Hydroxyapitite and Titanium

Materials and Design 55 (2014) 165–175

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Materials and Design

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Material processing of hydroxyapatite and titanium alloy (HA/Ti) composite as implant materials using powder metallurgy: A review

0261-3069/$ - see front matter � 2013 Elsevier Ltd. All rights reserved. http://dx.doi.org/10.1016/j.matdes.2013.09.045

⇑ Corresponding author. Tel.: +60 3 89216678; fax: +60 3 89259659. E-mail address: [email protected] (A.B. Sulong).

Amir Arifin a,b, Abu Bakar Sulong a,⇑, Norhamidi Muhamad a, Junaidi Syarif a, Mohd Ikram Ramli a a Department of Mechanical and Material Engineering, Universiti Kebangsaan Malaysia, 43600 Bangi, Selangor, Malaysia b Department of Mechanical Engineering, Sriwijaya University, 30662 Indralaya, Sumatera Selatan, Indonesia

a r t i c l e i n f o a b s t r a c t

Article history: Received 24 July 2013 Accepted 17 September 2013 Available online 25 September 2013

Keywords: Ceramic–metal composite Powder metallurgy Material processing Biocompatibility Mechanical properties

The bio-active and biodegradable properties of hydroxyapatite (HA) make this material a preferred can- didate for implants such as bone replacement in replacing natural tissues damaged by diseases and acci- dents. However, the low mechanical strength of HA hinders its application. Combining HA with a biocompatible material with a higher mechanical strength, such as a titanium (Ti) alloy, to form a com- posite has been of interest to researchers. A HA/Ti composite would possess characteristics essential to modern implant materials, such as bio-inertness, a low Young’s modulus, and high biocompatibility. However, there are issues in the material processing, such as the rheological behavior, stress-shielding, diffusion mechanism and compatibility between the two phases. This paper reviews the HA and Ti alloy interactions under various conditions, in vitro and in vivo tests for HA/Ti composites, and common pow- der metallurgy processes for HA/Ti composites (e.g., pressing and sintering, isostatic pressing, plasma spraying, and metal injection molding).

� 2013 Elsevier Ltd. All rights reserved.

1. Introduction

The rapid growth of the global population is leading to an in- creased demand for implants for bone dysfunction caused by dis- eases such as arthritis and cancer [1]. Such implants are necessary to repair or alter natural body tissues [2]. However, gi- ven the unique structures and mechanical properties of natural tis- sues such as bone tissue, repairing or changing them is challenging. Since the introduction of bioceramics as medical implants in the 1960s, metal implants such as titanium alloy, stainless steel, and cobalt–chromium alloys have been extensively used in medical applications [3]. In the early period of medical implant develop- ment, the only criteria for implant material suitability were appro- priate physical properties and non-toxicity [4]. Today, the criteria include the physical properties of the bone implant material and its ability to promote the growth of body tissue [5]. Metal-based implants have a higher Young’s modulus than bones, which leads to stress shielding. Metal implants also have poor biocompatibility, which is necessary to promote the growth of natural tissue. How- ever, metal-based implants have the beneficial mechanical proper- ties of strength and corrosion resistance. Hydroxyapatite (HA) is a bioceramic material with poor mechanical properties, especially for load-bearing applications. However, HA has a similar structure to bones and can promote the growth of natural tissues. Combining

a titanium alloy with HA creates a new biomaterial with excellent mechanical and biological properties. Thus, research on this mate- rial and its preparation process has been conducted [6–10]. This paper aims to review the interaction between titanium alloys and HA as a medical implant composite. Common methods of com- bining HA and titanium are also discussed.

2. Criteria for biomaterials applications

In the early period of implant material development, a material was considered suitable to replace natural tissue when it had min- imal or zero toxicity [11]. Later on, the ability to promote natural tissue growth was considered. Several studies have been con- ducted to achieve this goal in terms of the processing route, design and material modification [12–18]. Biomaterials for implants should not be cytotoxic. Cytotoxicity is caused by increased metal- lic ion content in the blood. Thus, biocompatibility itself can be translated simply as ‘‘do no harm’’ to the body and due to encour- age healing [19]. The roughness of an implant surface is conducive to the bonding between implant materials and tissues [20]. Porous structures and rough surfaces are necessary for facilitating bone in- growth and osteointegration [11,21–23]. Rough surfaces and por- ous structure able to control by its processing condition [24–29].

Stress shielding is a common problem of biomaterials [30,31]. This phenomenon arises when the Young’s modulus of bones (Fig. 1) and the implant material are different, which causes bone resorption. Most biomaterials based on metals and ceramics have

Fig. 1. Stress shielding mechanism [36].

Table 1 Comparison of the mechanical properties of implant materials [43].

E (GPa) YS (MPa) rf (MPa) %EL BF � 10�3

FeCrNiMo (316L) 210 450 250 40 1.2 CoCr (as cast) 200 500 300 8 1.5 CoNiCr (as wrought) 220 850 500 20 2.3 TiAl6VA 105 900 550 13 5.2 TiAl5Fe2.5 105 900 550 15 5.2 cp-Ti 100 300 200 40 1.8 cp-Ta 200 300 200 40 1.3 cp-Nb 120 250 150 70 1.3

YS = yield strength, E = Young’s modulus, rf = fatigue strength, BF = biofunctional- ity = rf/E, %EL = elongation.

Fig. 2. Density of selected materials [43].

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higher Young’s modulus than cortical and trabecular bones [32,33] (Table 1). Material selection and manufacturing processes have been performed to address this problem. Titanium alloys are po- tential implant materials because their Young’s modulus are closer to that of bone (10–30 GPa) compared to other metal implants [34]. Another approach involves controlling the Young’s modulus of bone by modifying its porosity [35].

3. Titanium and titanium alloys

Since the mid-1940s, titanium and its alloys have been widely used for medical applications, aerospace structures, and chemical industries [37,38]. Many researchers continue to develop titanium alloys for medical applications because of their unique properties, such as high specific strength, lightweightness and bioinertness [1,39–41]. Titanium and its alloys are classified as light materials with adequately high specific strength in engineering structures (Fig. 2). Titanium has a density of 4.5 g/cm3 and a melting point of 1668 �C. Titanium alloy has a tensile strength of approximately 1400 MPa at room temperature. At high temperatures, titanium becomes reactive with other materials. Meanwhile, titanium alloys have high corrosion resistance at room temperature because of the good biocompatibility and stable passivity of their surface oxide films [6]. Titanium alloys can be divided into a-, (a + b)-, and b-type alloys. b-type alloys have the lowest Young’s modulus (Table 1), which is closest to the Young’s modulus of bone (10–30 GPa) [42].

4. Hydroxyapatite

Hydroxyapatite theoretically exists as the hydroxyl end-mem- ber of apatite, which was suggested in 1912 [44]. Hydroxyapatite

is one of the apatite structures that were observed in rock, the apa- tite structure has the basic formula Ca10(PO4)6X2. X in the formula is the representation group member of apatite and refers to a hy- droxyl (OH) group for hydroxyapatite, a fluoride (F) group for flu- orapatite, and a chloride (Cl) group for chlorapatite [45]. Synthetic HA is widely applied as a substitute for the hard tissues of the human body damaged by disease or accident. HA as an im- plant can bond and promote natural tissue growth because of its chemical similarity to bone mineral [46–49]. At high temperatures, HA undergoes decomposition, which starts with dehydroxylation at approximately 900 �C in air and at 850 �C in a water-free atmo- sphere. The next stage is decomposition, which produces tetra cal- cium phosphate (TTCP) and tri calcium phosphate (TCP). TCP consists of b-TCP at <1200 �C and a-TCP at >1200 �C. HA loses many hydroxyl groups above 1300 �C and thus loses significant weight [45]. Other studies have revealed that at >1350 �C, the strength of HA drastically decreases [46].

5. Porous structure

A porous structure in an implant material has the function not only to decreasing the Young’s modulus of the material implant [50] but also of facilitating osteointegration. Moreover, characteris- tics of the porous structure play an essential role in the regenera- tion of bone [51–53]. Porous structures can be divided into open or

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closed cells. In an open cell condition, each cell in the structure has a connection. In a closed cell, no connection exists between cells on the structure [54]. The interconnection of each cell is critical for allowing blood and nutrition into the structure to encourage bone ingrowth and osteointegration. The pore size depends on the appli- cation and thus has no standard size. For bone ingrowth, the min- imum pore size is 50 lm. Large pores tend to be drained of blood and nutrient more than small pores. Thus, the pore size must be considered when examining the mechanical properties of an im- plant [51,55]. Powder metallurgy (PM) processes can be used to produce combined small and large pores by adjusting the powder size, pressure, sintering temperature, and method of adding the powder to the space holder [35,50,56–58].

Fig. 3. XRD pattern of Ti/10% HA composite at 1200 �C for 2 h under argon condition [65].

Fig. 4. XRD pattern of HA composite under an air environment [59].

6. Interaction of titanium alloys and hydroxyapatite

The interaction between titanium alloys and hydroxyapatite has been studied by several researchers [6,35,59–65]. Thermal sta- bility has been observed to affect the synthesis of HA [45], such that in the thermal processing of a HA/Ti system, TCP and TTCP are commonly produced after dehydroxylation and decomposition [66].

6.1. Argon environment

As previously discussed, dehydroxylation begins at approxi- mately 900 �C in air and at 850 �C in a water-free atmosphere. Pure HA is stable up to 1200 �C under an argon atmosphere. However, in HA/Ti composites, Ti ions react with the dehydrated water of HA to yield titanium oxide, thereby accelerating the dehydroxylation and decomposition of HA at approximately 800 �C. CaO and TTCP are the main products of HA decomposition in a Ti/HA system with in- creased sintering temperature. TTCP then decomposes to CaO at >1200 �C such that only two phases are formed, namely, CaO and an amorphous phase [64]. Balbinotti et al. reported that at approx- imately 1026 �C under an argon atmosphere, calcium titanate, TixPy, and TCP phases are the decomposition products, as observed in the XRD result in Fig. 3 [65]. These phases are formed at the boundary that covers the titanium particles.

6.2. Air environment

Egorov et al. investigated and characterized the interactions within Ti/HA composites in air at 700–1200 �C and observed that titanium is oxidized (Fig. 4) [59]. This result is supported by other studies on extreme oxidation at 1100 �C.

TiO2is produced at 800 �C, and only two crystalline phases exist: TiO2 and calcium titanate (CaTiO3). CaTiO3is produced by the inter- action between titanium and HA [59]. Berezhnaya et al. compared annealing results in air and argon [62]. Specimens annealed in ar- gon have a smooth surface, and those annealed in air have a sur- face covered with titanium oxide. Moreover, Berezhnaya et al. revealed that annealing in air reduces the biocompatibility because the optimum Ca/P ratio is not reached [45,62]. The optimum bio- compatibility of HA is achieved at approximately 1.5–1.7 Ca/P.

6.3. Vacuum environment

Yang et al. studied HA/Ti composites in a vacuum at 1100 �C, and observed the formation of TCP, TTCP, and calcium titanium oxide (Ca2Ti2O5) [60]. Fractured surface analysis revealed three types of surface morphologies of the HA/Ti composite, as shown in Fig. 5. Granular aggregation (calcium phosphate and Ca2Ti2O5), a porous structure, and a glassy phase are observed in HA, and a-TCP and TTCP are present in the composite. This morphology is

caused by the unequal melting points of the two materials as well as by the fusion between HA and molten Ti.

There is a possibility of reactions between hydroxyapatite and Ti under vacuum condition [60].

In the first stage, hydroxyapatite loses a hydroxyl chain in the following reaction:

Ca10ðPO4Þ6ðOHÞ2 ðin vacuum; > 800� CÞ ! Ca10ðPO4Þ6ðOHÞ2�2xOx þ xH2O ðgasÞ ð1Þ

H2O oxidized titanium became titanium oxide (TiO2)

Tiþ 2H2O ðgasÞ ! TiO2 þ 2H2 ð2Þ

Fig. 5. Scanning electron micrograph of the fracture morphology of Ti/HA composite in a vacuum: (a) 80% HA (1500� original magnification) and (b) 90% HA (1200� original magnification). (A) Granular aggregation, (B) porous area, and (C) glassy phase [60].

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CaTi2O5 was resulted from the reaction hydroxyapatite with TiO2

Ca10ðPO4Þ6ðOHÞ2 þ 2TiO2 ! 3Ca3ðPO4Þ2 þ CaTi2O5 þH2O ð3Þ

Under elevated temperature a-TCP and TTCP

Ca10ðPO4Þ6ðOHÞ2 ðin vacuum; > 800� CÞ ! Ca4P2O9 þ 2Ca3ðPO4Þ2 þH2O ð4Þ

Fig. 6. XRD spectra of HA–20 vol% Ti in a nitrogen environment for 60 min at (a) 1000 and (b) 1100 �C [67].

6.4. Nitrogen environment

The fabrication of a HA/Ti system in a nitrogen environment has been reported by several researchers [67,68].

Fig. 6 shows that at 1000 �C, only a small quantities of a-Ca3 (PO4)2, (a-TCP), and Ca4O(PO4)2 were detected. The decomposition phase of Ti/HA composites increase with increased sintering tem- perature, and no reaction occurs between HA and titanium [67]. TCP in HA/Ti composites degrades faster than HA, promoting new natural tissues and leading to unstable mechanical properties of the implant for natural tissue growth [8,67].

6.5. Diffusion mechanism between HA and Ti

At the HA/Ti interface, the Ti atoms at elevated temperature tend to oxidize to become TiO2, and normally, the TiO2 position on top of the titanium surface; however, parameter processing tends to alter the formation passivation of TiO2. In this situation, Ti atoms from the metallic bulk move and inter-diffuse on the film while oxygen atoms migrate to the titanium bulk [69,70]. The re- sult of this simultaneous process is titanium oxide in amorphous or crystalline form. The oxidation kinetics of titanium is deter- mined by the adsorption rate of oxygen, oxygen as an interstitial atom diffuses to the Ti lattice until reaching a saturation level, fol- lowed by oxidation of titanium. The diffusion rate of oxygen will decelerate when TiO2 occur [71]. Ye et al. and Chu et al. also re- ported that dehydroxylation of hydroxyapatite would be acceler- ated by the presence of Ti atoms, under this condition, oxygen diffuses to the Ti metallic bulk to form titanium oxides [64,67]. Furthermore, calcium and phosphorous ions diffuse into the Ti substrate to affect the Ca/P ratio of apatite. The phosphorous ion has the ability to migrate rapidly into the Ti substrate due to its smaller radii and lower activation energy. Other authors reported

that Ca, P and Ti have been detected on some regions but other phase that are common results of the reaction between HA and TiO2 were not observed. The presence of Ca, P and Ti atoms dem- onstrated that interdiffusion occurred in the HA/Ti composite [62,70].

Balbinotti et al. performed EDS line scans on the polished part of a HA/Ti composite; titanium and a small amount of calcium were detected on the polished area [65]. It was believed that the tita- nium formation phase titanium and calcium tend to be fragile and are partially lost because of the polishing mechanism. At the HA/Ti interface, titanium and phosphorous were detected, as ob- served in Fig. 7. The vertical arrow indicates the area where phos- phorus was detected; under this condition, Ti and P were detected in the formation TixPy phase which demonstrates that the phos- phorus ion has diffused to the titanium ion during sintering. Com- monly, differences in the thermal expansion of ceramic will lead to cracking due to nonuniform of residual stresses in the material [72]. Cracks have also been observed in HA/Ti composites that will degrade their mechanical properties. Shi et al. observed micro cracks in the HA/Ti composites, which were thought to be due to differences in the thermal expansion of Ti and HA [69]. Moreover, in addition, micro cracks were also found in the results using MIM products as reported by Thian et al. [63].

Fig. 7. SEM image of the polished surface of a HA/Ti composite (a) polished area and (b) EDS line scan analysis [65].

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7. Nanostructure of HA/Ti system

Nanomaterials are known as great potential candidates for im- plant material application [73]; Webster et al. reported that nano- structure materials (with grain sizes less than 100 nm) have the ability to increase function osteoblast adhesion, which supports the next stage of tissue growth [27,29,74], as shown on Fig. 8. Oth- ers researchers also reported that nanoscale topography played a critical role in increasing cell activity, which rapidly encouraged tissue growth compared with conventional materials [75–77]. The fabrication of HA/Ti composites on the nanoscale have the advantage of increasing consolidation between Ti and HA such as the hardness value, young’s modulus and corrosion resistance [14,78]. The combination of mechanical alloying and the powder metallurgical process has been proposed by Niespodziana et al. in the fabrication of nanostructure HA/Ti [78]. A unique method has been proposed by Farnoush et al., these researchers have com- bined friction stir processing and electrophoretic deposition to fab- ricate nanostructure HA/Ti [79].

8. In vitro and in vivo test

Tests of biological properties are important in the manufacture of implant materials [8]. These tests determine the suitability of an implant material. In vivo and in vitro tests are extensively used for biological evaluation. In vivo or ‘‘in the living organism’’ tests are performed inside the body of a living organism, and in vitro or ‘‘in glass’’ tests are performed in inanimate laboratory equipment [80]. The biocompatibility of a material is tested in vitro by immersing a specimen for a few weeks in simulated body fluid (SBF) whose conditions are similar to those in human blood plasma [81]. The ion concentrations in SBF and human blood plasma are listed in Table 2. Thian et al. reported that HA/Ti composites are tested through immersion in SBF solution to induce complete dis- solution of the secondary phase after 2 weeks [8,81]. Complete dis- solution of the calcium phosphate phase occurs during the initial stage of immersion, which results in the deterioration of the mechanical properties of the specimen. In the next stage, the mechanical properties of the HA/Ti composites significantly in- creases because of the precipitation of the apatite layer [8]. The combination of titanium and HA promotes apatite nucleation. In the secondary phase, Ti2O helps generate apatite on the composite surface [82,83]. In addition, the dissolution of CaO encourages a constructive location for apatite nucleation and growth [82].

In vivo tests of a HA/Ti composite have been conducted by Chu et al. [66,68]. using New Zealand white rabbits. These researchers cut the implant model into rectangular shapes 3.3 mm wide, 3.3 mm thick, and 5–6 mm long rectangular shapes, as illustrated in Fig. 9.

In vivo tests revealed that the HA/Ti composite has excellent biocompatibility that enables it to integrate with bone, which in- creases osteointegration and bonding strength with time after the initial stage of implantation [66,68,84]. According to Ning and Zhou, HA/Ti composites produce a bone-bonding interface with bone through an apatite layer. The bioactivity value of HA/ Ti compositesis determined by increasing the titanium content in a HA/Ti system [85]. These researchers also revealed that the in vivo and in vitro tests have similar results as confirmed by in vivo and in vitro investigations performed on other materials [86].

9. Powder metallurgy

PM is a manufacturing method in which powder metals are compressed with or without other materials, and then heated without a melting stage for solidification and strengthening [87]. PM produces excellent microstructures and compositions for man- ufacturing near-net-shaped products [88]. Several methods have been proposed to produce HA/Ti composites, such as conventional PM [61,65,89], non-conventional PM [90], and plasma spraying [91–93]. The conventional method or ‘‘pressing and sintering pro- cess’’ usually consists of mixing the powder, compacting the pow- der in a desired mold, and sintering [94]. Occasionally, during the compaction step, the mold temperature is raised to high tempera- tures and is classified as hot pressing [66,68]. Non-conventional processes include powder injection molding (PIM) and isostatic pressing, which can be further divided into hot and cold processes [94].

9.1. Pressing and sintering

Mixing is commonly required in PM and involves mixing the powder with other powders as well as with a binder. Failure in mixing irreversibly affects the next step. The mixing stage involves numerous parameters such as mixing time, mixing temperature, powder size, and powder shape, composition of powder, and com- position of the binder [95]. Thian et al. proposed the ceramic slurry approach for applying PIM and hot/cold isostatic pressure. Ti6Al4V

Fig. 8. Nanophase materials promote greater amounts of protein adsorption and encourage additional new bone formation compared with conventional materials [73].

Table 2 Concentration of ion simulated body fluid (SBF) and human blood plasma [81].

Species Ion concentration (mmol/l)

Blood plasma SBF

Ca2+ 2.5 2.2

HPO2�4 1.0 0.8

Na+ 142.0 140.3 Cl� 148.8 148.0 Mg2+ 1.5 1.3 K+ 5.0 5.3

SO2�4 0.5 0.5

HCO�3 4.2 4.0

Fig. 10. Scanning electron micrograph of HA/Ti6Al4V composite powder using the ceramic slurry approach [96].

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and HA powder were mixed with polyvinyl alcohol (PVA) as a bin- der for a certain time [96]. To remove the binder, the Ti6Al4V slur- ry was heated at 450 �C in air and the consolidation stage was performed at higher temperatures (�600–700 �C). Using this method, HA can be coated around the titanium core as shown in Fig. 10.

In conventional PM, uniaxial pressure is applied when the pressing powder is contained in a die. Removal of powder from the die can only be performed through the in-line movement of a punch from a no-cross-equipment mechanism [94]. Therefore, a complex geometry cannot be produced using this method.

Fig. 9. Implant model of a HA/Ti composite: (a) rectangular specimen and (b) cross-sectio

Moreover, the density of green compacts through conventional PM varies. However, despite these limitations, PM has numerous advantages such as low manufacturing cost, high tolerance, and minimal secondary machining process requirements.

nal view of specimen (DB = defective bone region; region for bone healing area) [66].

Fig. 11. Illustration of flexible mold movement in the isostatic pressing process [98].

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9.2. Isostatic pressing

Isostatic pressing is one well-known near-net-shape methods to produce parts with highly complex shapes [97]. Unlike conven- tional PM which uses uniaxial pressure, isostatic pressing powder makes use of a flexible mold ‘‘enveloped’’-contained powder with applied pressure through hydraulic pressure from all directions as shown in Fig. 11 [98]. Isostatic pressing can be divided into hot and cold types, wherein cold isostatic pressing operates at

Fig. 12. Cross section of the double layered capsu

Fig. 13. Illustration of plasm

room temperature using water and oil as the fluid pressure. In con- trast, hot isostatic pressing operates at elevated temperature [94].

Isostatic pressing is performed to induce maximum density [99]. In hot isostatic pressing, some parameters such as the sinter- ing temperature and pressure can be controlled to produce inter- connected porosity, good mechanical properties, and high permeability [100]. The resulting complete diffusion bond is one of the main advantages of this method. The complexity of the inner and outer surfaces of the product must be overcome to avoid the

le hydrothermal hot-pressing method [102].

a spray process [103].

Fig. 14. Stages of the metal injection process [111].

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negative effects of the large product size [97]. Under correct design and control, the mechanical properties of the product of hot iso- static pressing, i.e., near-wrought materials, and significantly im- prove [97,101].

Hot isostatic pressing requires necessary precision control for every stage to achieve optimum results. The range time per cycle is counted in hours to produce a specific component. However, producing many parts in one batch can decrease the running cycle. Sometimes, a product requires a high cooling rate to change its microstructure and mechanical properties. Thus, conventional hot isostatic pressing is the common option [97]. At high temper- atures, HA decomposes into a secondary phase [45], especially when the plasma-spraying method is used. Onoki and Hashida pro- posed a route to avoid this problem. These researchers used hot isostatic pressing to coat HA on titanium bars at low temperatures through the double-layered capsule hydrothermal hot-pressing method under hydrothermal isostatic pressure (Fig. 12) [102]. They successfully coated HA on titanium bars at 135 �C.

9.3. Plasma spray

Surface modification through plasma spraying may improve the biocompatibility and mechanical properties of metal implants [93]. Plasma spraying coats a material with other materials possessing specific properties (Fig. 13). For example, HA is used as a coating material and titanium alloy is used as a substrate. Metal implants with low biocompatibility are usually coated with a material of higher biocompatibility [91].

The structural integrity of bonds between a substrate and coat- ing under fatigue is determined by the suitability of the chemical and thermal properties. To avoid mismatches between a substrate and coating, an intermediate layer is used between them [104,105]. The mechanism of bonding between HA and titanium through plasma spraying that utilizes CaTiO3 as a bond layer is stronger than the bonding mechanism through mechanical inter- locking. HA reacts with TiO2 at 800–1000 �C on the Ti surface, where CaTiO3 and TCP are present as secondary phases [91]. Form- ing composite coatings with Ti increases the strength of the adhe- sive bonding coating. Consequently, Ti and HA can uniformly spread on the surface. A SBF testing also indicates no reduction in the bioactivity of a material [106]. The plasma-spraying method is known as an easy and safe method to coat HA onto Ti alloys. However, in biomedical applications, problems involving residual stress, a low level of crystallinity, a low level of porosity, and non-uniform distribution porosity are encountered [107,108] Microporosity includes sizes of approximately 10–300 nm caused by a significant decrease in the mechanical properties of a material [109]. Combining plasma spraying and isostatic pressing has been proposed by certain researchers to address this problem. Satisfac- tory results have been obtained in decreasing the microporosity and improving the mechanical properties of a material [109,110].

9.4. Metal injection molding

MIM is commonly used after modification and innovation at each stage of PIM [111], as shown in Fig. 14. PIM itself is a combi- nation of two methods: plastic injection molding and PM [95]. In MIM, the volume fraction of plastic for PIM is substituted by metal powder, which is mixed with plastic into pasta (feedstock). The feedstock is injected into the molding machine (green part), fol- lowed by an extraction binder and finally by sintering [112] If a material used is ceramic, it is called ceramic injection molding, whereas MIM denotes covering materials and ceramics [95]. The metal injection process offers reduced production costs, especially in producing complex components in large quantities [113,114]. In titanium processing using MIM, the small and complex products

are contaminated by carbon, oxygen, and nitrogen, which affects the other processes [115]. Contamination occurs in oxide form at the interface of a particle that initiates cracking. The contamination level of a material contributes to the fatigue properties of a product [116].

MIM is used to fabricate biomaterial components [8,12,113,117]. Developing composite structures containing tita- nium alloys and HA for medical applications using MIM has been extensively studied by Thian et al. [8,63,90,118]. HA/Ti composite powder is prepared through the ceramic slurry approach with

Table 3 Mechanical properties of HA/Ti composites based on various processes.

Process Sintering temp (�C) Composition E (GPa) HV r (MPa)

Press and sinter process 1300 HA (BHA and EHA) + Ti (5–10 wt%)

�22–30 GPa 166.4–235.44 50.47–53.29 [57]

Hot-pressing 1100 Ti–20 vol% HA 102.6 3.41 GPa 170.1 [62] Hot-pressing 1100 HA–20 vol% Ti 75.91 3.13 GPa 78.59 [63] Hot-pressing 1100 HA–40 vol% Ti 79.31 2.94 GPa 92.1 [64] Plasma-spray – HA and Ti – 384 28.6 [88] Plasma-spray – 50 wt% HA/50 wt% �37 (after in vitro test) �310 (after in vitro test) [76] Metal injection molding 1050–1150 50 wt% HA/50 wt% �22.5 GPa �1.20–1.26 GPa �15.9 [59] Metal injection molding 1050–1150 50 wt% HA/50 wt% �17.7–20 GPa (after in vitro test) – 23.5–25.8 (after in vitro test) [8]

E = Young’s modulus, r = bending strength, HV = vickers hardness.

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PVA as a binder. PVA is removed from the homogenized composite powder by heating. The powder is then crushed into small particles using a mortar and pestle [119]. The composite powder is mixed with commercial a multi-component binder system consisting of natural wax, fatty acid wax, stearic acid, poly-oxi-alkylen-ether, and olefin-hydrocarbons. A sigma blade mixer is used to mix feed- stock consisting of HA/Ti6Al4V composite powder and binder at a mixing temperature of 90 �C until homogeneity is achieved. The mixing temperature is determined by differential scanning calo- rimetry [118]. For 50 wt% Ti6Al4V and 50 wt% HA, Thian et al. found that a powder loading of 60 vol% is the most suitable for yielding homogeneous and moldable feedstock [90]. The effects of the sintering temperature, heating rate, and cooling rate on the density and porosity are determined. At a sintering tempera- ture of 1100 �C, the density, hardness, flexural strength, and flex- ural modulus increase at heating rate of 7.5 �C/min and cooling rate of 5 �C/min with a high flexural strength and modulus at 1150 �C [63].

At high sintering temperatures, a HA/Ti system decomposes at 800 �C [46,60]. Ye et al. [64] found that the presence of titanium in HA accelerates the dehydroxylation and decomposition of HA. The decomposition of HA increases with increased sintering tem- perature, and significant decomposition begins after the specimens are sintered above 1100 �C [63]. In vitro experiments on HA/Ti specimens have revealed their chemical content and surface mor- phology after immersion in SBF. After 2 weeks, the dissolution of secondary phases such as TCP, TTCP, and CaO is complete. At 2– 4 weeks, the mechanical properties of the specimens decrease be- cause of the dissolution of calcium phosphate layers. Afterwards, the mechanical properties are recovered by the precipitation of the apatite layer [8]. Some methods of fabricating HA/Ti compos- ites that approach the mechanical properties of bone and have Young’s modules ranging from 10 GPa to 30 GPa have been pro- posed, as shown in Table 3. However, the manufacturing cost, size, and geometry of the products must still be considered.

In the early weeks of in vitro tests, the dissolution of HA/Ti com- posites generally produce secondary phases such as TCP, TTCP, and CaO. The mechanical properties of implants significantly decrease and the mechanical properties increase to achieve the near-initial value after the next stage [8].

10. Conclusions

Issues concerning the material processing and metallurgy char- acteristics have been reviewed and discussed to meet the require- ments for medical implant applications using existing manufacturing process. The composition of titanium and HA deter- mines the effectiveness of the mechanical properties and biocom- patibility of HA/Ti composites. Moreover, the sintering parameters are critical factors in determining the phase of two materials formed during the diffusion process. In particular, the sintering

temperature plays a dominant role in the fabrication of HA/Ti com- posites because this parameter affects the thermal stability of HA. Generally, HA/Ti composites produce TCP, TTCP, and CaO in air, with the main phases being TiO2 and calcium titanate. In vivo and in vitro tests validate the ability of HA/Ti composites to form bonds with natural tissues, especially at the early stages of implan- tation. In conclusion, optimization of process parameters especially during sintering parameters (temperature, heating rate, time, gas condition) should be further investigated to determine a better working processing window for the manufacturing of HA/Ti composites.

Acknowledgments

This work was supported by Grant No. UKM-DLP-2012-027 from the Universiti Kebangsaan Malaysia’s Research University Grant and FRGS/1/2011/TK/UKM/02/20 from the Malaysia Ministry of Higher Education (MOHE).

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