H 2 can be usedto 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 inwhich a metal is oxidized, producing metallic ions within the uid environment.As a metal corrodes, it can effect both the overall implant propertiesresultingin premature failureand the body. In the latter case, the ions produced throughthe oxidation reaction can interfere with the normal physiologic process of thebody at either a local or a systemic level. It cannot be assumed that the ionsproduced will have no adverse reaction in the body if they are normally presentin trace amounts. For instance, iron, a mineral required for normal red bloodcell production, will prove to be toxic at elevated concentrations in the body,causing liver and pancreatic failure (Smith, 1983). Therefore, it is important toselect materials that will exist within their passivation or immunity regions inthe in vivo environment. Passivation is the creation of a protective, oxide or



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hydroxide layer on the surface of the metal. Immunity is the region in whichminimum oxidation is going to occur as the equilibrium concentration of metallicions is low. (The equilibrium concentration refers to the levels of ions in theimmediate, surrounding environment at which the oxidation reaction reaches

equilbirium, thus producing no further net release of ions.) Both passivation andimmunity require that less than 10 6 M of ions be produced (Park and Lakes,1992). Thus, in neither of these cases is oxidation eliminated for a metallicmaterial; however, the effect on the implants properties and the ion concentra-tion produced will be negligible. As a nal check of biocompatibility, it is im-portant to ascertain that the trace levels of ions produced by passivated orimmune metals will not produce toxic effects at even those low concentrations.

Pourbaix diagrams can only provide information on the ideal behavior of ametal within its immunity or passivation regions. As the limitation of corrosion

in these environments is due to the establishment of a chemical equilibrium, theionic environment of the implant surface and the surrounding physiologic uidmust remain undisturbed. Therefore, any damage to the passivation layer, or anyuid ow that removes the equilibrium ions from the region immediately adja-cent to the implant, may cause renewed corrosion. Cyclically loaded implantscould be expected to experience both surface damage, due to wear or contactwith other tissues, and movement of the uid 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 determinedfor pure water environments, and therefore do not include the role of additionalionic components in the physiologic uid. This would also be expected to inu-ence the corrosion behavior of a material. Therefore, the behavior predicted bythe Pourbaix diagrams may be taken as a base line expectation, but must beconrmed through both animal and clinical testing.

Some additional rules can be applied to minimize the occurrence of corrosionfor 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 elementalcomposition or processing, will create a galvanic potential when con-nected either directly or through an ionic uid. This can include partsmanufactured of the same material but by different companies or differenttechniques, 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 surfaceof 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 nal 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 interfacesbetween modular components of a total joint implant or the threads of abone screw, is likely to become anodic with respect to the surroundingmetaleven with respect to the remaining bulk of the metal. Corrosionwill therefore occur with the small, anodic region providing electrons to



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the entire cathode. If there is a discrepancy in the size of the regions, aswould normally be expected, then corrosion within the crevice can beaccelerated, leading to potential mechanical failure.

Leaching and Absorption

Polymers placed in a uid environment can experience two opposite phenomena.In leaching, unreacted monomer molecules, llers, or small chains of polymerscan diffuse from the bulk of the polymer to the surrounding uid. As in corrosionproducts, 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, signicant leaching may reduce thedensity of the polymer and consequently have an adverse effect on the propertiesof the structure. Absorption occurs when water molecules, proteins, or lipidsdiffuse from the uid into the mass of the polymer. The absorbed moleculesbecome 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 acrossthe surface of the implant. Ficks rst law of diffusion describes the rate of solute transfer across a permeable barrier:

C F D (4) x

where F is the rate of solute transfer per unit cross-sectional area of the surface, D is the diffusion coefcient, and C / x is the concentration gradient. The dif-fusion coefcient is dependent on the solute (the molecule that is being absorbedor released), the matrix through which it is moving (the remainder of the bulk polymer), and the type of diffusion that is occurring. Thus larger moleculesmoving through more tightly bonded matrices are likely to leach out at a lowerrate than small molecules diffusing through a more open, amorphous structure.

All materials, including metals and ceramics, can absorb moleculesparticularly waterfrom the surrounding environment. However, this occursmuch more readily in the relatively loosely bonded polymers. Absorption inpolymers can also result in swelling, due to their low elastic modulus, whichmay cause geometric changes that interfere with the performance of an implant.The strain that a polymeric object experiences due to swelling may induce cracksand may also reduce the ultimate strength of the object. This latter phenomenonoccurs because, due to the new baseline strain in the material, less stress isneeded to reach the materials ultimate strain. If the absorbed molecules aresmall, such as water, they will act as plasticizers and weaken the bonds betweenthe 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 apolymeric structure than absorption. However, as was the case in corrosion, localchanges in properties may occur as material is lost. In fact, if sufcient leachingoccurs to create adjacent or expanded voids within the chemical structure of thepolymer, these regions may act as stress risers. A high amount of leaching will



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act to increase the porosity of the polymer. In both of these cases, the elasticmodulus and the mechanical strength of the polymer would be affected.

In general, low-molecular-weight or highly amorphous polymers are moresusceptible to leaching or absorption processes than high-molecular-weight,

highly crystalline, or highly cross-linked polymers. As diffusion coefcients in-crease for smaller molecule sizes, in general, it can be expected that additivesand free monomers have a higher rate of leaching than the large polymer chainsthemselves, while the relatively small water molecule will be absorbed by hy-drophilic polymeric structures at a relatively high rate. The behavior of freemonomers and additives within polymers is of particular interest, as they mayhave 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 differentphysiologic 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 involvedand 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-nicantly larger (Willert and Semlitsch, 1996). These particles can have severaladverse effects on implant performance beyond the geometric changes discussedabove. 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 wearprocess. Second, the presence of particulate debris in the tissue surrounding thebone triggers an immune response that can result in signicant bone loss throughthe process of osteolysis. When this occurs without an associated infection, it istermed aseptic loosening. The presence of histiocytes, macrophages, and foreignbody giant cells in locations with both mild and severe osteolysis indicate thatthe process is inammatory 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 giantcells through phagocytosis and form granulation tissue, while macrophages reactwith the smaller particles. Normally, foreign debris on a small scale will beremoved through the lymphatic system. However, if the volume of debris pro-duced overloads the lymphatic system, then macrophages at the implant site mayrelease 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 cobaltchromium, for the femoral stem, head, and backing of the acetabluar cup. In asmaller number of designs, ceramics such as alumina (Al 2O3) and zirconia(ZrO 2) have been used for a modular femoral head component, both to reducefriction within the acetabulum and minimize the number of metallic componentsthat may exacerbate corrosion. Table 4 provides a summary of some of the



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Table 4 Summary of Mechanical Properties for Materials Commonly Used in CurrentTotal Joint Replacement Designs a

Property HMWPE Titanium Alloy CoCr Alloy Alumina Cortical Bone

Elastic Modulus[GPa]

2.2 110 220234 350400 1020

Comp. Strength[MPa]

4000 130280

Tensile Strength[MPa]

3 860 6001000 270 80160

Endurance Limit[MPa] at 10 7 cycles

620 500

Density [g/cm 3 ] 0.930.94 4.5 9.2 3.9 1.8Hardness [MPa] 3500 30004000 20,000a Titanium alloy, Ti6Al4V; CoCr alloy, wrought CoNiCrMo.Source: Data from ASTM-F136 (1998), ASTM-F562 (2000), ASTM-F648 (2000), Bernache-

Assolant (1991), Brunski (1996), Cowin (1989), and Park and Lakes (1992)

mechanical properties of these common materials and lists cortical bone forcomparison purposes.

In the selection of metals, cobaltchromium alloys with molybdenum havebeen preferentially chosen over those with nickel. Despite the fact that the me-chanical properties are slightly reduced in the molybdenum alloy compared tothe nickel variant, it has been shown to have improved wear properties. Titaniumis often chosen due to the fact that its elastic modulus is half of that seen incobaltchromium or stainless steel alloys, therefore reducing the stiffness of theimplant and the accompanying bone loss due to stress shielding.

During the past decade, substantial research has been conducted to developdifferent pairs of materials for the bearing surface in order to reduce the wearphenomenon commonly seen in HMWPE, which often leads to premature failureof an implant. Metal-on-metal and ceramic-on-ceramic headcup combinationshave shown promise in terms of long-term outcomes, especially when used inyounger patients for whom an implant life of greater than 20 years would bedesirable (Delaunay, 2000; Skinner, 1999; Wagner and Wagner, 2000). The ce-ramics typically selected have been alumina and zirconia, while the metal isgenerally a cobaltchromium alloy. Titanium does not lend itself to metalmetalbearing surfaces due to the fact that it tends to seize when in contact with othermetals. The success of these material combinations is due in large part to theimprovements in materials processing, machining, and polishing that have beenachieved in the last quarter century, allowing for the development of high-qualitymaterials, excellent geometric matches, and highly polished surfaces. This pro-gress also indicates that designs that were tried out at an earlier point in history,

and were perhaps abandoned, may deserve to be reexamined. Metal-on-metalbearings were originally used in total hip replacements in the 1950s, with limitedsuccess. Complete ceramic bearings were seen as early as the 1970s, but fractureof the ceramic components and high wear were attributed to problems in proc-essing that resulted in defects and inadequate grain sizes (Plenk, et al., 1992;Sedel, et al., 1991). In time, they were almost completely replaced in the clinicalworld by metalpolymer combinations. However, their revival in the past few



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Fig. 6 Schematic diagram of an artery showing the three tissue layers:adventitia (outer), media (middle), and intima (inner).

years, spurred on by developments in materials engineering, indicates that theoriginal developers had the right ideathey were just a few years ahead of theirtime.

3 BLOOD-CONTACTING BIOMATERIALS: VASCULAR PROSTHESES

When blood vessels are damaged through injury or disease, they often must bereplaced or bypassed in order to maintain adequate blood ow to and from theregions of the body. Disease-induced damage, such as atherosclerosis and an-eurisms, occurs more often in arteries than in veins, due in large part to thehigher working pressure of the blood within these vessels. Injury can occur toany blood vessel; however, collateral circulation typically eliminates the need toreplace small veins, and the low venous return pressure provides an environmentin the larger veins that is much more conducive to traditional surgical repair orautograft use. As a result, this section will focus on the selection of materialsfor the development of arterial prostheses.

Arteries are three-layer hollow tubes (Fig. 6) composed of a combination of elastin, collagen, and smooth musclewith the proportions of each componentdependent on the size of the artery. The size of an artery also varies substantiallyalong the arterial tree, from a typical diameter of about 4 cm for the aorta in anadult to a diameter of less than 1 mm for arterioles. The age and size of theindividual will also affect the dimensions of the arteries. Normal arterial pressureis approximately 120 mmHg during the systolic phase (contraction) of the car-diac cycle, decreasing to approximately 70 mmHg during the diastolic phase(relaxation). However, pressures can rise as high as 200 mmHg or more inindividuals with either transient or chronic hypertension, and hypertension itself is a risk factor for a number of the pathological processes that may requirearterial replacement or bypass.

Elastic arteries, such as the aorta, are expected to expand during systole asblood is forced into them by the heart. The elastic recoil of these vessels, duein large part to the high amount of elastin in the wall, acts to maintain arterialpressure during the relaxation phase of the cardiac cycle, further transportingblood away from the heart. Distributing arteries, which (as their name indicates)distribute blood from the larger, elastic arteries to the various regions of thebody, have more collagen than elastin in their walls. However, the physiologi-cally important component of these vessel walls is the smooth muscle, which



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3 BLOOD-CONTACTING BIOMATERIALS: VASCULAR PROSTHESES 1181

allows the distributing arteries to adjust their diameter and therefore their resis-tance. By controlling the arterial resistance to blood ow in various regions of the body through vasodilation and constriction, blood can be directed to regionsthat have the highest metabolic demandthe muscles during exercise or the

gastrointestinal system during digestion.Vascular grafts were rst introduced in the early twentieth century as solidtubular structures. Fabric grafts were developed in the 1950s and provide thebasis for the grafts that are currently in use or under development. An arterialgraft or arterial prosthesis has one obvious functionto transport blood fromvessels that are proximal to the graft to those that are distal to the graft. Inaddition, the vasodilation/constriction and elastic recoil capabilities of the var-ious types of arteries assist the heart with transport and direction of blood to thediverse tissues of the body. However, there are additional functional constraints

that are equally as important. First, the prosthesis must maintain hemostasisnot allow blood to leak from the artery into the surrounding tissues. Second,interaction of the blood cells with the vessel must not act to initiate coagulation,which could result in thrombus formation and embolisms, nor hemolysis, thedestruction of red blood cells. These functional constraints will be consideredin the discussion on biocompatibility.

3.1 Function

Blood Transport

To transport blood between two connecting vascular segments, a tubular structureis required. However, given modern material processing techniques, this require-ment does not itself limit the selection of materials substantially. The graft mustalso, however, be connected to the ends of the remaining vascular segments insome way, either through ligatures or sutures.

Using a maximum arterial pressure of 200 mmHg (30 kPa), the circumfer-ential and axial wall stresses within a graft (diameter of 1 cm, thickness of 0.5mm) can be approximated to reach 270 and 135 kPa, respectively. This is wellwithin the mechanical limits of most undegraded articial materials; however, itmay become a substantial constraint for vessels engineered from natural mate-rials in the future. In addition, the physiological environment, combined withthe cyclic loading seen by the vessel due to the normal vascular pressure vari-ations, will affect the properties of the graft material. As discussed above forhip replacement, corrosion, absorption, and leaching processes that occur whena material is placed in vivo will all negatively affect the ultimate strength of amaterial.

The expansion and recoil observed in elastic arteries must be taken into con-sideration when selecting a material for grafts at these locations. It is reasonable

to assume that a relatively short length of the graft does not itself need todynamically change dimensions to maintain blood ow through the segment of vasculature. However, any mismatch in the behavior of the graft to the connectedvessels will result in substantial stresses on the ligatures, sutures, or the vesselitself. In certain applications, the graft may replace a substantial portion of anartery, such as in the descending aorta. In this case, elastic recoil in the graftwill play an important role in maintaining the velocity of blood ow through



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Fig. 7 Schematic diagram showing the characteristic nonlinear forcedeformation relationshipof an artery. Initial region of high compliance is characteristic of the elastin component in the

walls, while the later, low-compliance region exemplies the deformation of the collagen bers.

the distal vessels, and the pulsatile ow may be important to maintaining func-tion in organs supplied by the arterial tree (Mergerman and Abbott, 1983). Thesame philosophy would apply to distributing arteries. At the current time, noarticial material exists that would respond to the physiologic control mecha-

nisms in order to dilate and constrict in conjunction with surrounding vessels.However, a graft that is less compliant than the attached vessel would result insubstantial stress concentrations at the junction between the articial and naturalmaterials. This elevation in stress could lead to failure of the sutures or ligaturesconnecting the structure, as well as accelerate fatigue processes within the graftends (Wilkerson and Zalina, 1994). In addition, a less compliant graft can alsoresult in a stenosis in the vascular tree during systole, when the proximal seg-ment of artery expands under pressure, but the blood then encounters the lengthof reduced diameter graft that acts to retard ow (Herring, 1983). These issues,

along with the fact that blood vessels experience signicant deformation due togeneral body motion, essentially dictate that a soft, exible material be used forthe graft.

An ideal graft would exactly match the compliance of the attached bloodvessel, allowing it to expand and recoil to the same extent as the natural struc-ture. This would require a nonlinear response, with an initial region of highlycompliant stretch, attributed to the unkinking and realigning of the collagenbres and a simultaneous stretching of the elastin. After the collagen in thevessel wall is straightened, its higher elastic modulus dominates the behavior of the vessel, and the compliance is signicantly reduced. It is this behavior (Fig.7) that is thought to limit the overexpansion of arteries during acute episodes of elevated blood pressure. The compliance of each segment of artery thereforedepends on the proportion of collagen, elastin, and smooth muscle (which hasa negligible effect on compliance in its passive state) within the vessel wall.



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Typical values for compliance (in units of percent radial change per millimeterof mercury) for cadaveric human arteries range between 10.9 at 60 mmHg and3.8 at 150 mmHg (Mergerman and Abbott, 1983).


The primary functional need of a blood vesseltransport of bloodcan easilybe met through general implant design. However, due to the delicate nature of blood cells and the ease at which the clotting cascade can be initiated, biocom-patibility issues place substantial limitations on material selection for this ap-plication. The natural vessel provides an optimal environment for blood ow,and the mimicking or replacing of its intimal layer is one of the underlying ideasin work to improve biocompatibility in vascular grafts.

Neointima Formation

Natural blood vessels consist of three layers: the outer adventicia, the middlemedia, and the inner intima. The intima has a layer of endothelial cells thatcontact the moving blood in order to minimize damage to the cellular elementsof the uid. Some articial materials that are implanted into the body as vasculargrafts have been shown to develop a neointima brin and broblasts that de-velop as a lining on the inside surface of the tube (Herring, 1983).

The rate of neointima formation and its constituents may depend on the ma-terial used in the vascular graft. A carbonceramic composite with a large degreeof surface porosity showed almost instantaneous ( 10 s) development of a brinmesh with limited platelet adhesion (Chignier et al., 1987). This developed intoa 5- to 7-cell thick layer of broblasts, collagen, and elastin by 15 days, withendothelial-like cells present and fully developed by 2 months. Herring (1983)provides a good description of the healing of vascular grafts and the devel-opment of the neointima. In most human vessels, complete endothelializationoccurs only in the vicinity of the anastomoses with the natural vessels, as cellsgrow in from the ends of the connected tissue. To expand the area of vessel thatexperiences complete healingsuch that the lining resembles that of a natural,healthy vesselresearch is being conducted into seeding the vessel wall withendothelial cells (Consigny, 2000; Herring, 1983). This can either be done duringthe preclotting process, in the operating room, using a small number of cellstaken from a vein that is exposed during the surgical procedure, or preoperativelyusing cells from the jugular vein that are cultured to provide a greater volumefor the seeding process. In both cases, the use of autologous (patient specic)cells is important to eliminate issues of rejection.

The geometry of the vessel also affects neointima formation. The initial in-teraction between the blood and the graft is the formation of a clot, or thrombus,on the inner surface of the vessel. The thickness of the implant wall is directly

proportional to the thickness of the thrombus formed on its surface (Park andLakes, 1992). As the thrombus must be remodeled to form the mature neointima,a smaller clot results in faster organization of the neointima. Neointima forma-tion must be limited, however, to prevent vessel occlusion. This is one of theissues with small diameter vascular grafts, where the brin layer may continueto grow past 1 mm in thickness (Wilkerson and Zalina, 1994). Finally, it isimportant to ensure that the neointimal layer that develops maintains its integrity



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and does not change in such a way, either through mineralization or adsorptionof additional organic components, that it no longer promotes the continued pa-tency of the vessel (Hufnagel, 1983).

Hemostasis

For vessels to adequately transport blood, the blood must remain within thelumen of the vessel and not leak into the surrounding tissues in any greatamount. The obvious answer to this problem is to employ solid materials forvascular grafts, nonporous structures that will adequately contain the blood.However, the development of a neointima within a vessel provides a seconddesign option, as it will act to seal a porous blood vessel against blood leakage.However, the time course for neointima formation is not short enough to providethe immediate seal needed during surgery and postoperatively. Therefore, many

vascular grafts are preclotted in the operating room before they are implantedby exposing both the inner and outer surfaces to the patients blood. This pro-vides an initial surface that serves as the basis for neointima formation.

Hemolysis

Hemolysis, or the damage and destruction of red blood cells, occurs continuallywithin the body. The cells are constantly replaced by new erythrocytes producedby the bone marrow. However, chronic damage to these cells and the release of their cellular contents into the plasma of the bloodstream can result in anemia,kidney failure, and other toxic reactions (Hershko, et al., 1998). Hemolysis dueto nonphysiologic mechanisms typically occurs due to high shear stresses. Ashear stress as low as 400 Pa can damage or rupture a red blood cell that comesin contact with another solid surface (Sallam and Hwang, 1984). Turbulent owcan also result in sufcient shear stresses to cause hemolysis. Therefore, twogoals of vascular graft design are to minimize the contact between red bloodcells and the graft surfaces and minimize the turbulence that may occur atbranches or divisions of blood vessels. The latter falls within the domain of uidmechanics, rather than biomaterial selection.

To minimize the contact between blood cells and graft walls that may leadto hemolysis, it is desirable to form a natural tissue layer between the two. Graftmaterials with high surface tensions tend to initially attract platelets and brinmolecules, which aggregate and form the natural boundary that is then remod-eled to become the neointima. The elements of the tissue layer also exhibit aslight negative electrical charge (Collins, 1983), which may act to repel thenegatively charged red blood cells, therefore additionally reducing the contactbetween the cells and the vessel wall.

Coagulation

Blood clotting, or coagulation, is a necessary physiologic process that allowsthe body to heal and maintain its blood pressure through hemostasis. However,the formation of stationary clots within blood vessels (thromboses) and themovement of those clots with the blood ow (emboli) can result in vascularblockage, tissue damage, and even death. Therefore, it is important to selectmaterials for graft use that will minimize the initiation of the clotting cascade.

The effect of a material on blood clotting is not easy to assess, particularlyas it is difcult to separate the effect of the material from the natural physiologic



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process of coagulation. Blood will also behave slightly differently in vivo thanit will in vitro and, like other tissues, can differ between individuals. Therefore,an in vitro coagulation test with blood from a dog may not be fully indicativeof how that material will behave when implanted into a human. Standards have

been developed, however, for coagulation tests (Bruck, 1980), and these progressfrom static, in vitro tests to dynamic, in vivo tests.Surface tension, surface charge, and surface roughness are properties of a

material that will affect the rate and amount of coagulation that takes place whenit is in contact with blood. Polymeric materials tend to adsorb a mixture of proteins to their surfaces in the rst 3060 s of contact with physiologic uid(Baier, 1975), the composition of which depends on the polar or nonpolar natureof the polymer (Herring, 1983). Adsorbtion and activation of key molecules fromblood, including Hageman factor, factor XI, and others, will trigger the clotting

cascade (Forbes, 1993). Rougher surfaces, including crimped grafts, have beenshown to increase the rate of coagulation that occurs when in contact with blood(Collins, 1983). This is probably due to the larger surface area that can come incontact with the blood. A rough surface may be desirable, however, in order topromote preclotting on the surface of a porous graft. In the case of crimping,the process also prevents kinking of the vessels during surgery or prolongedimplantation, which itself can lead to occlusion. Surface charge can help tominimize the contact between the graft and the blood elements. The formedelements of the bloodred and white blood cells and plateletshave beenshown to have a negative surface charge. As a result, a vascular surface with aslight negative charge, through the presence of either a neointima or a naturallyor articially induced surface charge, will act to repel the blood cells and plate-lets away from the vascular wall (Collins, 1983). When platelets do not comein contact with the vascular wall or foreign bodies, they are less prone to initiatethe clotting cascade.

The natural lining of blood vessels possesses unique properties that cannotbe easily mimicked. It was originally believed that the smooth surface and thenegative surface charge were the properties of an endothelial lining that pre-vented clot formation. However, studies have shown this to be more complex.In particular, the presence of endothelial cells allows for the secretion of para-crine agents that act to break down small thromboses and interfere with clotformation (Bruck, 1980). Smooth muscle and broblasts did not exhibit the samefunction, and in some cases precipitated platelet activation. Thus, as was thecase in bones resistance to fatigue, the living function of vascular tissue cannotbe fully replaced using current technologies and articial materials.

To reduce clotting in articial materials, several approaches have been taken.Heparin, a negatively charged polysaccharide that is commonly used to preventclottting in many clinical applications, has been coated on implants. Using the

same logic, anionic radicals have been included in an articial material to pro-duce the negative surface charge that only naturally can occur in polymers.Taking a different approach, materials with low surface tensions have been pro-posed, as they are less likely to attract the formed blood elements to the materialsurface and initiate a clotting cascade. This last tact is, of course, in contradictionto the suggested use of high surface tension materials to minimize hemolysis!

This latest dilemma is indicative of many decisions that must be made inselecting a material for use in biomedical applications. No materialbesides



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possibly the original tissuewill perfectly match all of the goals and constraintsof the design. It is generally necessary to weigh the benets and drawbacks of each material that makes the short list and to select the one that shows the bestbalance. While this is frustrating from a design aspect, it is what spurs material

development and allows for the continuing work in the area of material synthesisand implant design.

Property Degradation

Once a material is selected that does not seem to have a signicant, adverseeffect on the blood itself, it is necessary to evaluate how the properties of thegraft are affected during the expected contact with the physiological environ-ment. As discussed above for orthopedic implants, polymers can react to thephysiologic environment by either leaching smaller molecules into the surround-

ing tissue or absorbing water or other materials from the tissue. In the case of vascular grafts, the constant ow of blood past the implant surface will eliminatethe possibility to develop an equilibrium between the molecule concentrationsinside and outside of the implant material. As an example, nylon has been foundto absorb water from the surrounding environment when implanted (Edwards,1983). The water molecules act as plasticizers, which reduce the cohesion andbonding between the chains of the polymer. The nal result is a reduction instrength.

Whether property degradation will be a determining factor in material selec-tion for vascular implants depends largely on the planned time of implantationfor the device. For a graft that will permanently replace a section of blood vessel,it is extremely important to maintain the tissue properties at an appropriate level.However, many blood contacting applications involve short-term usefor in-stance intravenous catheters for administering blood or pharmaceuticals. In thesecases, the length of intended use is typically less than a few days, and thecatheter can be relatively easily replaced if necessary. Therefore, for short-termimplants, the prevention of blood clot formation and blood cell damage, alongwith ease of use and cost, become the determining factors in material selection.

3.3 Current Material SelectionSince the 1950s, polymer fabrics have been the primary material used for long-term vascular grafts. Nylon was introduced in 1955, but was withdrawn fromuse after the occurrence of aneurisms within the grafts indicated a loss of me-chanical integrity in vivo (Edwards, 1983). Teon (PTFE, Gore-Tex) and poly-ethylene terephthalate (PET, Dacron) have been shown to have acceptableamounts of property degradation when implanted for 20 years or more (Snyder,1983). Table 5 shows results on the loss of strength in various polymers thatwere implanted for 100 days. The Dacron showed an initial drop in strength that

then stablized at an acceptable level.Typical vascular grafts are constructed of a woven or knitted fabric and are

crimped, to both prevent kinking and to allow for longitudinal expansion (Fig.8). The response of the graft in vivo will depend not only on the constituentmaterial selected, but also on the weave of the fabric and on its processing(Sawye, et al., 1983). Grafts are typically tested to validate their tensile or burst-ing strength, and the values obtained for the constructed vessel will differ fromthose of the bulk material due to the knitted or woven nature of the fabric.



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Table 5 Average Changes in TensileProperties of Synthetic Grafts Implantedfor 100 Days

Material Loss of Strength (%)

Nylon 81Orlon 6.9Dacron 10.1Teon 3.2

Source: Data from (Edwards, 1983)

Fig. 8 Left and right: Examples of fabric arterial grafts, illustrating the crimped structure thatprevents kinking in vivo. Center: Extruded PTFE graft lined with pyrolytic carbon.

Pyrolitic or LTI (low-temperature isotropic) carbon has been found to haveexcellent anticoagulation properties (Bruck, et al., 1973), without the need forheparin coating of the material. The original LTI formulation required a solidsubstrate and so was more conducive for use on shunts or leaets of heart valves.A newer, ultra-low-temperature iostropic carbon (ULTI) can be vapor depositedonto fabrics as well (Bruck, 1980), allowing for its use in regions that requireexibility (Fig. 8). ULTI carbon, which can be deposited with a thickness of less than 1 m, has been shown not to affect the compliance of the underlyingfabric (Sharp, 1983), indicating that it will maintain its exibility when depositedon a fabric graft.

For short-term implants, generally in the form of catheters, silicone rubberhas become the standard. Its optical translucence makes it easy to monitor uidtransfer, it is easy to use, and it is highly biocompatible.

Fabric grafts and silicone tubes, however, are not adequate for the small-vesselreplacement that may be required during reconstructive surgery. As a result,



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tissue engineering techniques are currently being employed to construct small-diameter vessels from natural materials, such as chitosan (Chupa, et al., 2000).Research such as this is working to expand the possibilities for material selec-tion, not only for vascular prostheses but also for all biomedical implants.

4 SPACE-FILLING BIOMATERIALS: BREAST IMPLANTS

Implants designed to ll voids where tissue has either been removed or destroyedare commonly used in reconstructive surgery. Hard tissue, soft tissue, and com-bination materials have been utilized to replace or augment cranial defects, theloss of an ear, nose, or eye due to trauma or disease, and congenital facialabnormalities, among others. Perhaps the most familiar and infamous space-lling implant was, and is, the breast implant. Originally designed for recon-structive surgery following mastectomies, breast implants became increasinglypopular for cosmetic enhancement during the 1970s and 1980s. The question of whether the implants were the cause of the illnesses and disabilities reported inthe 1990s has never been answered to everyones satisfaction, and probablynever will be. It is an example of a biocompatibility issue that, if true, wentunnoticed for many years. It also serves notice to all biomedical implant de-signers that because an implant does not indicate any problems after 5 yearsdoes not mean that monitoring of potential complications should stop.

Breast implants were designed with one function in mindto replace tissuethat had either been removed through surgery or that an individual felt naturehad left lacking. There was no other physiologic role for the implant, whichmade its design much simpler. The rst breast augmentation utilized injectionsof silicone, parafn wax, or bees wax directly into the tissue surrounding thebreast. This method was banned by the Food and Drug Administration (FDA)in the 1960s as the injected material was seen to migrate and lose its shape. Inaddition, because of the large contacting surface area between the injected ma-terial and surrounding tissue, adverse tissue reactions were seen (Frisch, 1983).Since that time, breast implants have utilized conned volumes of materials(Figure 9).

4.1 FunctionSpace Filler

As stated above, the single role of a breast implant is to ll up a given volumeof space within the body. However, cosmetic appearance and feel have alsodictated much of the development of the implants. The density of the implantshould be similar to or less than the surrounding tissue, so that tissue damagedoes not occur due to increased weight. The consistency of the material shouldalso be somewhat similar to the composite of fatty and connective tissue that ithas replaced. Thus, metals, ceramics, and solid pieces of polymer would beinappropriate.


Capsule Formation

No articial material that is implanted into the body will be ignored completely.The degree of the reaction from the immune system will depend on a number



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4 SPACE-FILLING BIOMATERIALS: BREAST IMPLANTS 1189

Fig. 9 Examples of saline-lled breast implants showing the silicone elastomer envelope. Sur-faces can either be textured (upper) or smooth (lower). The textured design is thought to reduce

the incidence of capsular contraction.

of factors, including the chemical or electrical activity of the material. The initialreaction of the body to a foreign object is to wall it off. A brous capsulearesult of the normal wound healing response of the bodywill develop aroundthe structure, with the thickness of the capsule increasing in regions of highchemical activity, high electrical activity, or sharp corners. The presence of thisbrous capsule formation can be taken advantage of to assist with implant x-ation, however.

For objects with smooth surfaces, the brous capsule will simply grow aroundthe outer surface of the material. If the surface of the foreign body is porous,however, brous tissue will begin to grow into the interconnected pores of thestructure. For bony implants, this tissue can become calcied and develop intoa bony junction between the implant and the surrounding tissue. For soft tissueimplants, this brous capsule can become integrated with the surrounding con-nective tissue of the body to provide rm anchorage of the implant at a desiredlocation. Complete brous tissue integration with the implant material will be

most efciently achieved with a pore size ranging from 5 to 30 m (Rose andLitske, 1989; Wilkerson and Zalina, 1994).

The amount of desirable brous capsule integration with the implant must,however, be assessed and the area of porous surface determined as part of theoverall design. One breast implant design that failed because of an overabun-dance of brous capsule integration was the sponge design developed in the1950s (Dukes and Mitchley, 1962; Edgerton and McClary, 1958). This implant,



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which met the space-lling requirements and was composed of polyvinyl alcoholfoam, was found to experience complete brous ingrowth. Although the implantwas rmly xed in the correct anatomical position, the brous capsule tendedto contract and calcify, resulting in a condition known as marble breast syn-

drome.All breast implant designs have shown some incidence of capsular contrac-tion, which in severe instances has required the implant to be removed. This hasled to the FDA requirement in the United States that all implant manufacturersinform prospective patients of this potential complication and the fact that breastimplants should not be considered to be lifetime prostheses. While no implantdesign has been developed that eliminates this phenomenon, the area of xationpatches on the posterior surface has been reduced, textured (though nonporous)surfaces have been introduced for the silicone envelope, and clinical interven-

tions, including massage, have been prescribed in attempts to minimize the prob-lem.

Carcinogenesis and Immunological Complications

For both carcinogenesis and general immunological considerations, it is impor-tant to examine the response of not only the bulk material that is to be used butalso the components of the materials. This includes unreacted monomer, plas-ticizers, llers, and any products that can result from the metabolic breakdownof the bulk material in vivo . Any chemical compound that could be released byan implant is a potential hazard. The complexity of this analysis becomes readilyevident.

The exact mechanism, or mechanisms, by which healthy cells in the bodymutate into malignant cells and cause cancer is not fully understood. However,it is a well-accepted principal that all possible attempts should be made to avoidimplanting a material into the body that would increase the risk of cancer. Ascell mutations and tumor development can take many years to become evident,how can this be done practically for newly developed materials?

Carcinogenesis, or the production of cancer, can occur in the immediate vi-cinity of an implant, due to the presence of various chemicals, or at a remotelocation in the body, due to the transport of released materials. Materials in thebody can act in one of three ways to cause cancer: (1) as a complete carcinogen,resulting in cancerous changes by itself; (2) as a procarcinogen, a benign chem-ical that is metabolically modied into a carcinogen by the body; or (3) as a co-carcinogen that is not likely to itself cause cancer, but will increase the activityof a complete or procarcinogen with which it comes into contact. It is hypoth-esized that there is no threshold below which a carcinogen present in the bodyis completely safe; however, as concentrations of a chemical increase, the prob-ability that they will induce cellular changes that result in cancer also increase

(Black, 1999). A tumor may also develop in response to the presence of a solidforeign body made of a noncarcinogenic material, due to the disruption of thenormal environment of the surrounding tissues (Black, 1999).

Testing of materials for their carcinogenic potential ranges from cell culturestudies that examine whether a material causes cells to mutate (Forster, 1986),to in vivo animal tests and the development of databases to track the occurrence



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5 SUMMARY 1191

of tumors in implant recipients. There are drawbacks to all of these tests, andnone, with the exception of a longitudinal study of tens of thousands of patientswith the same implant, can denitively determine if there is a link to cancer foran implant or a material that does not initially appear to be carcinogenic. The

primary benet of these tests is to eliminate from consideration materials thatappear to have a high likelihood of causing tumors. General guidelines for basicbiocompatibility testing have been issued by the International Standards Orga-nization and other regulatory agencies (ISO-10993, 19922000).

Similar problems exist for the documentation of other systemic or remotetissue complications. Techniques have been developed to evaluate the immu-nological response to a material (Marritt, 1986); however, the link between localimmunological sensitivity and a systemic autoimmune reaction may not be easilydiscernable. It is important to determine whether the occurrence of a particular

disorder, whether it is lymphoma or rheumatoid arthritis, is occurring within theimplanted population at rate that is higher than the general population. One of the difculties in tracking the incidence of these occurrences is the fact thatmany implant materials and combinations change on a time scale that is shortcompared to the development of cancer and other immunological conditions. Asthese material combinations change, it introduces another variable that requiresa higher number of subjects before a statistical determination can be made.However, the establishment of databases at the national, or preferably interna-tional, level for tracking of all complications associated with biomedical implantswill allow for a more complete analysis and more rapid understanding of theproblems associated with this imperfect science.

4.3 Current Material Selection

Prior to 1992, the most common form of breast implant involved a small bag of silicone elastomer lled with silicone gel. Originally introduced in 1962 (Cronin,1983), this implant was popular among patients and physicians alike. While thesilicone elastomer of the bag has been utilized in biomedical implants for manyyears without signicant adverse effects, the breast implant scare of the 1990swas based on rumors that the silicone gel had a tendency to leak from thesurrounding bag and cause adverse reactions at both local and remote locations.Since that time, the majority of breast implants have utilized saline within thesame silicone rubber bag. If the saline does leak, or the implant does rupture,the reaction with the tissue will be completely benign. This design has been anoption since 1965 (Rubin, 1983); however, most patients opted for the siliconeversion due to its improved cosmetic appearance.

5 SUMMARY

There are few, if any, current implants that can be described as perfectly meeting

their design goals and constraints so that no further investigation of design ormaterial selection is warranted. As materials continue to be developed, whetherspecically for biomedical applications or in some different discipline, the se-lection of biomaterials for implants will remain a challenge in the design of theoptimum implant. The evolution of tissue engineering from a bench-top scienceto a clinically workable tool for new implant design will also open new doors



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for the development and use of biomaterials. In all of these cases, however, thesame principles apply to the selection of a material for a biomedical application.The selection process can be summarized in the following way:

1. Determine the functional requirements of the material for the particularapplication (preferably with an idea of the overall design in hand).

2. Select a group of materials that appear to meet those functional require-ments and ensure that all conrming tests are conducted in an environ-ment that simulates human physiology.

3. Determine the biocompatibility of the materials in terms of material deg-radation, tissue effects, blood compatibility, implant xation, and long-term physiologic consequences.

4. Complete the design and approval process, with mechanisms in place toobtain data on functional or material complications for many years afterclinical use is initiated.

As with all aspects of engineering design, selection of materials for biomedicalimplants is like solving a puzzlegiven all of the goals and constraints, whatis the one solution that best answers the question that has been posed? And if a current solution does not exist, can one be developed? Some may argue thatcareful selection of materials for medical applications is more important than

for other realms of engineering, due primarily to the overriding fear most peoplehave of disease or injury. Others may disagree. It is certain, however, that de-velopment, evaluation, and selection of biomaterials is one of the most chal-lenging areas of materials science. No two physiologic environments are exactlythe same; therefore, no two implants will respond in exactly the same way. It isalso much more difcult to predict the loading, chemical, and biological envi-ronment that an implant will see in a particular individual than it is to modelthe potential variations in loading of a car or a building. The science of bio-medical materials has improved tremendously since the days of George Wash-

ingtons wooden teethand the development of new materials has just begun.BIBLIOGRAPHY ASTM-F136, Standard Specication for Wrought Titanium-6 Aluminum-4 Vanadium ELI (Extra Low

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ASTM-F562, Standard Specication for Wrought Cobalt-35 Nickel-20 Chromium-10 Molybdenum Alloy for Surgical Implant Applications, American Society for Testing and Materials, W. Con-shohocken, PA, 2000.

ASTM-F648, Standard Specication for Ultra-High-Molecular-Weight Polyethylene Powder and Fab-ricated Form for Surgical Implants, American Society for Testing and Materials, W. Consho-hocken, PA, 2000.

Baier, R. E., Blood Compatibility of Synthetic Polymers: Perspectives and Problems. in Polymersin Medicine and Surgery , R. L. Kronenthal, Z. Osei, and E. Martin (eds.), Plenum, New York,1975, pp. 139159.

Bernache-Assolant, D., Bioceramics: Processing-Properties, in Biomaterials Degradation: Funda-mental Aspects and Related Clinical Phenomena , M. A. Barbosa (ed.), Elsevier Science Pub-lishers, Amsterdam, 1991, pp. 111168.

Black, J., Biological Performance of Materials: Fundamentals of Biocompatibility , Marcel Dekker,New York, 1999.



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Chupa, J. M., Foster, A. M., Sumner, S. R., Madihally, S. V., and Matthew, H. W., Vascular CellResponse to Polysaccharide Materials: In Vitro and In Vivo Evaluation, Biomaterials, 21(22),23152222 (2000).

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