Processing Methodologies for Polycaprolactone-Hydroxyapatite ...

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Copyright © Taylor & Francis Group, LLC. ISSN: 1042-6914 ... 1Department of Mechanical Engineering, The University of Akron, Akron, Ohio, USA. 2Department ...
Materials and Manufacturing Processes, 20: 211–218, 2006 Copyright © Taylor & Francis Group, LLC ISSN: 1042-6914 print/1532-2475 online DOI: 10.1081/AMP-200068681

Processing Methodologies for Polycaprolactone-Hydroxyapatite Composites: A Review Avinash Baji1 , Shing-Chung Wong1 , T. S. Srivatsan1 , Glen O. Njus2 , and Garima Mathur2 2

1 Department of Mechanical Engineering, The University of Akron, Akron, Ohio, USA Department of Biomedical Engineering, University of Akron, Akron General Medical Center, Akron, Ohio, USA

Biodegradable implants have shown great promise for the repair of bone defects and have been commonly used as bone substitutes, which traditionally would be treated using metallic implants. The need for a second surgery exacerbated by the stress shielding effect caused by an implant has led researchers to consider more effective, synthetic biodegradable graft substitutes. The hierarchical structures commonly designed are inspired by nature in human bones, which consist of minerals such as hydroxyapatite, a form of calcium phosphate and protein fiber. The bone graft bio-substitutes should possess a combination of properties for the purpose of facilitating cell growth and adhesion, a high degree of porosity, which would facilitate the transfer of nutrients and excretion of the waste products, and the scaffold should have high tensile strength and high toughness in order to be consistent with human tissues. Blending of polycaprolactone and hydroxyapatite has demonstrated great potential as bone substitutes. It is essential to identify a standardized processing methodology for the composite, which would result in optimum mechanical property for the biocomposite. In this study, biocomposites made of polycaprolactone (PCL) and hydroxyapatite (HAP) are reviewed for their applications in bone tissue engineering. The processing methodologies are discussed for the purpose of obtaining the porosity and pore size required in an ideal tissue scaffold. The properties of the composite can be varied based on the change in pore size, porosity, and processing methodology. This paper reviews and evaluates the methods to produce the hydroxyapatite-polycaprolactone scaffolds. Keywords Biocomposite; Hydroxyapatite; Polycaprolactone.

Tissue engineering using biodegradable, bioresorbable, and porous three-dimensional scaffold synthetic materials seeks to provide for the growth of capillaries and tissues. Such an approach provides a practical alternative for the treatment of damaged or malfunctioning organs and tissues [8–17]. In most recent years, blending experiments have been conducted to understand the use of biocompatible ceramics to provide good osteoconduction and good osteointegration properties in a biodegradable polymer. The bone consists of 70% hydroxyapatite, a form of calcium phosphate. Being the main mineral content of the bone, hydroxyapatites have the capability of stimulating bone growth for tissue scaffolds [1, 6]. Scaffold materials made of hydroxyapatite have been found to have a relatively slow rate of degradation and are intrinsically brittle. It is ideal to develop a scaffold material using hydroxyapatite and biodegradable polymers, which would allow for a balance of degradation and bone stimulus factors. Composites comprising poly (-caprolactone) (PCL) and hydroxyapatite (HAP) have found applications in the substitution, regeneration, and repair of bone tissues and other orthopedic usage [4, 7]. Hydroxyapatite takes the form of polygonal sintered coarse particles, which resemble the apatite in the natural bone. Poly (-caprolactone) is a hydrolytic polyester having appropriate resorption period and releases nontoxic byproducts upon degradation [9, 10]. Gross and Lorenzo [10], in an independent study, identified that pore sizes need to be greater than 100 µm to facilitate tissue growth and pore connectivity. The pores and connectivity primarily aid in the transfer of nutrients and oxygen into the scaffold and transfer the waste out of the scaffold. An ideal scaffold made of HAP and PCL requires

1. Introduction This paper reviews the potential applications and processing methodology of tissue scaffolds arising from hydroxyapatite (HAP)-reinforced polycaprolactone (PCL). Orthopedic surgeries using implants for replacing the diseased bone or stabilizing the fracture pose a medical challenge that often necessitates the need for a second surgical procedure to facilitate the removal of fractured components [1, 2]. The disadvantages in using metallic implants for the fixation of the fracture also include stress shielding in the region of the implant, inflammatory reaction, and foreign body reaction [3]. Large bone defects coupled with non-union, or misalignment, of the fractured components can cause pain and dysfunction, thereby necessitating the need for use of allograft or autograft as bone inductive agents for the purpose of bone repair [1–3]. The allograft utilizes frozen bones for the purpose of bone repair and regeneration and possesses intrinsic limitations such as disease transmission and immunological reactions. Though bone repairing capability is certainly better when using autografts, they essentially involve bone transplantation of the patient’s own bone to the diseased site, which leads to donor site morbidity [3–7]. Tissue engineering is an attractive alternative to the existing issues related to the use of techniques such as conventional metal implants, autograft, and allograft for the fixation of fractures.

Received February 22, 2005; Accepted March 18, 2005 Address correspondence to Shing-Chung Wong, Department of Mechanical Engineering, The University of Akron, Akron, OH 443253903, USA; E-mail: [email protected]

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the following attributes: • Three-dimensional porous interconnected structure to

facilitate cell growth and transport of nutrients and metabolic waste. • Biocompatible and bioresorbable with controlled degradation and resorption rates. • Mechanical properties quite similar to those of the neighboring tissues. • An intrinsic capability that would permit cell attachment and proliferation. The blending of the PCL and HAP to form a biocomposite scaffold can be performed using compounding techniques developed for engineering polymers. 2. Processing methodology for the PCL/HAP composites Studies have been performed to explore the ideal tissue scaffold materials possessing properties similar to that of the tissue site that could withstand the mechanical loads and concurrently exhibit the degradation time comparable to the healing time. The component material should be selected such that the scaffold will not severely influence the normal growth of neighboring tissues. Tissue growth and cellular behavior can be promoted by a judicious choice

of manufacturing processes and the use of growth factors in the scaffolds. In a study by Boyan and coworkers [2], composites were made from poly-D, L-lactide-co-glycolide copolymer (PLG) and recombinant human bone morphogenetic protein 2 (rh BMP-2). These composites had osteoinductive ability. The researchers processed the copolymers using dissolution/precipitation techniques coupled with heat and vacuum to produce the PLG rods having a porosity of 60–70% and pore size ranging from 50–250 µm. Saito and Takaoka [3] performed a similar study, wherein the BMP was used as a composite along with polylactic acid and its derivatives to induce entopic bone. In their study, they discussed three different methods, namely, cell therapy, gene therapy, and cytokine therapy, for the preparation of composites as shown in Fig. 1. The cell therapy uses the undifferentiated mesenchymal cells that are collected from the patient and cultivated to differentiate into osteogenetic cells. The BMP is added in vitro to these cells to reproduce new bone tissue and can then be implanted into the patient. For the gene therapy, the BMP vector can be added in vivo directly to the cells or ex vivo to the cells taken from the patient and cultivated. The cytokine therapy, on the other hand, uses rh-BMP along with the carrier materials that are implanted directly to the needed site. Extensive research was conducted on the effect of reinforcements in polymers for bone substitutes. Fillers such

Figure 1.—(A) Cell therapy. (B) Gene therapy. (C) Cytokine therapy.

PROCESSING METHODOLOGIES FOR POLYCAPROLACTONE

as polyacids, organic isocyanates, and silane were used with polymers as biocomposite and bone substitutes. It was found that some fillers gave rise to catatonic effects [4]. Hydroxyapatite is nontoxic, bioactive, and biocompatible and can be used as reinforcements in polymer to form a biocomposite. Azevedo and coworkers [4] illustrated two different methods for the processing of HAP/PCL biocomposite. Polycaprolactone MW = 80000 and sintered hydroxyapatite particles (particle size 38–53 µm) were initially dried in vacuum at 200 C for 2 days and then blended in an extruder at 130 C for 30 min. The other method consists of grafting PCL onto the surface of HAP particles. Caprolactone is dried with calcium hydride for 12 h and purged under argon flow at a reduced pressure. The HAP is dried in vacuum at 200 C for 2 days. The grafting could be achieved using ring-opening polymerization of caprolactone in the presence of HAP at 130 C for 3 days. It was observed that the composites prepared by blending had inadequate interaction between the polymer and the filler when compared to the grafted composites. Their report, however, did not provide information pertaining to shear flow and the degree of mixing in the extruder, which often influence the filler-matrix interfacial strength and thus the mechanical properties of the composite. The key observation is that the modulus increased with an increase in filler concentration, whereas the ultimate tensile strength (UTS  decreased with an increase in filler concentration. The observed decrease in tensile strength was attributed to non-optimization of the HAP/PCL interfaces. The degree of dispersion and homogenization of the hydroxyapatite particles in the polymer matrix remains unclear. In a comparative study by Calandrelli and coworkers [5], polycaprolactone:biological hydroxyapatite was mixed in the ratio 19:1 in an extruder. At the end of the preparation the mixture was cooled in an atmosphere of nitrogen. They observed the presence of biological hydroxyapatite improved the modulus while concurrently increasing the hydrophilicity of the polymeric substrate. It was also observed that the presence of hydroxyapatite stimulated osteoblasts attachment to the biomaterial and cell proliferation. Choi and coworkers [7] prepared composites of PCL/ HAP using 10%, 20%, 30%, and 40% wt HAP. H3 PO4 was added to tetrahydrofuran and a pH in excess of 11 was achieved by adding ammonium hydroxide. CaNO3 2 4H2 O was added to the solution to achieve a Ca/P ratio of 1.67 and was allowed to proceed for 24 h. The PCL was then dissolved in the solution and the resultant solution mixture was thoroughly stirred for 3–6 h. The pH of the solution was maintained at 10 by adding ammonium hydroxide, which facilitated the dispersion of the synthesized HAP particles. The solution was stirred until it turned viscous and then poured in petri dish to dry. Ammonium nitrate was removed from the dry composite by washing it with deionized water, thus yielding a porous composite. It was observed that the process resulted in homogenous distribution of HAP particles within the polymer matrix. This report discussed the method to create porosity but did not propose a viable method to control porosity. Percentage of porosity, shape of the pores, and size of the pores were not provided therein. This information is often required since they are essential

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criteria for an ideal tissue scaffold. The composite was immersed in water for 12 h to remove the impurities since traces of ammonium nitrate present in the composite can influence the mechanical properties. It was observed that tensile strength and elasticity decreased, whereas the elastic modulus increased with an increase in HAP content. The decrease in mechanical strength can be attributed to the formation and presence of microscopic pores upon removal of ammonium nitrate. The decrease in ultimate strength can also occur as a result of poor chemical interactions between the particles (HAP) and the matrix (PCL). It was observed that interfacial debonding prevailed at the HAPPCL interface. The ceramic is not chemically bonded to the PCL matrix, which explains the initiation of crack at the interface. In a study by Chen and coworkers [17], the PCL/HAP biocomposite was prepared by blending in melt form at 120 C until the torque reached equilibrium in the rheometer that was attached to the blender. Following blending, the sample was compression molded and cut into specimens of appropriate size for testing. It was observed that the composite containing 20 wt% HAP had the highest strength. The tensile modulus of the scaffold increased with an increase in the concentration of HAP. The mechanical properties of the scaffold depended on the conjoint and mutually interactive influences of the molecular weight of the polymer and the particle size of the hydroxyapatite. A higher yield strength and modulus was observed for smaller particles of HAP. This finding is interesting and ascribed to the larger interfacial surface area of the smaller HAP particles. Both the crystallization temperature (Tc  and melting temperature (Tm  were found to increase with an increase in the concentration of HAP, suggesting good interaction between PCL and HAP. The PCL matrix with a narrow molecular weight was observed to have stronger interfacial interaction and greater Tc and Tm . In a study by Marra and coworkers [11], a PCL/HAP biocomposite was prepared by initially dissolving polycaprolactone in chloroform at room temperature (7–10% weight/volume). Sieved NaCl (150–250 mm particle size), and HAP (∼10 mm particle size) were suspended in the solution and sonicated for 60 s. The solvent was evaporated and the scaffold was subsequently immersed in distilled water at room temperature for 24 h. The scaffolds were later dried and cut into discs of 12 mm in diameter. Upon evaporation of the solvent, the scaffold was weighed and immersed in distilled water. The thickness of the disc was controlled by applying a pressure of 6,000–10,000 psi using a hydraulic press. The 1-mm discs were immersed in distilled water to leach away the excess salt. The incorporation of HAP was studied using 0–50% of HA (w/w). Porosity was as high as 80% and controlled by the amount of NaCl. A similar study using the salt-leaching technique to obtain a biocomposite was performed by Dunn and coworkers [12]. The mechanical properties obtained using this technique was just one-third that of trabecular bone. Both the porosity and pore size resulting from this method can be matched with that of the trabecular bone, which allows for the regeneration of structurally equivalent trabecular bone within the biomaterial. The studies discussed thus far provide useful information

214 pertaining to processing methodology for the composites but fail to address properties of the composite with respect to pore size, porosity, and interconnectivity, which is essential in developing an ideal scaffold. Properties of composites made from polyesters and hydroxyapatite are influenced by particle diameter. Calandrelli and coworkers [5] demonstrated that the addition of HAP having a size of 60 µm or less improves the overall modulus of the composite. Mechanical properties of the composites were carefully determined from samples obtained by compression molding. They observed that an increase in the biological HAP concentration did increase both the modulus and yield stress of PCL. The results indicated good interfacial interactions between the biological HAP and PCL. Salt leaching is a simple technique and can be used to control the pore shape and connectivity. Coombes and Heckman [13] developed a gel-casting technique wherein the polymer is gradually dissolved in the solvent by heating. Cooling the solution results in a gel, which is subsequently dried in vacuum. Porosity in the scaffold is obtained by the removal of the solvent. The phase separation technique developed by Zhang and Ma [14] is quite similar to the gelcasting technique. In this technique the polymer is freezedried and the solvent is removed. Borden and coworkers discussed the factors contributing to high porosity and the reasons as to why the mechanical properties do not conform to the trabecular bone [15]. To overcome these inherent limitations they made use of microspheres to create porosity in the tissue scaffold. Using PLAGA microspheres, they developed a sintered matrix containing microsized pores, as shown in Fig. 2. Further, it was observed that the use of microspheres induced cell growth and adhesion as shown in Fig. 3. The PLAGA microspheres were thermally fused to form sintered three-dimensional porous matrix having pore diameters in the range of 100–300 µm and mechanical properties in the range of the trabecular bone. Wouterson and coworkers [16] examined the effect of dispersing

Figure 2.—Scanning electron micrographs demonstrating microsphere’s shape and size [15]. Reprinted with permission from Elsevier Science Ltd.

A. BAJI ET AL.

Figure 3.—Scanning electron micrographs demonstrating human osteoblasts adhesion and growth at day 16 [15]. Reprinted with permission from Elsevier Science Ltd.

microspheres in a polymer matrix on the deformation and fracture behaviors of the composites. The sample materials investigated included phenolic microspheres reinforced epoxy. It was demonstrated that the phenolic microsphere experienced deformation under loading as shown in Fig. 4. The observation is indicative of the toughening function in addition to the strengthening effect derived from the microspheres in polymers prior to composite failure. In our ongoing research study, we are exploring the use of PCL/HAP composites and microspheres for the purpose of enhancing porosity for cell growth. The HAP and PCL will be initially blended together using compounding techniques and subsequently sintered. Microspheres made by emulsion techniques will be introduced for the purpose of enhancing porosity of the scaffold. The HAP will be

Figure 4.—Deformation of phenolic micro sphere in a polymeric matrix [16].

PROCESSING METHODOLOGIES FOR POLYCAPROLACTONE

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Figure 6.—Weight loss vs. the immersion time for blended PCL/HAP (30%) [4].

Figure 5.—Schematic biocomposite scaffolds.

showing

designed

microstructure

for

nano-

dispersed in a PCL polymer matrix such that it forms hybrid nanocomposites structures as schematically shown in Fig. 5. It is surmised by Woutersun and coworkers that the two different orders of magnitude in reinforcement and toughening can provide synergistic interactions for the control and benefit of mechanical and fracture properties of novel composite scaffold materials. 3. Biodegradability of PCL, HAP, and their composites Controlled biodegradation is a critical factor in developing tissue scaffolds that can be gradually reabsorbed by and excreted from the body. Biodegradability generally depends on the following factors: 1) chemical stability of the polymer backbone, 2) hydrophobicity of the monomer, 3) morphology of the polymer, 4) initial molecular weight, 5) fabrication processes, 6) geometry of the implant, and 7) properties of the scaffold such as porosity and pore diameter. Biodegradation refers to a gradual chemical breakdown of the polymer as a result of the action of hydrolytic reaction and living organisms. The result is reflected in a noticeable change in mechanical, physical, and other material properties. The biodegradation of scaffolds under body environment should be precisely characterized prior to being implanted. The degradation of polymers can be assessed using thermogravimetric analysis (TGA) and differential scanning calorimeter (DSC). The degradation tests can be performed on a 0154 M NaCl aqueous isotonic solution having a pH = 74 that simulates the physiological fluids. Samples of the composite were immersed in NaCl and kept in a shaking water bath for a time period of 3, 7, and 16 days. It was proposed by Azevedo and coworkers [4] that weight loss and degree of water uptake provides an indication of degradation of the composite. This can be calculated using the relationship %S = Ma − Mi /Mi × 100

(1)

%W = Mi − Md /Mi × 100

(2)

where %S is the degree of water uptake, %W the weight loss, Ma the weight of the sample after immersion in the NaCl solution, Mi the initial weight of the sample, and Md the weight of the sample at the end of the experiment, after drying to achieve constant weight. It was observed that grafting leads to hydrolytically stable composites. The weight loss and water uptake as a function of immersion time in the solution for the PCL/HAP composites, prepared by grafted and blended techniques, is shown in Figs. 6 and 7.

4. Mechanical performance of PCL/HAP composites Mechanical properties of the composite can be characterized by loading the sample on a universal testing machine and obtaining the stress-strain relationships. Azevedo and coworkers [4] performed tensile tests on the composite on a Zwick Z020 universal material testing machine with a 500 N load cell, using a cross-head speed of 50 mm/min for a strain of 0.3%. Dumbbell-shaped samples (cross-sectional dimensions 05 mm × 4 mm) were cut from the compression-molded films. The cross-head speed was

Figure 7.—Water uptake vs. immersion days for blended PCL/HAP (30%) [4].

216 increased to 500 mm/min and the samples were deformed to failure. It was observed that modulus increased with an increase in HAP concentration. The ultimate tensile strength was observed to decrease with an increase in HAP concentration. Higher values of ultimate tensile strength could not be achieved due to the agglomeration of HAP in the PCL matrix. Improving the interaction between HAP and PCL coupled with good mixing of HAP and PCL will enhance the mechanical properties of the composite. Stress-strain curves for the composites containing 10%, 20%, 30%, and 40% wt of HAP were determined by Choi and coworkers [7]. Samples were thin strips 2 mm × 12 mm × 5 mm cut from the bulk composite and were loaded on a universal testing machine (UTM) using a 100 N load cell at a cross-head speed of 2 mm/min. The ultimate tensile strength decreased and elastic modulus increased with an increase in HAP content. The addition of HAP particles is effective in improving the elastic modulus of the PCL polymer. The stress-strain curves concerned are shown in Figs. 8A and 8B. The mechanical properties of the

A. BAJI ET AL. Table 1.—Comparison of mechanical properties of bones, HA, PCL, and scaffold materials.

Bone Parallel Normal Trabecular bone

Ultimate tensile strength (MPa)

Young’s modulus (MPa)

124–174 49

17.0–189 × 103 115 × 103

8

HA > 992% dense PCL HA/PCL composite (wt %) 10 20 30 40

Elongation (%)

50–100 80–110 × 103

1012 ± 097

15859 ± 1148

372 ± 033 281 ± 023 253 ± 021 239 ± 014

9198 ± 711 10635 ± 581 11192 ± 397 11841 ± 259

43836 ± 2604

1725 ± 108 749 ± 054 676 ± 043 577 ± 042

composite and the bone counterparts are given in Table 1. The elastic modulus and the ultimate tensile strength were found to decrease with an increase in the concentration of HAP. In a study performed by Marra and coworkers [11], the mechanical properties of the composite were determined from samples of size: 12 mm × 1 mm × 6 mm length × thickness × width. The tensile strength and elastic modulus were measured by an Instron mechanical test machine at a cross-head speed of 2 mm/min. The addition of HAP to the scaffold increased the modulus. However, there was no significant increase in tensile strength of the scaffold with the addition of HAP. Basically, the mechanical properties of the scaffolds can be enhanced by improving the interfacial characteristics between the HAP and PCL matrix. Calandrelli and coworkers [5] determined the mechanical properties of the tissue scaffold by testing the samples in tension. The addition of biological hydroxyapatite increased modulus and decreased peak stress and strain at rupture. They observed the smaller particles of HAP exhibited good interfacial interaction with the PCL matrix. Similar results were reported by Chen and Sun [17], who suggested that the smaller particles of HAP could act as an effective nucleating agent and possess greater yield strength and tensile modulus. The molecular weight distribution of the PCL also exerts an influence on the PCL-HAP interface and thereby the resultant mechanical properties of the composite. Overall, it is clear that the mechanical properties of the scaffold depends on the interactive parameters such as pore size, porosity, processing methodology used, particle size of the HAP, and distribution of the molecular weight of the PCL matrix.

Figure 8.—Stress-strain curves for pure PCL and composites containing 10%, 20%, 30%, and 40% wt of HAP.

5. Bone fracture and issues related to trauma According to Wolff’s law [18], the bone will constantly remodel itself based on an externally applied load. It was observed that a compressive loading resulted in an osteogenic response of an isolated bone segment. The response is related to the strain the tissue undergoes during

PROCESSING METHODOLOGIES FOR POLYCAPROLACTONE

loading. In adults, fracture of the bone can be healed by the presence of osteoblasts (bone-forming cells) adhering onto the surface of the bones. The biological steps through which healing occurs goes through the following steps: (1) An inflammatory phase, which serves to immobilize the fractured bone while concurrently activating the cells responsible for repair. This phase lasts for 3–7 days. (2) The reparative phase whereby the periosteal and medullar calluses are formed by osteoblasts that come from periosteum and bone marrow. This phase lasts for about 1 month. (3) The remodeling phase: This phase of healing occurs once union of the bones has been achieved. Modeling and remodeling of the fracture site occurs during this phase for the purpose of restoring the original shape and internal structure of the bone. Essentially four biomechanical stages can be defined during the healing of the fracture. 1. Stage 1 corresponds to one where the fracture site has low stiffness and the bone fails through the original fracture line. 2. Stage 2 corresponds to one where the bone fails through the original fracture line with a higher stiffness. 3. In Stage 3 the bone fails partially in the original site and partially through the original bone having a high stiffness. 4. Stage 4 corresponds to the phase where the failure site is not related to the original fracture line and occurs in the high stiffness region. A comprehensive understanding of fracture mechanics and fatigue damage in bone is desirable for the purpose of designing tissue scaffolds targeted at reducing stress concentration and fatigue. A normal bone possesses an ample number of naturally occurring microscopic cracks. Crack propagation in the cortical bone occurs through the formation of a microcrack zone, which subsequently develops into a wake. Fracture toughness is defined as the measure of the intrinsically inherent resistance of a material to crack extension. Crack propagation depends on the interactive influences of crack-growth rate per loading cycle and stress intensity factor. Nalla and coworkers [19] recently determined the propagation of a crack of macroscopic dimension in human cortical bone. The crack propagation occurs preferentially along the cement lines. It was observed that under conditions of sustained loading the cement lines fracture ahead of the advancing crack tip. Toughening arises from the uncracked ligaments that bridge the microcrack attempting to link with the main crack. Crack-tip blunting is favored when the crack velocity is slower and the inducing stresses are low. Under fatigue loading conditions, blunting is normally accompanied by crack growth. The process of crack-tip blunting and resharpening during each fatigue cycle results in an advancement of the fatigue crack [19]. The fatigue strength of the bone is determined by subjecting the bone to cyclical loading and determining the number of cycles to failure. The expected fatigue life for the cortical bone containing a crack/flaw is made based on the number

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of loading cycles required for the crack to propagate to failure. Controlled fracture behavior for bones using implants helps provide a positive effect on healing and promotion of tissue growth. Excessive motion across the fracture plane would result in a non-union of the fractured fragments. Though some motion across the fracture plane is beneficial, there exists an inverse relationship between the amount of motion and the formation of callus. Hence, it is important for the fracture to be stabilized and immobilized, and the transfer of load across the fracture zone would serve to strengthen the formation of callus. The fractures can be stabilized by the use of implants that would permit load transfer and some mechanical motion. Though surgical procedures are necessary in bone fracture treatment, the disadvantages of using metallic implants are the following: (1) stress shielding in the region of the implant, (2) inflammatory reaction, (3) hardware loosening, and (4) resurgery and carcinogenic potential from osteo-synthesis of the metals. It would be ideal to develop biodegradable implants that can address the drawbacks of metallic implants for fracture stabilization. Such biodegradable scaffolds form a critical objective in the development of novel HAPpolymer biocomposites. 6. Promise of PCL/HAP as a bone scaffold material Control of pore size, porosity, pore architecture, and degradation time are necessary for selection of scaffold materials of high fracture toughness. Few studies have focused on PCL/HAP composites for high toughness tissue scaffold applications. PCL is known to be nontoxic and possesses mechanical properties that are one-third of the trabecular bone. HAP is biocompatible and can be used as reinforcement for the PCL or other polymer matrices. The HAP component can stimulate bone growth by providing a surface for seeding osteoblasts. Also, HAP is known to have osteoconductive properties and has been successfully used in various studies as filler in tissue engineering. The PCL/HAP composites can be fabricated in a porous form that enhances the healing process, which includes regeneration of the bone tissues. The mechanical property of the tissue scaffold can be controlled by (a) varying the compositions of components in the scaffold, (b) inducing porosity by microspheres, and (c) control of pore size using emulsion techniques. Carefully adjusting the molecular weight of the polymers can also alter the degradation rates. These are attributes of the PCL/HAP composites and they are potentially useful for tissue engineering. The scaffold materials development will certainly open up new possibilities for orthopedic advances. Conclusions In this paper, the processing methodology and properties of HAP/PCL biocomposites were reviewed. Emphasis was placed on the mechanical properties and the practical demands for an ideal scaffold. It was generally agreed that the particle dimensions of filler material influences the mechanical properties of composites made of biodegradable polyesters and hydroxyapatite. Interaction between PCL and

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HAP determines the properties of the scaffold and can be controlled by selecting small particle sizes of HAP and narrow molecular weight distribution of PCL. 1. It was reported that the presence of a rigid filler phase often enhances the mechanical strength and stiffness of the biodegradable polymers regardless of the processing methodology used. The mechanical properties are influenced by the molecular weight and the moduli of the constituent materials. 2. It was reported that pores in a scaffold need to have sizes greater than 100 µm to facilitate the essential tissue growth and pore connectivity. 3. An understanding in the processing methodologies used to produce porosity is lacking. Future work will focus on introducing microscopic pores of controlled sizes for dispersion in HAP-reinforced PCL to generate hybrid nanocomposites. Such an approach shows potential for producing a material that offers a concomitant combination of high strength, high degree of porosity of controlled sizes, and high toughness for bone replacements and healing. References 1. Goto, T.; Kojima, T.; Iijima, T.; Yokokura, S.; Kawano, H.; Yamamoto, A.; Matsuda, K. Resorption of synthetic porous hydroxyapatite and replacement by newly formed bone. Journal of Orthopaedic Science 2001, 6, 444–447. 2. Boyan, B.; Lohmann, C.; Somers, A.; Neiderauer, G.; Wozney, J.; Dean, D.; Carnes, D.; Schwartz, Z. Potential of porous polyD,L-lactide-co-glycolide particles as a carrier for recombinant human bone morphogenetic protein-2 during osteoinduction in vivo. Journal of Biomedical Material Research 1999, 46, 51–59. 3. Saito, N.; Takaoka, C. New synthetic biodegradable polymers as BMP carriers for bone tissue engineering. Biomaterials 2003, 24, 2287–2293. 4. Azevedo, M.; Reis, R.; Claase, M.; Grijpma, D.; Feijen, J. Development and properties of polycaprolactone/hydroxyapatite composite biomaterials. Journal of Materials Science: Materials in Medicine 2003, 14, 103–107. 5. Calandrelli, L.; Immirzi, B.; Malinconico, M.; Volpe, M.; Oliva, A.; Ragione, F. Preparation and characterisation of composites based on biodegradable polymers for “in vivo” application. Polymer 2000, 41, 8027–8033. 6. Khan, Y.; Katti, D.; Laurencin, C. Novel polymer-synthesized ceramic composite–based system for bone repair: an in vitro evaluation. Journal of Biomedical Material Research 2004, 69A, 728–737.

7. Choi, D.; Marra, K.; Kumta, P. Chemical synthesis of hydroxyapatite/poly(e-caprolactone) composites. Materials Research Bulletin 2004, 39, 417–432: Gorna, K.; Gogolewski, S. Preparation, degradation, and calcification of biodegradable polyurethane foams for bone graft substitutes. Journal of Biomedical Material Research 2003, 67A, 813–827. 8. Hao, J.; Yuan, M.; Deng, X. Biodegradable and biocompatible nanocomposites of poly(e-caprolactone) with hydroxyapatite nanocrystals: thermal and mechanical properties. Journal of Applied Polymer Science 2003, 86, 676–683. 9. Kweon, H.; Yoo, M.; Park, I.; Kim, T.; Lee, H.; Lee, S.; Oh, J.; Akaike, T.; Cho, C. A novel degradable polycaprolactone network for tissue engineering. Biomaterials 2003, 24, 801–808. 10. Gross, K.; Lorenzo, L. Biodegradable composite scaffolds with an interconnected spherical network for bone tissue engineering. Biomaterials 2004, 25, 4955–4962. 11. Marra, K.; Szem, J.; Kumta, P.; Dimilla, P.; Weiss, L. In vitro analysis of biodegradable polymer blend/hydroxyapatite composites for bone tissue engineering. Journal of Biomedical Material Research 1999, 47, 324–335. 12. Dunn, A.; Campbell, P.; Marra, K. The influence of polymer blend composition on the degradation of polymer/hydroxyapatite biomaterials. Journal of Materials Science: Materials in Medicine 2001, 12, 673–677. 13. Coombes, A.; Heckman, J. Gel casting of resorbable polymers. Biomaterials 1992, 13, 217–224. 14. Zhang, R.; Ma, P. Poly(-hydroxyl acids)/hydroxyapatite porous composites for bone-tissue engineering. I. Preparation and morphology. Journal of Biomedical Materials Research 1999, 44, 446–455. 15. Borden, M.; Amin, S.; Attawia, M.; Laurencin, C. Structural and human cellular assessment of a novel microsphere-based tissue engineered scaffold for bone repair. Biomaterials 2003, 24, 597–609. 16. Wouterson, E.M.; Boey, FYC; Hu, X.; Wong, S.C. Specific properties and fracture toughness of syntactic foam: effect of foam microstructures. Composite Science and Technology 2005, 65, 1840–1850. 17. Chen, B.; Sun, K. Poly (-caprolactone)/hydroxyapatite composites: effects of particle size, molecular weight distribution and irradiation on interfacial interaction and properties. Polymer Testing 2005, 24, 64–70. 18. Martin, R.; Burr, D.; Sharkey, N. Skeletal Tissue Mechanics; Springer: New York, 1998. 19. Nalla, R.; Kruzic, J.; Kinney, J.; Ritchie, R. Aspects of in vitro fatigue in human cortical bone:time and cycle dependent crack growth. Biomaterials 2005, 26, 2183–2195.