Fundamentals of MedicalImplant MaterialsSoumya Nag and Rajarshi Banerjee, University of North Texas

OVER THE LAST SEVERAL DECADES,an increase in longevity and life expectancyhas raised the average age of the world’spopulation. Among the countries currently clas-sified by the United Nations as more developed(with a total population of 1.2 billion in 2005),the overall median age rose from 29.0 in 1950to 37.3 in 2000 and is forecast to rise to 45.5by 2050 (Ref 1). This worldwide increase inthe average age of the population has, in turn,led to a rapidly increasing number of surgicalprocedures involving prosthesis implantation,because as the human body ages, the load-bear-ing joints become more prone to ailments. Thishas resulted in an urgent need for improvedbiomaterials and processing technologies forimplants, more so for orthopaedic and dentalapplications.

Need for Prostheses

The first question to ask while undertakingsuch a study is “What is the need for implants toreplace or fix human joints during traumatic con-ditions?” Human joints are complex and delicatestructures capable of functioning under criticalconditions, and it is a great challenge for doctorsas well as scientists to develop site-specificimplants that can be used in a human body toserve a specific purpose for orthopaedic, dental,ophthalmological, cardiovascular, cochlear, andmaxillofacial applications.Synovial joints such as hips, knees, and

shoulders perform due to the combined effortsof articular cartilage, a load-bearing connectivetissue covering the bones involved in the joints,and synovial fluid, a nutrient fluid secretedwithin the joint area (Ref 2–4). However, thesejoints are more often than not prone to degener-ative and inflammatory diseases that result inpain and joint stiffness (Ref 5). Apart fromthe usual decay of articular cartilage due toage, there are illnesses such as osteoarthritis(inflammation of bone), rheumatoid arthritis(inflammation of synovial membrane), andchondromalacia (softening of cartilage). An

astounding 90% of people above the age of 40suffer from such degenerative conditions. Thestructure of a normal bone is distinctly differentwhen compared to a bone that is suffering fromosteoporosis, with the bone cell density beingsubstantially lower for the osteoporotic boneas compared to the normal bone. Such prema-ture joint degeneration may arise mainly fromthree conditions: deficiencies in joint biomate-rial properties, excessive loading conditions,and failure of normal repair processes (Ref 2).Although minor surgical treatments are doneto provide temporary relief to numerouspatients, there is a consensus that the ultimatestep is to replace the dysfunctional naturaljoints for prolonged pain relief and mobility.Thus, the field of arthroplasty has become pop-ular in the surgical world and, according to themedical term, means surgical repair of joints(Ref 2). Currently, one of the main achieve-ments in the field of arthroplasty is total jointreplacement (TJR), where the entire load-bearing joint (mainly in the knee, hip, or shoul-der) is replaced surgically by ceramic, metal, orpolymeric artificial materials. As stated earlier,the problem is that not all artificial materialscould be used for such purposes, only the onesthat fulfill certain broad specifications.In comparison, the human tooth, consisting of

enamel, dentin, pulp, and cementum, is a highlyspecialized calcified structure used to breakdown food. It is a site where most surgical proce-dures in humans are performed, requiringimplants of a subperiosteal (in contact with exte-rior bone surface) or endosteal (extending intothe bone tissue) nature (Ref 6). The fixtures canbe either fixed or removable, which reallydepends on the type of employed prostheses, amajority of which involve complete or partialdentures. In any case, the biomaterial interactionand tissue reaction of these implants, along withother intraoral devices, is critical for the stabilityand sustainability of dental prostheses.The following section outlines some of the

selection criteria that must be kept in mindwhen choosing an implant material for a spe-cific purpose.

Implant Properties

The property requirements of a modern-dayimplant can broadly be categorized into threeequally important features (Ref 7):

� The human body must be compatible withthe material used for the prosthesis. Whileit is understandable that there is bound tobe some amount of tissue reaction due tothe introduction of a foreign substance, theresulting changes in mechanical, physical,and chemical properties within the localizedenvironment should not lead to local delete-rious changes and harmful systemic effects.

� The implant should have the desired balanceof mechanical and physical properties neces-sary to perform as expected. The specific opti-mization of properties such as elasticity, yieldstress, ductility, time-dependent deformation,ultimate strength, fatigue strength, hardness,andwear resistance really depends on the typeand functionality of the specific implant part.

� The device under question should be rela-tively easy to fabricate, being reproducible,consistent, and conforming to all technicaland biological requirements. Some of the con-straints could include the techniques to pro-duce excellent surface finish or texture, thecapability of the material to achieve adequatesterilization, and the cost of production. Therepair of such implants in case of failure is alsovery important. It has been noted that for anydental prostheses or TJR surgery, the revisionsurgery of an implant is more difficult, haslower success rates, and may induce addi-tional damage to the surrounding tissues(Ref 8). Unfortunately, in vivo degradation,primarily due to the higher wear rates asso-ciated with artificial implant materials andthe consequent adverse biological effect ofthe generated wear debris, results in a shorterlifetime for these artificial implants whencompared with their natural counterparts.Thus, it is imperative to account for the physi-cal stability of the foreign material once it isplaced inside a human body.

ASM Handbook, Volume 23, Materials for Medical DevicesR. Narayan, editor

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Apart from these factors, the selection ofthe implant material itself is the principal crite-rion for proper functioning. No amount ofdesign changes can help if the material is notbiologically and mechanically compatible.That, along with the surgery location anddesired functioning of the artificial joint, deter-mines what material should be used. For exam-ple, smaller implants used for cochlear anddental prostheses are manufactured using aplastic or ceramic material. However, formaking total hip replacements and total kneereplacements, metals are considered the bestcandidate due to their higher tensile load-bearing capabilities. The various parts of hipand knee implants require different property char-acteristics. Thus, it is understandable that for bestresults, modern-day implants such as the Trape-zoidal-28 (T-28) hip (Ref 9, 10), the Burstein-Lane (B-L) knee (Ref 11), or the Total CondylarProsthesisKnee (Ref 12) are assembled by joiningthe various componentsmade ofmetals, ceramics,and/or polymers to form one unit.For metallic implants, casting and forging

metallic components is still one of the mostaccepted techniques in the implant fabricationarea, even though minuscule cracks and inho-mogeneous composition of parts provide majorhurdles for the process (Ref 13). Along withthat, the fabrication technique itself has someimpact on implant performance. For example,preparing rough, serrated implant surfaces helpsin better cell adhesion, differentiation, and pro-liferation (Ref 14, 15). Having porous implantshas shown to help in the growth and attachmentof bone cells.Ceramic devices are manufactured by a vari-

ety of techniques. Typical powder-metallurgy-based routes follow compaction and solid-statesintering of powdered ceramics (alumina andcalcium phosphate) or metal-ceramic compo-sites (CermeTi, Dynamet Technology, Inc.)(Ref 16). Depending on the property require-ment, the heating schedules can be variedto determine grain size and crystallinity. Forexample, sintering at a higher temperature(liquid-phase sintering or vitrification) is oftendone to produce a combination of fine-grainedcrystalline matrix with reduced porosity (Ref6). Other materials, such as hydroxyapatite,are used as coatings on various biomaterialsand are plasma sprayed onto the material. Con-ventional casting routes are adopted to producebioceramic glasses. For this, one must ensurethat the solidification process in this method isslow enough to prevent crystallization (other-wise, polycrystalline products will form). Thisfrustrated nucleation leads to the formationof glasses below the glass transition tempera-ture (Ref 6). If required, these glasses canbe annealed at higher temperatures to nucleateand grow crystalline phases in the glassymatrix, commonly forming a new class of mate-rials called glass-ceramics.Polymeric implants can be divided into two

broad categories: natural and synthetic poly-mers. Natural polymers are made of an

extracellular matrix of connective tissue, suchas tendons, ligaments, skin, blood vessels, andbone. However, they are very difficult to pro-cure and reproduce on a regular basis. Gener-ally, synthetic polymers are synthesized bypolymerization and condensation techniquesto form long chains of the desired shape andproperty (Ref 6). Other synthetic materials,such as fibers and biotextiles, are prepared bymelt spinning and electrospinning, while hydro-gels are prepared by simply swelling cross-linked polymeric structures in water or otherbiological fluids.

Development of Implant Materials

Metallic Implants

In the early days of arthroplastic surgery,stainless steel was considered a viable implantmaterial mainly because of its availability andprocessing ease. Alloying additions of chro-mium, nickel, and molybdenum were made tothe ferrous matrix to prepare alloys such316L, also known as ASTM F138 (Ref 2). Theywere primarily used to make temporary devicessuch as fracture plates, screws, and hip nails.However, as TJR surgery became popular,it was evident that the very high modulus ofstainless steel (�200 GPa, or 29 � 106 psi)was a deterrent (Table 1). Also, researchersstarted looking for alloys that were more bio-compatible and corrosion and wear resistant.Cobalt-base alloys came into the picture wherewrought alloys were used to fabricate prostheticstems and load-bearing components. Eventhough they offered excellent corrosion resis-tance, wear resistance, and fatigue strength,these Co-Cr-Mo alloys (ASTM F75 and F799)still had higher modulus (�210 GPa, or 30 �106 psi) (Table 1) and inferior biocompatibilitythan what was desired for implant materials(Ref 2). After the early 1970s, titanium alloysstarted to gain much popularity due to theirexcellent specific strength, lower modulus,superior tissue compatibility, and higher corro-sion resistance (Ref 17). Commercially puretitanium (ASTM F67) was the first to be usedbecause its oxide (titanium in atmosphere read-ily forms a nascent oxide layer) had excellentosseointegration properties; that is, human bonecells bonded and grew on the titanium oxidelayer quite effectively. However, due to its lim-ited strength, the implants were confined to spe-cific parts, such as hip cup shells, dental crownand bridges, endosseous dental implants,

pacemaker cases, and heart valve cages(Ref 18). To improve the strength for load-bearing applications such as total joint replace-ments, the alloy Ti-6Al-4V ELI (ASTM F136,the extra-low interstitial, or ELI, alloy com-posed of titanium, 6 wt% Al, and 4 wt% V)was chosen. This Ti-6Al-4V alloy was origi-nally developed for aerospace applications andhad superior performance in the field of avia-tion, with an elastic modulus of approximately110 GPa (16 � 106 psi) (Table 1), only half thatof 316L stainless steel. It was used for TJR sur-gery with modular femoral heads and for long-term devices such as pacemakers. However, itwas soon discovered that the presence of vana-dium caused cytotoxicity and adverse tissuereactions (Ref 19, 20). Thus, niobium and ironwere introduced, replacing vanadium, todevelop alloys such as Ti-6Al-7Nb (Ref 21)and Ti-5Al-2.5Fe (Ref 22). Other alloys withaluminum additions, such as Ti-15Mo-5Zr-3Al(Ref 23) and Ti-15Mo-2.8Nb-3Al (Ref 2), weretried. Further studies showed that the release ofboth vanadium and aluminum ions from thealloys may cause long-term health problems,such as peripheral neuropathy, osteomalacia,and Alzheimer diseases (Ref 24, 25). Thus,Ti-6Al-4V somewhat lost its importance as themost viable orthopaedic alloy.These circumstances led to an urgent need

to develop newer and better orthopaedic alloys.This required the researchers to first identifythose metallic elements that were completelybiocompatible and could be alloyed with tita-nium. The ideal recipe for an implanted alloyincluded excellent biocompatibility with noadverse tissue reactions, excellent corrosionresistance in body fluid, high mechanicalstrength and fatigue resistance, low modulus,low density, and good wear resistance. Unfortu-nately, only a few of the alloying elementsdo not cause harmful reactions when plantedinside the human body (Ref 26). These includetitanium, molybdenum, niobium, tantalum, zir-conium, iron, and tin. Of these, only tantalumshowed an osseocompatibility similar to thatof titanium. However, its high atomic weightprevented tantalum from being used as aprimary alloying addition. In fact, the biocom-patibility of higher amounts of tantalum andpalladium additions was only tested for dentaland craniofacial prostheses where implantweight would not be of much concern(Ref 27). For other types of load-bearingimplants, several molybdenum- and niobium-base alloys were analyzed. Investigations onternary Ti-Mo-Fe alloys were carried out,

Table 1 Comparison of mechanical properties of commonly used orthopaedic alloys

Alloy

Modulus Yield strength Ultimate tensile strength

GPa 106 psi MPa ksi MPa ksi

Stainless steel 200 29 170–750 25–110 465–950 (65–140)Co-Cr-Mo 200–230 29–33 275–1585 40–230 600–1795 (90–260)Commercially pure Ti 105 15 692 100 785 115Ti-6Al-4V 110 16 850–900 120–130 960–970 140–141

Fundamentals of Medical Implant Materials / 7

where the strengthening effect of the iron addi-tion was studied in a Ti-7.5Mo alloy (Ref 28,29). Guillermot et al. conducted tests on Ti-Mo-Fe-Ta alloys with hafnium additions(Ref 30). The early works of Feeney et al.focused on one of the most promising quater-nary molybdenum-base b-titanium alloys,Ti-11.5Mo-6Zr-4.5Sn, also known as bIII(Ref 31). The phase transformations occurringin these alloys were found to be similar tothat of binary titanium-molybdenum alloys.At room temperature, the as-quenched bIIIalloy showed low yield strength, high ductility,and high toughness. The effects of iron intitanium-molybdenum alloys (Ref 28) and thesuperior properties of bIII (Ref 31) were finallycombined together to develop Ti-12Mo-6Zr-2Fe (Ref 32, 33), which recorded superior yieldstrength and modulus values. A parallel, if notbetter, effort was made to develop niobium-base b-titanium alloys. Karudo et al. (Ref 34)and Tang et al. (Ref 35) developed some alloysbased on the Ti-Nb-Ta, Ti-Nb-Ta-Zr, Ti-Nb-Ta-Mo, and Ti-Nb-Ta-Sn systems. Of thedifferent alloys that were chosen, the tensilestrength and elongation of Ti-29Nb-13Ta-4.6Zr alloy were found to be greater than orequivalent to those of conventional titaniumalloys for implant materials (Ref 36–38).On comparing the hardness values of thequaternary alloys, it was evident that the homo-genized samples had higher hardness thanthe air- or water-quenched samples. Finally,the dynamic moduli were observed to be lowestat 5 at.% Zr and a niobium/tantalum ratio of12.0, which was attributed to the preferred siteoccupancy of niobium, tantalum, and zirconiumwithin the body-centered cubic unit cell and itseffect on the nature of bonding (Ref 35, 39).The alloys that possessed the lowest moduliwere Ti-35.5Nb-5.0Ta-6.9Zr and Ti-35.3Nb-5.7Ta-7.3Zr.Based on this research, a number of contem-

porary and prospective alloys were developed,such as Ti-12Mo-6Zr-2Fe (Ref 32, 33); Ti-15Mo-3Nb-0.3O (Ref 40); interstitial oxygen,also referred as TIMETAL 21 SRx; Ti-13Nb-13Zr (Ref 41); and Ti-35Nb-7Zr-5Ta (Ref 42).Interestingly, all these alloys were primarily

b-type titanium alloys. This shift in the searchfor better biomaterials from a/b-titanium tob-titanium alloys could be explained by the factthat the latter fit in very well with the tightmechanical property requirements of orthopae-dic alloys. Two of those important propertiesinclude yield strength and elastic modulus.Yield Strength. The yield strength deter-

mines the load-bearing capability of theimplant. For example, in the case of TJR sur-geries where a high load-bearing capability ofthe implant is essential, one ideally needs anappropriately high yield strength value of thealloy. Thus, the orthopaedic alloys should havea sufficiently high yield strength value withadequate ductility (defined by percentage elon-gation or percentage reduction of area in a stan-dard tensile test). Table 2 lists the yield strengthand ultimate tensile strength values of someof the common titanium alloys. Interestingly,some of the metastable b-titanium alloys doexhibit very high values in comparison to thea- or a/b-titanium alloys.Elastic Modulus. A number of experimental

techniques have been used to determine theelastic properties of solids (Ref 43). Thereis always a concern for the relatively highermodulus of the implant compared to that ofthe bone (�10 to 40 GPa, or 1.5 to 6 � 106

psi) (Ref 2). Long- term experiences indicatethat insufficient load transfer from the artificialimplant to the adjacent remodeling bone mayresult in bone reabsorption and eventual loosen-ing of the prosthetic device (Ref 44, 45). Ithas been seen that when the tensile/compressiveload or the bending moment to which theliving bone is exposed is reduced, decreasedbone thickness, bone mass loss, and increasedosteoporosis occur. This is termed the stress-shielding effect, caused by the difference inflexibility and stiffness, which is partly depen-dent on the elastic moduli difference betweenthe natural bone and the implant material(Ref 46). Any reduction in the stiffness of theimplant by using a lower-modulus materialwould definitely enhance the stress redistribu-tion to the adjacent bone tissues, thus minimiz-ing stress shielding and eventually prolongingthe device lifetime. In an attempt to reduce

the modulus of the implant alloys to match thatof the bone tissue, Ti-6Al-4V and related a/balloys were considered to be inferior. Theb-titanium alloys have a microstructure pre-dominantly consisting of b-phase that exhibitslower overall moduli. Table 2 shows that Ti-15Mo-5Zr-3Al, Ti-12Mo-6Zr-2Fe, Ti-15Mo-3Nb-0.3O (21SRx), and Ti-13Nb-13Zr haveelastic moduli ranging from 74 to 88 GPa(11 to 13 � 106 psi), which is approximately2 to 7 times higher than the modulus of bones.Fatigue. Variable fatigue resistance of the

metallic implants is also a cause of concernwhile developing an alloy. The orthopaedicimplants undergo cyclic loading during bodymotion, resulting in alternating plastic deforma-tion of microscopically small zones of stressconcentration produced by notches and micro-structural inhomogeneities. Standard fatiguetests include tension/compression, bending,torsion, and rotation-bending fatigue testing(Ref 2).There were several advantages and disadvan-

tages of the various alloys that were researched,and many more will probably be developed andtested in the near future. Two of the mostpromising alloys appear to be the Ti-35Nb-7Zr-5Ta (often referred to as TNZT) andTi-29Nb-13Ta-4.6Zr (often referred to asTNTZ) compositions, mainly because thesealloys exhibit the lowest modulus valuesreported to date—�55 GPa (8 � 106 psi) inthe case of TNZT, almost 20 to 25% lower thanother available alloys (Ref 2, 42). While TNZTwas developed at Clemson University by Racket al. (Ref 42), TNTZ was developed at TohukuUniversity, Sendai, Japan, by Niinomi et al.(Ref 34). TNZT is now commercially sold byAllvac in the United States as TiOsteum andTiOstalloy. Its low yield strength value (547MPa, or 79 ksi) was increased by adding inter-stitial oxygen; thus, Ti-35Nb-7Zr-5Ta-0.4Oshowed a strength of 976 MPa (142 ksi) anda modulus of 66 GPa (9.6 � 106 psi) (Ref 42).

Ceramic Implants

Ceramics, including glasses and glass-ceramics, are used for a variety of implant

Table 2 Mechanical properties of orthopaedic alloys developed and/or used as orthopaedic implants

Alloy designation Microstructure

Elastic modulus Yield strength Ultimate tensile strength

GPa 106 psi MPa ksi MPa ksi

Commercially pure Ti {a} 105 15 692 100 785 115Ti-6Al-4V {a/b} 110 16 850–900 125–130 960–970 140–141Ti-6Al-7Nb {a/b} 105 15 921 135 1024 150Ti-5Al-2.5Fe {a/b} 110 16 914 130 1033 150Ti-12Mo-6Zr-2Fe {Metastable b} 74–85 10–12 1000–1060 145–155 1060–1100 155–160Ti-15Mo-5Zr-3Al {Metastable b} 75 10 870–968 125–140 882–975 130–140

{Aged b + a} 88–113 13–16 1087–1284 160–190 1099–1312 160–190Ti-15Mo-2.8Nb-3Al {Metastable b} 82 12 771 110 812 115

{Aged b + a} 100 14 1215 175 1310 190Ti-13Nb-13Zr {a0/b} 79 11 900 130 1030 150Ti-15Mo-3Nb-0.3O (21SRx) {Metastable b} + silicides 82 12 1020 150 1020 150Ti-35Nb-7Zr-5Ta {Metastable b} 55 80 530 75 590 85Ti-35Nb-7Zr-5Ta-0.4O {Metastable b} 66 9 976 140 1010 145

Source: Ref 2

8 / Introduction

applications in dental and orthopaedic pros-theses. Implanting ceramics in the body canpresent a number of different scenarios. Thebioceramic-tissue attachment can occur due tophysical attachment or fitting of inert ceramicto the tissue (morphological fixation), boneingrowth and mechanical attachment intoporous ceramic (biological fixation), chemicalbonding of bones with the dense, nonporousceramic (bioactive fixation), or temporaryattachment of resorbable ceramic that is finallyreplaced by bones (Ref 6).One of the most commonly known groups of

bioactive ceramics is the calcium phosphates.They are naturally formed in minerals as wellas in the human body. These bioceramics canbe further classified in terms of their calcium-phosphorus ratios. For example dicalciumphosphate, tricalcium phosphate, and tetra cal-cium phosphate have calcium-phosphorusratios of 1, 1.5, and 2, respectively. In the caseof hydroxyapatite (HA, or 3Ca3(PO4)2�Ca(OH)2), which is considered a bioactive mate-rial, the calcium-phosphorus ratio is 1.67, andthis ratio must be accurately maintained. Other-wise, during heat treatments the compound candecompose to more stable products such as a-or b-tricalcium phosphate. To prevent such anoccurrence, many efforts have been directedtoward the development of fabrication routesfor HA that mainly involve compaction fol-lowed by sintering. Despite the enhancedefforts toward better processing routes,abnormalities such as dehydration of HA andformation of defects and impurities continueto arise, and such defects can be characterizedby x-ray diffraction, infrared spectroscopy,and spectrochemical analyses. Calcium-phos-phate-base materials can be used for bioactiveas well as bioresorbable fixations in non-load-bearing parts and for coatings on metallicimplants via sputtering techniques such asplasma spraying. Other commonly known pro-cessing routes are based on electrophoresis,sol-gel, and electrochemical processing (Ref47). Recently, laser-induced calcium-phos-phate-base surface coatings have been success-fully deposited to obtain desired biologicalproperties in terms of cell adhesion, differentia-tion, and proliferation (Ref 48). These coatingsare ideally expected to be of desired thickness,have excellent adhesion strength, and preventbiodegradation. They are also used for makingbone cements, a calcium-deficient HA-basedproduct for anchoring artificial joints by fillingin the space between prosthesis and bone. Suchanchoring with soft tissues and bone can alsobe achieved by using glasses of certain propor-tions of SiO2, Na2O, CaO, and P2O5. Ideally,such glasses are processed so that they containless than 60 mol% of SiO2, a high Na2O andCaO content, and a high CaO/P2O5 ratio (Ref6). Depending on the relative amount of theaforementioned oxides, they can be bioactive(form an adherent interface with tissues) orbioresorbable (disappear after a month ofimplantation).

Among bioinert implant materials, alumina(Al2O3) is the most commonly known ceramic,used for load-bearing prostheses and dentalimplants. It has excellent corrosion and wearresistance and high strength. In fact, the coeffi-cient of friction of the alumina-alumina surfaceis better than that of metal-polyethylene sur-faces (Ref 6). It also has excellent biocompati-bility that enables cementless fixation ofimplants. Purer forms of alumina with finergrain sizes can be used to improve mechanicalproperties such as strength and fatigue resis-tance, as well as increase the longevity of theprosthetic devices (Ref 6). Despite these advan-tages, the primary drawback of using alumina-base ball-and-socket joints is the relatively highelastic modulus of alumina (>300 GPa, or 44 �106 psi), which can be responsible for stress-shielding effects. However, much of this issolved by using zirconia (ZrO2)-base productsthat have lower elastic modulus (�200 GPa,or 29 � 106 psi). Again, while both alumina-or zirconia-ceramic femoral heads offer excel-lent wear resistance, these ceramics do not havethe same level of fracture toughness as theirmetallic counterparts, leading to problems suchas fracture of these heads in use. This has evenled to the recall of hip implants using zirconiafemoral heads (Ref 49). Furthermore, theuse of a ceramic femoral head attached to ametallic femoral stem also leads to an undesir-able abrupt ceramic/metal interface in the hipimplant. These are outstanding issues in termsof optimized implant design and must beaddressed. There are some efforts toward devel-oping the concept of a unitized implant thatuses a laser-based processing technique to fab-ricate a monolithic functionally-graded implant,the details of which are discussed in the section“Functionally-Graded Implants: Hybrid Proces-sing Techniques” in this article.

Polymeric Implants

Polymers are the most widely used materialsfor biomedical devices for orthopaedic, dental,soft-tissue, and cardiovascular applications, aswell as for drug delivery and tissue engineering.They consist of macromolecules having a largenumber of repeat units of covalently bondedchains of atoms (Ref 50). The polymers caninclude a range of natural materials, such ascellulose, natural rubber, sutures, collagen, anddeoxyribonucleic acid, as well as syntheticallyfabricated products, such as polyethylene (PE),polypropylene (PP), polyethylene terephthalate(PET), polyvinyl chloride (PVC), polyethyleneglycol (PEG), polycaprolactone (PCL), polyte-trafloroethylene (PTFE), polymethyl methacry-late (PMMA), and nylon (Ref 6).Natural polymers are often pretty similar to

the biological environment in which they areused, because they are basically an extracellularmatrix of connective tissue such as tendons,ligaments, skin, blood vessel, and bone(Ref 6). Thus, there is a reduced chance of

inflammation and risk of toxicity when intro-duced into the body. The natural set of poly-mers perform a diverse set of functionsin their native setting; for example, polysac-charides such as cellulose, chitin, and amyloseact as membrane support and intracellular com-munication; proteins such as collagen, actin,myocin, and elastin function as structural mate-rials and catalysts; and lipids function as energystores (Ref 51). Also, it is a great advantagethat these naturally occurring implants caneventually degrade after their scheduled “task”is complete, only to be replaced by the body’sown metabolic process. This degradation ratecan be controlled to allow for the completionof the specific function for which the implantwas introduced. The main problem with thesenaturally formed polymers is their reproducibil-ity; the material is very specific to where andwhich species they are extracted. Also, due totheir complex structural nature, synthetic prepa-ration of these materials is very difficult.The synthetic polymers can be prepared

by addition polymerization (e.g., PE, PVC,PMMA) or condensation polymerization (e.g.,PET, nylon). In the former process, the mono-mers go through the steps of initiation, propaga-tion, and termination to reach a desired lengthof polymeric chain. In contrast, the condensa-tion polymerization process usually involves areaction of two monomers, resulting in elimina-tion of small molecules such as water, carbondioxide, or methanol. These materials exhibita range of hydrophobic to hydrophilic proper-ties and thus are used for specific applicationsonly. For example, soft contact lenses that arein constant contact with human eyes are prefer-ably made of materials that are hydrophilic,such as poly 2-hydroxyethyl methacrylate(polyHEMA).Table 3 lists the applications of various bio-

polymers (Ref 6).The mechanical and thermalproperties of polymers are dictated by severalparameters, such as the composition of back-bone and sidegroups, structure of chains, andmolecular weight of molecules (Ref 50). Inthe case of polymers, structural changes at hightemperature are determined by performing dif-ferential scanning calorimetry experiments.The deformation behavior of polymers can beanalyzed by dynamic mechanical analyses aswell as via normal tensile testing of dog bonesamples. It should be noted here that, comparedto metals and ceramics, polymers have muchlower strength and modulus, but they can bedeformed to a much greater extent beforefailure.A relatively new area of research has focused

on biodegradable polymers that do not need tobe surgically removed on completion of theirtask. They are used for five main types of deg-radable implant applications: temporary supportdevice, temporary barrier, drug-delivery device,tissue scaffold, and multifunctional implants(Ref 6). The additional concern while design-ing this type of implant is the toxicity of thedegradation products, along with the obvious

Fundamentals of Medical Implant Materials / 9

biocompatibility issues of all implant materials.The terms biodegradation, bioerosion, bioab-sorption, and bioresorption are all looselycoined in the medical world to indicate thatthe implant device would eventually disappearafter being introduced into the body (Ref 6).The successful use of a degradable polymer-based device depends on understanding howthe material would lose its physicochemicalproperties, followed by structural disintegrationand ultimate resorption from the implant site.Despite their potential advantages, there areonly a limited number of nontoxic materialsthat have been successfully studied andapproved by the U.S. Food and Drug Adminis-tration as degradable biopolymers. Theseinclude polylactic acid, polyglycolic acid, poly-dioxanone, polycaprolactone, and poly(PCPP-SA anhydride), along with the naturally occur-ring collagen, gelatin, and hyaluronic acid(Ref 6).Several totally or partially biodegradable

self-polymerizing composites are being usedfor orthopaedic surgery and dental applications.For fixation of endoprostheses, self-curingacrylic resins based on blends of PMMA parti-cles and MMA monomer or a copolymer ofMMA with styrene are often used (Ref 52).The slurry containing the aforementionedblends can be introduced into the bone cavity.The use of biodegradable composites is greatlyencouraged for making bone cements and beadsfor drug-delivery applications. This is primarilydue to their minimal release of residual mono-mers along with their ability to control theresorption rates in conjunction with boneingrowth. Biodegradable antibiotic-loadedbeads have the advantage of releasing the entireload of drug as they degrade (Ref 52). Compo-sites based on polypropylene fumarate andPMMA are being used extensively for thesepurposes. For example, partially resorbablepolymeric composites with bioactive propertieshave been prepared by adding aqueous a- tri-calcium phosphate dispersions to PMMA bonecement (Ref 53). The resulting composite wasa suitable bone substitute with a polymeric

porous body and bioactive inorganic phase con-fined inside the pores. Again, PMMA/PCLbeads formulated with partially biodegradableacrylic cements were used for delivery of drugssuch as antibiotics, analgesics, or antiinflamma-tories (Ref 53).

Functionally Graded Implants—Hybrid Processing Techniques

As amply evident from the discussions in thepreceding sections, the need for prosthesisimplants, ranging from dental to orthopaedicapplications, is increasing at an alarming rate.While currently-existing implants functionappropriately, they do not represent the bestcompromise of required properties. Further-more, the present manufacturing of implants islargely via subtractive technologies involvingsubstantial material waste, leading to increasedcosts and time of production. Therefore, animperative need exists for functionally-gradedimplants representing a better balance of prop-erties and manufactured via novel additivemanufacturing technologies based on near-netshape processing.Some specific problems associated with cur-

rently-used implant manufacturing processesand the consequent compromise in propertiesare listed as follows:

� The manufacturing is based on conventionalcasting and forging of components, followedby material-removal steps via subtractivetechnologies such as precision machining.These technologies not only involve sub-stantial material waste but are also limitedto monolithic components without any com-positional/functional changes within thesame component.

� Diverse property requirements at differentlocations on an implant are satisfied by join-ing different components (e.g., femoral stemand femoral head) made of different materi-als in a total hip replacement system. This

always leads to the formation of chemicallyabrupt interfaces that are detrimental to theproperties of the implant. For example, itwas a standard convention of using titaniumalloy stems for orthopaedic applications tobe fitted with more wear-resistant cobaltalloy for the head. However, some of thedesigns showed significant fretting corrosioneffects due to micromotion between thesecomponents (Ref 54).

The current manufacturing route for implantsdoes not allow custom designing for specificpatients with rapid turnaround times. Conse-quently, instead of custom designing theimplant, the surgeon is often forced to adaptthe pre-existing design to fit the patient’srequirements. This can become particularlychallenging if the required physical dimensionsof the implant differ substantially from those ofthe standard manufactured ones, for example,implants to be used for children.To get around this problem, a novel proces-

sing technique called Laser Engineered NetShaping (LENS) (Sandia National Labora-tories) shows great promise. Similar to rapidprototyping technologies such as stereolithogra-phy, the LENS process (Ref 55, 56) begins witha computer-aided design file of a three-dimen-sional component, which is sliced into a seriesof layers electronically. The information abouteach of these layers is transmitted to themanufacturing assembly. A metal or alloy sub-strate is used as a base for depositing thecomponent. A high-power laser (capable ofdelivering several hundred watts of power) isfocused on the substrate to create a melt poolinto which the powder feedstock is deliveredvia an inert gas flowing through a multinozzleassembly. The powder-feeder system of theLENS system consists of multiple hoppers. Bycontrolling the deposition rates from individualhoppers, it is possible to design composition-ally-graded and, consequently, functionally-graded materials, as demonstrated in a numberof previous papers on laser-processed composi-tionally-graded titanium alloys (Ref 57–59).The nozzle is designed such that the powderstreams converge at the same point on thefocused laser beam. Subsequently, the substrateis moved relative to the laser beam on a com-puter-controlled stage to deposit thin layers ofcontrolled width and thickness. There are fourprimary components of the LENS assembly:the laser system, the powder-delivery system,the controlled-environment glove box, and themotion-control system. A 750 W neodymium:yttrium-aluminum-garnet (Nd:YAG) laser,which produces near-infrared laser radiation ata wavelength of 1.064 mm, is used for all thedepositions. The energy density is in the rangeof 30,000 to 100,000 W/cm2. The oxygencontent in the glove box is maintained below10 ppm during all the depositions. The powderflow rates are typically 2.5 g/min, whilethe argon volumetric flow rate is maintainedat 3 L/min. The LENS offers a unique

Table 3 Applications of various biopolymers

Polymer Application

Polyethylene (PE) Catheters, acetabular cup of hip jointPolypropylene (PP) SuturesPolyvinyl chloride (PVC) Tubing, blood storage bagPolyethylene terephthalate(PET)

Tissue engineering, fabric tubes, cardiovascular implants

Polyethylene glycol (PEG) Drug deliveryPolylactic and polyglycolic acid Tissue engineeringPolytetrafloroethylene (PTFE) Vascular graftsPolymethyl methacrylate(PMMA)

Hard contact lenses, bone cement for orthopaedic implants

Polyacrylamide Swelling suppressantPolyacryl acid Dental cement, drug deliveryPolydimethyl siloxane (PDMS) Heart valves; breast implants; catheters; insulation for pacemaker; ear, chin, and nose

reconstructionCellulose acetate Dialysis membrane, drug deliveryNylon Surgical sutures

Source: Ref 6

10 / Introduction

combination of near-net shape manufacturingand rapid solidification processing that can beparticularly useful for manufacturing orthopae-dic implants. A schematic representation ofthe LENS process is shown in Fig. 1. Fromthe viewpoint of making implants based onmetallic, ceramic, or even hybrid materials,compositional gradation can be particularlybeneficial because it will enable the develop-ment of custom-designed orthopaedic implantswith site-specific properties. Furthermore, engi-neering functional gradation in these implantswill allow for a single unitized component tobe manufactured without any chemically orstructurally abrupt interfaces, leading toenhanced properties and performance.Surface engineering of the near-net shape

laser-processed implant can be carried out usinga number of related processing techniques.Examples of these include the addition of bio-ceramic surface layers via a different type oflaser deposition to improve osseointegration ofthe implant, and the addition of wear-resistantcoatings via sputter deposition or other physicalvapor deposition techniques.

Functionally Graded Hip Implant

Fabricating implants for total joint replace-ment surgeries such as total hip replacement(THR) is rather challenging because the

property requirements at different locations ofthe monolithic implant are quite different. Asdiscussed earlier, the LENS technique withmultiple powder feeders is a viable tool toproduce a functionally graded implant in itsnear-net shape form with site-specific proper-ties. The basic core structure (hollow or similarto the prototype shown in Fig. 2) of the femoralstem and head assembly can be fabricated usingLENS.However, instead of conventional alloys such

as Ti-6Al-4V, the material of choice for thecore of the femoral stem and head assemblycould be based on one of the newer-generationlow-modulus, biocompatible beta-titaniumalloys, such as those based on the Ti-Nb-Zr-Ta system. To achieve an optimal balance ofmechanical properties, both solid geometriesin terms of femoral head and stem, along withinternal cavities, can be processed together.Because the surface of the femoral stem isrequired to exhibit excellent osseointegrationproperties, additional roughness can be intro-duced on the surface of the stem by laserdepositing lines of the same alloy or even bio-compatible coatings such as calcium phosphatein the form of a grid (Ref 48). The pattern ofthese lines/grids can be optimized for achievingthe best potential of osseointegration based ontrial in vitro studies.In contrast, the femoral head material must

possess excellent wear resistance, especially inthe regions that rub off against the internal sur-face of the acetabular cup made of ultrahigh-molecular-weight polyethylene (UHMWPE).Thus, ceramic-based materials (e.g., ZrO2) aregenerally preferred over titanium and its alloys.As mentioned earlier, one drawback of usingceramic materials is that they exhibit poor frac-ture toughness, and thus, the joint betweenthe head and the stem (made of titanium alloy)creates a weak interface between two dissimilar

materials. The tendency for high-impact frac-ture makes these materials fragile. A moreappropriate approach may be to manufacturethe core of the femoral stem and head assemblyin the form of a single monolithic componentand use surface engineering to improve thewear resistance of the base titanium alloylocally in the femoral head section. The LENSprocess could be quite handy in implementingsuch an idea, where the core of the femoralhead could be made of tough b-titanium alloy(such as Ti-Nb-Zr-Ta), and there could bea radial gradation of optimal amounts ofboride (or carbide) precipitates dispersed withinthe matrix. This comes from the idea thathard titanium boride (or titanium carbide) pre-cipitates within the soft beta-titanium matrixcan enhance the wear resistance of thesemetal-matrix-based hybrids quite substantially(Ref 60).The basic parts of the functionally-graded hip

implant that can be manufactured by using theLENS process include:

� Femoral stem: Made of beta-titanium-baseTi-Nb-Zr-Ta alloys

� Femoral head: Made of metal-matrix-basedhybrid materials with radial gradationsfrom Ti-Nb-Zr-Ta- to Ti-Nb-Zr-Ta-rein-forced borides (or carbides)

Subsequent to LENS processing of the femo-ral implant, surface engineering strategies canbe used to enhance the osseointegration of thefemoral stem. For example, laser-based directmelting techniques may be used to simulta-neously synthesize a physically textured surfaceinvolving a substrate (such as Ti-6Al-4V) and acoating (such as calcium phosphate) (Ref 48).Such a process can help in systematic organiza-tion of the calcium-phosphorus coating byeffectively controlling the thermophysical inter-actions. Furthermore a metallurgically-bondedinterface can be obtained by controlling thelaser processing parameters, because both thecoating and substrate material are melted andsolidified at very high cooling rates (�104 to108 K/s). Again, laser-induced surface modifi-cation techniques can be used to increase thewear resistance of the femoral head in hipimplants. Reinforcing the soft matrix of metal-lic components (such as new-generation b-titanium alloys) with hard ceramic precipitatessuch as borides offers the possibility of substan-tially enhancing the wear resistance of thesecomposites (Ref 61). The wear resistance seemsto further improve when lubricious ZnO coatingis sputter deposited on the surface of theseboride-reinforced composites.

Host Response to Biomaterials

The first thing that happens to a living organ-ism after a foreign implant material is intro-duced into the body is its interaction withproteins such as fibrinogen, albumin, losozyme,

Fig. 1 Schematic representation and image of theLaser Engineered Net Shaping (LENS) laser

deposition system

Fig. 2 Schematic diagram illustrating the functionallygraded femoral head and stem of a hip

implant that can be fabricated using the LENS system

Fundamentals of Medical Implant Materials / 11

high- and low-density lipoprotein, and manyothers. These proteins are present in large num-bers within body fluids such as blood, saliva,and tears. Within seconds, the implant surfacebecomes coated with these proteins that, inturn, play a vital role in determining the tis-sue-implant interaction. In fact, the preferentialadsorption of proteins (at much higher concen-trations than the bulk) onto the biomaterial sur-face makes it a biologically recognizablematerial. This not only affects subsequent bloodcoagulation and inflammation, but it is what thecells “see” and respond to. The type of proteinas well as the nature of the biomaterial surfaceis responsible for the aforementioned factors.For proteins to react more readily with the sur-face, they must be larger in size and should beable to unfold at a faster rate. Surfaces, on theother hand, play their part depending on theirtexture, nonuniformity, hydrophobicity, andcomposition (Ref 50). Typically, proteins arebrought onto the surface of the foreign bodyvia diffusion and/or convection. Variables suchas concentration, velocity, and molecular sizeare important factors that determine such amovement. At the surface, the protein mole-cules selectively bind with the substrate at dif-ferent orientations (to minimize repulsiveinteractions) via intramolecular forces such asionic bonding, hydrophobic interactions, andcharge-transfer interactions (Ref 50). In termsof kinetics of adsorption of proteins, the pro-teins initially attach quite rapidly to the largelyopen surface. However, at later stages, it is verydifficult for the arriving proteins to find and fitinto the empty spaces. This causes conforma-tional changes so as to increase their contactpoints with the surface, by molecular spreadingor by increasing the concentration of thesemolecules. Most adsorbed proteins are irrevers-ibly attached to the surface, meaning that oncethey are attached, it is very difficult to detachthem. In fact, desorption of protein moleculeswould require a simultaneous dissociation of allinteractions between the molecule and the sur-face (Ref 50). Nevertheless, at longer times, theadsorbed molecules can eventually exchangewith other competing protein molecules thathave stronger interaction with the surface (Vro-man effect). For example, in blood, which hasmore than 150 proteins, albumin dominates theinitial interaction with the surface, primarilydue to its high concentration and mobility. How-ever, in due course, other proteins such as immu-noglobulin G and fibrinogen, which have muchless mobility but a higher affinity with the sur-face, can exchange with the albumin moleculesand form a stable coating (Ref 50).During implantation of a biomaterial, knowl-

edge of the aforementioned processes is veryimportant, because bleeding and injury of cellsis a part of the wound-healing process. (Whileblood is a mixture of plasma, red blood cells,white blood cells, and platelets, cells jointogether to form tissue, muscle, nerves, andeven the epithelium.) During injury, the endo-thelial cells and collagen fibers are exposed to

blood. Fibrinogen, a protein present in blood,reacts with enzymes such as thrombin to formpolymerized fibrin threads or clots.The cells, on the other hand, constantly adapt

themselves to the changes in environmentsaround them. Due to the presence of implants,the cells can face trauma by way of physical,chemical, or biological agents, causing inflam-mation. If the tissue injury is minimal or if itcan regenerate (e.g., skin tissue forming due toproliferation and differentiation of stem cells),then after complete healing, the normal func-tionalities can be achieved. However, scarringcan occur, causing permanent loss of cell func-tionality if the injury is extensive or the tissuecannot regenerate, for example, heart musclecells. Inflammation is an important part of thewound-healing process following implantation,often resulting in swelling and redness accom-panied by heat and pain. It involves the migra-tion of cells to the injury site via a processcalled chemotaxis, removing the dead cellularand tissue material from the site, destroying orquarantining all the harmful biological andchemical substances, and making sure that tis-sue rebuilding can start. During inflammation,blood flow to the injury site is increased bydilation of blood vessels (vasodilation) alongwith increased vascular permeability. Increasein blood circulation and metabolism rate causesthe region to feel warm. Also, the endotheliumbecomes more adhesive, thus trapping andretaining the leukocytes, present in eosinophil,neutrophil, and basophil, at the trauma site.These cells are responsible for killing theinvading pathogens, such as viruses, bacteria,and fungi, and consuming the foreign objects,such as debris from biomaterials, damaged tis-sue, and dead cells (Ref 50). While basophilsand eosinophils are responsible for releasingharmful chemicals to kill the attacking foreignparasites, the neutrophils help in phagocytosisof foreign particles. Phagocytosis at all levelsinvolves identifying or marking the foreignobjects, surrounding and engulfing them, andfinally releasing harmful chemicals to destroythem. Phagocytosis is also the primary functionof macrophages that form during the phase ofchronic inflammation, leave the blood stream,and attach themselves to tissues at the site ofimplantation (biomaterial surface). These takesome time to form, unlike the neutrophils thatresult from an immediate response to injury(active inflammation). The macrophage cellsusually span close to 50 to 100 mm, can livefrom a couple of months to a year, and are con-sidered the first line of defense. These macro-phages can also fuse together and formmultinucleated foreign body giant cells thatcan phagocytose even larger particles.Along with inflammation, there may be

undesired colonization of tissue by bacteria,fungi, or viruses during biomaterial implanta-tion. This is called infection, which is a seriouscause of concern during surgery. It should benoted that infection can result in inflammation,but the reverse may not be true. One of the

indications of infection is pus formation, whichis a result of neutrophils and macrophages thatdie after killing the foreign parasites. To pre-vent the spread of infection, fibrous tissues usu-ally form around the pus. If these form at thesurface of the skin, as in the case of superficialimmediate infection, the region stretches until itbursts open or is surgically drained (Ref 62).The second type of infection, called deepimmediate infection, is the primary effect ofthe implantation procedure and is caused by air-borne or skin bacteria that are introduced intothe body involuntarily. The biggest threat tosmooth surgery is deep late infection, whichoccurs several months after the procedure.It may be caused by a longer incubation periodof the bacteria or even by slower developmentof the infection.As mentioned previously, tissue injury due to

implantation can affect the morphology, func-tion, and phenotype of the cells (Ref 50). Thechange in environment due to the presence ofimplants can temporarily or permanently alterthe functionality of surrounding tissues, forexample, bone loss due to the stress-shieldingeffect. When blood vessels are damaged, bloodclotting provides essential time for cell migra-tion and proliferation to start the rebuilding pro-cess. This involves inflammation, which is alsoa necessary step toward successful would heal-ing. The cells, promoted by growth factors, syn-thesize extracellular matrix proteins from thepoint of the wound inward (Ref 50). Nextcomes the formation of new blood vessels thatare necessary to aid the newly formed tissues.Due to the macrophages, endothelial cells, andseveral growth factors, the environment at thewound site is made suitable for developmentof endothelial cells to form capillaries. Thenewly formed tissue, called granulation tissue,consists of smaller blood vessels and is verydelicate. At a later stage of wound healing, theremodeling phase begins, where the newlyformed tissue may have the structural and func-tional characteristics of the original tissue. Oth-erwise, remodeling may cause scar tissueformation with reduced functionality.How long does it take for the entire wound

healing to successfully take place? It dependson several factors, from the proliferative capac-ity of cells to the type of tissue and from theseverity of wounds to the health condition andage of the patient (Ref 50). Typically after sur-gery, the clotting of blood occurs almost instan-taneously. The migration of leukocyte cellstakes place within a couple of days. In contrast,the macrophages migrate to the trauma sitewithin a week. The tissue rebuilding and repaircan last for several weeks, with the remodelingphase extending for as much as a couple ofyears (Ref 50).

Implant Failure

From the time an implant is introduced intothe body via invasive surgery, the biomaterial

12 / Introduction

remains in prolonged contact with the cells andtissues. There are four different types of tissueresponses to the biomaterial:

� The material is toxic, and the surroundingtissue dies.

� The material is nontoxic and biologicallyinactive, and a fibrous tissue forms.

� The material is nontoxic and active, and thetissue bonds with it.

� The material is nontoxic and dissolves, andthe surrounding tissue replaces it (Ref 6).

The implant material, its geometry, and theenvironmental conditions around it play a veryimportant role in governing how the proteinsinteract with the foreign substance, and theensuing cell adhesion occurs via receptors inthe cell membrane. The various steps of thephysiological wound-healing process, includingthe activation of neutrophils and macrophages,and the differentiation and proliferation of theadhered cells in turn determine the structureand functionality of the neighboring tissues. Infact, evidence has shown that macrophagesand foreign body giant cells can be present atthe implant-tissue interface throughout theentire period of implantation, with all of thisbeing a part of foreign body reaction (Ref 50).In an ideal situation, one would expect thatthe wound-healing process should be smoothlycompleted and proper function of the tissuesaround the implant should be restored. Also,one would hope that the implant should beintegrated with the body, with no short- orlong-term repercussions (failure of implant,infection, etc.).However, in real life a number of processes

can delay and complicate the wound-healingprocess. Sometimes, the cells around theimplant completely reject it, leading to chronicinflammation, which can lead to removal ofthe biomaterial. This type of situation may arisefor various reasons, such as inappropriate heal-ing (too little or overgrowth of the tissue) andstructural failure or migration of the implant(wear, fracture, stress shielding) (Ref 6). Wear-ing or fracture of implants can cause debris for-mation, which is more often observed inmetallic materials. During dynamic loading/unloading of load-bearing implants (hip, knee,etc.), constant friction between two articulatingparts can lead to the release of small particu-lates that are detrimental to the body. Stressshielding is also a serious problem for load-bearing implants, where unequal distributionof the load between the implant and the bonearound it may lead to reabsorption. Other phys-iological reasons for implant rejection could beleaching out of ions (metals), degradation ofmaterial due to interaction with enzymes (poly-mers), and inadequate encapsulation of theimplant surface via proteins and cells (Ref50). Although introduction of biodegradablepolymers can minimize issues in this aspect,upon completion of their scheduled task, thesematerials degrade into by-products that are

degradable and/or can be removed by existingmetabolic pathways (Ref 50). Another compli-cation of implant surgery is the formation offibrous tissue around the surface. The thicknessand texture of fibrous tissues formed around theimplant can depend on the type of implantmaterial, size and shape of the implant, site ofsurgery in terms of functionality, and type oftissue that needs to be healed. The problemoccurs when such layering prevents the normaloperation of the implant in terms of mechanicalfunction and drug delivery (Ref 50). Theseissues are further complicated if superficialand deep infections occur due to the coloniza-tion of bacteria, fungi, and viruses around theimplant. Medications may not work due to thepresence of an impervious fibrous layer, andthe removal of the implant may be the onlyoption.To prevent such failures, several precautions

are usually taken before using an implant for aprosthesis. The first step is to evaluate the mate-rial itself, because it is known to be the mostcommon reason for implant failure. Unsuitablematerials used for a prosthesis would mean thatits physical, chemical, and biological propertiesare not suitable or compatible for the specificimplant application. The design of the deviceis the next critical item. Information from pastfailures as well as upcoming experimental andmodeling results must be incorporated to arriveat better designs. For example, while designingorthopaedic load-bearing implants, data fromfinite-element analyses of stress-concentrationpoints as well as stress-strain results from walksimulators are taken into account. During fabri-cation, the implant should be free of any defectsand inclusions that may lead to implant failure.In some cases, the sterilization process itselfmay cause changes in the structure and propertyof the prosthetic device. On the other hand,incomplete sterilization can also lead to infec-tion, as with improper packaging and shipping.After the implant is fabricated, both mechanical(tensile, wear, fatigue) and biological (in vitro,in vivo) testing must be conducted to determineits feasibility. While the mechanical testing pro-cesses have been discussed in the section“Development of Implant Materials” in thisarticle, a brief discussion of the biological testsfollows.InVitroAssessment of Tissue Compatibility.

This usually involves performing cell culturesfor a wide variety of materials used in medicaldevices. Three different cell culture assays areused for in vitro study: direct contact, agar dif-fusion, and elution. In all the tests, experimen-tal variables such as cell type (usually L-929mouse fibroblast), number of cells, duration ofexposure, and test sample size are kept constant(Ref 6). Positive and negative controls are oftenused during the assay test to determine the via-bility of the test. In all cases, the amount ofaffected or dead cells in each assay provides ameasure of the cytotoxicity and biocompatibil-ity of the biomaterials. In the direct contact test,the material is placed directly on the cell

culture medium. After the test, the cells arestained by hematoxylin blue, and the toxicityis evaluated by the absence of stained cells,because dead cells do not stain. The main prob-lem with this type of test is cell trauma anddeath due to movement of the sample or theweight of highly dense materials. This is over-come by the agar diffusion test, where the agar(colloidal polymer from red algae) forms alayer between the test sample and the cells. Inthis assay, the healthy cells are stained red ascompared to dead or affected cells. The mainproblem with this type of test is the risk of thesample absorbing water from the agar, thuscausing dehydration of the cells. The third typeof test, elution, is conducted in two separatesteps: extraction of fluid (0.9% NaCl orserum-free cultural medium) that is in contactwith the biomaterial, and biological testing ofthe medium with cells. Although this type oftesting is time-consuming, it is very effective.It is universally observed and accepted thatmaterials found to be nontoxic in vitro are non-toxic in in vivo assays as well (Ref 6).In Vivo Assessment of Tissue Compatibility.

This type of test is conducted to determine thebiocompatibility of a prosthetic device and alsoto assess whether the device is performingaccording to expectations without causing harmto the patient. It provides valuable data aboutthe initial tissue response to the biomaterial,which in turn helps in selection and design ofthe device. Some tests, such as toxicity, carci-nogenicity, sensitization, and irritation, deter-mine if the leachable products of the medicaldevice affect the tissues near or far from theimplant site. Other tests, such as implantationand biodegradation, study the postsurgerychanges in the implant material itself and theirensuing effect on the body. Overall, there maybe an array of tests that must be conductedand evaluated, depending on where and why aspecific device is used, before certifying animplant. For conducting the actual in vivo tests,animal models (sheep, pig, rat) are usuallyselected after weighing the advantages and dis-advantages for human clinical applications.As evident from the previous discussions, a

variety of tests can be conducted before a pros-thetic device is considered suitable for implan-tation. The choice of test depends on thespecific application of the implant under con-sideration. Sometimes, it is very difficult to rep-licate the exact test, even while performing invivo tests. For example, there is no adequateanimal model to study the inflammatory reac-tion to wear debris near hip joints (Ref 6).After successful completion of all these

steps, the implant is finally ready to be surgi-cally inserted into a patient. The prostheticdevice is chosen based on the patient and thesite of implantation. Each patient is different.He/she may have different allergic reactionsto implant materials, may have previoushealth conditions unsuitable for the prosthesis,or may even have different immunologicalresponses to fighting infection. Failures can

Fundamentals of Medical Implant Materials / 13

also be caused by the misuse of implants. Forexample, a patient who has undergone totalhip replacement surgery can cause severe dam-age and loosening of the implant by excessiveexercising before proper healing takes place(Ref 6). All of these factors can single-handedlyor jointly contribute to improper functioningand eventual failure of implant devices.The implant structure should be carefully

analyzed postfailure to determine the exactcause and mechanism by which it failed. Thiscan lead to the improvement of processing tech-niques and materials used for fabricatingthe device, can help in bettering the designand testing mechanisms used for these pro-ducts, and can provide enough insight intoadopting alternate surgery procedures and drugtherapy postimplantation (Ref 6). Thus, implantretrieval and evaluation is a vital study todetermine the safety and biocompatibility ofimplants. Along with the implant material,examination of the tissue must be conductedto assess the implant-tissue interface. At first,the overall implant and tissue specimen can beanalyzed by light microscopy and cell culture,respectively. Consequently, specific aspects ofthe material can be studied using techniquessuch as scanning electron microscopy, trans-mission electron microscopy, energy-dispersivespectroscopy, contact-angle measurement,Fourier transform infrared spectroscopy, andscanning ion mass spectroscopy. Similarly,studies of proteins and genes can be conductedon the tissue sample. Compilation of these datacan aid in the development of next-generationimplants. A good example of the benefits ofimplant evaluation is the modern use ofUHMWPE as a polymeric cup instead of syn-thetic fluorine-containing resin, with whichsome biological problems were encountered.In the case of dental implants, the integrationof bone with the metal was far better under-stood after evaluation of failed fixtures.

Summary

According to D.F. Williams in 1987, biocom-patibility can be defined as the capability of amedical device (implant) to perform with anappropriate host response in a specific applica-tion (Ref 63). Various parts of the device maybe individually assessed, or every part maybe considered separately. While the former isthe biocompatibility of the device, the latteris the biocompatibility or bioresponse of indi-vidual materials (Ref 6). Either way, it isimportant to note that no material is suitablefor all biomaterial applications. Nevertheless,implant science has been developing new tech-nologies for implant devices as well as improv-ing cell tissue interactions with biomaterials,some of which is discussed next.Over the years, the use of polymers as bioma-

terials has increased considerably. Scientists aretaking a very active interest in the developmentof newer stimulus-responsive smart polymers.

In the presence of various physical, chemical,and biological stimuli, the polymers exhibit dif-ferent responses, such as gelation, surfaceadsorption, and collapse of hydrogel. All of theseresponses are reversible processes, and, in theabsence of the stimuli, the response is alsoreversed (Ref 6). These smart polymers are evencombined with biomolecules to enhance theiruse beyond implant devices. These types of poly-mers, in combination with proteins and drugs,can be used in solution, in surfaces, and as hydro-gels for applications such as drug delivery,removal of toxins, and enzyme processes (Ref6). To improve the osseointegration properties,the polymer matrices are filled with HAs. Thesematerials contain approximately 50 vol% HA ina polyethylene matrix and are used to makeimplants for ears (Ref 64).Similarly, the osseointegration properties of

ceramics andmetals can be improved by introdu-cing porosity on those surfaces that are in directcontact with the bone. The growth of bones intothese pores would also ensure good mechanicalstability of both load-bearing and non-load-bear-ing implants. Research has shown that optimalengineered porosity, fabricated by laser-deposi-tion processes such as LENS with pore sizes inthe range of 100 to 150 mm, can promote thegrowth of osteoblast cells within these pores(Ref 65). Figure 3 shows one such study, whereTi64 samples containing different sizes of engi-neered porosity were fabricated by LENS. Theimplant in this case acts as a scaffold for boneformation. For smaller pores, the fibrous tissueoccupies the void space, because an extensivecapillary network for osteogenesis does notoccur (Ref 50). The degree of porosity of thesematerials has a big impact on bone integrationand modulus, with substantial reductions inmodulus with increasing porosity. However,with increase in porosity, the strength of theimplant material reduces drastically, whichcould be a big challenge for load-bearingimplants. Other surface-modification techniquesare currently being studied to promote bonegrowth, including increasing the surface rough-ness of the device, using nanograined materials(Ref 66) to increase the surface area, and coatingthe implant with bioactive materials such as

calcium phosphate (Fig. 4) (Ref 48, 67). Laser-engineered textured materials can also promotedirectional growth and movement of cells. Otherphysicochemical methods are also being used tochange the surface composition as well as thebiochemical properties of the surfaces. The latterapproach uses the organic components of bone toaffect tissue behavior by introducing peptidesand proteins. Many bone growth factors couldbe used to influence the growth and differentia-tion of osteoblasts (Ref 50). Even tissue engi-neering approaches have been used to stimulateprecise reactions with proteins and cells at amolecular level. As mentioned previously, thecell and tissue response to implantation is greatlydependent on what they “see” on the surface ofthe foreign device. Even the interaction of cellswith the extracellular matrix (ECM) dependson the rigidity of the substrate. Adhesive sur-faces created by targeted use of proteins, pep-tides, and other biomolecules help inmimicking the ECM environment (Ref 6). Evenchemically patterned surfaces, aided by techni-ques such as photolithography, could be used tocontrol cell adhesion at certain specific regions.In the future, several of the aforementioned sur-face-modification techniques could be combinedto design devices with chemically engineeredsurfaces and controlled scaffold architecture thatcould manipulate specific cell growth, which inturn develops into specialized tissues.Using this same concept has led to the design

of multifunctional devices. Their application isthought to be a combination of a variety offunctions that require the design of materialswith specific properties. For example, thedevelopment of biodegradable bone nail canprovide mechanical support to fracture sites aswell as ensure growth on new bone at implantsites (Ref 6). As mentioned previously, biode-gradable materials are used more and more forthis purpose, and the use of other functionalcombinations involving tissue-engineering scaf-folds is also being actively considered. In thefuture, these scaffolds could provide structuralstability as well as serve as a means for drugdelivery (Ref 6).Similar to cell-biomaterial interaction, another

cause of concern is the blood- biomaterial

Fig. 3 Scanning electron microscopy backscattered electron images of Ti64 samples that were deposited via theLENS system using hatch widths of (a) 0.89 mm (0.035 in.), (b) 1.5 mm (0.06 in.), and (c) 2.0 mm (0.08

in.). The images show three different-sized scales of engineered porosities, resulting in different elastic moduli andcell responses.

14 / Introduction

interaction for implants such as vascular graftsand heart valves. Ongoing efforts have beendirected toward producing blood-compatible bio-materials that have properties similar to the endo-thelium (Ref 50). Usually, these materials arehydrophilic in nature, which reduces plateletadhesion and coagulation. Even anticoagulantsapplied along blood-contacting surfaces orincorporated in the chemical structure of poly-mers have shown promise in terms of reductionof thrombus formation.Another major concern about introducing a

foreign material into the body, apart from thebody’s normal foreign body reaction, is theundesired colonization of bacteria, viruses, orfungi, causing short- and long-term infections.Many researchers have been working to designbiomaterials that discourage germ adhesion andgrowth. Some biomaterials have been designedto release antibiotics via diffusion or dissolutionof material (Ref 50). However, it is difficult topredetermine the dose without prior knowledgeof the type and extent of germs that can affectthe implantation site. In addition, there are vari-ous concerns, such as patients becoming sensi-tive to the antibiotic dose and germs mutatingto develop antibiotic-resistant strains. Thus it is

clear that more research needs to be conductedbefore successfully implementing this concept.The design and development of new types of

implant devices and their targeted applicationwill also dictate newer test protocols to analyzeand evaluate these biomaterials. With newerdevices such as smart polymers and bioactiveglasses, more studies have been focusedon active tissue-biomaterial interactions. Invivo testing and assessment of the targetedbiological response of a tissue-engineereddevice would, in turn, provide pivotal informa-tion toward research and development of thedevice (Ref 6). The ultimate goal would be touse these devices universally with minimuminflammatory and reactive response from thepatient, quick healing of tissue (with no fibroustissue formation), and successful integration ofthe device within the body, with desired perfor-mance and no long-term repercussions.

ACKNOWLEDGMENTS

The authors take this opportunity to expresstheir gratitude to a number of people withoutwhose support this work would not have

been possible. First and foremost, the authorsgratefully acknowledge the support, encourage-ment, and, guidance from Professor HamishFraser at The Ohio State University for hiswork on laser deposition of orthopaedic bioma-terials. The authors would also like to acknowl-edge Dr. Narendra Dahotre from the Universityof North Texas (UNT) for his guidanceand support in the laser-deposited calciumphosphate work. The authors would also liketo acknowledge the excellent work by graduatestudents, both past and present, working on thisprogram over a period of time and at differentinstitutions. These include Peeyush Nandwanaand Antariksh Singh at UNT.

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