PLGA-CaP composites for orthopedic applications: A review

Acta Biomaterialia 8 (2012) 1999–2016

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Review

Fabrication aspects of PLA-CaP/PLGA-CaP composites for orthopedic applications: A review Huan Zhou a,⇑, Joseph G. Lawrence a, Sarit B. Bhaduri b,c a

Department of Bioengineering, University of Toledo, Toledo, OH 43606, USA Department of Mechanical, Industrial and Manufacturing Engineering, University of Toledo, Toledo, OH 43606, USA c Department of Surgery (Dentistry), University of Toledo, Toledo, OH 43606, USA b

a r t i c l e

i n f o

Article history: Received 9 September 2011 Received in revised form 14 December 2011 Accepted 25 January 2012 Available online 2 February 2012 Keywords: PLA PLGA CaP Composite Fabrication

a b s t r a c t For several decades, composites made of polylactic acid-calcium phosphates (PLA-CaP) and polylactic acid-co-glycolic acid-calcium phosphates (PLGA-CaP) have seen widespread uses in orthopedic applications. This paper reviews the fabrication aspects of these composites, following the ubiquitous materials science approach by studying ‘‘processing–structure–property’’ correlations. Various fabrication processes such as microencapsulation, phase separation, electrospinning, supercritical gas foaming, etc., are reviewed, with specific examples of their applications in fabricating these composites. The effect of the incorporation of CaP materials on the mechanical and biological performance of PLA/PLGA is addressed. In addition, this paper describes the state of the art on challenges and innovations concerning CaP dispersion, incorporation of biomolecules/stem cells and long-term degradation of the composites. Ó 2012 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.

1. Introduction Bone, a critical component of the skeletal system in the human body, works as the site for load bearing and mineral metabolism. It is a dynamic system with physiological remodeling, removal, and replacement capabilities throughout life [1]. Due to this self-healing property, small defects can be easily treated by bone itself. However, in the case of extensive bone defects, such as bone fracture, this self-healing mechanism for the reconstruction of large bone segments is not feasible. Nowadays, autografts, allografts and xenografts are commonly used for the treatment of bone fracture. Even though these bone grafts are reported to achieve good clinical results under certain conditions, their applications are limited due to infection, disease transfer, non-unions, rejection, lack of dependable graft supplies and high harvesting cost [2–4]. In recent years, synthetic grafts have been developed as substitutions of bone grafts, which are primarily used for treatment of bone fracture [5,6]. They are readily available and can be processed into different shapes for specific orthopedic applications with excellent mechanical and biological performance with no risk of infection and disease transfer [4–10]. Natural bone is a complex composite, mainly constituted of biological apatite and molecules of collagen. In this system, collagen serves as the matrix for cell growth and tissue repair, while apatite ⇑ Corresponding author. Tel.: +1 419 530 6081; fax: +1 419 530 8076. E-mail address: [email protected] (H. Zhou).

serves as the inorganic phase to improve mechanical strength and regeneration of bone [11]. The higher tensile strength and fracture toughness of bone are attributed to the tough and flexible collagen fibers reinforced by apatite crystals. In recent years, fabrication of inorganic–organic composites mimicking the composite nature of real bone has attracted research interest in the field of bone regeneration. In such composites, inorganic additives with good compressive strength reinforce the elastic polymer phase to generate bioactive materials with improved mechanical performance and degradation profiles [7,12]. Materials used for such bone-mimicking composites are usually biopolymers and inorganic ceramics. Biopolymers can be natural (collagen, hyaluronic acid, etc.) or synthetic (polyesters, etc.). Natural biopolymers occur in nature and can be extracted; and synthetic biopolymers are industrially synthesized for biomedical applications. Even though natural biopolymers favor bone cell attachment and biochemical response in specific situations, there are concerns regarding antigenicity and immunogenicity, potential disease transmission, sourcing, poor mechanical properties and lack of controlled biodegradability. As compared to natural biopolymers, synthetic biopolymers can be tailored to control their properties, such as degradation rate and mechanical strength. They can be easily processed into composite scaffolds with desirable morphological features conducive to tissue ingrowth [8,13,14]. Inorganic ceramics mainly include calcium phosphate (CaP) ceramics and bioactive glasses [15]. Most of these are bioactive and can induce strong bone bonding between implanted materials

1742-7061/$ - see front matter Ó 2012 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved. doi:10.1016/j.actbio.2012.01.031

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H. Zhou et al. / Acta Biomaterialia 8 (2012) 1999–2016

and bone tissues in order to withstand shear and distraction loads [15]. The family of CaP ceramics constitute the predominant bonesubstitute materials. The main driving force behind the use of CaP ceramics is their chemical similarity to the mineral component of hard tissue in the body [11,16]. The complete list of important CaP materials has been well summarized by Dorozhkin et al. [11,17]. Unfortunately, as for any ceramic material, CaP ceramics lack the necessary mechanical and elastic properties of the calcified tissues and exhibit high brittleness, and thus their mechanical performance after implantation is limited [7,13]. Incorporation of CaP phase into biopolymer matrices such as polylactic acid (PLA) and its co-polymer poly (lactic acid-co-glycolic acid) (PLGA) can improve mechanical performance as compared to CaP alone [7,18,19]. Investigations on composites composed of CaP phases and synthetic biopolymers for orthopedic applications appear to be a popular topic in biomaterials research. The efforts include variation in composition, fabrication methods and clinical evaluation [7,13,15]. In this paper, we focus on and emphasize the fabrication aspects of PLA-CaP and PLGA-CaP composites from a materials engineering perspective. (Hereinafter these two individual composites will be abbreviated as PLA-CaP/PLGA-CaP composites.) The motivation for this paper comes from the fact that PLA-CaP/ PLGA-CaP composites are a rather narrow but important focus area in biomedical engineering. It is noted that research on PLA-CaP/ PLGA-CaP composites has been vigorous. A diverse range of fabrication techniques has resulted in a wide range of desirable properties and microstructures. Yet, to the best of our knowledge, a comprehensive review on the fabrication aspects is lacking. In preparing this paper, the authors are guided by the ubiquitous materials engineering approach of considering ‘‘processing–structure–property’’ correlations. Emphasis is placed on how a fabrication process dictates the development of the microstructure, which, in turn, results in enhancement of properties, leading to potential applications. The paper, therefore, includes all of the above topics: beginning with materials issues and ending with biomedical applications.

2. Material aspects of PLA-CaP/PLGA-CaP composites Among the important synthetic biopolymers, polyesters play the most important role. Fully biodegradable synthetic polyesters such as PLA, polyglycolic acid (PGA), polycaprolactone (PCL) and polyhydroxybutyrate (PHB) are commercially available. Among these polyesters, studies on PLA and PGA are predominant (Table 1). PLA (Fig. 1a) is a thermoplastic, biodegradable, biocompatible, synthetic polymer with high strength and modulus that can be made from renewable resources for use in either the industrial packaging or medical device markets [20]. It can be easily processed into shapes such as screws, pins and plates for orthopedic applications and fabricated into scaffolds for replacement and regeneration of tissues, or devices for controlled delivery of biomolecules

Table 1 Search results in Science Citation Index expanded when using different key words. Key words

Literature No.

Polylactic acid Polyglycolic acid Polycaprolactone Polyhydroxybutyrate Polylactic acid + bone Polyglycolic acid + bone Polycaprolactone + bone Polyhydroxybutyrate + bone

13,916 9500 7977 1360 1652 1469 717 95

Fig. 1. Structure of (a) PLA, (b) PGA and (c) PLGA.

[1,14,21,22]. Its stereochemical structure can easily be modified by polymerizing a controlled mixture of the L- or D-isomers (lactate) to yield high molecular-weight amorphous or crystalline polymers [20]. PLA has a slow degradation rate of between 10 months and 4 years due to the hydrophobic methyl group in the backbone. It is degraded by simple hydrolysis of the ester bond at a rate that depends on the molecular weight, isomer ratio and crystallinity. PGA (Fig. 1b) is a rigid thermoplastic material, with high crystallinity, modulus and strength. PGA can be synthesized via polymerization of glycolic acid. It has a faster degradation rate (6–12 weeks) as compared to PLA. Due to its high degree of crystallinity, it is insoluble in many common organic solvents, making the solution processing of PGA non-trivial [23]. Therefore, many known solvent-based composite fabrication techniques are not widely used to prepare PGA-CaP composite. As an alternative, PLGA (Fig. 1c), the co-polymer of PLA and PGA, is used in order to combine the advantages of both polymers. It is synthesized by means of random ring-opening co-polymerization of monomeric glycolic acid and lactic acid. The degradation rate of PLGA is much faster than that of PLA due to the component glycolic acid in the backbone, and in addition the degradation rate can be adjusted by varying the amounts of glycolic acid and lactic acid [24]. PLGA can be easily fabricated into different structures for bone implant applications [14,23]. It is important to note that PLA and PLGA are the only synthetic and biodegradable polymers with an extensive FDA approval history [7]. More detailed information on PLA, PLGA and PGA is available in the literature [20,24,25,14,26]. An important drawback with these biopolymers is that degradation of PLA and PLGA generates acidic products that can lower of the local solution pH. The resultant acidic medium accelerates further degradation in an autocatalytic manner, which triggers inflammatory and foreign body reactions in vivo [27–29]. In addition, their hydrophobic surface property lowers cell adhesion and eventually their poor mechanical strength limits their application in load-bearing situations [30,31]. These limitations can be overcome by incorporating additional inorganic fillers or coatings in their matrix. CaP materials, on the contrary, are bioactive and osteoinductive; some are even quite biodegradable. Even so, the major limitation to the use of CaP materials as bone implant is their brittleness. Their poor mechanical behavior is even more pronounced when they are fabricated into scaffolds with high porosity to mimic bone structure. Therefore, CaP materials are commonly used as fillers or coatings with biopolymers rather than being used on their own. The principal data on various phases of the CaP family used for making PLA-CaP/PLGA-CaP composites is outlined in Table 2. Among this family, hydroxyapatite (HA, Ca10(PO4)6(OH)2) and tricalcium phosphate (TCP, Ca3(PO4)2) are the most commonly used CaP ceramics in fabricating PLA-CaP/PLGA-CaP composites [6–8,12]. HA is the principal inorganic constituent of hard tissues and is one of the most stable and least soluble of all CaP materials [17].

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H. Zhou et al. / Acta Biomaterialia 8 (2012) 1999–2016 Table 2 Information on various CaP materials used to manufacture PLA-CaP/PLGA-CaP composites [7,11,32]. Compound Ca/P ratio

Chemical formula

Solubility at 25 °C, log (Ks)

pH stability

Space group

Unit cell parameter

HA

1.67

Ca10(PO4)6(OH)2

116.8

9.5–12

Monoclinic P21/b or hexagonal P63/m

CDHA

85.1

6.5–9.5

a-TCP

1.5–1.67 Ca10 x(HPO4)x(PO4)6 x(OH)2 (0 < x < 1) 1.5 a-Ca3(PO4)2

25.5

N/A

Monoclinic P21/b or hexagonal P63/m Monoclinic P21/a

a = 9.84,214(8), b = 2a, c = 6.8814(7) Å, c = 120° (monoclinic) a = b = 9.4302(5), c = 6.8911(2) Å, c = 120° (hexagonal) N/A

b-TCP

1.5

b-Ca3(PO4)2

28.9

N/A

Rhombohedral R3Ch

DCPD

1.0

CaHPO4
2H2O

6.59

2.0

Monoclinic Ia

DCPA

1.0

CaHPO4

6.9

N/A

Triclinic Pı¯

ACP

1.0–2.2

CaxHy(PO4)z nH2O, n = 3–4.5; 15–20%
H2O

N/A

5–12

N/A

x

HA is usually synthesized via wet-chemical methods or semisolid-state reactions [17,33]. Wet-chemical methods use aqueous solutions of compounds containing Ca2+ and PO43- ions with pH > 7, and are influenced by many factors including the concentrations of the starting salts, mixing sequence and rate, solution pH, reaction temperature, and holding time. Semi-solid-state reactions refer to hydrothermal and hydrolysis of other calcium orthophosphates to synthesize HA. In addition, other techniques such as water-free synthesis, electrochemical synthesis and hydrothermal treatment of coral have also been developed for HA synthesis [17,33–36]. There are variations to the HA synthesis process, including the reduction of particle size from microscale to nanoscale in order to provide a high surface area to volume ratio to improve osteoconductivity and solubility. This reduction in particle size improves the dispersion of HA particles in the PLA or PLGA matrix, which eventually enhances the mechanical performance [37,38]. Another variation calls for the preparation of calcium-deficient hydroxyapatite (CDHA, Ca10 x(HPO4)x(PO4)6 x(OH)2 x (0 < x < 1)) instead of HA [11,17]. The ion vacancies in the lattice structure of CDHA enables other ions to enter the crystal structure. The extent of this substitution depends on the counter-ions from the chemicals used for the preparation (such as Na+, Cl , etc.). These ion-substituted CDHAs (Na+, K+, Mg2+, Sr2+ for Ca2+; CO32 for PO43 or HPO42 ; F , Cl ,CO32 for OH ), when combined with water, form biological apatite which is the main inorganic part of natural bone [39,40]. This makes CDHA a very promising compound which can be used as bone substitutes. Simultaneous addition of calcium- and orthophosphate-containing solutions into boiling water followed by boiling the suspension for several hours, hydrolysis of precursors and biomimetic coating are all methods that can be used to synthesize CDHA materials [17,41–46]. TCP, another CaP material with higher dissolution properties is widely used in the medical field instead of HA [7,11]. TCP can be further classified as a-TCP and b-TCP. b-TCP can be prepared via calcination of precursors (CDHA with Ca/P ratio = 1.5) at temperatures above 800 °C [11,17]. In combination with HA, b-TCP forms biphasic calcium phosphate (BCP) that can also be used as bone substitution material and filler for composite fabrication [47,48]. At temperatures above 1125 °C, BCP transforms into a-TCP, which is more reactive in aqueous systems than b-TCP and can be hydrolyzed to HA [11,17]. The solubility of TCP materials is much higher than HA. However, from a crystallographic point of view, HA is more similar to natural bone tissue apatite than TCP [49]. Other CaP materials such as brushite (DCPD, CaHPO4
2H2O) and monetite (DCPA, CaHPO4) also have the potential of being used to

a = 12.887(2), b = 27.280(4), c = 15.219(2) Å, b = 126.20(1)° a = b = 10.4183(5), c = 37.3464(23) Å, c = 120° a = 5.812(2), b = 15.180(3), c = 6.239(2) Å, b = 116.42(3)° a = 6.910(1), b = 6.627(2), c = 6.998(2) Å , a = 96.34(2)°, b = 103.82(2)°, c = 88.33(2)° N/A

fabricate PLA-CaP/PLGA-CaP composites. However, there are few relevant reports in this area of research in the literature [7,50,51]. One reason is the acidic property of HPO4- in these materials, which in turn increases the local pH during the degradation process, thus limiting their use in the PLA/PLGA matrix. Amorphous calcium phosphate (ACP) is another potential CaP candidate for PLA-CaP/PLGA-CaP composites. ACP has attracted increasing interest due to its high solubility and excellent remineralization ability, as compared to other CaP materials. It has been widely regarded as the metastable precursor phase for the subsequent formation of CaP in biological organisms, playing a critical role in the process of tissue mineralization. Generally, ACP can be synthesized via two routes: one using a low-temperature method based on the double decomposition of a calcium and phosphate salts in aqueous medium, water/alcohol or alcoholic medium [52–56], and the other using a high-temperature method involving rapid quenching of molten CaPs [52,57–59]. However, using the high-temperature route it is mot possible to control the composition of synthesized ACP materials and impurities such as contaminated elements or other crystalline phases [52]. Therefore, the low-temperature route is commonly preferred in order to control the composition and phase characteristics of the as-synthesized ACP.

3. Fabrication aspects of PLA-CaP/PLGA-CaP composites Bone can be divided into cortical bone (5–13% porosity) and cancellous bone (30–90%) based on porosity of the structure [60]. Human cortical bone exhibits mechanical and physical properties such as a compressive strength of 130–180 MPa), a flexural strength of 135–190 MPa, a tensile strength of 50–151 MPa and an elastic modulus of 12–18 GPa. Cancellous bone, in comparison exhibits a compressive strength of 4–12 MPa, a tensile strength of 1–6 MPa and an elastic modulus of 0.1–0.5 GPa [60–68]. Careful consideration of the bone type and mechanical properties are needed for synthetic bone implants. For cortical bone applications, a suitable stiffness of the implant is needed in order to provide stability for load bearing but not to cause strain shielding. On the other hand, for cancellous bone applications, the 3-D structure is more important for aesthetic reasons [7]. There are a number of techniques available for the fabrication of composites composed of biopolymers and CaP materials. The composites can be in the form of microspheres, film, disk, coatings or porous scaffold for cortical or cancellous bone applications. The

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Fig. 2. Fabrication techniques can be separated into two groups for cortical or cancellous bone applications based on related product porosity.

details have been illustrated in a number of published reviews [7,8,12]. In this section, we focus only on the fabrication of PLACaP/PLGA-CaP composites and the related structure–porosity correlation (Fig. 2). Fabrication techniques can be separated into two distinct groups for cortical or cancellous bone applications based on related product porosity. The various fabrication techniques are presented and discussed in the subsequent paragraphs.

These fabricated PLA-CaP/PLGA-CaP microparticles show the properties necessary to work as fillers or delivery systems for bone regeneration [47,48,70,77,78,80,81]. However, they cannot be directly used for load-bearing applications since they comprise loose particles. This limitation can be overcome by incorporating these particles into a structure which has more integrity [82–86]. 3.2. Solvent casting

3.1. Microencapsulation process The microencapsulation process, in which the removal of the hydrophobic polymer solvent is achieved by evaporation, has been widely reported in recent years for the preparation of microspheres and microcapsules based on biodegradable polymers and copolymers of hydroxy acids [69]. It is widely applied to entrap agents such as drugs into polymer structures for long-term release. As an alternative to standard drugs, CaP materials can also be incorporated into PLA/PLGA microspheres as cores or fillers to form PLA-CaP/PLGA-CaP composites. The most widely used technique exploits the properties of emulsions. The protocol calls for adding phase(s) containing CaP and PLA/PLGA to stabilize aqueous medium to form oil (o)/water (w) or w/o/w emulsion systems with stirring to from microspheres [70–72]. Polyvinyl alcohol (PVA) is the usual stabilizer used to control the size distribution of fabricated microspheres. As an alternative to conventional emulsions, new approaches have been developed to produce PLA-CaP/PLGA-CaP microspheres. Nagata et al. developed surfactant-free emulsion solvent evaporation method by adding (NH4)2HPO4 solutions to a mixture of Ca(CH3COO)2 and PLA [71,73,74]. In this study, the COOH end-groups of PLA play an important role as nucleation sites for calcium phosphate precipitate at the o/w interface and the resultant calcium phosphate plays a role as a stabilizer for microsphere fabrication [71,73,74]. In contrast, fabrication of HA-coated PLA microspheres using nano-HA particles instead of precursors ((NH4)2HPO4 and Ca(CH3COO)2) as particulate emulsifiers was reported by Fujii et al. [75]. In their process, nano-HA was initially dispersed in water at pH 6.5 and then the HA-dispersed aqueous solution was handshaken with PLA-CH2Cl2 solution. The HA-coated PLA microspheres were prepared by the evaporation of CH2Cl2 from the emulsion (Fig. 3). Xu et al. fabricated negatively charged PLGA microspheres using the anionic surfactant sodium dodecyl sulfate (SDS), followed by depositing HA coating onto microspheres via a dual constant composition method [76]. The coating process is sufficiently controllable and efficient (1–6 h) to avoid potential PLGA degradation. As an alternative to conventional emulsion to produce microspheres, PLGA-coated BCP particles using a solvent/non-solvent system were produced by Uskokovic´’s group [47,48,77–79]. In their experiment, chemically synthesized granules or nanoparticles of BCP were added into completely dissolved PLGA. After mixing, methanol was added to the solution. After solvent evaporation with stirring, BCP particles were fully covered by PLGA.

Solvent casting is a conventional technique to fabricate polymer film based on organic solvent evaporation, and can also be applied to prepare polymer-based composite structure. In this process PLA/ PLGA and CaP are mixed or self-assembled in solvent(s), and then cast into molds such as dishes, glass slides or plates after solvent evaporation with CaP randomly dispersed in the polymer matrix (Fig. 4) [87–91]. Furthermore, salt leaching can be combined with solvent casting to tailor pores on the surface. Kothapalli et al. reported the fabrication of PLA/HA composites with 90% porosity and interconnectivity using a solvent casting/salt leaching method [92]. However, salt leaching assisted solvent casting can only be used to fabricate 2-D or very thin 3-D structures. Otherwise, salt particles may remain trapped in the polymer matrix without exposure to aqueous environment to create porosity. There are two concerns as regards using the solvent casting method to produce PLA-CaP/PLGA-CaP composites: (i) this process uses toxic organic solvents; and (ii) during solvent evaporation, CaP particles can spontaneously precipitate from the polymer solution due to poor affinity and can cause non-uniform dispersion of CaP in PLA/PLGA matrix. 3.3. Phase separation The phase-separation technique can be applied to produce 3-D porous scaffolds. This process, first reported in 1995, is based on the fact that once a thermal treatment is induced, a homogeneous polymer solution becomes thermodynamically unstable and tends to separate into a multiphase system including a polymer-rich phase and a polymer-lean phase to lower the free energy [88]. After removal of the solvent, the polymer-rich phase solidifies to form the structure, while the polymer-lean phase becomes pores [93]. CaP particles can be added to the polymer system, and dispersed in the matrix after solvent processing (Fig. 5). Nejati et al. fabricated PLA/HA matrix via phase separation using needle-like nanosized HA particles as filler; they achieved a scaffold porosity of up to 85% with macropore diameters in the range of 64–175 lm (Fig. 6) [41]. Fabrication of PLGA/nano-HA composite scaffold via freezedrying phase separation using different PLGA solution concentrations and HA contents was reported by Huang et al. [94]. In their report, a high PLGA solution concentration decreased the porosity of the fabricated scaffold because the increase in viscosity caused by the high concentration of PLGA solution prevented phase separation; additionally, the pore structure became increasingly

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Fig. 3. (a) SEM image of the HA nanoparticle-coated PLLA microspheres prepared after evaporation of CH2Cl2 from the CH2Cl2 solution of PLA-in-water emulsion. White arrows indicate free HA nanoparticles. (b) Magnified SEM image showing a single microsphere. Inset shows the surface morphology of the microsphere. (c) TEM image of an ultrathin cross-section of the microspheres. (d) Magnified TEM image of the ultrathin cross-section showing HA nanoparticles in coated PLLA microspheres that form the shell [75]. From Fujii S et al. Langmuir 2009;25:9759–9766. Reproduced with kind permission from the American Chemical Society.

3.4. Electrospinning

Fig. 4. Schematic of solvent casting: PLA-CaP/PLGA-CaP is mixed and casted in molds, after solvent(s) evaporation; CaP particles are randomly dispersed in the polymer matrix.

irregular and some pores also closed with increasing HA content. Yang et al. fabricated porous PLGA/b-TCP scaffold using phase separation and porogen leaching [95]. It was observed that in the scaffold, the process of porogen leaching generated pores with diameters of 75–400 lm; pores with diameters of less than 20 lm were produced by phase separation. Additionally, the toxic solvent chloroform was extracted by ethanol to reduce scaffold toxicity. Work done by Wang et al. illustrated the fabrication of nanofibrous PLA/HA composites (100–750 nm) with interconnected microporous structure (1–10 lm), high porosity up to 90% and homogeneous HA dispersion via phase separation in order to mimic the architecture of natural extracellular matrix (ECM) for tissue engineering applications [96]. Similar to solvent-casting method, residual toxic organic solvent is still a concern in this solvent-based technique.

Electrospinning is a simple and versatile technique capable of generating non-woven fibrous mats directly from a variety of polymers and composite materials [97]. The applications of electrospun fibers include filtration, textile manufacturing, catalysts and medical fields. The use of electrospun fibers in biomedical applications has dramatically increased of late because these fibers offer a range of attractive features such as large surface areas, high porosities and ease of incorporation of functional components (drug, gene, enzyme, etc.), making them ideal candidates for tissue engineering applications [97–99]. In the process of electrospinning, an electric force is applied to overcome the surface tension forces of a biopolymer solution loaded in a capillary tube to generate a jet ejected from the tip of the capillary tube, which moves in the direction of the external electric field, elongates according to external and internal electrical forces, and finally gets deposited on a conductive substrate in the form of fibers (Fig. 7) [93,94]. A number of processing parameters (applied voltage, flow rate, working distance, solution conductivity, solution viscosity, solvent(s) property, etc.) can greatly influence the features of electrospun fibers such as surface porosity, fiber dimension, etc. [98,100]. The ability to co-spin PLA/ PLGA with CaP additives offers the possibility to fabricate fibrous PLA/CaP or PLGA/CaP composite scaffolds via electrospinning with high surface area and interconnected channels suitable for tissue regeneration applications. For example, Jeong et al. fabricated PLA/HA nanofibrous scaffolds via electrospinning, and reported that the incorporation of HA can shrink the fiber size from 365 ± 83 to 135 ± 13 nm due to the interaction between HA and PLA and the impact of HA on the mixture viscosity [101]. PLA/ CDHA electrospun fibers were prepared by Zhou et al., who also

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Fig. 5. Schematic of phase separation, the system is composed of thermal treatment for phase separation and solvent removal for phase solidification.

Fig. 6. SEM micrographs of pure PLA and PLA/nano-HA scaffold fabricated via the phase-separation technique: (a and b) pure PLLA and cross-section; (c and d) nHAP/PLLA and cross-section [41]. From Nejati E et al. Mater Sci Eng C 2009;29:942–949. Reproduced with the kind permission from Elsevier SA.

Fig. 7. A typical electrospinning set-up, composed of syringe, needle, power supply, collector and a pump.

reported fiber size shrinkage related to CDHA incorporation [102]. Additionally, they also showed that the surface feature of the fibers can be adjusted from smooth to porous by controlling the solvent composition. In these composites fabricated via co-spinning, CaP particles are usually embedded along the polymer fiber matrix. As an alternative, Cui et al. fabricated PLA fibers with calcium

nitrate entrapment and incubated fibers in phosphate solution to form CaP on the PLA matrix instead of co-spinning PLA with CaP to improve CaP dispersion uniformity in the PLA matrix [103]. In addition, Touny et al. recently introduced a novel reactive electrospinning process to fabricate PLA/monetite fibers [50]. In their process, PLA was initially mixed with H3PO4 in organic phase, and

H. Zhou et al. / Acta Biomaterialia 8 (2012) 1999–2016

Ca(OH)2 powders were well dispersed in another organic phase and then slowly poured into the PLA + H3PO4 mixture with vigorous stirring. This mixture was finally used to fabricate fibers via electrospinning. Monetite was produced via the reaction of precursors during electrospinning, which were uniformly dispersed in the PLA matrix. In addition, the heat and water produced during the reaction induced pore formation on PLA surface to further improve the surface area of the fibers. The limitations of the electrospinning process are the use of toxic solvent(s) that may cause adverse effects to cells if not removed completely and difficulty in controlling pore size and pore shape [8]. Also, there are several processing parameters involved in the electrospinning process that need to be optimized in order to obtain the desired composite. This optimization process is cumbersome and time consuming when compared to solvent casting or other fabrication techniques. 3.5. Coating technique The coating technique can be categorized into two types: depositing CaP coating on PLA/PLGA matrix, or depositing PLA/ PLGA on CaP scaffold. The best-known method of CaP deposition is the biomimetic coating process which involves soaking the matrix in simulated body fluid (SBF) to deposit bone-like apatite (CDHA) on the targeted material’s surface (Ti, biopolymer, bioceramics, etc.) [44,104]. Zhang et al. soaked PLA scaffold into SBF to deposit CDHA coatings on PLA surface to enhance osteoinductive properties [43]. Labberzadeh et al. incubated PLGA scaffold in SBF to deposit a CDHA layer, as a site for protein adsorption and release studies. They observed that the protein release pattern was similar for PLGA and mineralized PLGA scaffolds, but precipitation of the CDHA layer on PLGA led to enhanced protein adsorption and slower protein release [81]. A drawback of this process is that the biomimetic coating is time consuming (about 21 days). As an alternative, production of nano-HA-coated PLA composite by soaking alkali-treated PLA matrix into nano-HA/ethanol suspension for only 1 h was achieved by Yanagida et al. [105]. In their report, carboxyl groups were introduced to the PLA surface by alkali treatment and the interactions of these groups with calcium ions on HA

2005

facilitated HA coating deposition on the PLA surface (Fig. 8). The aerosol deposition method (ADM) is also a novel way to deposit HA film on PLA substrates [106,107]. In this process, PLA substrates are coated with HA particles that have been deposited by means of high-velocity helium carrier gas. Subsequently, HA-coated PLA substrates are heat treated to prevent HA coating from being peeled through washing with water. The process of depositing PLA/PLGA to CaP scaffold involves simply soaking scaffolds in PLA/PLGA-containing organic phases followed by drying to remove the solvent. Miao et al. deposited PLGA coating to HA/b-TCP scaffold, and observed that the macroporosity of the scaffolds after infiltration and coating with the PLGA was slightly decreased due to the observed thin polymer coating present on the strut surfaces [18]. Zhao et al. deposited PLA/HA composite coatings on HA macroporous scaffold using an alternative method [19]. They injected PLA/HA solution into the porous HA scaffolds by a vacuum pump and eventually removed extra solutions by centrifugation to ensure interconnectivity of the scaffolds. The porosity of the scaffold slightly decreased from 85.0 ± 3.6% to 82.0 ± 2.7% after coating while both the mechanical and biological properties increased. The results demonstrated that centrifugation showed great potential to prepare polymer-lined scaffolds with improved 3-D interconnectivity and reinforced adhesion force between the polymer coating and the substrate. This technique does not involve optimization of too many parameters when compared to either solvent casting or electrospinning. However, fabrication of different sizes and shapes may still be limited. 3.6. Supercritical gas foaming Supercritical gas foaming is a solvent-free fabrication technique to produce scaffolds with controlled architecture and properties. The process utilizes the nucleation and growth of gas bubbles (internal phase) dispersed throughout a polymer (continuous phase) by sudden depressurization. The porous structure of polymer is formed when the dispersed gas phase is removed from the continuous phase of polymer [108–110]. In the process of fabrication of PLA-CaP/PLGA-CaP scaffold, a homogeneous mixture of

Fig. 8. SEM images of the surfaces of (a and b) non-treated and (c and d) alkali-treated PLA fabrics immersed in HA ethanol suspension: (a and c) lower magnification and (b and d) higher magnification; H: HA nanocrystal [105]. From Yanagida H et al. J Biosci Bioeng 2009;108:235–243. Reproduced with the kind permission from the Society for Biotechnology, Japan.

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Fig. 9. Schematic of supercritical gas foaming: melted polymer–CaP–gas solution with a saturation pressure and temperature above the melting temperature of the polymer undergoes sudden gas release and cooling to produce porous polymer matrix with CaP dispersion.

CaP particles and biopolymers (PLA/PLGA) is initially prepared with thermal treatment (usually melt-extrusion) in a pressure chamber followed by CO2 diffusion to form a melt polymer–gas solution with a saturation pressure and temperature above the melting temperature of the polymer (Fig. 9). Foam expansion is then carried out by sudden gas release with additional cooling, which increases polymer viscosity and creates a permanent foam architecture by solidification and recrystallization of the polymer. The rate of cooling was reported to have a significant effect on the porous structure. Cooling too rapidly fixed small, closed pores, whereas cooling too slowly did not allow freezing of the structure, resulting in the structure collapsing [109,110]. Mathieu et al. fabricated PLA/b-TCP and PLA/HA scaffolds via supercritical gas foaming in order to mimic natural bone with a porosity of over 80% [110]. The characterization results demonstrated that the as-fabricated 3-D scaffold was morphologically anisotropic with pores oriented in the foaming direction, which making the axial modulus of the scaffolds up to 1.5 times greater than the transverse modulus. This composite also showed viscoelastic behavior similar to natural bone with increased modulus for higher strain rates. These results indicated that supercritical gas foamed composite scaffolds have suitable architecture and properties for bone tissue regeneration applications (Fig. 10). Comparison of supercritical gas foamed PLGA/HA composite and solvent-cast PLGA/HA composite was conducted by Kim et al. [111,112]. First, supercritical gas foaming eliminated the application of toxic organic solvent. Second, the foamed PLGA/HA composite showed interconnected porous structures without a skin layer and exhibited superior enhanced mechanical properties when compared to the scaffolds that were fabricated via solvent casting. Third, foamed PLGA/HA composite had abundant exposed HA particles on scaffold surface, which were expected to promote bone regeneration. 3.7. Melting Melt extrusion is based on the low melting point of PLA/PLGA and the high melting point of CaP materials (Fig. 11). For example, Damadzadeh et al. used different HA contents and particle sizes to produce PLA/HA and PLGA/HA composites via extrusion [113]. As alternatives to melt extrusion, processes such as forging, pressing, sintering and microwave irradiation based heating were also applied to produce dense composites. Shikinami et al. extruded PLA/HA granules into a thick billet, which was forged at 103 °C into thin billets and cut into devices with different shapes and size such as screws, pins and plates [114]. The composites satisfied initial mechanical strengths and maintained these for as long as required

(up to 24 weeks) for internal bone fixation devices. Mohn et al. prepared composites of PLGA and spherical amorphous TCP nanoparticles via blending directly or through a two-step approach, including solvent casting and extrusion followed by hot pressing. The latter route yielded good particle dispersion while blending alone led to inhomogeneous mixtures [115]. Ignjatovic´ et al. developed a cold/ hot pressing approach to manufacture PLA/HA composites with different densities, porosities, compressive strengths and elastic moduli by controlling pressing parameters such as temperature, pressure and time [116–120]. A sintering method to produce PLA/ calcium metaphosphate (CMP) composite for bone regeneration was reported by Jung et al. [121]. In their process, mixtures of PLA, CMP and salt particles were pressed into discs followed by heat treatment at 210 °C for 30 min to generate a scaffold with a homogeneously interconnected porous structure without a skin layer. This resulted in a narrower pore size distribution, higher mechanical strength and higher exposure of CMP to the surrounding environment in comparison with scaffolds made by a solvent-casting method. Sintered microsphere scaffolds have also been reported since 2001. These involve a multifunctional scaffold with a connected pore system and mechanical properties in the mid-range of cancellous bone [122]. In this process, pre-prepared PLGA microspheres were poured into a mold and heated to above the glass transition temperature of PLGA for 24 h. Upon heating, the polymer chains were activated to intertwine with neighboring polymer chains and thus formed contacts between neighboring microspheres. Following heating, the matrix was allowed to cool down to room temperature, which eliminated the mobility of the polymer chains. As a result, those contacts formed with adjacent spheres during high-temperature polymer fusion became fixed to provide mechanical support. PLGA/CaP sintered microsphere scaffold has also been fabricated using PLGA/CaP microspheres [82–84]. As the sintering temperature or time increased, the compressive mechanical properties significantly increased while the porosity of the scaffolds decreased. Consequently, the spherical morphologies of the microspheres were lost [82]. Jin et al. introduced a novel method to generate PLGA/b-TCP composites through polymerization using microwave energy [123]. Initially, L-lactide and glycolide were melted completely at 140 °C and b-TCP particles were dispersed into polymer solution with stannous octanoate as an initiator. The mixture was then polymerized using microwaves for 2 h with an irradiation power of 100 W (150 °C) followed by subsequent cooling. In these composites, the b-TCP powders were well dispersed in the PLGA matrix up to 30 wt.% content. This behavior was attributed to the in situ heating induced by the microwave energy. 3.8. Solid free fabrication/rapid prototyping Solid free fabrication (SFF), or rapid prototyping (RP), is a series of digital information controlled 3-D scaffold fabrication techniques, which converts computer-aided design (CAD) into a specific cross-sectional format, expressing the 3-D model as layers. The digital layered 3-D model is then manufactured layer by layer using a SFF machine with each layer representing the cross-section of the digital model at specific levels. The applications of SFF include design visualization, metal casting, architecture, geospatial, healthcare and entertainment/retail. SFF has been applied to fabricate complex 3-D scaffolds with delicate architecture for tissue engineering applications [8124,125]. In general, SFF methods can be grouped into three categories: melt-dissolution deposition, particle bonding, and indirect [126]. The melt-dissolution deposition method is a process wherein each layer is created by extrusion of a strand of material through an orifice as it moves across the plane of the layer cross-section. Once the deposited material cools, it solidifies and attaches to the previous layer [127]. In the particlebonding method, particles are selectively bonded to a thin layer

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Fig. 10. Similarity of cancellous bone structures and composite foam macrostructures [110]. From Mathieu LM et al. Biomaterials 2006;27:925–916. Reproduced with the kind permission from Pergamon Press.

Fig. 11. Schematic of melting extrusion, melted materials are pushed through a die of the desired cross-section.

of powder material, which are then bonded in multilayers to form a complex 3-D scaffold. During fabrication, the surrounding unprocessed powders support the object of interest. After completion of all layers, the object then is removed from the bed of unbonded powder [127]. The indirect SFF method involves fabrication of a negative mold based on the scaffold design. The scaffold is then cast using the desired materials. Finally, the mold is sacrificed after the scaffold is finally obtained [125,128]. Such techniques enable the user to control both the external and the internal morphology of the final construct. In addition, it also requires less raw scaffold material while increasing the range of materials that can be used. The original material properties are conserved since it does not require further heating [126]. Several review articles and books have summarized and compared SFF scaffold fabrication approaches in detail [124–127,129]. Therefore, this section will focus mainly on recently reported PLA-CaP/PLGA-CaP composite scaffolds fabricated via SFF/RP. Low-temperature deposition manufacturing (LDM) is a type of melt-dissolution deposition (Fig. 12). Its scaffold-building cycle is

Fig. 12. Schematic of low-temperature deposition manufacturing: the computerdriven system builds the scaffold layer by layer directly using a digital model in a low-temperature environment under 0 °C.

performed in a low-temperature environment under 0 °C, so that it can preserve the bioactivity of scaffold materials because of its benign processing conditions [130]. PLA/TCP scaffold has been successfully fabricated via LDM. The LDM system uses a pump nozzle to deposit PLA/TCP dioxane slurry to build scaffolds layer by layer based on 3-D digital models [130]. The as-fabricated scaffolds have high porosity up to 89.6% and mechanical property values close to those of human spongy bone (Fig. 13). Multi-nozzle deposition manufacturing (MDM) is an improved version of LDM. Scaffolds with different properties are fabricated in a low-temperature environment under 0 °C by not only a single-nozzle deposition process, but also by bi-nozzle and tri-nozzle deposition processes [131].

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Fig. 13. SEM images of the cross-section of the PLLA/TCP composite scaffold: (a) low magnification and (b) high magnification [132]. From Sherwood JK et al. Biomaterials 2002;23:4739–4751. Reproduced with the kind permission from Pergamon Press.

Fig. 14. Schematic of 3-D printing: powder particles are delivered to the powder bed and assembled with a liquid binder to form solid 3-D objects one layer in each cycle.

Therefore, it can easily create functionally graded scaffolds with a range of materials and morphologies, where bioactivity are customized and preserved. Yan et al. prepared a bone morphogenetic protein (BMP)-loaded PLGA/TCP scaffold using a tri-nozzle deposition manufacturing [131]. It was composed of two regions: spongy parts with BMP and compact parts. The compact part of the scaffold comprised of small pore size without the BMP which was expected to degrade more slowly and produce more limited bone regeneration than the spongy part with larger pore size. The compact part had superior mechanical properties and provided better support during the process of bone repair. It took a relatively long time of about 1 year for the compact part to be resorbed by the body before the regenerated bone resumed its full function by the tissue regeneration triggered by BMP. 3-D printing (3DP) is a particle-bonding technique, which selectively binds powder particles with a liquid binder to form solid 3-D objects one layer in each cycle (Fig. 14). After the binder is dried, the completed scaffold is retrieved and unattached particles are removed [125]. Sherwood et al. used PLGA/PLA/TCP/NaCl mixture to fabricate 3-D scaffolds via 3DP [132]. In this process, liquid CO2 was used to remove the residual binder. Finally, salt leaching created the micropores. The resulting scaffold comprised three regions with different functions: the upper cartilage region had 90% porosity and was composed of PLGA/PLA, with macroscopic staggered

channels to facilitate homogeneous cell seeding. The lower, cloverleaf-shaped bone portion had 55% porosity and consisted of PLGA/ TCP composite, designed to maximize bone ingrowth while maintaining critical mechanical properties. The transition region between these two sections contained a gradient of materials and porosity to prevent delamination. Roy et al. implanted 3-D printed PLGA/b-TCP scaffold in rabbit calvarial defect for 8 weeks to promote bone healing and observed scaffolds with macroscopic channels [133]. These scaffolds had a controlled gradient of porosity and pore size. The results showed that it had higher percentages of new bone area as compared to scaffold without porosity [133]. Selective laser sintering (SLS) is another particle-bonding technique, which uses a deflected laser beam selectively to scan over the powder surface following the cross-sectional profiles carried by the slice data (Fig. 15) [129]. The interaction of the laser beam with the powder elevates the powder temperature to reach the glass-transition temperature, causing surfaces in contact to deform and fuse together. PLGA/carbonated HA (CHA) scaffold was prepared by Zhou et al. using SLS [86]. It was observed that the degree of fusion of the PLA/CHA composite powder was lower than that of the pure PLA powder, making it easier to remove excessive powder from pores. The increased viscosity of the composite material could explain this. Also, the CHA nanoparticles on the powder surface might act as a barrier against fusion.

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Fig. 15. Schematic of selective laser sintering: powder particles are delivered to the powder bed and assembled using a laser beam to form solid 3-D objects one layer in each cycle.

Fig. 16. PLA-HA scaffold fabricated via indirect SFF: (a) as-fabricated scaffold; (b) 3d design [134]. From Taboas JM et al. Biomaterials 2003;24:181–194. Reproduced with the kind permission from Pergamon Press.

There are few reports on indirect SFF fabrication of PLA-CaP/ PLGA-CaP scaffolds [134]. A sintered HA ceramic scaffold base was fabricated containing a primary and secondary pore network. The primary network was designed for tissue growth into the ceramic body, while the secondary network served as the polymer mold. The secondary network was contiguous with the primary pore network in the manufactured scaffolds. PLA was melt cast into half of the base as outlined in the global pore melt-casting techniques. Ceramic-free regions within the scaffold were created by selective application of a ceramic solvent, yielding a non-blend, discrete composite with mechanical interdigitation (interlocking) of the ceramic and polymer phases. Finally, porous discrete composites that include regions of pure sintered HA, PLA and combinations of the two in the same scaffold were produced (Fig. 16). The polymer wall surrounding the periphery of the polymer half serves to entrap seeded chondroblasts and prevent migration of osteoblasts into the polymer region. Thus, it prevents bone formation in the region of cartilage growth. Conversely, osteoblast migration and bone production are occurring in the HA region. 4. Performance of PLA-CaP/PLGA-CaP composites In PLA-CaP/PLGA-CaP composites, PLA/PLGA is usually referred to as the continuous phase or matrix, while CaP materials are referred to as the dispersed phase or fillers. The continuous phase is responsible for supporting the dispersed materials by maintaining their relative positions. The dispersed phase is usually responsible

for enhancing the physical and biological properties of the matrix [7]. For enhancement of the mechanical properties of the PLA-CaP/ PLGA-CaP composites, CaP works as a reinforcing material to meet performance criteria for a particular application or implant [92,135]. For biocompatibility and bioactivity of the PLA-CaP/ PLGA-CaP composites, incorporation of CaP materials into PLA/PLGA matrix greatly promotes cell proliferation and tissue regeneration [111,136]. 4.1. Mechanical performance As seen in Table 3, the influence of one type of composite manufacturing compared to another on the implant’s mechanical performance is difficult to access. This is because different papers focus on selected mechanical properties (compressive modulus, tensile strength, Young’s modulus, compressive strength, etc.) and many variables (porosity, ratio of CaP, CaP type, etc.). The strengthening effect of CaP in PLA/PLGA matrix can be explained by the fact that the PLA/PLGA matrix is a load transfer medium and thus transfers the load to the intrinsically rigid CaP crystals. However, the incorporation of CaP can result in a decrease in tensile strength of PLA/PLGA matrix due to the brittleness of CaP materials [91]. Another issue is the content of incorporated CaP in the composite. When the CaP content is low, the PLA/PLGA matrix is continuous, which ensures good interfacial bonding with CaP particles and thus results in improved compressive strength. In composites with greater CaP content, the amount of PLA/PLGA

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Table 3 Physical properties of PLA-CaP/PLGA-CaP composites fabricated using different techniques. Techniques

Materials

Architecture

Mechanical Properties

Ref.

Solvent casting

PLA, amorphous-TCP (0–30 wt.%) PLA, HA (0–50 wt.%)

Film

Young’s modulus: up to 2800 MPa (30 wt.% amorphous-TCP)

[87]

Scaffold with porosity 86–92%

[92]

PLGA, HA (0–15 wt.%)

Film

PLA, CDHA (0– 10.5 vol.%) PLGA, b-TCP (0– 60 wt.%) PLA, HA (0–30% wt.%) PLA, HA (0 to 16.7 wt.%) PLA, HA (N/A)

Film

Compressive modulus: up to 9.87 ± 1.8 MPa (50 wt.% HA) Yield strength: up to 0.44 ± 0.01 MPa (50 wt.% HA) Tensile strength: up to 3.9 ± 0.37 MPa (10 wt.% HA) Young’s modulus: up to 16.56 ± 5.96 MPa (10 wt.% HA) Tensile modulus: up to 2.47 GPa (10.5 vol.% CDHA)

Scaffold with porosity 83.5%-91.4%

Compressive modulus: up to 6.64 ± 1.03 MPa (30 wt.% b-TCP)

[95]

Fibrous scaffold with porosity up to 90% Fibers

[96] [101]

PLA, CDHA PLGA, HA, b-TCP

Scaffold with porosity up to 95% Scaffold with porosity up to 87%

PLA, HA PLGA, HA (50 wt.%)

Scaffold with porosity 82.0 ± 2.7% Scaffold with porosity 91 ± 3%

PLA, HA/b-TCP (0– 10 wt.%) PLA/PLGA, HA/b-TCP (0–30 wt.%)

Scaffold with porosity 74–81%

Compressive modulus: up to 0.63 MPa (20 wt.% HA) Young’s modulus: up to 4.711 ± 0.019 MPa (16.7 wt.% HA) Tensile strength: up to 0.262 ± 0.007 MPa (16.7 wt.% HA) Young’s modulus: up to 118 ± 10 MPa (with HA) Tensile strength: up to 2.86 ± 0.24 MPa (with HA) Compressive modulus: approximately 10 MPa (with CDHA) Compressive modulus: up to 665 ± 50 kPa (with PLGA) Compressive strength: up to 66 ± 7 kPa (with PLGA) Compressive strength: approximately 1.7 MPa (with PLA) Compressive modulus: 4.5 ± 0.3 MPa Young’s modulus: 26.9 ± 0.2 MPa Young’s modulus: up to approximately 180 MPa (10% b-TCP)

[113]

Forging

PLA, HA (0–50 wt.%)

Billet

Cold/Hot Pressing

PLA, HA (N/A)

3-D block

Sintering

PLGA, HA (20 wt.%)

Scaffold with porosity 30%

SFF

PLA, TCP (50 wt.%) PLGA, TCP (23– 33.3 wt.%)

Scaffold with porosity up to 89.6% Scaffold with porosity from 50% to 90% in different regions

Young’s modulus: up to 4.5 ± 0.2 GPa (30 wt.% HA and 70 wt.% PLA) Flexural strength: up to 136 ± 5 MPa (10 wt.% HA and 90 wt.% PLGA) Flexural strength: up to 270 ± 4.1 MPa (40 wt.% HA) Flexural modulus: up to 12.3 ± 0.2 GPa (50 wt.% HA)Tensile strength: 154.1 ± 0.8 MPa (PLA) Young’s modulus: up to 2.4 ± 0.9 GPa (50 wt.% HA) Compressive strength: up to 115.3 ± 3.9 MPa (PLA) Compressive modulus: up to 6.5 ± 0.2 GPa (50 wt.% HA) Compressive strength: up to 140 MPa Compressive modulus: up to 10 GPa Compressive strength: up to approximate 4 MPa Compressive modulus: up to 200 MPa Compressive modulus: 60.11 MPa Tensile strength: up to 5.7 ± 1 MPa (33.3 wt.% TCP) Young’s modulus: up to 233 ± 27 MPa (23 wt.% TCP) Compressive strength: up to 13.7 ± 0.8 MPa (23 wt.% TCP) Compressive modulus: up to 450 ± 79 MPa (23 wt.% TCP)

Phase separation

Electrospinning

CaP coating PLA/PLGA coating

Supercritical gas foaming

Melt extrusion

Fibers

Rod-like profiles with a diameter of 2– 2.5 mm

matrix surrounding the CaP particles is lower. This results in poor integrity of the composite, causing low compressive properties [95,137]. From the bone-mimicking perspective, it is very hard to produce PLA-CaP/PLGA-CaP composite or any biocomposites that perfectly match the mechanical properties and highly organized structure of cancellous or cortical bone. Hence, further in-depth research is required on the manufacturing aspect of PLA-CaP/PLGACaP composites in order to make them mimic the skeletal bone structure.

[89] [90]

[138] [43] [18] [19] [111,112] [110]

[114]

[119] [83] [130] [132]

apatite formation on treated PLA surface after 7 days in SBF solution as compared to the flat surface of untreated PLA (Fig. 17) [140]. Zhou et al. also reported a similar phenomenon when incubating PLA/CDHA and PLA electrospun fibers in SBF for 7 days [102]. The results of using PLA-CaP/PLGA-CaP composites in vitro and in vivo are presented in Table 4, which illustrates the increased biocompatibility and efficient bone-forming capacity of PLA-CaP/ PLGA-CaP composites as compared to controlled PLA/PLGA groups. All results confirm that PLA-CaP/PLGA-CaP composites have greater potential for clinical applications.

4.2. Biological performance Bioactivity is referred to the ability of materials to posses direct bone-bonding capabilities once implanted. SBF has been widely accepted as a powerful test to evaluate the bioactivity of the target material [139]. PLA and PLGA have the ability to form apatite once incubated in SBF, because their hydrolysis during biodegradation results in –COOH which can bind to Ca2+, causing further apatite nucleation and growth [43,85]. On considering the degradation rate difference between PLA and PLGA, PLGA is more favored for apatite deposition. However, it is possible to modify the surface of PLA/PLGA to increase the level of the functional group –COOH [101]. The incorporation of highly bioactive CaP can intensify the deposition of bone-like apatite from SBF. Kim et al. bonded HA particles to PLA surface, and reported a considerable level of

5. Challenges 5.1. Mechanical integrity of the composite As mentioned before, uniform dispersion of CaP particles in the composite structure is always a concern during fabrication, especially when high-ratio CaP materials are used as fillers. CaP particles tend to combine together and form aggregates via interlocking, electrostatic attraction, van der Waals forces, liquid bridges and solid bridging [144]. In addition, CaP shows low affinity towards PLA/PLGA. Therefore, CaP particles are poorly dispersed or are easily agglomerated, resulting in uneven distribution within the polymeric matrix. In addition, in solvent-based fabrication processes,

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Fig. 17. SEM surface and cross-section images of the samples after immersion in SBF for 7 days: (a and b) HA-sandblasted PLA and (c and d) untreated PLA. Apatite crystallites were significantly formed in a thin layer on the sandblasted PLA, while no such formation was observed on PLA [140]. From Bae JY et al. Mater Lett 2011;65:2951–2955. Reproduced with the kind permission from Elsevier BV.

Table 4 Results of in vitro and in vivo experiments using different PLA-CaP/PLGA-CaP composites. Composites

Method

Cell culturing or implantation

Period

In vitro experiment PLGA/HA [89]

Solvent casting

PLGA/HA [94]

Phase seperation

PLA/b-TCP [141] PLGA-HA [83]

Electrospinning Sintering

Chondrocytes isolated from joint of porcine rear leg Mesenchymal stem cells from rabbit tibia Human adipose stem cells Human mesenchymal stem cells

6 days (cell number is 1.3 times higher than control group) 2 weeks (significantly higher cell growth and alkaline phosphatast (ALP) activity) 2 weeks (high DNA content and ALP activity) 3 weeks (enhance cell proliferation at early point, but differentiation and mineralization were promoted at later point)

In vivo experiment PLGA/BCP [47]

Emulsion

Implantation in alveolar bones

Emulsion Electrospinning Supercritical gas foaming with salt leaching Supercritical gas foaming Cold/hot pressing SFF

Implantation in rabbit femur Implantation in rabbit calvarial bone Implantation in rat skulls

24 weeks (significantly higher bone density as compared to control group) 4 weeks (mineralized bone with repairing of soft tissue) 4 weeks (mineralized bone) 8 weeks (mineralized bone with lamellar structures and osteoid formation) 12 months (mineralized bone) 12 weeks (good growth in surrounding connective tissue) 8 weeks (mineralized bone)

PLA/b-TCP [70] PLGA/TCP [136] PLGA/HA [111] PLA/b-TCP [142] PLA/HA [143] PLGA/b-TCP [133]

Implantation in sheep femur and tibia Implantation in mice femur Implantation in rabbit calvarial bone

phase separation between CaP and polymer solution occurs when a high content of CaP is used. This is due to lack of affinity of CaP to organic solvents, which results in difficulties during processing and produces uneven dispersion of CaP in the composite structure. The agglomeration of CaP in the polymer matrix can cause a decrease in mechanical strength, since these agglomerations start to act as stress concentration points giving rise to brittle failure with a lower ultimate tensile strength [137]. Application of strong ultrasonication to destroy CaP agglomerates is the most common approach to preparing PLA/CaP or PLGA/CaP mixtures in solvent(s) [135,145,146]. As an alternative, Takamatsu et al. used ball milling to decrease CaP aggregates [147]. In their process, PLA/b-TCP solution was ball-milled at a velocity of 80 rpm at ambient temperature for periods of up to 48 h; the mixtures were then dried and hot

pressed. The Young’s modulus and bending strength of the composites increased with increase in ball-milling time. This is a result of the absence of large b-TCP aggregates in ball-milled composites. In addition, it was reported that the choice of dispersant also influences CaP dispersion in polymer solution. The ability of various conventional organic dispersants (chloroform, dimethylformamide, tetrahydrofuran, acetone, ethanol) to form homogeneous and stable PLA/HA solution was compared by Deng et al. [148]. It was observed that dimethylformamide is an effective dispersant to deaggregate HA particles in solvent-based PLA/HA composite fabrication. The application of surfactant is another approach to improve CaP dispersion. Kim et al. introduced a surfactant, 12-hydroxysteric acid (HSA), to address the problem of HA dispersion and agglomeration [149]. TEM analysis of the as-fabricated PLA/HA electrospun fibers

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revealed uniform distribution of HA powders throughout the PLA polymeric matrix, whereas products using commercial HA powders without HSA showed severe agglomeration of HA in PLA matrix. This resulted in separation of the PLA-rich part from HA agglomerates. The size and morphology of CaP fillers are also important to the composite mechanical strength. Nejati et al. used both needle-like nano-HA and microsized HA to fabricate HA-filled PLA composite scaffolds via phase separation. The mechanical testing results showed that the compressive strengths of the nanocomposite scaffolds were increased to 8.46 MPa. Compared to that, compressive strengths of pure PLA and HA microcomposite scaffolds were 1.79 and 4.61 MPa, respectively [41]. This is attributed not only to the larger interfacial area between nano-HA and PLA matrix, but also to uniform dispersion of nano-HA. On the other hand, it may also be argued that irregular CaP particles are preferred because PLA/PLGA liquid can penetrate into CaP particle surface and can form a mechanical interlock with the particle as it transits back to solid form during composite fabrication [137]. To further promote the adhesion between CaP and PLA/PLGA, formation of strong chemical bonding between PLA/PLGA and CaP has been studied [74,150–160]. The most commonly applied technique is to modify the CaP particle surface. In some polymer/ HA composites coupling agent molecules such as silane, zirconyl salts, hexamethylene, etc., were grafted on HA particles surface through covalent bonds by chemical reaction with hydroxyl groups on the surface of HA [150–152]. Rakovsky et al. modified the surface of CDHA by reaction with hexamethylene diisocyanate (HDI) in a non-aqueous suspension [153]. Composites of CDHA–HDI with PLA were consolidated under high pressure (2.5 GPa) at room temperature; the composites had a compressive strength of approximately 300 MPa. Hong et al. made HA particles surface grafted with PLA via ring-opening polymerization of L-lactide in the presence of stannous octanoate (Sn(Oct)2) as catalyst. The results showed that the grafted PLA molecules greatly improved the adhesive strength between HA and PLA matrix as interconnecting molecules [154,155]. To further enhance bonding, the surface morphology of CaP can also be roughened. Hong et al. produced lenticular carbonated HA particles 30–40 nm in diameter and 100–200 nm in length via precipitation; these particles had a coarse surface with many global protuberances, which could favor interaction with the polymer matrix [156]. However, the grafting rate is limited due to a reduced amount of hydroxyl groups on HA particles. In order to increase the number of hydroxyl groups, HA particles were first grafted with L-lactic acid and then with PLA. Chemical bonding between Ca atoms on HA surface and carboxyl groups of L-lactic acid or PLA were formed with grafting ratios of up to 40% [157]. These grafted CaP particles were also reported to show homogeneous dispersion at the nanoscale in PLA/PLGA matrix such as electrospun fibers as compared to conventional CaP particles [157,158]. In another approach, the surface of HA was initially coated using phosphonic acid–based agent to introduce more reactive hydroxyl groups on the HA particles for PLA bonding [159]. The diametric tensile strength of PLA/HA composite prepared from PLA-grafted HA was found to be over twice that of the composite prepared from the non-modified HA. Moreover, the tensile strength of the improved composite was 23% higher than that of PLA alone. On the other hand, PLA/PLGA can also be modified to improve its adhesion to CaP particles. Alkali treatment of PLA can induce carboxyl groups formation on PLA surface, which can then interact with calcium ions on HA to drive the HA coating deposition on PLA surface [74]. Yet another approach is to use poly(a-methacrylic acid) (PMMA) to introduce carboxyl groups to PLA surface via photo-oxidization and UV-induced polymerization; the modified PLA can then act as a template to manipulate the nucleation and growth of HA crystals [160].

5.2. Incorporation of proteins/drugs/genes and stem cells There is significant scope in the loading of proteins/drugs/genes to further promote tissue regeneration. However, this does not allow for extreme temperature ranges or extremely aggressive chemical conditions during processing. The incorporation of proteins, drugs, genes and stem cells in the composites will be discussed in detail in the subsequent paragraphs. In PLA/PLGA matrix, CaP materials can also work as loading sites for proteins for further enhancing bone regeneration. Woo et al. examined the adsorption of serum protein and adhesion of molecules on PLA and PLA/HA scaffolds prepared by a thermalinduced phase separation process [161]. PLA/HA scaffolds adsorbed much higher amounts of serum proteins, adhesion proteins and peptides (fibronectin, vitronectin, RGD and KRSR). Furthermore, b1 and b3 integrins and phosphorylation of Fak and Akt proteins in the cells on the PLA/HA scaffolds were significantly more abundant than those on PLA scaffolds, which suggested that the enhanced adsorption of serum adhesion proteins to PLLA/HAP scaffolds protects the cells from apoptosis possibly through the integrin–Fak–Akt pathway. Therefore, proteins (e.g. growth factor BMPs) can be directly loaded onto composite surface by soaking composite scaffolds in biomolecule-containing aqueous solution [162–164]. Instead of direct absorption of proteins to composite surface, they can be also be combined into the inner matrix of the composite. Kwon et al. prepared rhGDF-5 (recombinant human growth and differentiation factor-5)/b-TCP/PLGA composites by blending rhGDF-5-coated b-TCP with PLGA solution. rhGDF-5 is a growth/differentiation factor for bone growth and this in vivo study indicated that this composite greatly promotes bone healing as compared to control group (PLGA/b-TCP) [165]. For drug delivery applications, emulsion is the most common approach to produce microspheres for drug delivery. Shi et al. used both o/w and w/o/w emulsion methods to fabricate PLGA/HA microspheres loaded with alendronate to study the release kinetics [80]. Alendronate is a bisphosphonate, one of the primary agents countering osteoblast-mediated bone loss. Bisphosphonates also have high affinities towards HA crystals [166]. It was reported that in a w/o/w emulsion system, the affinity between HA (in organic phase) and alendronate (in aqueous phase) was inhibited due to limited HA exposure to alendronate caused by PLGA-containing organic phase, resulting in an encapsulation efficiency of alendronate below 40%. On the other hand, PLGA/HA-alendronate microspheres fabricated via o/w emulsion showed approximately 90% encapsulation efficiency. All PLGA/HA-alendronate systems exhibited controlled release without a remarkable initial burst release. As an alternative to loose microparticles, drugs are also loaded to 3-D implant composites. Chu et al. fabricated a PLGA/5-fluorouracil/ HA scaffold via SFF as an implantable drug delivery system; 5fluorouracil is an anti-cancer drug [167]. The relationship between drug release rate and the amount of HA showed a bilinear kinetics and an HA content of about 5–20% seemed to be a reasonable amount of additive to control the release rate of 5-fluorouracil. Xie et al. added icaritin to PLGA/TCP slurry to fabricate PLGA/ TCP/icaritin composite scaffold material with slow release of icaritin during scaffold degradation for enhancing bone repair [168]. DNA can also be loaded to composites in order to induce biomolecule generation in vivo and enhance bone regeneration. PLGA/HA electrospun composite scaffold loaded with DNA was constructed by Nie et al. via three different approaches including naked DNA coated outside, DNA-loaded chitosan nanoparticles coated outside, and DNA-loaded chitosan nanoparticles encapsulated inside the fibers [169]. A human marrow stem cell culture experiment showed that the scaffolds with encapsulated DNA/ chitosan nanoparticles had higher cell attachment, higher cell viability and desirable DNA transfection efficiency.

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PLA-CaP/PLGA-CaP composite scaffolds have been shown to be good candidates for stem cell loading in order to induce cooperative cell–cell and cell–matrix interactions, thus promoting bone regeneration. Li et al. transplanted PLGA/HA electrospun composites cultured with/without MSCs into cartilage defects in rats; both scaffolds had better tissue morphology, structural integrity, matrix staining and much thicker newly formed cartilage than the control group. The histological score for PLGA/HA-MSC is better than that for PLGA/HA, showing that the MSCs play an important role in tissue repair [170]. Montjovent et al. seeded human fetal bone cells onto supercritical gas foamed PLA/b-TCP scaffold and implanted the composite in rat. Composite scaffold resorption as well as cortical and trabecular bone repair were assessed. The composite scaffold degradation was coupled to the rate of tissue regeneration, leading to a structural integrity of the constructs. A complete bone ingrowth induced by PLA/b-TCP composite scaffold with bone cells was observed as compared to blank PLA/b-TCP composite scaffold [171]. Bone marrow stromal cells were also seeded to PLA/b-TCP composite scaffold and implanted into bone defects [172]. It was observed that the mineralization in the defect sites occurred after 30 days and a callus was formed around the periphery of the implants and along the adjacent host bone after 90 days; however, the empty bone defects had not fully healed after 90 days and the bone regeneration did not progress after 60 days.

5.3. Long-term degradation The degradation behavior of PLA/PLGA matrix is influenced by the degradation properties of combined CaP fillers. It has been reported that the incorporation of HA can greatly inhibit the degradation and water uptake of PLA matrix, while acting as a physical barrier to block off the entry of water, thus slowing down PLA degradation kinetics [146]. As an alternative, TCP promotes the degradation of PLA/PLGA matrix by intensifying water uptake. In an experiment reported by Loher et al., pure PLGA matrix only had 0.5% mass loss and 2% water uptake after 6 weeks, while 30% amorphous TCP incorporation resulted in 7% mass loss and 14% water uptake after 6 weeks. It is possible that the high degradation rate of amorphous TCP generated pores in the composite, thus resulting in a larger polymer surface area being exposed to the surrounding liquid to promote PLGA degradation [87]. The inflammation caused by degradation of PLA/PLGA is one of the reasons to incorporate CaP materials into PLA/PLGA matrix [79,96,101]. However, in long-term investigations, the inflammation problem is still a concern in the application of PLA-CaP/PLGA-CaP composite due to the difference in degradation rates between CaP fillers and PLA/ PLGA matrix. Ignatius et al. implanted PLA or PLA/TCP scaffolds into female merino sheep [173]. A mild inflammatory response was observed in the first 12 months but a strong inflammatory reaction was seen after 24 months. The incorporation of a-TCP can improve the biocompatibility of the scaffold to avoid inflammation reaction in vivo. After a total degradation of a-TCP after 6 months, PLA degradation can cause a potential inflammation reaction in vivo. After 2 years, the strength of this reaction was significantly decreased. A better tailoring of the contents and compositions of PLA/PLGA matrix and CaP fillers is necessary for successful long-term clinical applications. On the other hand, biodegradation of the composite and deposition of apatite are two competing process in vivo. Ideally, it is desired that there is a smooth transition between a primarily polymer-based composite to a mineralized bone material in the bone regeneration process. Careful tailoring of the CaP material content and composition, the composite structure, and the degradation rate of PLA/PLGA may provide a smooth transition from composite to mineralized bone.

2013

Designing and implementing an in vitro system to mimic longterm in vivo degradation of PLA-CaP/PLGA-CaP composite scaffold is a new approach as an alternative to the more expensive in vivo animal experiments. However, such long-term studies focusing on PLA-CaP/PLGA-CaP composite are rare. Yang et al. developed an in vitro dynamic system for the long-term evaluation of composite degradation, which showed potential for investigation of the long-term mechanical and chemical transition of composite scaffolds [174]. In contrast to simple incubation in phosphate-buffered saline (PBS) at 37 °C, PLGA/b-TCP composite scaffolds were incubated in weekly refreshed PBS at 37 °C with dynamic loading to mimic the in vivo environment. The results indicated that the cyclic loading under dynamic conditions accelerated the degradation of the PLGA component in the composite. This system could provide fundamental experimental results for the degradation behavior of biodegradable polymeric scaffolds used in bone tissue engineering in vivo. More work is still required to develop a long-term study. 6. Summary In order to substitute bone grafts, the development of composites made of PLA/PLGA and CaP materials for bone tissue engineering applications has attracted much interest. The biological response, mechanical properties, composite architecture and degradability of PLA/PLGA-CaP composites can be influenced by fabrication techniques, properties and contents of raw materials, and the interaction between matrix and fillers. PLA/PLGA-CaP composites can be tailored to meet different bone tissue engineering application requirements. However, there are still a number of issues that remain to be improved (e.g. inflammation, long-term degradation, scale-up, mechanical strength improvement, etc.) for an even wider application of PLA/PLGA-CaP composites in clinical applications. Acknowledgment This work was partially supported by NSF Grant 0700306. Appendix A. Figures with essential colour discrimination Certain figures in this article, particularly Figs. 2, 5, 7, 9, 11, 12 and 14–16 are difficult to interpret in black and white. The full colour images can be found in the on-line version, at doi:10.1016/ j.actbio.2012.01.031. References [1] Buckwalter JA, Glimcher MJ, Copper RR, Recker R. Bone biology. II: Formation, form, modeling, remodeling, and regulation of cell function. Instr Course Lect 1996;45:387–99. [2] Keating JK, McQueen MM. Substitutes for autologous bone graft in orthopaedic trauma. J Bone Joint Surg Br 2001;82B:3–8. [3] Younger EM, Chapman MW. Morbidity at bone graft donor sites. J Orthop Trauma 1989;3:192–5. [4] Logeart-Avramoglou D, Anagnostou F, Bizios R, Petite H. Engineering bone: challenges and obstacles. J Cell Mol Med 2005;9:72–84. [5] Hutmacher DW, Schantz JT, Lam CXF, Tan KC, Lim TC. State of the art and future directions of scaffold-based bone engineering from a biomaterials perspective. J Tissue Eng Regen Med 2007;1:245–60. [6] Rezwan K, Chen QZ, Blaker JJ, Boccaccini AR. Biodegradable and bioactive porous polymer/inorganic composite scaffolds for bone tissue engineering. Biomaterials 2006;27:3413–31. [7] Dorozhkin SV. Calcium orthophosphate-based biocomposites and hybrid biomaterials. J Mater Sci 2009;44:2343–87. [8] Liu C, Xia Z, Czernuszka JT. Design and development of three-dimensional scaffolds for tissue engineering. Chem Eng Res Des 2007;85:1051–64. [9] Porter JR, Ruckh TT, Popat KC. Bone tissue engineering: a review in bone biomimetics and drug delivery strategies. Biotechnol Prog 2009;25:1539–60. [10] Hench LL, Wilson J. Surface-active biomaterials. Science 1984;226:630–6.

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