Biodegradable Polymers and their Bone Applications: A Review

Transcription

1 International Journal of Basic & Applied Sciences IJBAS-IJENS Vol: 12 No: Biodegradable Polymers and their Bone Applications: A Review Saiful Izwan Abd Razak, Noor Fadzliana Ahmad Sharif and Wan Aizan Wan Abdul Rahman*, Polymer Engineering Department, Faculty of Chemical Engineering, Universiti Teknologi Malaysia, Skudai, Johor, Malaysia. Abstract Several biodegradable polymers were reviewed in regard of their properties to be used as bone substitute. Highlights were made on poly (glycolic acid) (PGA), the stereoisomers forms of poly (lactic acid) (PLA) and their copolymer poly (lactic-co-glycolide) (PLGA), poly(ε-caprolactone) (PCL) including natural biodegradable polymers; chitosan and collagen. Their composites, forms (scaffold and dense), drawbacks and potential usage were also discussed in detail. Index Terms Synthetic biodegradable polymers, Bone substitute, Scaffold. B I. INTRODUCTION one is a complex living tissue which has an elegant structure at a range of different hierarchy scales. It is a composite comprise of an organic phase (base on collagen) by which calcium containing inorganic crystals are embedded (Fig. 1) [1]. Specifically, bone is a natural composite material, which by weight contains about 60% mineral, 30% matrix and 10% water [2]. The matrix bone is primarily consisting of collagen. This organic component of bone is predominantly responsible for the tensile strength. The mineral component of bone is of calcium phosphate, this mineral component gives rise to the compressive strength [3]. In the body, bone serves as number of functions, such as providing the cells found in the marrow, that differentiate into blood cells, and also acting as a calcium reservoir [4-6]. Nevertheless, its primary purpose is to provide mechanical support for soft tissues and serve as an anchor for the muscles that generate motion. Fig. 1. Actual vision inner part of natural bone There are two types of bone, compact or corical, and cancellous or trabecular (spongy) bone. The anisotropic structure of bone leads to mechanical properties that exhibit directionally. This directionality results from the fact that bone has evolved to be both tough and stiff, two competing properties which are optimized in bone but with an inherent loss in isotropy [7]. In fact, bone exhibits extraordinary mechanical properties, displaying both elastic and semi-brittle behavior [2, 8]. Compact bone has a compressive strength in the range of MPa, and a Young s modulus ranging from GPa [2, 6], while compressive strength and Young s modulus for trabecular bones are 5 10 MPa and MPa, respectively [2, 6]. Bone plays an important role in the movement, support and protection of vital organs. However, it is susceptible to fracture as a result of physiological resorption, dental losses, trauma injuries, bone pathology, infection, aging, population and bone disease [9-11]. The ways of bone repair improve year by year, parallel to the development of high technologies. Traditional biological methods of bone defects management includes: 1) Autogeneous bone graft which harvested from different side of the patient s body, it is considered as the ideal bone graft substitute [12]. While the use of autograft material is the preferred technique, there are some limitations such as donor site morbidity, limited donor bone supply, anatomical and structural problems and elevated levels of resorption during healing [13] 2) Allograft bone which is obtain from The corresponding author can be contacted via *w.aizan@cheme.utm.my

2 International Journal of Basic & Applied Sciences IJBAS-IJENS Vol: 12 No: individuals of the same species as the receiver [9]. Allograft has the disadvantage of eliciting an immunological response due to the genetic differences and the risk of inducing transmissible disease [14, 15]. This necessities a thorough sterilization procedure that not only damages the graft s biological and mechanical properties, but also expensive [16, 17]. Reconstruction of bone with fully synthetic artificial bone graft is another way to provide and create lost bone mineral. It was started past years, the terms of innovation of ceramic for bone repair and reconstruction become an interest in the major advance of medical applications [18]. Calcium phosphate is highly promising as a bone substitute in orthopaedics among the other ceramics [19-21]. This group of material exhibit high biocompatibility, bioactivity, self setting characteristic, low setting temperature, adequate stiffness and easy shaping in complicated geometries [22, 23]. However, several drawbacks do exist causing some problems that hinder clinical use of calcium phosphate. The problem is that calcium phosphate lack of macropores and low porosity making its resorbing rate rather slow and mechanically brittle, thus restricts its utilization in medical applications [24-26]. For the last two decades of the twentieth century, a paradigm shift occurred from the biostable materials to biodegradable materials (hydrolytically and enzymatically degradable) for the medical and related applications [27-29]. Current trend predicts that in the next couple of years, many of the permanent prosthetic devices used for temporary therapeutic applications will be replaced by biodegradable devices that could help the body to repair and regenerate the damage bone [30]. A biomaterial can be defined as a material intended to interface with biological system to evaluate, treat, augment or replace any organ or function of the body [31]. The essential prerequisite to qualify a material as a biomaterial is its biocompatibility, which is the ability of the material to perform with an appropriate host response in a specific application. The host response to an implant depends on factors ranging from the chemical, physical and biological of the materials to the shape and structure of the implant. Biodegradable polymers are materials that fulfill the above requirement. The chemical, physical, mechanical polymers will vary with the time and degradation products produced have different level of bony compatibility. Notwithstanding, biodegradable polymers sometimes fall short of achieving satisfactory results in bone fixation [32]. These results should not necessarily be taken as a negative factor, since they are available to be reinforced with other materials like ceramics, metals, or clay in order to improve their mechanical properties and meet the promising requirements in bone substitute. Often, from the medical point of view, biodegradable polymer with sufficient mechanical strength and optimized lifetime in the body, which could finally be replaced by bone, is most desirable. Biodegradable polymers lack in mechanical properties. It is desirable to reinforce it with other biocompatible substance to enhance its properties. The best partner in reinforcing with ceramic is calcium phosphate, especially hydroxyapatite (HA) since their condition mimic natural bone composition [3]. Facing a complex biological and sensitive system as the human body, the requirement for bone repair using biodegradable polymers are manifold and challenging. Some of the important properties of biodegradable polymers can be summarized as follows [33]: 1) The material should not evoke or toxic response upon implantation in the body. 2) The material should have acceptable shelf life. 3) The degradation time of the material should match the healing or regeneration process. 4) The polymers should have appropriate mechanical properties for the indicated application and the variation in mechanical properties with degradation should be compatible with the healing process. 5) The degradation products should be nontoxic, and be able to metabolized and cleared from the body. 6) The polymers should have appropriate permeability and processability for the intended application. Biodegradable polymers offer a number of advantages for developing bone at defect sites [34]. The key advantages include the ability to tailor mechanical properties and degradation kinetic. Biodegradable polymers can differ in their molecular weight (M W ), polydispersity, crystallinity, structure and thermal transition, allowing different absorption rates. By means of porosity, biodegradable polymers can be used to impart macroporosity to the cement as polymers degrade and expose macropores to bony ingrowth. The strengthening of the graft from the bone growth and the deposition of the new bone should offset the weakening of the graft due to polymer degradation [35]. In bone application, biodegradable polymer or its composite can be prepared in two form, scaffold and dense. Usually, scaffold is used to associate with cell seeded in purpose of cell growth to form bone. Scaffolds are able to promote cellular interactions and process uniformly interconnected pores with adequate structural integrity. The interaction between cell and substrate is related to the osteogenic cells attachment, adhesion and spreading and its quality will influence cell proliferation and differentiation [36]. Cell can be obtained from calvarine [37, 38], trabecular bone [39], human embryonic stem (hes) cells and bovine osteoblasts (bob) [40]. Dense form of biodegradable polymer can be obtained from forging, hot or cold pressing [41]. Formation of bone in this particular form is usually by immersing the sample in simulated body fluid (SBF). SBF is the most favored model solution simulating the inorganic part of the blood plasma. The ions concentrations of simulated body fluid are nearly equal to those of blood plasma (Table 1) by maintaining the ph at 7.25 for the apatite nucleation [42].

3 International Journal of Basic & Applied Sciences IJBAS-IJENS Vol: 12 No: TABLE I COMPARISON OF THE IONIC CONCENTRATION IN BLOOD PLASMA AND SBF (MM/L) [42] Na + K + Mg 2+ Ca 2+ Cl - HCO3 - HPO4 2- SO4 2- Blood plasma SBF Though various studies have been done throughout the scientific community on biodegradable polymers, there is yet a systematic report on the application of biodegradable polymers on bone engineering. The purpose of this review is to observe various parameters (porosity, particle size, molecular weight, temperature, crystallinity, filler characteristic and etc) that possibly affect the mechanical properties and degradation rates of biodegradable polymers, whether in pure or mixing condition and also observation of bone growth after in vivo and in vitro implantation. Furthermore, biocompatibilities of biodegradable polymers in bone engineering were also investigated II. SYNTHETIC BIODEGRADABLE POLYMERS A. Polyesters / poly (α-hydroxy acids) Naturally occurring hydroxyl acids, such as glycolic, lactic and ε-caproic acids have been utilized to synthesize an array of useful biodegradable polymers for a variety of medical product applications. Especially in the implantable devices (internal fixation devices for orthopedic repair) made from these polymers are becoming part of standard surgical protocol [43-45]. It was started 30 years ago where polyesters increased the popularization of polymers as material for sutures and bone fixation [46] and remain among the widely used synthetic degradable polymers. Although all polyesters are theoretically degradable through the method of esterification, it is a chemically reversible process, only aliphatic chains between ester bonds can degrade over the time frame required for most of the biomedical applications [47]. Poly (α-hydroxy acids) comprise one of the earliest and most extensively investigated class of biodegradable polymers. It can be developed from a variety of monomers via ring opening and condensation polymerization routes depending on the monomeric units. The ester bond of the poly (hydroxyl acids) is cleaved by hydrolysis, which results in a decrease in the polymer M w (but not mass) of the implant [47]. The degradation rate of polyester is determined by initial M w, exposed surface area, crystallinity, and (in case of the copolymers) ratio of the hydroxyl acid monomers [48]. Among the class of poly (α-hydroxy acid)s, the most extensively used polymers are poly (glycolic acid) (PGA), the stereoisomers forms of poly (lactic acid) (PLA) and their copolymer poly (lactic-co-glycolide) (PLGA) [49-52]. These polymers are candidate in bone tissue engineering since it does promote biocompatibility and biodegradability which would allow cell growth in human body [30]. The chemical properties of these polymers allow hydrolytic degradation through de-esterification. Once they degraded, the monomeric components of each polymer are removed by natural pathway [53, 54]. The human body already contains highly regulated mechanism for completely removing monomeric components of lactic and glycolic acids. Due to these specialty polymers, such degradable material used in bone substitute have been approved by the US Food and Drug Administration (USFDA) [51]. With the exception of PGA, the polymers in this family are soluble in many common organic solvents and thus it can be processed by a variety of thermal and solvent-based methods [54, 55]. Due to several problems, it is advisable not to apply polyesters alone when dealing with bone substitution, since it can caused some drawbacks in the implementation in vitro and in vivo such as the degradation products reduce the local ph value, which in turn, may accelerate the polyesters degradation rates [56] and induce an inflammatory reaction, which are often ascribed to acidosis caused (chemically unavoidable) by the release of acidic degradation products (monomeric or oligomeric hydroxycarboxylic acids) [57-59]. Another disadvantage is that the mechanical properties of the highly porous scaffolds made from polyesters are relatively weak, which limits their use for bone-tissue engineering, especially the in vivo implant site [60]. B. Poly (lactic acid) PLA has received the most attention due to its renewable resource [61], biocompatibility, biodegradation, excellent thermal and mechanical properties, and superior transparency of the processed materials [62]. It is a linear aliphatic thermoplastic polyester and is readily biodegradable through hydrolytic and enzymatic pathways [61, 63-65]. Since lactic acid is a chiral molecule, it exists in two stereoisomeric forms that give rise to four morphologically distinct polymers that is D-PLA, L -PLA, D,L -PLA and meso-pla. D -PLA and L -PLA are the two stereoregular polymers, D,L -PLA is the racemic polymer obtained from a mixture of D- and L- lactic acid, and meso-pla can be obtained from D,L -lactide. Generally, L -PLA is more frequently employed than D -PLA, since the hydrolysis of L-PLA yields L(+)-lactic acid, which is the naturally occurring stereoisomer of lactic acid. On the other hand, the semicrystalline solids of L -PLA are preferred in the orthopedic surgeries application where high mechanical strength and toughness are required [66-69]. In bone tissue engineering, the scaffolds require appropriate material properties, controlled porous architechture, proper porosity and appropriate mechanical properties [70]. Poly ( L - lactic acid) (PLLA) satisfies many of those material requirements and already has been fabricated into scaffolds for bone tissue engineering [71-73]. Typical shaping technologies of those scaffolds include phase separation technology [74, 75], salt leaching technology [76] and gas-induced foaming technology [77, 78]. Those technologies can easily form controlled porous polymers below 200 µm size. But they can

4 International Journal of Basic & Applied Sciences IJBAS-IJENS Vol: 12 No: hardly control the porous from µm in size which also very important for bone regeneration and vascularization. A thermally induced phase separation technique (TIPS) was used by Ruiyun and Peter, 1999, in order to create highly porous composite scaffolds for bone tissue engineering [60]. PLLA was mixed with hydroxyapatite (HA) using dioxane as solvent. Freeze drying procedure allows to produce hard and tough foam with a co-continuous structure of interconnected pores and HA/PLLA composite skeleton. The composite mixture was proved to be in good bone bonding [79, 80]. Their work showed that foam porosity could be reached as high as 95% while pore sizes ranging from several microns to few hundred microns were obtained. Mechanical properties of HA/PLLA scaffold was compared to those corresponding pure PLLA and it was found out that the compressive modulus of composite (11 MPa) was significantly higher than the pure PLLA (6.4 MPa). Same technique and material was implemented by Nejati and co-worker [81], instead of using HA particles, they used nano and micro sized HA as filler in scaffolding. Both nano and micro composite scaffolds showed some molecular interactions and chemical linkages between HA particles and PLLA matrix, affecting the interfacial behavior and mechanical properties of the HA/PLLA composites. The porosity of scaffold was up to 85% and the average micropore diameter was in the range of µm. The compressive strength of nano-composite scaffold was up to 14.9 MPa while pure PLLA and micro-ha/plla scaffolds were found to be 1.79 MPa and MPa, respectively. This study demonstrated the positive effect nano-ha and micro-ha on improving the mechanical properties of the PLLA matrix, also indicating that the values are comparable to the high end strength of cancellous bone [82]. Furthermore, Mesenchymal stem cells (MSCs) for cell culture of scaffolds revealed that the cell affinity and biocompatibility of nano-ha/plla was found to be higher than the other two scaffolds. Xiong and co-worker [78] fabricated PLLA scaffold using precise extrusion manufacturing (PEM). The measured value of scaffold porosity was only 60% but the pore size values produced could not even reached by TIPS method, which was ranged from 200 µm to 500 µm. This happened due to the well prepared column inside the PEM with the additional of compressed air and nozzle. Effectiveness of the porosity formed increased the mechanical properties of the PLLA scaffolds up to 8.32 MPa. From the observation of Table 2, the compression strength with the presence of filler was much higher than non-filler, indicating that most of the particles bonded to PLLA matrix were on the surfaces of the pores, and it showed good adhesion between them [79, 80]. Some molecular interaction and chemical bonding which were presented between HA and PLLA influenced the interfacial behavior and mechanical properties of the composites [83, 84]. It is obvious that among pure PLLA and composite scaffolds, nano-composite scaffolds showed the highest compressive strength because of some reasons: 1) Particle sizes may attribute to the large interfacial area between nano HA and PLLA matrix. 2) More uniform distribution of nano HA particles in polymer matrix. Xiong and his co-workers successfully used PEM in order to achieve high compression strength and modulus. However, it is still inconvenient since there are other alternative methods with reasonable cost [78]. A new artificial structural elastic, partially degradable bone graft for supporting weak bone in the proximal femur was presented by Lagoa and co-workers [34]. The implant materials consist of a functionally graded combination of titanium, PLA, HA and calcium carbonate. The implant has an outer elastic layer in contact with the bone, consisting of D,L -PLA, HA and calcium carbonate, inner layer was made by manual dip-coating of the metal into solutions of PLLA with suspended calcium salts. The purpose of D,L -PLA being placed at the outer layer is due to its fast degradation rate which gave an opportunity for bone on and ingrowth when it degrades. PLLA has slow degradation rate and provides a biocompatible interface between the living tissue and the bioinert metal while the bone consolidates. The artificial strut grafts strengthen its implant by inserting titanium which offered mechanical stability for bone growth.

5 International Journal of Basic & Applied Sciences IJBAS-IJENS Vol: 12 No: References [60] Initial PLLA M w (kda) n.a. TABLE 2 COMPARISON MECHANICAL PROPERTIES OF SCAFFOLD COMPOSITES (PLA/FILLER) Material characteristics Mechanical properties Filler characterist ics None Porosity (%) Average pore size n.a. Methods Process condition Compression Thermally induced phase separation Stirred: 50 o C, 2hrs Solution: dioxane Immersed in liquid nitrogen. Freeze Dry: o C, 4 days Strength 0.24 Modulus 6.4 [81] 140 (0.5g) 374 HA, None ( µm) (64 µm) Thermally induced phase separation Stirred: 4hrs* Solution: dioxane Freeze Dry: C, 72 hrs micro-ha [78] nano-ha None ( µm) Precise extrusion 160 o C

6 International Journal of Basic & Applied Sciences IJBAS-IJENS Vol: 12 No: References Initial PLA M w (kda) TABLE 3 COMPARISON MECHANICAL PROPERTIES OF DENSE COMPOSITES (PLA/FILLER) Material characteristics Mechanical properties Filler characteristics Methods Process condition Tension (MPa) Compression (MPa) Bending (MPa) strength modulus strength modulus strength modulus [91] 220 none Hot pressing 103 o C HA, 50 wt%, Hot pressing 103 o C , µm [92] 100 HA, 80 wt%, Hot pressing 176 o C, mm 98.1 MPa nanoha, 80 Hot pressing 176 o C, wt%, 100 nm 98.1 MPa 430 HA, 80 wt%, Hot pressing 184 o C, mm 98.1 MPa nanoha, 80 Hot pressing 194 o C, ,000 wt%, 100 nm 98.1 MPa [93] 160 none Hot pressing 180 o C, MPa HA fibers, Hot pressing 180 o C, 40 11, wt% 40 MPa [94] 300 none Hot pressing 175 o C, MPa Grafted Hot pressing 175 o C, nanoha, 4 wt% 10 MPa [85] 104 none Hot pressing 160 o C, 56 nanoha, 25 wt%, D=2.1 µm nanoha, 50 wt%, D=2.1 µm [86] n.a none i) mixed ii) pressing mmmt, 1.5 i) mixed wt% ii) pressing mmmt, 3.0 i) mixed wt% ii) press mmmt, 6.0 i) mixed wt% ii) press 300 MPa Hot pressing 160 o C, 300 MPa Hot pressing 160 o C, 300 MPa i) 25 o C, 24h ii) 200 o C i) 25 o C, 24 ii) 200 o C i) 25 o C, 24h ii) 200 o C i) 25 o C, 24h ii) 200 o C

7 International Journal of Basic & Applied Sciences IJBAS-IJENS Vol: 12 No: Besides using separation technique to produce scaffold composite at a range of porosity, there is a way to prepare the mixture of the same composite (nano-ha/plla, 0:100, 20:80 and 50:50) by using hot pressing, dense samples were investigated rather than porous scaffolds [85]. In addition, annealing treatment was performed with the hope to enhance the crystallinity of the polymers. However, annealing treatment does not influence the physical properties of the composites (particle sizes, porosity and mechanical properties, see Table 3) since the polymer has reached its ultimate crystallinity at the end of the cooling step, and yet, only remaining chloroform at the surface of the composites thoroughly eliminated after annealing treatment. Dispersion of HA nanoparticles into a PLLA matrix was a success whereby compressive strength of over 100 MPa have been achieved, this value expressed that the developed nano HA/PLLA properties was comparable to those of cortical bone. Nano-HA/PLLA composites are thus promising; to be develop as bioresorbable porous bone substitutes showing superior mechanical performance. Tzong and Cheng [86] used solution mixing process of PLA reinforced with montmorillonite (MMT), the MMT modified by chitosan function was aimed to improve the chemical similarity between PLA and MMT. Polymer/clay composites have received significant attention in academic and industrial area in recent years due to the excellent enhancement in physical and/or chemical properties relative to the neat polymer matrix [87-89]. The results showed that the interaction between PLA and MMT was successful for the used in biomedical application and the detail is shown in Table 3. Mechanical resistance was also evaluated in order to determine the best operating conditions of HA/PLA scaffolds to produce implants which offers optimized properties for use as bone substitutes [90]. A pore-forming agent (salt) was carried out in the forming of scaffolds which restricted to 70 wt%, whereby resistance as high as 2.2 MPa could distort the composites mechanical properties. It was shown that the addition of 30 % to 40 % of filler allows a satisfactory dispersion of the mineral within the matrix. A high percentage of filler is unfavorable for processing of porous scaffolds, as it prevents the matrix from playing its role as a binder. Table 3 displays several examples of the composite according to the initial M w, the nature of filler and the process conditions. Presence of nano-particles as filler obviously improves the mechanical strength and stiffness of PLA composites. The maximum strength that can be reached was at 139 MPa [92]. Nano-particles have gained much recognition as bone dealer not only due to their composition and structural similarity with natural bone but also because of their unique functional properties such as larger surface area and superior mechanical strength than those of single or pure constituents [95]. The effect of micro-fiber demonstrated by Kasuga and co-workers [93] achieved a very high bending modulus but showed the vice versa effect on the bending compression. The compressive strength increased by increasing the quantity of filler (%wt) and the initial M w. Nano-particles obtain better results in terms of mechanical strength and moduli. In addition, composite in the dense form more preferable compare to scaffold due to their capability to perform in high load bearing applications. C. Poly (glycolic acid) PGA has the simplest structure. It can be considered as one of the first biodegradable synthetic polymer investigated for biomedical implant. PGA is a highly crystalline polymer (45-50 % crystallinity) and therefore exhibits high tensile modulus with very low solubility in organic solvents. The glass transition temperature (T g ) of the polymer ranges from 35 to 40 o C and the melting point is greater than 200 o C. It shows excellent mechanical properties, also due to its high crystallinity. A self reinforced foam composed of PGA is stiffer than any other degradable polymeric system used clinically [96] and has been shown to exhibit a modulus approximately 12.5 GPa [97]. Due to its good initial mechanical properties, PGA has been investigated as bone internal fixation devices. However, PGA is also known to lose its strength in 1-2 months when hydrolyzed and losses mass within 6-12 months. Thus, this indicates that the degradation rate is very high, which in the end would limit the mechanical properties and biomedical applications for PGA. Therefore, as reported by Linhart and his co-workers [37], they used PGA reinforced with amorphous carbonate apatite to investigate any potential performance of PGA in bone substitute and found out that the composite could serve only as bone substitution materials for small or non-load bearing indications. D. Poly (lactide-co-glycolide) The combination of lactic and glycolic side formed PLGA amorphous polymer. Both L- and DL- have been used for copolymerization. In the composition range of %, PLGA forms amorphous polymers. Miller and his co-workers have shown that the 50-50% PLGA is very hydrolytically unstable and the resistance to hydrolytic degradation was found to be more pronounced at either end of the copolymers composition range [98, 99]. PLGA was used in a number of clinical application [100] but few in the orthopedic area. Most of the PLGA appears in scaffold due to its favorable properties. However, the surface chemistry of PLGA does not fully promote cell adhesion for enhancing the bone ingrowth and proliferation due to its hydrophobic nature [101]. The mechanical properties and biodegradation rate of PLGA can be manipulated to some extent by controlling the molecular weight of the copolymer and the unit ratio of lactide to glycolide in the copolymer. PLGA used in bone repair or substitute applications have shown to be biocompatible, nontoxic and non-inflammatory [102, 103]. However, PLGA cannot be utilized as load bearing applications either due to its low mechanical strength, too flexible characteristic or lack of osteogenic bioactivity.

8 International Journal of Basic & Applied Sciences IJBAS-IJENS Vol: 12 No: Fei and co-workers created a novel biodegradable PLGA for bone repair [104]. The moldable, degradable bone substitute with osteogenesis materials composed of PLGA loaded with bone morphogenetic protein-2 (BMP-2) microspheres incorporated with calcium phosphate cement (CPC) for bone growth. Lactic to glycolic ratio was 50:50. The scaffolds only appeared when the composites immersed in the phosphate buffer saline (PBS) in order to induce degradation towards BMP-2 which then formed PLGA/CPC scaffolds. The results showed that the CPC composites compressive strength decreased from MPa to 8.26 MPa when PLGA was added during setting. This happened because the structure of interlocking CPC crystal needles, which produce the cement strength, was inhibited when microspheres were added into CPC [105]. Although the mechanical properties of the composites decreased after PLGA was added, better degradation and stable composite was produce and the results were equivalent to the cancellous bone. Another PLGA pellets with a lactide/glycolide ratio of 85:15 with the pore sizes range of µm were used by William and his co-workers in their investigation [106]. In this study, the process of mineralization in tooth and bone development was directly mimicked through the use of SBF as the mineral growth solution. In other words, PLGA was incubated in SBF at different times to demonstrate the growth of a continuous bonelike apatite layer within the pores of the polymer. In addition, the process was a single step and occurred in the room temperature. It resulted in a growth of a continuous layer on much of the inner pore surface, which significantly increased the compressive modulus up to 320 ± 60 kpa after 16 days of treatment, without a large decreased in total porosity. The mineral grown was a carbonated apatite mineral layer, similar to the mineral present in bone, and hence is expected to have a high degree of bioreactivity. Such a layer could be of great importance in the field of bone tissue engineering, providing the advantages of increased osteoconductivity and enhanced mechanical properties. Also, it is an efficient and relatively inexpensive mineral growth process. Surface grafted of HA in the PLGA matrix deposition (g- HA/PLGA) showed excessive improvement on interface between the polymers and CPC rather than non-grafted HA/PLGA [107]. Three dimensional porous scaffolds of grafted HA/PLGA was fabricated using solvent casting/particulate leaching method to compare with conventional HA/PLGA scaffolds and to investigate its application in bone replacement [107]. In their report, in vivo mineralization and osteogenesis were investigated by replacement for repairing radius defects of rabbits. Table 4 analyzes the advantages of g-ha/plga and HA-PLGA. Despite of decreased in calcium exposure when implantation, surface grafted-ha/plga showed some advantages and good requirements in bone fixation. Unfortunately, no mechanical properties have been reported for comparison, but it can be said that non-grafted composites showed a better achievement after 20 months of post surgery. Other than that, both composites offered different advantages which make no differences in in vivo implantation where both showed similar ability to enhance in vivo mineralization and allowed rapid mineralization conjugation with neighboring bone better than neat PLGA. Aligned nanofibrous scaffolds based on PLGA and nano- HA were synthesized by electrospinning for bone engineering [108]. At higher concentration (> 10wt %), a reduction in mechanical properties occurred because of the agglomeration of HA in the PLGA matrix fiber, whereby the presence of nano-ha resulted in fiber breaking (Table 5). In electrospinning, the major factors that have a direct effect on the fibers diameters (size particles) were solution concentration and electric field. At higher voltages, needle tip inside the machine produced unstable sources of fibers. The collection of fiber scaffolds on a high speed rotator resulted in a highly oriented fibers. The diameter of the scaffolds from this method can be affected by the choice of the solvent. It was suggested that the used of more polar and hydrophilic solvents can reduce the degree of aggloromeration of the HA powder and can thereby find a good dispersion of particles in the polymer matrix [109]. Lee and his co-workers [40] aimed to develop a composite scaffold for bone regeneration that would meet criteria by hybridizing bladder submucosa matrix (BSM) as a natural bioactive material with synthetic PLGA polymers. BSM alone was not suitable for bone graft applications due to its small pore size, poor interconnectivity and inability to maintain structural integrity. To overcome these limitations, a design of natural-synthetic polymer scaffold that would possess an interconnected network of pores and sufficient mechanical and chemical properties that would maintain structural integrity, thus preventing collapse during the handling and implantation process. PLGA scaffold with incorporation of BSM did not alter structural changes during the fabrication process. From Table 5, the compressive strength of hybrid natural-synthetic scaffold was lower than the pure PLGA due to the natural behavior of hydrophilic characteristic by BSM itself. Although physically weaker than other scaffolds, the result after cell seeded in the scaffolds was promising after all. BSM-PLGA allowed 90% cell adhesion, while 30% by pure PLGA.

9 International Journal of Basic & Applied Sciences IJBAS-IJENS Vol: 12 No: TABLE 4 ADVANTAGES OF GRAFTED HA FILLER AND UN-GRAFTED HA FILLER IN PLGA MATRIX [107] g-ha/plga HA/PLGA More uniform distribution of granules on the pore walls surface. After 20 months post surgery, calcium content that attached to the composite was 10 wt%, higher than grafted composite (8 wt %). Improve the microstructure stability of the composite scaffolds Better osteogenesis after implantation TABLE 5 MECHANICAL PROPERTIES OF SCAFFOLD COMPOSITES (PLGA/FILLER) Material characteristics Mechanical properties Ref. Initial PLGA M w (kda) Filler characteristics Lactic:glycolic ratio Porosity (%) Average pore size [40] 110 none 50: ( µm) 110 Collagen, BSM 50: Methods Process condition Storage modulus (MPa) Solvent casting/ particulate leaching PLGA: NaCl 1:10, Air dried 24h, vacuum dried 48h µm Compressive strength (MPa) Young modulus (MPa) ( µm) [107] 196 none 80: Solvent casting/ particulate leaching HA, 10 wt% 80: Grafted HA, wt% 80:20 [106] 15 None 94.0 Solvent casting/ 85:15 particulate leaching 63 none 85: electrospinning HA, 1 wt% 85: HA, 5 wt% 85: HA, 10 wt% 85: HA, 20 wt% 85: PLGA/HA:sucrose 1: µm PLGA:NaCl µm kv, rotating speed=6000 rpm, solvent=hfp [104] 24 CPC, 40 wt% 50:50 Molding

10 International Journal of Basic & Applied Sciences IJBAS-IJENS Vol: 12 No: PLGA scaffolds reinforced with calcium phosphate like HA as filler showed better mechanical properties compared to pure PLGA, due to the HAs more similar characteristics to bone apatite, its ability to promote the attachment of cultured osteoblasts and to improve metabolic activity [ ]. Addition of HA in polymer matrix had been approved to increase mineralization in matrices in vitro osteoblasts culture [113]. It provide more favorable surface for cell attachment and the ability of the material to support bone ingrowth [114]. In addition, calcium phosphate additives in PLGA matrix provides surface roughness which enhanced bone attachment and proliferation [115]. The advantage of PLGA is that the strength of the composite could be controlled by the amount of PLGA microspheres, which makes it possible for the surgeons to modify the material to obtain an adequate strength. The particles sizes also play a significant role in bone implantation, whereby particle sizes ranging from µm have been found to be optimal for bone ingrowth [116]. However, previous studies noted that the size of interconnection in the scaffolds is the main limiting factor for bony ingrowth rather than the size of pores themselves [117]. Whether the method used are solvent casting, electrospinning or BMP loaded in PLGA matrix and degrade when immerse in the PBS solution, the mechanical properties produced (shown in Table 5) were all in the range of potential in forming cancellous bone, (compression strength, 10 MPa, modulus, MPa). PLGA in the dense form was poorly studied and hardly found in the previous work, it could maybe due to the amorphous characteristics of PLGA whereby unknown melting temperature renders PLGA unsuitable to be prepared by melt blending or hot pressing. highly processable as it is soluble in a wide range of organic solvents. The uncommon things of PCL is its high thermal stability, with decomposition temperature (T d ) of 350 o C, whereas others aliphatic polyesters are between 235 o C and 255 o C [118]. In bone engineering area, PCL can be categorized as promising biocompatible and biodegradable polymers since it is being used to enhance bone ingrowth and regeneration in the treatment of bone defects [119, 120], however, PCL is poorly studied due to the slow degradation [121]. One of the work done by Chen and Sun [122] was by reinforcing PCL with HA by melt blending technique (Table 6). This method requires the exact melting temperature of the biodegradable polymers to ensure that the polymer fully melted and provide fillers with spaces for fine dispersion in the polymer matrix. Smaller HA particles (3-8 µm) distributed higher compressive strength than HA (20-80 µm), indicating that size of particle affect the improvement of the composites mechanical properties [122]. The smaller the particle diameter the greater the specific surface and interfacial areas within the matrix. Their results can simply be plotted as seen in Fig. 2. Fig. 2 indicates that the addition of HA increased the compression of composite towards bone fixation, but, until some level, the failure mechanism of the composites changes from plastic to brittle (easily rupture) and hence lowering the mechanical properties of PCL and other biodegradable polymers. E. Poly (ε-caprolactone) Most research work has been directed at PLA, PGA and PLGA copolymers. The success of these for pharmaceutical applications has further led to the evaluation of aliphatic polyesters such as poly(ε-caprolactone) (PCL). PCL is an aliphatic polyester that has been intensively investigated as a biomaterial. It appears as semicrystalline polyester and is Fig. 2. Compression strength versus weight of HA in PCL [122].

11 International Journal of Basic & Applied Sciences IJBAS-IJENS Vol: 12 No: Material Characteristics TABLE 6 MECHANICAL PROPERTIES OF COMPOSITES (PCL/FILLER) Mechanical properties Ref. Initial PCL M w (kda) Filler characteristics Composite form Methods Process condition Compressive strength (MPa) [122] 80 None dense Melt blending 61.1 o C, 10 min, Pressing mould 120 o C HA, 3-8 µm, dense wt% HA, 3-8 µm, dense wt% HA, 3-8 µm, dense wt% 80 None dense Melt blending 62.3 o C, 10 min, Pressing mould 120 o C HA, µm, dense wt% HA, µm, dense wt% HA, µm, dense wt% [123] 50 None scaffold impregnating 600 o C 0.27 HA scaffold 600 o C 0.57 TABLE 7 DEGRADATION RATE OF BIODEGRADABLE POLYESTERS Polymers Glass transition point, T g ( o C) Biodegradable time (months) Ref lower upper DL-PLA [128, 131, 132] L-PLA > 24 [128, 132] PGA [128, 131, 133] PLGA Adjustable: [52] 1-12 PLGA 1-2 [134] (50/50) PLGA 4-5 (75/25) PLGA 5-6 (85/15) PCL > 24 [135] F. Biodegradation of polyesters Degradation could be characterized in many ways, like loss of M w, loss of mass or loss of mechanical strength, chemical composition, configurational structure, stress and strain, crystallinity, porosity, chain orientation, dispersion of chemically reactive compounds within the matrix [124, 125]. The polymers degrade by random chain scission [126]. Since only the monomer degradation products are soluble, a very large reduction in both M w and mechanical strength typically occurs before a decrease in the weight of the sample is observed. In the case of implantation in vivo, the body already contains highly regulated mechanism for completely removing monomeric components of lactic and glycolic acids. PGA is converted to metabolites or eliminated by other mechanisms, while PLA can be cleared through the tricarboxylic acid cycle. The crystallinity of L -PLA samples has been observed to increase during degradation and particles of highly crystalline material can be found in vivo long after the bulk of the implant has disintegrated [56]. Biodegradable polyester degradation occurs by uptake of water followed by the hydrolysis of ester bonds. PLGA has a wide range of degradation rates since the ratio of lactic over glycolic varied, the degradation kinetics being governed by both hydrophilic/hydrophobic balance and crystallinity. Composition of chain (lactic or glycolic units) determines the degradation rates of PLGA copolymers. Blends that contains higher amount of PGA have been shown to degrade faster [124]. PCL degradation proceeds with two stages: random hydrolytic ester cleavage and weight loss through the diffusion of oligometric species from the bulk. PCL with high M w could take several years to degrade in vivo [ ]. In general, the sequence of polyesters degradation rates decreases in the following order [54]: PGA > PDLLA > PLLA > PCL The kinetic patterns of polyesters degradation are shown below [130]: --COOR HOOC-- + H 2 O 2COOH + ROH Table 7 shows the degradation rate of some biopolymers. The degradation rate, physical and mechanical properties of PLA and PGA are adjustable by varying their M w and copolymers. However, from one review of responses towards resorbable pins made from either PGA or copolymer, over 500

12 International Journal of Basic & Applied Sciences IJBAS-IJENS Vol: 12 No: patients, 7.9 % developed a late noninfectious inflammatory response during healing process. It has been suggested that the consequences occurred due to the release of acidic degradation products (PGA and PLA) [136]. Böstman gave a solution to these problems; for orthopaedic applications, it requires the development of a polymer that is more hydrophobic than PLA and PGA and does not release acidic degradation products on hydrolysis [136]. New era with new technologies come out with better theories and materials. Linhart and his co-workers [37] announced that the used of calcium phosphate blended with polyesters could overcome the drawbacks. These composites showed good biocompatibility because of: 1. Compensation of acidic release from the polymer through the alkaline calcium phosphate. 2. Participation of carbonated apatite in the biological remodeling process. Less stability of apatite has higher solubility under conditions of acidic resorption [137, 138]. 3. Physicochemical and biological transformation of calcium phosphate into biological mineral as present in bone. 4. Good adhesion of cells on aliphatic polyesters [139]. 5. Better distribution of degradation products through the microporous structure, thereby avoiding accumulation of acidic or harmful products. G. Benzyl ester of hyaluronic acid Benzyl ester of hyaluronic acid, also known as HYAFF-11, has shown good degradation rate and the products of degradation are non-toxic and non carcinogenic [140]. It undergoes hydrolytic degradation via ester bonds in the absence of enzymatic activity with degradation times varying from 1-2 weeks to 2-3 months, depending on the degree of esterification. The de-esterified HYAFF-11 is more hydrated, soluble and resembles native hyarulonic acid [141, 142]. HYAFF-11 has been studied for the use in vascular grafts [143, 144] and for bone tissue engineering [145, 146]. Hyaluronic acid containing fibroblast growth factor is undergoing clinical trial as a synthetic bone graft to accelerate bone fracture healing [30]. Sanginario and co-workers [147] worked on HYAFF-11 reinforced with α-tri calcium phosphate in the quantity of (14/86 wt%, respectively) in the form of hydrogel. The highest compression strength was in the range of MPa, compared to pure HYAFF-11 (3.0 MPa). This values are in the range of cancellous bone, thus, HYAFF-11 can be utilized bone fillers in orthopaedic and dental applications. A. Chitosan III. NATURAL BIODEGRADABLE POLYMERS Chitosan derived from chitin, is a natural biopolymer originating from crustacean and cell wall of fungi. It is known as linear polysaccharide, composed of glucosamine and N- acetyl glucosamine ratio being referred as the degree of deacetylation, which was found in marine environment. Depending on the source and the processing method, its M w may range from 300 to 1000 kda. Chitosan is normally insoluble in water and aquoues solutions above ph 7, but, in diluted acid which ph less than 6, the protonated free amino group on glucosamine facilitate solubility of that molecule [ ]. Chitosan has been found to be non-toxic after oral administration in humans and has FDA approval. Enzymes, such as chitosanase, lysozyme and papain are known to degrade chitosan in vitro [163]. The in vivo degradation of chitosan is primarily due to lysozyme and takes place through the hydrolysis of the acetylated residues. The rate of degradation of chitosan inversely depends on the degree of acetylation and crystallinity of the polymer [164]. The highly deacetylated form exhibits the lowest degradation rates and may last several months in vivo. Apart from this, chemical modification of chitosan can significantly affect its solubility and degradation rate [164]. Chitosan is biocompatible, biodegradable, and can be molded into porous structures (allows osteoconduction) [165]. Several studies have focused on the use of chitosan /calcium phosphate composite for this purpose. Chitosan /nanocrystalline calcium phosphate scaffolds characterized by a relatively rough surface and approximately 20 times greater area/unit mass than chitosan scaffold indicate increased adsorption, improved cell attachment for bone regeneration [166]. In the development of chitosan/nano-ha scaffold conducted by Thein-Han and his co-workers [167], high M w chitosan scaffolds were examined by having superior mechanical properties compared to medium M w chitosan scaffolds (Table 8). In addition, the significant increased in compression modulus of the nano-ha/chitosan composite was attributed to the good dispersion of nano-ha and strong interaction between chitosan and nano-ha to formed a composite. In the nanocomposite scaffold, the organic chitosan network uniformly covered the HA nanoparticles, interactions may have occurred similar to those occurring components of bone. Possible interactions involve Ca and PO 4 charged + species of HA with NH 3 of chitosan. However, the mechanical properties of the composites obtained were inadequate for bone regeneration. Chitosan with incorporation of collagen (CCS) composite microgranules were fabricated as bone substitutes for the purpose of obtaining high bone forming efficacy. The microgranules have the flexibility to fill various types of defect sites with closer packing. The interconnected pores formed spaces between the microgranules, which allowed new bone ingrowth and vascularization [168, 169]. Like any other biodegradable polymers, mechanical strength of CCS significantly improved after adding HA or any other calcium phosphate components (Table 8) [170]. Other than composite scaffold, chitosan also appeared in the form of membrane which was fabricated with silica xerogel [171]. Their potential applications in guided bone regeneration (GBR) were investigated in terms of bone regeneration ability. The in vivo study in a rat calvarial model demonstrated significantly enhanced bone regeneration using the chitosan /xerogel

13 International Journal of Basic & Applied Sciences IJBAS-IJENS Vol: 12 No: membrane compared to that of using the pure chitosan one. Histomorphometric analysis performed 3 weeks after implantation revealed a fully closed defect in the hybrid membrane, whereas there was only 57% defect closure in the chitosan membrane. Another scientific studies of modified chitosan structure, sulfated chitosans with varied sulfate group; the M W and sulfur content were tailored to investigate the effects on the bioactivity of BMP-2 [172]. Low dose of synthetic sulfated chitosan (structure shown in Fig. 3) proved to stimulate osteoblast differentiation induced by BMP-2 in vitro and ectopic bone formation in vivo. Another foundation of chitosan modification is to allow water solubility in chitosan matrix, done by Gomez d Ayala and co-workers [173] (structure shown in Fig. 4). The transformation of chitosan into N-succinylchitosan incorporated with 50 wt% of hemihydrates calcium sulphate resulted in compression strength and Young s Modulus of 0.80 MPa and 14.7 MPa, respectively. Fig. 4. Succinylation reaction of chitosan [173]. Fig. 3. Chemical structures of chitosan modified with different sulfate groups. 2-N-Sulfated chitosan, 2SCS; 6-O-Sulfated chitosan, 6SCS; 2-N,6-O-Sulfated chitosan, 26SCS [172].