Materials for joint replacement

4

Materials for joint replacement

K S K A T T I , D V E R M A and D R K A T T I , North Dakota State University, USA

4.1

Introduction

The materials and devices used in orthopedic applications are designed to sustain the load bearing function of human bones for the duration of the patient's life. Orthopedic applications include numerous products for the rehabilitation and reconstruction required as a result of various diseases of the musculoskeletal system as well as aging. The worldwide market for materials used for orthopedic applications is estimated to be $14 billion in 2002. Also, about $12 billion is spent on joint replacements (Hallab et al., 2004). This chapter provides an overview of the current and future materials used for joint replacement. Key physical and mechanical properties are discussed in addition to mechanics, degradation, and biocompatibility issues associated with specific materials. New advances in the use of novel nanocomposite systems and natural materials are also discussed.

4.2

Materials criteria for total joint replacement

The primary function of orthopedic materials is to bear load and provide structural integrity to the human body. Table 4.1 shows the mechanical properties of bones. Structural integrity implies a combination of fracture toughness, strength, ductility, and hardness, and also time-dependent properties and fatigue resistance. In addition, the human body provides a fairly corrosive environment (Ratner et al., 2004) and thus biocompatibility and corrosion resistance are also important requirements of these materials. In addition, the deterioration products of the orthopedic materials such as from a joint replacement implant should not adversely affect the bodily environment. Thus, the combination of biocompatibility requirements over and above the mechanical properties expected under the corrosive environment of the human body makes for very stringent requirements on the materials design of orthopedic materials. Wear resistance and corrosion resistance properties of various biomaterials are shown in Tables 4.2 and 4.3.

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Joint replacement technology Table 4.1 Mechanical properties of bones (adapted from Black, 1992; Currey, 1984; Fratzl et al., 1998; Katti, 2004) Tissue

Tibia Femur Radius Humerus Cervical Lumbar

Compressive strength (MPa)

Tensile strength (MPa)

Elastic modulus (GPa)

159 167 114 132 10 5

140 121 149 130 3.1 3.7

18.1 17.2 18.6 17.2 0.23 0.16

Table 4.2 Wear rate of materials used in orthopedics Materials

Wear rate (mm3/million cycles)

UHMWPE/zirconia (n ˆ 3) Cobalt chrome/cobalt chrome (n ˆ 3) Alumina/alumina (n ˆ 3) Alumina/UHMPE Alumina/crosslinked UHMWPE CoCrMo/CoCr/Mo

References

31  4.0 1.23  0.5

Tipper et al. (2001) Tipper et al. (2001)

0.05  0.02 51  11 5.62  3.5

Tipper et al. (2001) Essner et al. (2005) Essner et al. (2005)

6.30  10.3

Essner et al. (2005)

* UHMWPE ˆ ultra-high moleculat weight polyethylene.

Table 4.3 Electrochemical properties of implant metals in 0.1 M NaCl at pH 7. Higher corrosion potential, lower passive current density, and higher breakdown voltage represent better corrosion resistance (adapted from Ratner et al., 2004) Alloy

Stainless steel CoCrMo CPTi Ti±6Al±4V Ti±5Al±2.5Fe Ni±45Ti

Corrosion potential (mV)

Passive current density (mA/cm2)

Breakdown potential (mV)

ÿ400 ÿ390 ÿ90 to ÿ630 ÿ180 to ÿ510 ÿ530 ÿ430

0.56 1.36 0.72±9.0 0.9±2.0 0.68 0.44

200±770 420 >2000 >1500 >1500 890

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4.3

83

History of materials used in joint replacement

One of the earliest treatments to address painful hip joints consisted of simply removing the acetabular and femoral diseased bone. These procedures were attempted as early as the 1820s. During the 1830s to 1880s, wooden blocks and animal soft tissues were also placed between the acetabular and femoral components in order to ease pain in the hip joints. A prosthetic replacement for hip joints was first attempted in 1890 using a carved ivory structure to replace the femoral head with the use of plaster of Paris and pumice-based bone cement. Prosthetic replacement of femoral heads using ivory and rubber was extensively popularized in the late 1800s and early 1900s. Using various materials such as wood, gold foil, and animal soft tissues as an interpositional membrane continued until the early 1900s. These procedures were not very successful in relieving pain, and a quest for a solution to the replacement of joints continued. One of the earliest attempts at replacement of hip joints with synthetic materials was only as recent as 1925 wherein mold arthroplasty was attempted. This method used a molded piece of glass in the shape of a hollow hemisphere that fitted over the ball of the hip joint. This attempt was made by Dr SmithPeterson, a surgeon at Massachusetts General Hospital in Boston. The primary reason for the failure of this device was the poor mechanical performance of glass. Several attempts at using other materials, with similar biocompatibility properties as glass but superior mechanical properties such as stainless steel, were fabricated into hollow hemispheres. In the quest for better materials that were more suitable, the next breakthrough in biomaterials was in 1936 with the fabrication of cobalt±chromium alloys. Many attempts were made in using these new alloys in mold arthroplasty but these devices did not adequately satisfy the need to cure a variety of painful deformities of the hip resulting from arthritis and other conditions. The next major breakthrough was the type of hip replacement called hemiarthroplasty which consisted of replacing the entire ball of the hip but not the socket. This procedure of hip replacement consisted of a long metal stem placed in the femur connected to a metal ball that sat in the hip socket. This was the state of the art in the 1950s. This procedure often resulted in loosening of the implant. New bone cement fixation techniques were also pursued around the same time. One of the pioneers of total hip replacement was the surgeon Dr John Charnley who first attempted replacement of the diseased hip socket. He used Teflon and polyethylene. In the late 1950s Dr Charnley performed many successful hip replacement surgeries, which resulted in his eventual knighting by Queen Elizabeth II. He used a steel femoral component and a plastic socket cup. The use of replacement of both the hip socket and femoral heads has since been extensively popularized and a variety of materials such as polymers, metals, composites, and biological materials are being used by surgeons and many more studied by researchers in the quest for a replacement of

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a hip joint. The primary advantage of this method is the ability to design anatomically sized femoral heads, stems, and sockets. Attempts to replace the knee joint were also being made concurrently with those of the hip joint. The earliest attempts consisted of hinges that were fixed to the bones through the hollow bone marrow cavity. In the 1950s metal spacers were placed between the knee bones to prevent the rubbing of bone against bone. Later in the 1960s, Frank Gunston, an orthopedist at the Sir John Charnley Hip Center, designed a metal on polymer knee joint that was attached using a bone cement. But the first total knee replacement was attempted by Dr John Insall in New York in 1972, which consisted of replacement for surfaces of all three surfaces of the knee, the femur, tibia, and kneecap. This method remains a prototype for current knee replacement methodologies.

4.4

Traditional materials

4.4.1

Metals

Metallic implants are the primary biomaterials used for joint replacement and becoming increasingly important. The metallic implants used for orthopedic applications can be categorized as stainless steel, CoCr alloys, and Ti and Ti alloys. These metallic materials have several properties such as high strength, high fracture toughness, hardness, corrosion resistance and biocompatibility, which make them an excellent choice for total joint replacement. The disadvantage with metallic implants is their high elastic modulus, which causes stress shielding. Toxic effects caused by ions released from metallic implants are also a major concern. Stainless steel alloys were the first metals to be used for orthopedics. Stainless steel alloys contain carbon, chromium, nickel, molybdenum, and manganese, phosphorus, sulfur, and silicon as trace elements. These components affect the mechanical properties of steel by alteration of its microstructure. A high nickel content (10±14%) in stainless steel can cause toxicity. This has prompted research in the development of Ni-free stainless steel alloys. Cobalt-based alloys are the other metallic implants used for joint replacement. CoCrMo and CoNiCrMo are the two main cobalt based alloys generally used in orthopedics. Especially for joint replacement where low frictional resistance is desired, CoCrMo alloys are preferred over CoNiCrMo alloys. Commercially pure Ti (CPTi) and Ti±6Al±4V are two dominant Ti-based materials used in joint replacement. Ti-based implants have excellent corrosion resistance and biocompatibility. And the credit goes to the oxide layer, which spontaneously forms in the presence of oxygen. It has been shown that Ti-based alloys promote osteoblast activities. Therefore for uncemented joint replacement, Ti-based alloys are preferred over other metallic implants. One drawback with Ti alloys is that they are relatively softer than stainless steel alloys and cobalt-based alloys. This makes them more susceptible to wear where articulation is required.

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Zirconium is classified as a refractory metal because of its high temperature resistance properties. It also provides high chemical resistance because of the formation of a highly stable oxide layer, which is about 5 m in thickness. Several zirconium-based alloys have been developed for use in joint prostheses. One such alloy is Oxinium. Oxinium possesses excellent wear resistance (approximately 10 times that of Co and Ti-based alloys). Also, Zr-based implants have the advantage of both ceramic and metallic implants. Their metallic core provides high fracture toughness and the oxide layer provides excellent wear resistance and biocompatibility. The only disadvantage associated with these alloys is the high cost of forming and machining. Implants interact with the body through their surfaces. Further, wear and corrosion are initiated at the surface. Surface characteristics of an implant decide its fate in the body. Therefore, for proper surface response, several strategies have been proposed recently (Lappalainen and Santavirta, 2005). Surface response depends on both surface topography and chemical composition. Surface topography can be modified by grit or sand blasting or plasma treatments. Large varieties of coatings have been used, such as hydroxyapatite (Capello et al., 1998), titanium oxide and nitride (Teresa Raimondi and Pietrabissa, 2000), zirconium oxide (Patel and Spector, 1997) and diamond-like carbon coatings (Affatato et al., 2000; Lappalainen et al., 1998) to improve surface characteristics. Several studies have also discussed coating implants with growth factor (Cole et al., 1997; Lind et al., 2000), collagen (Roehlecke et al., 2001), RGD peptides (De Giglio et al., 2000), and fibronectin (Degasne et al., 1999). Moreover, implants are also being coated with osteoblast cells (Frosch et al., 2003). Table 4.4 shows the mechanical properties of various metals and alloys used in orthopedic applications.

4.4.2

Ceramics

Ceramic materials possess several useful properties, which make them excellent materials for orthopedic implants. They exhibit high stiffness, inert behavior under physiological environment, and superior wear resistance as compared with metallic and polymeric bearing surfaces. One limiting property of ceramic materials is their brittleness. Since the mechanical properties of ceramic materials are highly dependent on their density, small voids left in the implant during processing severely affect their longevity. Alumina was the first ceramic material used for joint replacement (Boutin, 1972). Table 4.5 shows the mechanical properties of various ceramics used in orthopedic applications. The wear rate of alumina is reported to be 20 times lower than that of ultra-high molecular weight polyethylene (UHMWPE). Femoral heads for hip replacements and wear plates in knee replacements have been fabricated using alumina. One of the concerns with alumina implants was its low fracture toughness, which was overcome later by increasing purity, lowering porosity, grain size and

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Joint replacement technology Table 4.4 Mechanical properties of alloys in total joint replacement (Davidson and Georgette, 1987; Davis, 2003; Katti, 2004; Long and Rack, 1998) Material

Ti±Zr Ti±6Al±4V Ti±6Al±7Nb Ti±5Al±2.5Fe Ti±3Al±2.5V Ti±13Nb±13Zr Ti±15Mo±5Zr±3Al Ti±12Mo±6Zr±2Fe Ti±15Mo±2.8Nb±3Al Ti±35Nb±5Ta±7Zr (TNZT) Ti±15Mo±2.8Nb±0.2Si±0.3O Ti±35Nb±5Ta±7Zr±0.4O Ti±15Mo Ti±16Nb±10Hf CPTi (>>98% Ti) Co±Cr±Mo Co±Cr alloys Stainless steel 316L

Tensile strength (MPa)

Modulus (MPa)

900 960±970 1024 1033 690 1030 882±975 1060±1100 812 590 1020 1010 795 851 785 600±1795 655±1896 465±950

± 110 105 110 100 79 75 74±85 82 55 82 66 78 81 105 200±230 210±253 200

improving manufacturing techniques (Boehler et al., 2000; Fritsch and Gleitz, 1996). In hip replacements, alumina is also used as the femoral head with a metallic femoral stem and UHMWPE as an acetabular cup opposing articulating surface. In February 2003, the United States Food and Drug Administration Table 4.5 Mechanical properties of ceramic materials used in orthopedics (Davis, 2003; Katti, 2004; Ramakrishna et al., 2001; Schmitt, 1985) Ceramic

Zirconia Alumina Bioglass C (graphite) C (vitreous) C (low-temperature isotropic carbon (LTI) pyrolytic) C (silicon alloyed LTI) C ultra-low temperature isotropic carbon (ULTI) Hydroxyapatite Apatite-Wollastonite (AW) glass ceramic

Compressive strength (MPa)

Modulus (GPa)

2000 4000 1000 138 172 900

220 380 75 25 31 28

± ±

28±41 14±21

600 1080

117 118

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(FDA) approved the first ceramic-on-ceramic articulated hip implant for marketing in the United States. Later, several studies focused on other materials as an alternate for alumina. The first paper on zirconia as a potential material for application in orthopedics was published in 1969 (Helmer and Driskell, 1969) and the first publication on design of zirconia ball heads for total hip replacement was reported in 1988 (Christel et al., 1988). Because of its better mechanical properties over alumina, zirconia has attracted considerable research interest. Zirconia femoral heads have been gaining market share, and more than 300 000 zirconia femoral heads have already been implanted (Chevalier et al., 1997; Hench and Wilson, 1993). One problem with zirconia implants is low-temperature degradation (Yoshimura et al., 1987). Zirconia ceramics are polycrystals of tetragonal phase, stabilized by yttria, and commonly referred to as yttria-stabilized±tetragonal-zirconia polycrystals (Y-TZP). The Y-TZP material slowly undergoes phase transformation to monoclinic form at room temperature, which is accompanied by deterioration in its mechanical properties. Composites of ceria-stabilized± zirconia-tetragonal polycrystal (C-TZP) with alumina polycrystals have shown improved resistance to low-temperature degradation (Tanaka et al., 2002, 2003). However, this composite showed no improvement in bone bonding ability. Recent studies have shown that bioactivity of these composites can be improved by surface chemical treatments (Takemoto et al., 2005; Uchida et al., 2002). Zirconia-toughened alumina prostheses have also shown superior properties over currently used alumina implants for hip replacement (Insley and Streicher, 2004; Wang and Stevens, 1989). Also, calcium phosphate materials and bioglass ceramics are being investigated as alternatives to poly(methylmethacrylate) (PMMA) for bone cement applications. Their osteophilic characteristics make them excellent candidates for orthopedic applications. A better match between the bulk material properties of the implant and the bone it replaces can decrease some of the problems such as stress shielding currently associated with metallic implants. This is often achieved by coating the metallic implants with bioactive materials such as hydroxyapatite (HAP), tricalcium phosphate (TCP), and bioglass. Tricalcium phosphate (TCP) (Ca3(PO4)2) and HAP (Ca10(PO4)6(OH)2) are both biocompatible materials and have the ability to bond directly to bone. Several researchers have attempted to develop high-strength consolidated HAP bodies (Bagambisa et al., 1993; Hench and Wilson, 1993). Bending strength as high as 90 MPa has been achieved by colloidal processing of HAP (Hench and Wilson, 1993). However, these materials still suffer from poor mechanical properties such as low strength and limited fatigue resistance, which limit their application as load bearing biomaterials. Mechanical properties of ceramic biomaterials are shown in Table 4.5. Several methods have been developed to coat implants with hydroxyapatite and other calcium phosphate materials. Among them, thermal spraying has produced most promising results

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(Chu, 2007; Heimann, 2006). Simply soaking an implant in simulated body fluid to coat it with apatite has also been investigated (Li, 2003). Initially, the formation of amorphous calcium phosphate layer at the interface of implant was causing coating failure (Ji et al., 1992; Park and Condrate, 1999; Park et al., 1998). However, using a `bond coat' on implant surface or annealing the implant after coating beyond 900 ëC improved adhesion of coating significantly (De Groot et al., 1987; Gross et al., 1998). Improvement in adhesion due to annealing occurs with the formation of a several micrometer thick layer of Ca± Ti±oxide at the interface. The important parameters that control the success of a coating are composition and crystallinity (Fazan and Marquis, 2000). Crystallinity also affects the dissolution of the apatite. Although dissolution of apatite is necessary for bonding of coating with the surrounding bone tissues, dissolution in excessive amounts may cause inflammation due to changes in local pH (Chou et al., 1999; LeGeros et al., 1991).

4.4.3

Polymeric materials

Several polymers have been used for orthopedic applications such as acrylic, nylon, silicone, polyurethane, UHMWPE, and polypropylene (PP) (Davidson and Georgette, 1987). Mechanical properties of these polymers are shown in Table 4.6. UHMWPE is one of the most preferred polymers as an orthopedic implant because of its high mechanical strength, low wear rate, and biocompatibility (Costa and Brach del Prever, 2000; Kelly, 2002). Although UHMWPE has been used for over 30 years, osteolysis caused by wear debris is still a concern (Goldring et al., 1983; Sinha et al., 1998; Willert and Semlitsch, 1977). Several studies have been conducted to understand the wear mechanism and also the osteolysis caused by the wear debris (Ingham and Fisher, 2005; Ren et al., 2006; von Knoch et al., 2005; Wedemeyer et al., 2007). Osteolysis is the Table 4.6 Mechanical properties of polymeric materials in orthopedics (Katti, 2004; Ramakrishna et al., 2001; Schmitt, 1985) Polymer

Tensile strength (MPa)

UHMWPE Polyacetal Polysulfone Polyurethane Silicone tubber Polyetheretherketones (PEEK) Polytetrafluoroethane (PTFE) Polyethylene terephthalate (PET) Poly(methylmethacrylate) (PMMA)

21 67 75 35 7.6 139 28 61 21

Modulus (GPa) 1 2.1 2.67 0.02 0.008 8.3 0.4 2.85 4.5

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resorption of bone surrounding the implant, which occurs in association with the formation of vascularized granuloma at the interface of implant and the bone (Granchi et al., 2005). The formation of granuloma is a body response to clean up the wear particles. The response to wear particles is dealt with in detail by Revell in Chapter 15. Since the main problems associated with the use of UHMWPE as acetabular cups is not the degradation in mechanical properties of cups but weakening of the interfacial adhesion between tissue and implant (due to wear debris) considerable efforts have been made to improve the wear resistance of UHMWPE (Affatato et al., 2005; Bell et al., 2001; McEwen et al., 2005). It has been observed that increasing crystallinity and crosslinking density improve wear resistance of UHMWPE (Endo et al., 2001). The crosslinking of UHMWPE is usually achieved by exposing implant to irradiation. Although, crosslinking improves wear resistance, but at the same time also degrade tensile strength, fracture toughness and fatigue crack propagation resistance (Baker et al., 2003; Gomoll et al., 2002). Increasing crystallinity of UHMWPE also improves its wear resistance, elastic modulus and resistance to crack propagation (Champion et al., 1994). Recently, it has been suggested that crosslinking of UHMWPE in combination with higher crystallization can improve wear resistance and fatigue fracture resistance (Simis et al., 2006). One other reason for failure of UHMWPE in implants is its oxidation during the sterilization process (Fisher et al., 2004; Premnath et al., 1996). Oxidation of UHMWPE is minimized by sterilizing the implant in an inert atmosphere and adopting other sterilization procedures such as gas plasma and ethylene oxide (Kurtz et al., 1999), but exposure to high-intensity radiation causes formation of free radicals in the crystalline phase of UHMWPE. These free radicals react with dissolved oxygen and cause oxidative embrittlement and subsequently hamper the mechanical properties of the implant. To prevent oxidative embrittlement of UHMWPE, addition of vitamin E is suggested as an antioxidant (Parth et al., 2002; Reno and Cannas, 2006; Tomita et al., 1999). Polyurethanes are another class of polymers that has been considered for joint replacement materials. Recent studies have shown that polyurethane provides a lower coefficient of friction than UHMWPE bearings (Quigley et al., 2002). There are different types of polyurethane and they are identified from the type of linkages they have, such as polyesterurethanes (Coury et al., 1984; Mandarino and Salvatore, 1960) which incorporate ester linkages, polyetherurethanes (PEUs) (Lamba et al., 1998; Zdrahala, 1996) which incorporate ether moieties, and polycarbonateurethanes (PCUs) which incorporate carbonate linkages (Hoffman et al., 1993; Lemm, 1984). Polyesterethane and polyetherurethane are prone to hydrolytic degradation. So for biomedical applications such as joint replacement where long-term stability is required, polycarbourethanes are being investigated (Gunatillake et al., 2003; Khan et al., 2005). Other materials which have been studied as an artificial cartilage are water swollen hydrogels (Oka et

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al., 2000), poly(vinyl alcohol) cryogel (PVA-c) (Szkowski et al., 2006), hyaluronan esters (Zhang and James, 2005) and multilayer polyelectrolyte films (Pavoor et al., 2006). Multilayer polyelectrolyte films showed 33% reduction in wear compared with UHMWPE bearings.

4.5

Bone cement materials

The primary functions of bone cement are to secure the orthopedic implants to bone and transfer mechanical loads from the implant to the bone. The femoral stem and acetabular cups are cemented, screwed or press fit into place. Approximately 50% of all orthopedic implants utilize bone cement to achieve implant fixation. PMMA is the most commonly used bone cement. PMMAbased bone cements are mainly two-part formulations. The first part contains pre-polymerized PMMA, an initiator and a radiopacifier. The second part contains mostly liquid MMA, an accelerator and a stabilizer to prevent premature polymerization. These bone cements have shown high success rate, averaging 90% after 15 years (Murray et al., 1995; Nafei et al., 1996). Seven main drawbacks associated with PMMA-based bone cements have been identified (Lewis, 1997), as follows: 1. Local tissue damage, which occurs due to the exothermic nature of the cement setting reaction (Liu et al., 1987). The temperature goes up as high as from 60 to 120 ëC depending on formulation of the cement (Kindt-Larsen et al., 1995; Wang et al., 1995). 2. The release of the unreacted MMA, which causes chemical necrosis of the bone (Kindt-Larsen et al., 1995). 3. The high shrinkage of the cement after polymerization which is about 21% (Thompson et al., 1979). 4. The stiffness mismatch between bone and the cement. 5. The cement does not bond chemically with either of bone and implant and acts in such as way as to bring about `weak link zones' (Bragdon et al., 1995; Harrigan et al., 1992). 6. Cement particle mediated osteolysis of the bone. 7. Bacterial infection is also associated with bone cements. Several studies have focused on solving the problems associated with PMMA-based cements outlined above. Partial replacement of MMA with 2,2bis [4(2-hydroxy-3-methacryloxypropoxy) phenyl] propane caused significant improvement in volume shrinkage (Vallo and Schroeder, 2005). Efforts have also been made to improve the mechanical and biological properties of PMMAbased cements. Several studies have investigated the effect of additives such as carbon (Friis et al., 1996), graphite (Knoell et al., 1975), aramid (Pourdeyhimi et al., 1986), titanium (Timmie Topoleski et al., 1992), UHMWPE (Gilbert et al., 1994; Pourdeyhimi and Wagner, 1989). To improve the bioactivity of PMMA-

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based cements, additives such as hydroxyapatite have also been included (Kim et al., 2004; Serbetci et al., 2004; Vallo et al., 1999). Bacterial infection issues have been addressed by incorporation of antibiotic within the bone cements. Some studies also suggested use of combination of antibiotics. Bone cements are also dealt with in detail in Chapter 9.

4.6

Composite materials and new nanocomposite systems

In spite of tremendous success achieved by currently used bone implants for joint replacement, there is still need for development of materials that are more biocompatible and last longer in the body. Ceramic polymer composites have shown more superior properties than either of their components used as total hip replacement materials (Wang et al., 1996). Specifically, HAP-based polymer composites have received significant research interest. Table 4.7 shows the mechanical properties of composites used in orthopedic applications. HAP is a natural component of bone and is thus highly biocompatible, with superior bondforming ability. Hence several studies have been conducted on polymer composite where HAP is used as the ceramic filler component (Boanini et al., 2006; Boduch-Lee et al., 2004; Causa et al., 2006; Higashi et al., 1986; Kikuchi et al., 1997; Yoshida et al., 2006; Zhitomirsky and Pang, 2006). Use of nanosized HAP particles and various techniques for modifying HAP±polymer interfaces have been explored (Sinha et al., 2007; Verma et al., 2007). Specially, the HAP particles having high aspect ratio (whisker or fiber) significantly improves the modulus with a lower loading wt%. Thus, several attempts have been made to synthesize whisker-like HAP particles (Converse et al., 2007; Roeder et al., 2003; Viswanath and Ravishankar, 2005; Yue and Roeder, 2006; Zhang et al., 2002). One study showed ultimate strength, elastic modulus and elongation at break of composite based on poly( -hydroxyalkanoates) (PHA) with HAP similar to similar to bone and is being investigated as a potential material for total hip replacement (Galego et al., 2000). Carbon fibers have good biocompatibility and excellent mechanical properties. They have been used to reinforce ultra high molecular weight polyethylene in total hip replacement components. The composites of carbon fiber with PMMA (Woo et al., 1974), polypropylene and polysulphone (Christel et al., 1980; Claes et al., 1997), polyethylene, polybutylene terephthalate, and PEEK (Gillett et al., 1986; Jockisch et al., 1992; Rushton and Rae, 1984) have all been investigated for potential applications in load-bearing applications. Multilayered laminated composites of carbon fibers with PMMA and PEEK have also been investigated (Fujihara et al., 2003; Sorrell et al., 2000). HAP is a bioactive material. Several studies have focused on HAP-containing composites to improve bioactivity and mechanical properties of composites for orthopedic applications: metal and ceramic fiber reinforcement of HAP (Ehsani et al., 1995;

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Table 4.7 Mechanical properties of composites for orthopedics Composites

Strength (MPa)

Modulus (GPa)

Reference

Poorly crystalline carbonateapatite + tetracalcium phosphate + collagen

6.08±11 (tensile)

0.66±2.24

Du et al. (2000)

Direct mineralized collagen composite (0±39% calcium phosphate)

34±53 (tensile)

0.44±2.82

Wahl and Czernuszka (2006)

Decalcified bone composite (10±15% calcium phosphate)

44.87 (tensile)

0.68

Wahl and Czernuszka (2006)

67

2.52

Galego et al. (2000)

Polyacrylic acid/HAP (40±70% HAP)

20±60 (compressive)

1±1.8

Katti et al. (2008)

UHMWPE±collagen± hydroxyapatite (23±40% HAP)

11.0±17.0 (tensile)

0.11±0.17

163 (compressive)

2.06

Verma et al. (2007)

Chitosan/hydroxyapatite composite (50% HAP)

74.08

1.02

Verma et al. (2007)

Chitosan/hydroxyapatite (70% HAP)

120 (compressive)

Zhang et al. (2007)

Self-hardening chitosan/ hydroxyapatite

26.2 0.88±4.29 (compressive)

Lu et al. (2007)

PHB/HAP (30% HAP)

Chitosan±polygalacturonic acid-hydroxyapatite (50% HAP)

Chemically coupled PE/HAP 18.34±20.67 Biphasic calcium phosphate/polylactic acid Polylactic acid/HAP

30±60

Roy Chowdhury et al. (2007)

Wang and Bonfield (2001) 0.296±2.48

Bleach et al. (2002)

0.66±2.24

Ignjatovic et al. (1999a)

Ruys et al., 1991), HAP/polyethylene (Bonfield, 1993; Wang et al., 1994), HAP/ polyethyl ester (Liu et al., 1997), HAP/polyphosphasone (Reed et al., 1996), HAP/polylactide (Ignjatovic et al., 1999a) and HAP/alumina composites (Li et al., 1995) have all been described. To improve the mechanical properties and bone-bonding properties of PMMA bone cements, the composites of PMMA with HAP and bioglass have been investigated. The addition of these materials showed enhancement of

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osteogenic properties of the implants as well as improvement in mechanical properties (Moursi et al., 2002; Vallo et al., 1999). In general the polymer/HAP interfaces are known to have an important role on the resulting mechanical properties (Ignjatovic et al., 1999b). Interfaces play an important role in deciding the overall mechanical properties of a composite. A weak interface may lead to a deterioration of the mechanical properties of the composite. Some studies have shown that chemical modification of an interface results in improvement in mechanical properties of composites. For example chemically coupled HAP±polyethylene composites (Wang and Bonfield, 2001), chemically formed HAP±Ca poly(vinyl phosphonate) composites (Greish and Brown, 2001) and polylactic acid HAP fiber composites (Kasuga et al., 2001) have shown improvement in mechanical properties.

4.7

Natural materials

A new class of materials has been developed which have a similar structure or composition to bone or are synthesized by following basic principles of biomineralization. These are called biomimetic materials. A key step in composite synthesis is the growth of minerals on an organic matrix in aqueous media (Du et al., 2000). Several polymers of both synthetic and natural origin have been used for synthesis of biomimetic HAP/polymer composites (Bakos et al., 1999; Bigi et al., 1998; Itoh et al., 2002; Katti et al., 2008; Teng et al., 2006; Wan et al., 2006; Zhao et al., 2002). This strategy has shown significant potential for development of materials for bone substitution. Collagen and calcium phosphate minerals, being the natural component of bone, are the natural choices for the development of these composites (Lawson and Czernuszka, 1998). Collagen constitutes 20% of bone and provides toughness. Collagen is also known for its bone formation ability. Several recent studies have focused on the development of collagen/hydroxyapatite composites (Wahl and Czernuszka, 2006). An additional advantage with collagen/hydroxyapatite composite is that it can easily be remodeled by the body (Du et al., 1998). The mechanical properties of these composites lie between those of cancellous and cortical bone (Clarke et al., 1993; Mathers and Czernuszka, 1991). A recent study indicated comparable wear resistance of hydroxyapatite±collagen±hyaluronic acid with UHMWPE (Roy Chowdhury et al., 2007). Chitosan-based composites prepared by biomimetic methods have also gathered significant research attention for possible application in load-bearing applications (Lu et al., 2007; Zhang et al., 2007). Chitosan, a polysaccharide, is biocompatible, biodegradable and exhibits antigenic properties. Incorporation of polyanionic polymer to chitosan-based composites has shown improvement in mechanical responses (Li et al., 2007; Rodrigues et al., 2003; Verma et al., 2007; Zhang et al., 2004).

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4.8

Joint replacement technology

Summary

Consistent with the current advancements in materials science and engineering, especially in the realm of advanced materials design, the medical community has significantly benefited by the applications of many advanced materials and composites for orthopedic applications, especially for joint replacements. This chapter describes an overview of the various different polymeric, ceramic, metallic, composite, and natural-based materials used for joint replacement applications. This review also describes current and new advancements in composites research with the use of nano-reinforcements for use for implant applications. Mechanical properties of these materials and their merits and demerits for implant applications are also described here. The current applications of materials for orthopedic applications have relied heavily on experimental research. Many testing-based experimental studies have been the basis for the advancement in the field until recently. Only recently with the availability of fast, large and expansive parallel computing capabilities is materials design starting to be aided by computational simulations. Simulationbased design is certainly the way of the future in advanced materials research and it is to be expected that medical research will benefit from such advancements in simulations of tissue±biomaterial interactions over the lifetime of implants at length scales ranging from the molecular to the macroscopic.

4.9

Acknowledgments

This work is supported in part with a grant from National Science Foundation (NSF CAREER # 0132768). D Verma would like to acknowledge support from North Dakota State University, Graduate School doctoral dissertation award.

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