Biodegradable ceramic-polymer composites for ...

Materials Science and Engineering C 71 (2017) 1175–1191

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Review

Biodegradable ceramic-polymer composites for biomedical applications: A review Michal Dziadek a,⁎, Ewa Stodolak-Zych b,⁎, Katarzyna Cholewa-Kowalska a,⁎ a b

AGH University of Science and Technology, Faculty of Materials Science and Ceramics, Department of Glass Technology and Amorphous Coatings, 30 Mickiewicza Ave., 30-059 Krakow, Poland AGH University of Science and Technology, Faculty of Materials Science and Ceramics, Department of Biomaterials, 30 Mickiewicza Ave., 30-059 Krakow, Poland

a r t i c l e

i n f o

Article history: Received 2 August 2016 Received in revised form 18 September 2016 Accepted 13 October 2016 Available online 14 October 2016 Keywords: Silica Bioglasses Wollastonite Calcium phosphate ceramics

a b s t r a c t The present work focuses on the state-of-the-art of biodegradable ceramic-polymer composites with particular emphasis on influence of various types of ceramic fillers on properties of the composites. First, the general needs to create composite materials for medical applications are briefly introduced. Second, various types of polymeric materials used as matrices of ceramic-containing composites and their properties are reviewed. Third, silica nanocomposites and their material as well as biological characteristics are presented. Fourth, different types of glass fillers including silicate, borate and phosphate glasses and their effect on a number of properties of the composites are described. Fifth, wollastonite as a composite modifier and its effect on composite characteristics are discussed. Sixth, composites containing calcium phosphate ceramics, namely hydroxyapatite, tricalcium phosphate and biphasic calcium phosphate are presented. Finally, general possibilities for control of properties of composite materials are highlighted. © 2016 Elsevier B.V. All rights reserved.

Contents 1. 2. 3. 4.

5. 6.

Introduction . . . . . . . . . . . . . . . . Biodegradable polymer matrices . . . . . . . Silica based composites . . . . . . . . . . . Bioglass based composites . . . . . . . . . . 4.1. Silicate bioactive glasses . . . . . . . 4.2. Borate and borosilicate bioactive glasses 4.3. Phosphate glasses . . . . . . . . . . Wollastonite based composites. . . . . . . . Calcium phosphate ceramics based composites

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Abbreviations: ABTS•+, 2,2′-azinobis(3-ethylbenzothiazoline-6-sulfonic acid) radical cation; ACP, amorphous calcium phosphate; ALP, alkaline phosphatase; BBG, borate-based glass; BCP, biphasic calcium phosphate; BET SSA, Brunauer-Emmett-Teller specific surface area; BG, bioactive glass/bioglass; BMP-2, bone morphogenetic protein 2; BSP, bone sialoprotein; CAM, cylinder chorioallantoic membrane; CaP, calcium phosphate; CBF-alpha-1, core-binding factor alpha-1; CMC, carboxymethylcellulose; CS, chitosan; DEX, dexamethasone; DPPH•, 2,2diphenyl-1-picrylhydrazyl radical; DSC, differential scanning calorimetry; ECM, extracellular matrix; EDX, energy dispersive X-ray spectroscopy; FN, fibronectin; FTIR, Fourier transform infrared spectroscopy; FTIR-ATR, attenuated total reflectance Fourier transform infrared spectroscopy; gHAp, grafted hydroxyapatite; HAp, hydroxyapatite; hAT-MSCs, human adipose tissue-derived mesenchymal stem cells; hBMSCs, human bone marrow mesenchymal stem cells; HCA, carbonated hydroxyapatite; HOC, human osteoblastic cells; MSCs, mesenchymal stem cells; MSNs, mesoporous silica nanoparticles; mWS, mesoporous wollastonite particles; nHAp, nano-sized hydroxyapatite; NMR, nuclear magnetic resonance; OC, osteocalcin; ON, osteonectin; OPN, osteopontin; Osx, osteoblast-specific transcription factor Osterix; PBG, phosphate glass; PBS, phosphate buffer saline; PBSu, poly(butylene succinate); PCL, poly(ε-caprolactone); PDGF, platelet-derived growth factor; PDLLA, poly(D,L-lactide); PEG, polyethylene glycol; PEI, polyethylenimine; PEO, poly(ethylene oxide); PGA, poly(glycolic acid); PHB, poly(3-hydroxybutyrate); PHBV, poly(3-hydroxybutyrate-co-3-hydroxyvalerate); PIXE, proton-induced X-ray emission; PLA, poly(lactic acid); PLDLA, poly(L-lactide-co-D,L-lactide); PLGA, poly(lactic-co-glycolic acid); PLLA, poly(L-lactide); PRP, platelet-rich plasma; rBMSCs, rat bone marrow mesenchymal stem cells; RT-PCR, reverse transcription polymerase chain reaction; RUNX2, runt related transcription factor 2; SBF, simulated body fluid; SBG, silicate bioactive glass; SCPL, solventcasting particulate leaching; SF, silk fibroin; TCP, tricalcium phosphate; TGF, transforming growth factor; TIPS, thermally induced phase separation; VEGF, vascular endothelial growth factor; WAXS, wide-angle X-ray scattering; wHAp, hydroxyapatite whiskers; WS, wollastonite; XRD, X-ray diffraction. ⁎ Corresponding authors. E-mail addresses: [email protected] (M. Dziadek), [email protected] (E. Stodolak-Zych), [email protected] (K. Cholewa-Kowalska).

http://dx.doi.org/10.1016/j.msec.2016.10.014 0928-4931/© 2016 Elsevier B.V. All rights reserved.

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6.1. Hydroxyapatite based composites . . . . . . 6.2. Tricalcium phosphate based composites . . . 6.3. Biphasic calcium phosphate based composites . 7. Summary. . . . . . . . . . . . . . . . . . . . . Acknowledgments . . . . . . . . . . . . . . . . . . . References. . . . . . . . . . . . . . . . . . . . . . .

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1. Introduction On account of rapid development of novel biomedical technologies, including tissue engineering, regenerative medicine, gene therapy and controlled drug delivery, new materials are being developed to meet specific requirements of these fields. Conventional single-component ceramic or polymer materials cannot satisfy them. In addition, in order to fully meet the basic requirements such as biocompatibility, biodegradability, appropriate mechanical properties, there is a need to obtain materials fulfilling several advanced functions at once. For example, a multifunctional material for bone tissue regeneration should induce formation of new bone tissue without an addition of organic bone growth factors (e.g. BMP-2), degrade progressively at a rate matching the regeneration of new bone, induce new blood vessels formation and exhibit antibacterial and anti-inflammatory activity. Therefore, key material and biological features can be achieved by design and development of multi-component materials, including selection of matrix and modifier materials, their parameters (e.g. shape, distribution, content), as well as fabrication techniques of composites. In particular, this work deals with the introduction of ceramic modifiers into biodegradable polymer matrices to obtain composites with specific properties for biomedical applications, especially tissue engineering and regenerative medicine. The most widely used ceramic modifiers including bioactive glasses and calcium phosphates, as well as less widespread silica and wollastonite, with focus on their beneficial effects on material and biological properties of the composites will be discussed. A table placed at the end of the review (Table 3) provides an overview of polymer matrices, ceramic modifiers, methods of fabrication and forms of composites discussed in this work, as well as their physical properties.

2. Biodegradable polymer matrices Many types of biodegradable polymeric materials have already been used as matrices of ceramic-modified composites for tissue engineering applications. These materials can be classified into two major groups based on their origin, namely natural-based polymers, including proteins (soy, collagen, fibrin gels, silk) or polysaccharides (starch, alginate, chitin/chitosan, hyaluronic acid derivatives) and synthetic polymers, such as poly(lactic acid) (PLA), poly(glycolic acid) (PGA), poly(εcaprolactone) (PCL), poly(3-hydroxybutyrate) (PHB) [1–2]. Generally, biodegradation of polymeric biomaterials involves cleavage of enzymatically (natural polymers) or hydrolytically (synthetic polymers) sensitive bonds in the polymer structure leading to the polymer erosion [3]. Naturally-derived polymers possess the ability to biological recognition, including presentation of receptor-binding ligands, that may support cell adhesion, migration, differentiation and proliferation. However, the rate of their in vivo degradation depends on availability and concentration of the enzymes at the site of implantation, therefore it is difficult to predict. The use of natural polymers alone is often restricted because of potential immunogenic reactions, possibility of disease transmission and relatively poor mechanical properties [1–3]. Synthetic polymers have more predictable properties including degradation kinetics that can be controlled by chemical composition and configurational structure, molecular weight, polydispersity, crystallinity, material morphology (e.g. porosity, surface area), chain orientation

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and overall hydrophilicity [1]. Synthetic polymers possess relatively good mechanical strength and their properties (e.g. porosity, shape) can be tailored for specific applications. However, the surfaces of synthetic polymers are hydrophobic and lacking in cell-recognition sequences [1–2]. In addition to the biodegradability, the composites take advantage of high formability of polymer matrices. Many methods have been developed to fabricate composite materials for biomedical applications, including solvent casting [4–6], tape casting [7], particulate leaching [8], freeze drying [9], phase separation [10], thermal processing [11−12], gas foaming [13], electrospinning [14] and rapid prototyping [15]. These methods allow obtaining the required properties, especially the form and microstructure of the composites.

3. Silica based composites Reports on silica (silicon dioxide, SiO2) as a filler or a nanofiller of polymer matrix composites showed many advantages of composite or nanocomposite materials based on biodegradable polymers [16–55]. Silica particles, incorporated into polymer matrices, cause higher biocompatibility and bioactivity of the materials and/or implants. These compositions stimulate biological properties interesting for bone tissue applications, such as given bioresorption rate and porosity, as well as the ability to induce formation of calcium phosphate similar to the one present in bone on biomaterials surfaces, and the introduction of biologically active agents [16–18]. Since silica reinforcements are frequently used in the form of nanoparticles, understanding of mechanisms of dissolution and metabolism of these particles in the organisms have also deserved considerable attention [17]. Nanometric particles of SiO2 in polyesters provide many advantages compared to other nanofillers and exhibit many properties associated with an ideal material for grafting and scaffolding [19]. Due to silanol groups (SiOH) present on the surface of silica, covalent bonds can be formed between macromolecular chains and the fillers [20]. During the process of nanocomposite synthesis, silane coupling agents play important roles in connecting the interfaces of organic (polymer chain) and inorganic phases (nanofillers, SiO2 particles). It is because they can be functionalized at the interface to create a chemical bridge between the reinforcement and the polymer matrix, and thus improve the stability, adhesion and mechanical properties such as strength, Young's modulus or wear resistance of the nanocomposites [21–24]. Some investigations on nuclear magnetic resonance spectra (13C NMR) of nanocomposite materials i.e.; PBSu/SiO2 (poly(butylene succinate)/SiO2) confirmed a reaction between the surface silanol groups from SiO2 nanoparticles with hydroxyl end groups of the polymer (PBSu) leading to formation of covalent bonds [20]. These covalent bonds resulted in a substantial improvement of mechanical properties of PBSu, even in cases where amounts of SiO2 lower than 2.5 wt.% were used [20,25]. Nano-SiO2 in different forms i.e.; nanotubes, nanoparticles present in the polymer matrix changes also physicochemical properties of the nanocomposite surface e.g.; some of the hydroxyl groups are exposed on the surface and increase hydrophilic character of the surface of materials. This phenomenon can explain a higher hydrolysis rate of nanocomposites based on biodegradable polymers [26]. A similar accelerating effect of SiO2 nanoparticles on a PLA hydrolysis rate was reported by many authors [27–30]. They found


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Fig. 1. Durability of poly(L-lactide-co-D,L-lactide)/SiO2 (PLDLA/SiO2) membrane after 8-week incubation in H2O/37 °C/PBS.

that PLA/SiO2 nanocomposites degrade faster than the pure PLA, which indicate that incorporation of silica enhances biodegradation of PLA in the nanocomposites. The addition of silica may lead to facilitated attack of water and enzyme molecules on ester groups of the PLA chains, because of high hydrophilicity of silica. The effect of degradation was much more visible (Fig. 1) in case of porous nanocomposite membrane after 8-week in vitro degradation [30]. This experiment also showed a decreasing molecular weight of the polymer and disintegration of the materials. The incorporation of silica nanoparticles into the polymer matrix stimulated osteoblast-like cells interaction with natural tissue after contact with the surface of material i.e. cell viability increased since silicon (at critical concentrations) is able to stimulate proliferation of osteoblast-like MG-63 cells [31]. Silicon can also be involved in bone formation and mineralization [32], whereas orthosilicate acid (Si(OH)4) at physiological concentration of 10 μmol was shown to stimulate formation of type I collagen in human osteoblastic cells (HOC) and to stimulate cell differentiation [33]. Even though no significant effects of SiO2 inclusion on cells behavior could be detected, the results were in agreement with reports from the literature showing silica materials used for bone regeneration. Si-OH groups can induce apatite formation and thus strong bone bonding [34]. Furthermore, silica nanotube meshes were shown to support the adhesion of murine osteoblast-like MC3T3-E1 cells and expression of alkaline phosphatase (ALP), even without BMP-2 supplements [35]. Mesoporous silica characterized with regular, parallel pores of 2 to 50 nm diameter, high specific surface area, low density, high biocompatibility shows capacity for encapsulation of drugs and biological agents [36]. These particles are widely used in pharmaceutical applications, in most cases in drug delivery systems. In order to control drug release from the mesoporous silica, it was mainly mixed with biodegradable polymers such as; poly(lactic-co-glycolic acid) (PLGA), polyethylene glycol (PEG), PCL, chitosan (CS), alginate, gelatin, polyethylenimine (PEI) etc. [37–40]. Drugs such as vancomicine, gentamicine, fluoxetine, lidocaine, morphine, nifedipine, paracetamol, tetracyclin, ibuprofen etc. were used in these applications [41–43]. The drug could be incorporated in the silica xerogel during the sol–gel process instead of a adsorption method. In some cases the drug can be incorporated into the polymer matrix as well, which may have an effect on the release properties of the composite [39]. Ahola et al. showed that increase of an initial drug loading increased an amount of toremifene citrate released from PCL matrix [40]. The release rate was found to be directly proportional to the load of toremifene citrate. The drug release kinetics was controlled by the slowest process i.e. the release from the polymer matrix which was a diffusion controlled process [44–45]. Deviations from linearity were observed after 70% of the drug had been released. This may be due to a large amount of toremifene

citrate used, which enhanced degradation of the matrix, thus leading to an increased release rate. A new method of carrying drugs into a damaged tissue is an application of core-shell structures. Formation of firm bonds between silica and polymer (PEG, PEI, PLGA) on the silica core improves controlled release of a drug or other molecules i.e. molecules that can regulate biological behavior of osteoblasts/osteoclasts [46–47]. Silica nanoparticles with a drug incorporated inside their pores could be used as fillers of a polymeric matrix used to prepare a scaffold or a membrane, which provided an additional degree of control of drug release, independent from the polymer material itself [48]. Additionally, scaffolds based on mesoporous silica characterized by macro-, meso- as well as micropores were identified as important contributors to bone formation. The micropores may serve as nucleation sites for mineralization, causing local supersaturation and subsequent nucleation of apatite forms, the mesopores can be used as reservoirs of nutrients and growth factors or carry therapeutic agents, and macropores can promote cells infiltration. Such implants could by fabricated by traditional techniques such as; salt leaching, freeze drying or new methods like electrospinning or rapid prototyping. Polymer scaffolds modified by mesoporous silica nanoparticles (MSNs) were often called multifunctional scaffolds [48–50]. Porous 2D or 3D structures fabricated from a polymer modified with nanometric silica using the electrospinning technique can be used to obtain a superhydrophobic layer on the implant surface [51]. This method enables control of nanoroughness of the surface by the presence of polymer fibers with nanometric or submicrometric diameters containing silica nanoparticles which in turn influence wettability and surface free energy (Table 1). CS fibers were successfully electrospun with poly(ethylene oxide) (PEO) as a co-blending polymer and with nanometric silica (CS(PEO)/SiO2) by Suzuki and Mizusima [52]. These organic-inorganic hybrid biomaterials earlier proved to increase oxygen permeability, biocompatibility and biodegradability. Other works showed how to control a process of manufacturing of CS(PEO)/SiO2 nanofibers and the ability of these hybrid nanofibers to favor cell attachment of murine osteoblast-like 7F2 cells. A hybrid fibrous mesh has a capability to be modified by incorporation of calcium ions resulting in formation of bioactive carbonated hydroxyapatite (HCA) crystals

Table 1 Diameters of the PCL and the PCL/SiO2 nanocomposite fibers with physicochemical properties of the layer. Material

Fiber diameter (nm)

PCL 152 ± 19 PCL/SiO2 178 ± 53

Water contact angle (°)

Surface free energy [mN/mm]

151 113

44.8 34.7

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[53]. The incorporation of silica particles into natural hydrogels such as CS or alginate creates another possibilities for cells e.g.; bone-like cells such as human osteoblast-like Saos-2 cells and RAW 264.7 mouse macrophage cells differentiated into osteoclast-like cells were embedded in silica-containing Na-alginate-based hydrogel. During the test biocompatible hydrogels with Saos-2 cells retained their capacity to synthesize the hydroxyapatite (HAp) crystals [54] and increasingly express the gene encoding for osteoprotegrin [55]. In another study when CS was used as a matrix a significantly increased protein adsorption and improved apatite deposition were observed by the addition of nano-silica into the composite scaffold [54]. To conclude, nanosilica present in the polymer matrix composite works not only as a filler enhancing its biological properties and guaranteeing increase of its biomineralization capability but it also increases stiffness of the polymer composite without decreasing its mechanical strength.

4. Bioglass based composites Bioactive glasses (BGs, bioglasses) are very attractive materials for modifying a polymer matrix of composites dedicated for tissue engineering applications. One of the key reason that makes these materials relevant composite modifiers is the possibility of controlling a wide range of biological and chemical properties. The structure and chemistry of glasses can be tailored at a molecular level by varying composition, a preparation method (melt-quenching or sol-gel), and manufacturing conditions [1,56–57]. It provides a simple way to control properties of glass-containing composites. BGs are commonly produced by a melt-quench route or a sol-gel method [56,58]. Gel-derived BGs usually exhibit improved bioactivity and cellular response related to a larger volume fraction of mesopores [59–61], a larger concentration of silanols on the surface [59–60] and also larger specific surface area [58,60–61] than the melt-derived glasses. Furthermore, an advantage of the sol-gel process is the ability to control the phase composition, structure, texture, microstructure and surface chemistry of the glasses by the synthesis conditions (e.g. type of a catalyst) and thermal treatment [61–62]. The sol– gel process allows obtaining mesoporous micro- and nanoparticles, simply by changing pH of the process, which can be directly used as composite modifiers [61]. To improve or induce desired biological response and adjust the surface reactivity, as well as solubility in biological environment, BGs can be doped with trace elements and other therapeutic oxides [17,58,63,64]. The research works mainly focus on developing silicate and borate/borosilicate BGs, as well as

resorbable phosphate glasses and their application in tissue engineering. 4.1. Silicate bioactive glasses Silicate bioactive glasses (SBGs) are the most extensively researched glasses for biomedical applications and also the most frequently considered as fillers of polymer-based composites for bone tissue engineering. The first melt-derived silicate SBG of the quaternary SiO2–Na2O–CaO– P2O5 system – Bioglass® 45S5, synthesized by Hench et al., is widely reported as a modifier of biodegradable synthetic polymers and copolymers: PLA [65–66], PGA [67], PCL [4,8,10,68], PHB [69], PLGA [65,70–71], as well as natural polymers: collagen [72] and CS [73]. Moreover, other SBGs based on both binary SiO2–CaO [74] and ternary SiO2–CaO–P2O5 [4–6,8,71,75] systems are also used for modifying a polymer matrix. Bioglass® 45S5 and other SBGs exhibit an ability to bond to soft and hard tissues which is attributed to the formation of a HCA layer on the glass surface in contact with the body fluids [57–58]. Apart from the bioactivity, it was observed that ionic dissolution products (e.g. Si, Ca, P) from some of BGs stimulate expression of several genes of osteoblastic [76] and stem cells [71,77] involved in bone growth, show ability to stimulate angiogenesis [78], while possibly exhibit antibacterial [79] and anti-inflammatory actions [80]. Furthermore, release of calcium (Clotting factor IV) being considered a reason for SBG haemostatic properties [81]. All these specific properties and advantages of BGs can be conferred to a polymer matrix by the glass filler incorporation. Probably the most important reason for developing polymer-glass composites for bone tissue engineering is the possibility of achieving a bioactive properties of the polymer matrix [1]. The degree of bioactivity of a composite can be controlled mainly by the chemical composition [8, 71,82], volume fraction [9–11,65–66,83], size [5,11,68–69,73] and distribution [5,68–69] of a bioactive phase. It was shown that the rate and intensity of apatite-like layer formation on the surface of PCL and PLGA-based composites containing SBG particles increased with an increasing CaO:SiO2 molar ratio of the glass [8,71] (Fig. 2.). Other results indicated that lower P2O5 content in SiO2–CaO–P2O5 SBG nanoparticles enhanced the bioactivity of (poly(L-lactide)-based (PLLA) scaffolds [82]. Furthermore, an increased volume fraction [9–11,65–66,83] and higher surface area to volume ratio of the glass modifiers (e.g. the incorporation of nano-sized particles instead of microparticles) [5,11,68–69,73] improved the composite bioactivity. The incorporation of SBG particles in a biodegradable polymer matrix can be advantageous as it imparts osteoinductivity to the composite. It was demonstrated that the addition of gel-derived SBG particles into a PLGA matrix promoted the expression of early osteogenesis-related transcription factors (runt related transcription factor 2 (RUNX2) and

Fig. 2. In vitro bioactivity of PCL/SBG films after 7-day incubation in simulated body fluid (SBF). No morphological changes on pure PCL film (a) after immersion in SBF are observed. On the surface of a composite film containing silica-rich bioglass (S2 (mol%): 80SiO2–16CaO–4P2O5) (b) clusters of calcium phosphate precipitates occur, while the surface of a film with calciumrich glass (A2 (mol%): 40SiO2–54CaO–6P2O5) (c) is fully covered with a thick, continuous layer of HAp.


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Fig. 3. The real-time reverse transcription polymerase chain reaction (RT-PCR) analyses of RUNX2 and Osx mRNA levels in; (a) hBMSCs cultured for two days on PLGA, PLGA/S2 and PLGA/ A2 without osteogenic inducers and BMP-2, osteopontin and collagen I mRNA levels (b-d) in hBMSCs cultured for ten days without osteogenic inducers (NONE) or in the presence of recombinant human BMP-2 or dexamethasone (DEX). Fold changes in mRNA levels (mean ± SD) are expressed relative to PLGA [77]. © IOP Publishing. Reproduced with permission. All rights reserved.

osteoblast-specific transcription factor Osterix (Osx)) and osteogenic genes (BMP-2, osteopontin (OPN), osteocalcin (OC) and collagen I) in human bone marrow mesenchymal stem cells (hBMSCs). As with bioactive properties, osteogenic potential depends on chemical composition of glass modifiers [77] (Fig. 3.) and their content in composites [71]. The experiment during which rat mesenchymal stem cells (MSCs) were cultured in the presence of PLGA/SBG porous scaffolds (but physically separated from them) showed that the dissolution products from the materials stimulated osteogenesis alone [84]. Apart from the effect of solution-mediated factors, a three-dimensional structure and surface properties of the scaffolds also strongly influenced the osteogenic response [8,84]. A recent trend is the incorporation of different elements into the composition of SBGs to enhance their biological benefits, especially osteogenic and angiogenic potential [63]. Therefore, Ren et al. [14] and Poh et al. [15] demonstrated that the addition of strontiumsubstituted SBG into a PCL matrix enhance osteoblast differentiation, collagen deposition and also matrix mineralisation compared to pure PCL scaffolds. In turn, Oh et al. showed that PLDLA-based membranes with zinc-substituted SBG possessed significantly improved osteogenic and matrix mineralization potential of the rat bone marrow mesenchymal stem cells (rBMSCs) comparing to zinc-free glass-containing membranes and pure polymer membranes [85]. The addition of 45S5 SBG particles into poly(D,L-lactide) (PDLLA) and PLGA matrices, as well as a PGA mesh significantly enhanced vascular endothelial growth factor (VEGF) in vitro secretion of fibroblasts [67, 86–87] and also in vivo blood vessel formation [67,86]. These results indicated a beneficial effect of bioglass inclusion into a polymer matrix on induction of neo-vascularisation and rapid vascular ingrowth necessary for delivery of nutrients and oxygen into the developing tissues, and thus the possibility of using glass-containing composites, not only for hard tissue regeneration, but also for the soft one [67], and soft-hard tissue interfaces [86]. Other experimental study showed that the addition of a limited concentration (10 wt.%) of 45S5 SBGs nanoparticles to collagen films induced an early angiogenic response evaluated using the cylinder chorioallantoic membrane (CAM) model [72]. Quinlan et al. developed collagen-glycosaminoglycan-based porous scaffolds containing cobalt-substituted SBG which show ability to stimulate both angiogenesis and vascularisation through the release of Co2 + ions which are known for their potential to mimic hypoxia [88]. SBGs present in a bioresorbable polymer matrix can also affect the polymer degradation behavior by modification of its surface and bulk properties. On one hand, the degradation process of the composite can be accelerated because hydrophilic SBG particles improve surface wettability of the composite, allowing the scaffold to absorb more water [9,12,83,89] and also increase the surface area of a hydrolytic attack [4]. Furthermore, water can more easily diffuse through the interface between the two phases by capillarity and microcrack transport mechanisms [12]. In particular, the greater amount and the smaller particle

size of SBG are, the more rapid degradation rate of the poly(εcaprolactone-co-D,L-lactide)/SBG [12] and the higher water absorption of the PHB/SBG composites are [69]. On the other hand, a delayed degradation rate in the composite can be achieved by the dissolution of alkaline ions (e.g. Ca2+, Na+) from the bioactive glass structure. It results in neutralisation of the released acidic by-products of the polymer degradation and thus prevents its autocatalytic degradation process [83]. Furthermore, a buffering effect of the glass modifiers can be considered as an advantage, because it helps to avoid a possible inflammatory response, due to acidic degradation of the polymer [4,8–9]. It was shown that the degradation behavior of PLGA, PLLA and PCL composites can be controlled not only by tailoring the concentration of BG [9,70–71] but also SBG composition, namely the amount of alkaline glass network modifiers (e.g. CaO) [4,8,71]. The introduction of SBG particles into a polymer matrix contributed to increase stiffness of the composites [4,10,71,75,86]. In particular, nano-sized SBG particles had a significant enhancing effect on the Young's modulus of CS/SBG [73], PHB/SBG [69] and PCL/SBG [68] composites when compared with SBG microparticles. Mechanical properties of a composite can be controlled mainly by the amount [4,10,70–71,75, 86] textural properties [4,90], and phase composition [4,68] of glass fillers. It was shown that porous microstructure of gel-derived 58S SBG particles compared to dense melt-derived 45S5 glass [90] and also larger mesopores of gel-derived calcium-rich SBG (~ 13.3 nm) with respect to smaller pores of silica-rich BG (~3.0 nm) [4–5] resulted in an improvement of mechanical strength of PDLLA and PCL-based films, respectively. In turn, partially crystallized SBG particles could have different interfacial bonding strength with PCL as compared with the fully amorphous particles, which would affect strength of the composites [4–5,68]. The presence of SBG particles in a polymer porous scaffolds generally induces an increase of their compressive strength [10,77, 86]. On the other hand, decrease of tensile strength is a frequent effect of modification of polymer films with SBG fillers [4,71,75]. In order to improve the interface compatibility between SBG particles and a polymer matrix, surface modification of the SBG particles is used. The modification would make the particles disperse within the matrix more homogeneously [91]. Furthermore, molecules grafted onto the particles could interact with the polymer matrix, thus the final mechanical properties could be improved. It was shown that surface modification of SBG particles by esterification reaction [90] or surface-grafting with diisocyanate [91] improved the tensile strength of PDLLA and PLLAbased composites, respectively. In case of composites consisting of semicrystalline polymers (e.g. PCL, PLLA), their properties strongly depends on crystallinity of the polymer matrix. It is well known that a degree of crystallinity, size and distribution of crystallites in semicrystalline polymers have a large effect on their mechanical properties [92]. Furthermore, amorphous regions of a semicrystalline polymer degrade prior to crystalline

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Table 2 Degree of crystallinity (χc), melting temperature (Tm) (mean ± SD) and crystallite size (L110 and L200) of pure PCL and SBG-PCL composite films. The composite materials contained 12 and 21 vol.% of gel-derived silica-based bioglass microparticles with different compositions - S2 and A2.

χc (%)a Tm (°C)a L110 (nm)b L200 (nm)b a b

PCL

12S2-PCL

21S2-PCL

12A2-PCL

21A2-PCL

42.69 ± 1.16 62.33 ± 0.09 30.4 23.6

36.22 ± 0.68 60.17 ± 0.14 26.5 19.1

20.79 ± 0.79 59.00 ± 0.14 22.1 17.2

37.25 ± 0.72 60.67 ± 0.13 24.7 20.5

20.84 ± 0.85 59.20 ± 0.10 19.1 15.5

Evaluated by the differential scanning calorimetry (DSC). Evaluated by the wide-angle X-ray scattering (WAXS).

domains, thus initial crystallinity of the polymer determines its degradation behavior [89]. Recent studies showed that different polymer crystallinity strongly influences adhesion and proliferation of osteoblasts and fibroblasts [93]. Some studies indicated a possibility of modulation of crystalline properties of a polymer matrix by the addition of glass particles. On one hand, microparticles of SBG introduced into PCL and PLLA matrix caused a reduction of the polymer crystallinity, melting temperature [4–5,8] and crystallite size [4,94], which were affected by the filler amount (Table 2.). On the other hand, the incorporation of nano-sized SBG into PLLA resulted in an increase of a nucleation rate and crystallinity of the polymer. In addition, PLLA grafting onto the surface of the nanoparticles with diisocyanate enhanced this effect [91]. Microstructure and porosity of a material play a critical role in tissue engineering because they determine cell migration, cell-cell interactions and molecular transport of nutrients, wastes, and biological mediators within a scaffold. There are various methods for controlling porosity and pore architecture of polymer and composite scaffolds. In particular, it is possible by varying the amount and size of porogen particles in the solvent-casting particulate leaching (SCPL) method [8] or by using various solvents [10], polymer concentrations [10,65] and cooling temperatures [65] in the thermally induced phase separation (TIPS) process. It turned out that when TIPS method was used to obtain composite scaffolds based on PCL [10], PDLLA [65–66,83], PLGA [65,70] and PLLA [9], porosity of the materials decreased and the pores became irregular in shape with increasing content of SBG particles. According to the authors, the probable reason for these phenomena is a hindering effect of randomly distributed solid inorganic fillers in the polymer/solvent solution on the growth of solvent crystals during the phase separation process. On the other hand, the enlargement of pore size in PLLA-based scaffold upon the addition of SBG particles can be attributed to the effect of released ions from the glass surface on the TIPS thermodynamics [94]. Surface properties of biomaterials, such as wettability, roughness, surface energy, surface charge determine cell adhesion, proliferation and differentiation [95], protein adsorption, bioactivity [69], as well as degradation behavior. It was shown that SBG particles, as a hydrophilic material, remarkably improve hydrophilicity of PCL [4,68,75] and PHB [69] matrices. Furthermore, the decrease in water contact angle was more visible for PHB-based composites containing SBG nanoparticles than for microparticles due to the fact that more of the nano-sized SBG particles were exposed on the surface [69]. Some authors observed that the excess over the certain content of BG in a composite did not result in a further reduction of the contact angle. It can be attributed to the notable increase of surface roughness [4,69,75]. Our recent study indicated that PCL/SBG composite films are capable of serving as a carrier of natural polyphenols extracted from sage (Salvia officinalis L.) [6]. Polyphenols-loaded composites showed potential antioxidant activity against the ABTS•+ and DPPH• radicals, and antiproliferative activity against human malignant melanoma WM266–4 cells. In contrast to PCL film, composites containing gel-derived and melt-derived SBG particles exhibited a reduced initial burst release of drug in the first day and also excellent in vitro bioactivity. Reduction in the initial fast release was particularly pronounced for a film with gelderived fillers which can be attributed to much larger specific surface area of these particles (BET SSA 101.6 m2 g− 1 ) compared to the

melted ones (BET SSA 0.3 m2 g− 1) and also to the presence of –OH groups on the surface of the gel-derived glass that can bind polyphenol molecules. 4.2. Borate and borosilicate bioactive glasses Borate-based glasses (BBGs) are known for their high reactivity and bioactivity, however their biocompatibility still remains debatable [58, 96]. It is due to cytotoxicity of high concentration of boron released into the solution as borate ions (BO3)3− related to high solubility of borate glasses [97–99]. Probably for this reason, borate-based glasses are not as common choice for tissue engineering applications as the silicate ones. Under static in vitro culture conditions, some BBGs showed significant reduction in cell proliferation, in turn, the cytotoxicity was inhibited under dynamic culture conditions [100]. In vitro response study of scaffolds consisting of borate glass 13-93B3 (B2O3–CaO–K2O– Na2O–MgO–P2O5 system) exhibited a cytotoxic effect, whereas the same scaffolds showed the ability to support in vivo tissue infiltration [101]. The conversion mechanism of BBG to HAp is similar to that of silicate glass, with the formation of a borate-rich layer, similar to the silicarich layer of the SBGs [58]. Partial replacement of B2O3 in BBG with Al2O3 [102] or SrO [96] was shown to reduce the dissolution rate of the glass. Furthermore, the presence of strontium oxide induce the adhesion of osteoblast-like cells, thus significantly increasing biocompatibility [96,102]. Based on the available literature, there are only a few reports on the use of borate-based glasses as polymer matrix modifiers [7,103–105]. Because of high degradation rate, BBGs are considered as modifiers of a biodegradable polymer matrix to obtain bioactive and drug-releasing scaffolds. Zhang and Jia et al. developed teicoplanin-loaded CS/BBG composites for the chronic osteomyelitis (bone infection) treatment. The materials did not cause any toxic injury to local or systemic tissues, while they showed ability to cure chronic bone infection within 12 weeks of implantation in a rabbit tibia osteomyelitis model. Furthermore, the composites exhibited a multifunctional role including an antibiotic release function, sufficient load-bearing capacity, biodegradability and the ability to support bone regeneration [103–105]. A significantly higher degradation rate of borate-based glasses compared to the silicate ones makes them more suitable for soft tissue applications. Furthermore, by altering composition of the glass fillers, namely SiO2 and B2O3 content, physical and chemical properties of a polymerbased composite can be easily tailored to meet tissue regenerative requirements. Mohammadkhah et al. showed that composites containing 50 wt.% PCL and (1) 50 wt.% 13-93B3 BBG or (2) 50 wt.% 45S5 SBG or (3) a blend of 25 wt.% 13-93B3 and 25 wt.% 45S5 glass particles showed a modified degradation rate. It was also demonstrated that faster conversion of 13–93B3 fillers to HAp may improve the effectiveness of the PCLbased material as a nerve guide conduit when compared to silicate glass-containing material [7]. 4.3. Phosphate glasses Phosphate glasses (PBGs), containing P2O5 as a network-forming oxide, have a wide range of solubility which can be controlled by modifying their composition [106]. In the basic P2O5-CaO-Na2O system, a


lower CaO content causes higher dissolution rate [107]. Furthermore, the dissolution rate of PBGs can be tailored by the addition of several oxides such as; Al2O3, B2O3, Fe2O3, Ga2O3, TiO2, ZnO, MgO, SrO, NiO, MnO and CuO [17,58,106,108]. Therefore, PBGs may be used as an inorganic filler of a bioresorbable polymer matrix (e.g. PCL, PLA) to obtain fully resorbable composite materials with a controlled resorption profile [106, 109]. Moreover, PBGs are considered as materials for bone tissue engineering due to their chemical similarity to the inorganic phase of human bone [17]. It was also observed that dissolution products of PBGs from P2O5-CaO-Na2O system with a lower dissolution rate (28 and 40 mol% CaO) enhanced proliferation of osteoblastic cells and also expression of the bone-associated proteins (bone sialoprotein (BSP), osteonectin (ON) and fibronectin (FN)) in contrast to highly soluble glasses (8 and 16 mol% CaO) which showed inhibitory action [107]. Another work showed that the incorporation of TiO2 (5 mol%) alone, as well as a mixture of TiO 2 (5 mol%) and ZnO (1, 3 and 5 mol%) in PBG of the basic system enhanced proliferation of osteoblastic cells compared with the basic glass and PBG containing ZnO only, which can be attributed to the significant reduction in a degradation rate of these glasses and also beneficial effect of Zn2 + ions. Moreover, the transcription level of osteogenesis-related genes (ALP, ON and core-binding factor alpha-1 (CBF-alpha-1)) of osteoblastic cells seeding on glasses containing a mixture of TiO2 and 1, 3 mol% ZnO was comparable or even significantly higher compared to a positive control (Thermanox®) [108]. The presented studies showed that, by tailoring composition of PBGs, it is possible to modulate their biological properties. Most of the works concerning the modification of a polymer matrix with phosphate bioglass are devoted to degradation studies. It was shown that the dissolution of glass fillers directly influences the degradation behavior of the polymer-based composites, as with pure phosphate glasses, the degradation rate of the composites increased with decreasing CaO content in PBG fillers [110−111]. Mohammadi et al. obtained two kinds of PCL-based composites containing PBGs modified with SiO2 and/or Fe2O3: (1) composite films containing PBG fibers of the (mol%) 50P 2O 5 –40CaO– (10-x)SiO 2 – xFe2O3 system, where x = 0, 5 and 10 [109], as well as (2) composite films containing 40 vol.% glass particles of the (mol%) 50P2O 5 – 40CaO–10SiO 2 or 50P2O 5 –40CaO–10Fe2 O 3 systems or blends of these glasses (0/40, 10/30, 20/20, 30/10, 40/0) [106]. The authors showed that the increasing SiO2 content enhanced the degradation rate of the composites. It was manifested by higher weight loss, increase in the release of PO4 3 − , Ca2 +, Fe3 + and Si4 + ions, formation of pores, decrease in mechanical properties and more rapid reduction in pH during the degradation test [106,109]. In turn, Georgiou et al. showed that the degradation behavior of PLLA-based composites can be also controlled by tailoring the concentration of glass particles: the higher content of the (mol%) 46CaO– 4Na2O–50P2O5 PBG particles in the composite is, the higher is the degradation rate of the materials [13]. Polymer-based composites containing phosphate glasses are considered as drug delivery systems. It was shown that the higher the CaO content in glass is, the lower a vancomycin release from the PCL-based composites is. Furthermore, the addition of glass with higher CaO content into PCL matrix reduced the initial burst effect and prolongs release of a drug. It is probably related to high affinity of hydrophilic drugs to the hydrophilic surface of glass particles and lower solubility of highCaO glasses. These results indicated that it is possible to control degradation rate and drug release of the PCL/PBG composites depending on the glass composition while maintaining excellent biocompatibility [110]. Navarro et al. showed that the incorporation of PBG particles from the (mol%) 44.5P2O5–44.5CaO–6Na2O–5TiO2 system in a PLA matrix induced formation of calcium phosphate precipitates at the composite surface after immersion in SBF. The X-ray diffraction (XRD) and Fourier transform infrared spectroscopy (FTIR) of the CaP precipitation showed

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the presence of an amorphous phase. The Ca/P ratio obtained from the energy dispersive X-ray spectroscopy (EDX) confirmed the presence of amorphous calcium phosphate (ACP) [112]. To conclude, besides calcium phosphate ceramics, bioactive and resorbable glasses are the most widely used and researched materials for modifying biodegradable polymer matrices. It is due to the possibility of modulating a wide range of BG properties and thus, by its incorporation into a polymer matrix, possibility of imparting and controlling, not only physicochemical or mechanical properties, but also a number of biological effects of the obtained composites. 5. Wollastonite based composites Wollastonite (WS, CaO-SiO2, CaSiO3) is a typical example of calciumsilicate-based ceramics. WS is a naturally occurring calcium silicate, which has been widely used as a filler in polymers and cements to fabricate composites with improved mechanical properties [113–115]. Because of its excellent bioactivity and degradability, WS has been proposed as a potential material for bone tissue regeneration [116– 118] with the same or better possibilities than classic ceramic particles such as TCP or HAp. Nevertheless, the extensive use of WS, TCP, HAp or BG is still limited by their brittle nature. Most of literature positions showed composite materials based on synthetic or natural polymers in which WS was used as a filler to produce composites with improved mechanical and bioactive properties [1,119–120]. Synthetic biodegradable matrices such as; PLGA, PLDLA, PCL or poly(3-hydroxybutyrate-co3-hydroxyvalerate (PHBV) guarantee also short-term durability under in vitro/in vivo conditions, this property of materials can be controlled by modification with WS. Results of studies of Li and Chang [121−122] showed influence of WS particles on PHBV matrix. Thermoplastic polyester such as PHBV gains bioactivity in similar way as polymer composites modified with BG particles (with CaO-SiO2 formula). According to this mechanism, silicate ions released from WS particles are attached to Si–OH groups inducing nucleation of apatite on the surface. At the same time, release of calcium ions from WS particles increases the ionic activity product of SBF with respect to apatite and accelerates the apatite nucleation. Finally, the apatite nuclei formed on the surface grow spontaneously by consuming Ca and P ions from the fluid [123–124]. A layer formed by apatite crystals was observed on the surface of PHBV/20 or 40 wt.% WS composite after soaking in SBF for 14 days. Additionally, pH of the immersion medium (SBF) was stable until 21 days of soaking which means that the WS incorporated into the PHBV could neutralize acidic by-products and stabilize pH of the SBF solution [121–122]. It suggested that WS/PHBV composites could gradually release Ca and Si ions, and the release rate depended on content of WS in the composites (20 wt.% or 40 wt.%). The released Ca and Si ions were able to form basic hydrates and this can possibly explain the ability of pH-stabilization showed by WS. A rapid degradation process of WS particles was confirmed in long-term observations and in the other immersion medium i.e.; phosphate buffer saline (PBS) where pH was also stable after 14 days of incubation [122]. The alkaline ions neutralized acidic products of degradation of PHBV and had a buffering effect on the acidification caused by degradation of PHBV. This effect could also be identified in the other aliphatic polyesters matrices such as PGLA or PLLA [125–126]. Both properties i.e.; bioactivity and biodegradability of composite materials based on WS are results of high hydrophilicity of WS particles. The suitable amount of inorganic material i.e. N 20 wt.% reduces hydrophobic character of polymers surface and facilitates local supersaturation, which are necessary for apatite nucleation, and hydrolysis of a polymer chain, which is a beginning of its degradation process. A good example of possibility of the water-uptake capacity was characterized by a composite scaffold based on silk fibroin (SF) formed by the freeze-drying technique [124]. The first reason of high water absorption is high porosity (80–90%) and then wettability of composite materials e.g. water contact angle decreases from 65o to 40o with an increasing amount of WS.

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Fig. 4. The dispersion of WS in the PCL/nanoWS composite determined with the PIXE technique. Pure PCL – reference sample. FTIR-ATR confirming chemical interaction between a polymer chain and nanofillers; a) nano-WS, b) pure PCL, c), d), e) the nanocomposite with different contents of nano-WS particles; 1%, 4%, 6%wt.

Another aspect of modification of a biodegradable polymer matrix by WS is its influence on mechanical properties of composite materials i.e.; porous scaffolds. The addition of WS to the SF matrix resulted in a structure with higher compressive strength than the pure SF e.g. compressive strength of SF/40 wt.% WS composite was two times higher than strength of the pure SF [124]. Compressive strength of PHBV/WS composite scaffolds was higher and reached 0.28 MPa when weight ratio of WS in the composite was 40% [122]. The same effect was observed in the case of collagen/WS composites; where tensile strength and bioactivity of the composite scaffolds were improved by increasing WS content [126]. The same author showed that composite systems based on polymer and wollastonite (PGLA/WS, PBVH/WS) might be used as bioactive drug delivery systems which were not only able to enhance bone growth, but

also could release antibiotics to minimize infection problems [123–124]. The most popular model of this study is gentamicin – an antibiotic used in inflammatory therapeutic. It was proved that WS in polymer-based composites such as PGLA/WS/gentamicin or PBVH/WS/ gentamicin retarded release of gentamicin and this might be due to formation of an apatite layer on the surface of the composite. This effect strongly depends on composition of an ions-reached immersion medium such as SBF or PBS which could stimulate bioactivity process [127–128]. Mesoporous particles of wollastonite (mWS), similarly to the mesoporous silica, possess high specific surface area and high pore volume. These properties significantly improve the kinetics of process of apatite formation on their surface thereby exhibiting better bioactivity and possibility of being a drug carrier [129–130]. Wie et al. confirmed that an

Fig. 5. Morphology of the nanocomposite fibers; PLA modified with 1% wt. TCP nanoparticles (a), morphology of the PLA/TCP nanocomposite fibers after 7 days of incubation in SBF and EDX analysis confirming apatite composition (b). Republished with permission of Trans Tech Publications Inc., from [184]; permission conveyed through Copyright Clearance Center, Inc.

Table 3 Review of biodegradable ceramic-polymer composites, their form, preparation method, potential application and mechanical properties. Polymer matrix

PCL

Ceramic modifier

Scaffold (porosity 55–92%, pore size b 300 μm) Film (thickness ~ 90 μm)

Solvent casting particulate leaching; Solid-liquid phase separation; Phase inversion Solvent casting

Bone tissue engineering Bone tissue engineering

10.5–14.5 (T)

Scaffold (porosity 87–92%, pore size 10- some hundreds μm) Scaffold (fiber diameter 46.1 μm, fiber spacing 222.1 μm) Scaffold (porosity ~75%, pore size b 2.05 mm) Film (thickness 65 μm) Film (thickness 60 μm)

Solid–liquid phase separation

Bone tissue engineering

92–214·10−3 132–251·10−3 [10] (C)

Melt-electrospinning

Bone graft substitutes

3D printing, melt extrusion-based additive manufacturing technology Solvent casting


Particles/melting/10 wt.%

Film

Solvent extraction and thermal pressing

Fibers/melt–draw spinning/diameter 10–20 μm/~18 vol.%

Film (thickness ~ 200 μm)

Thermal pressing

Particles/sol-gel/b50 μm/21 vol.%

Particles/sol-gel/b40 μm/12, 21 vol.%

Particles/melting/b38 μm/10 wt.%

45S5 Bioglasss® Sr-SBG 46SiO2-24Na2O–20SrO–7CaO–3P2O5 (mol%) SBG 60SiO2–63CaO–4P2O5 (mol%) BBG 53B2O3–20CaO–12K2O–6Na2O–5MgO–4P2O5 (wt.%) 45S5 Bioglasss® PBG 45P2O5–xCaO–(55-x)Na2O, x = 20, 30, 40, 50

Particles/ b38 μm/10 wt.%

WS

SBG 80SiO2–16CaO–4P2O5 40SiO2–54CaO–6P2O5 (mol%)

45S5 Bioglasss®

Particles/sol-gel template method/50–150 nm/10, 20, 30 wt.% Particles/b20 μm/50 wt.%

Particles/microwave-hydrothermal Scaffold (porosity method/80-100 nm/50–70 wt.% 71–82%, pore size 100–500 μm,) Particles/ 2.3–2.7 nm/10, 30 wt.% Scaffolds (porosity 11–55%, pore size 4.1 nm-39.5 μm) Particles/sol-gel/20, 40, 50 wt.% Scaffold (porosity 57–74%, pore size 1–500 μm) Particles/sol-gel/b50 μm/12, 21, 33 Film (thickness ~ vol.% 110 μm /scaffolds (porosity 88–95%, pore size 90–300 μm) Particles/melting/b5 μm/10, 25, 50 Scaffold (porosity wt.% N90%, pore size 10–100 μm)

Tape casting

Bone tissue engineering Neural tissue engineering

[8]

0.6–1.3·103

[4]

[14]

48–59

[15]

15–19 (T)

200–383

[75]

40–50 (T)

150–190

[7]

Tissue regeneration and wound-healing 35–45 (F) Bone fracture fixation devices

Solvent casting particulate leaching


1.0–2.2 (C)

Supercritical carbon dioxide (scCO2)-assisted foaming/mixing Solvent casting particulate leaching

Hard tissue engineering Hard tissue repair

5–12 (C)

Solvent casting /solvent casting particulate leaching


24–55 (T) films

Thermal induced phase separation

Hard and soft tissue engineering

[110]

~1.8·103

[109]

6.5–20.5·103

[189]

10–175

[48]


Particles/melting/b45 μm/25, 50 wt.%

Sr-SBG 46SiO2-26Na2O–17SrO–6CaO–5P2O5 (mol%)

SiO2

PLGA

Potential application

SBG 80SiO2–16CaO–4P2O5 40SiO2–54CaO–6P2O5 (mol%) SBG 80SiO2–16CaO–4P2O5 40SiO2–54CaO–6P2O5 (mol%) 45S5 Bioglasss®

PBG 50P2O5–40CaO–(10- x)SiO2–xFe2O3, x = 0, 5, 10 (mol%) BCP

Reference

Composite preparation method

Form/preparation method/size/amount

Compressive (C), tensile (T) flexural (F) strength (MPa)

Young's modulus (MPa)

Composite form

Type

[129]

2–4·103 films

[71,77]

22–27

[65,70]

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(continued on next page)

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Table 3 (continued) Polymer matrix

PLLA

Form/preparation method/size/amount

BCP 63/37 HAp/β-TCP

Particles/ heating bovine bone/b100 nm/ 10–50 wt.%

HAp

Particles/~100 nm/50 wt.%

SBG 55SiO2–40CaO–5P2O5 30SiO2–60CaO–10P2O5 68SiO2–28CaO–4P2O5 (mol%) 45S5 Bioglasss®

Particles/sol-gel template method/20 nm/25 wt.%

Composite preparation method

Potential application

Compressive (C), tensile (T) flexural (F) strength (MPa)

Young's modulus (MPa)

Reference

Scaffold (porosity N89%, pore size 60–150 μm) Scaffold (porosity 86–91%) Scaffold (pore size 10–300 μm)



0.16–0.55 (yield strength) (C)

3.2–15.0

[192]

Gas foaming, Particulate leaching Thermal induced phase separation


2.3–4.5

[146]

Scaffold (porosity 48–52%, pore size b 40 μm) Scaffold (pore size 27–150 μm) Scaffold (porosity ~81%, pore size 200–400 μm) Cylindrical profile

Particles/melting/10 μm/10, 30, 50 wt.% Particles/melting/32–1000 μm/5, 10, 20 wt.%

HAp

Particles/0–50 μm, 5 μm, b200 nm/10, 20, 30 wt.%

β-TCP

Particles/microwave accelerated wet method/100–300 nm/60 vol.%

β-TCP

Particles/in the presence of PEG/290 nm/10–30 wt.%

45S5 Bioglasss®

Particles/melting/b5 μm/5, 10, 25, 40, 50 wt.%

45S5 Bioglasss®

Particles/melting/0.1–25 μm/5, 10, 20 wt.% - films, 20 wt.%- scaffolds Particles/flame spray synthesis/35–40 nm/5, 10, 20 wt.% - films, 20 wt.%- scaffolds Particles/sol-gel/b45 μm/30 wt.% Membrane (thickness ~ 150 μm)

Solvent casting/solvent casting particulate leaching

Solvent casting

Guided bone regeneration

Particles/melting/b80 μm/50 wt.%

Disk (thickness ~ 1.5 mm) Disk (thickness 2 mm) Film (thickness 120–140 μm)

Solvent casting and thermal pressing Compression moulding

Bone tissue engineering Bone substitution Bone tissue engineering

Scaffold (porosity 73–78%, pore size 30–300 μm)

Compression moulding particulate leaching

SBG 70SiO2–30CaO Zn-SBG 70SiO2–25CaO–5ZnO PBG 44.5P2O5–44.5CaO–6Na2O–5TiO2 (mol%) Poly(e-caprolactone-co-DL-lactide) SBG 53SiO2-23Na2O–20CaO–4P2O5 (wt.%) PHB 45S5 Bioglasss®

WS

Particles/b45 μm, 90–315 μm/40, 60, 70 wt.% Particles/melting/b5 μm/10, 20, 30 wt.% Particles/flame spray synthesis/29 nm/10, 20, 30 wt.% Particles/chemical coprecipitation/98–154 μm/20, 40 wt.%

Scaffold (porosity 50%, pore size 200–450 μm) Scaffold (porosity N93%, pore size 0.5–300 μm, fiber diameter 70–300 nm) Scaffold (porosity N90%, pore size 10–100 μm) Film (thickness 25–40 μm)/scaffolds (porosity 85–93%,

[82]

Solid–liquid phase Bone tissue separation/freeze-extraction engineering

SBG 33CaO–28P2O5–16MgO–23SiO2 (wt.%) PBG 46CaO–4Na2O–50P2O5 (mol%)

PLDLA

PHBV

Particles/melting/50–63 μm/25, 50 wt.%

Composite form

[9]

Thermal induced phase separation Melt extrusion and supercritical CO2-foaming


0.62–0.75 (C) 6.61–10.88

[94]

80.0–116.5

[13]

Melt extrusion

Various medical application Bone tissue engineering

100–136 (F)

3.3–4.3·103

[151]

Particulate leaching

1–5 (C)

[168]

Thermal induced phase separation + particulate leaching


0.3–0.8

[171]


Hard and soft tissue engineering Regeneration of hard-soft tissue defects

21

[65–66,83]

Solvent casting


0.38–1.59 (C) 0.36–1.80

[86]

54.2–56.3 (T)

[85]

[112] [11–12] 0.8–1.6·103

[69]

0.20–0.28

[121]


PDLLA

Ceramic modifier Type

[124] 1.6–2 Freeze-drying Scaffold (porosity 81–85%, pore size ~ 100 μm) Particles/chemical precipitation/10, 20 wt.%

Thermal induced phase separation Scaffold (porosity N 98%) Particles/melting/38, 100 μm/

SF

SBG 49SiO2-26Na2O–23CaO–1P2O5 Co-SBG 49SiO2-26Na2O–19CaO–1P2O5-4CoO (mol%) WS Collagen-glycosaminoglycan

Particles/20–30 nm/10, 20 wt.% 45S5 Bioglasss® Collagen

Particles/~50 nm/0.5, 1, 2 wt.%

Scaffold (pore size 60–125 μm) Film (thickness ~ 40 μm)

Thermal induced phase separation Compression moulding

Treatment of chronic osteomyelitis Bone tissue engineering Hard and soft tissue engineering Bone tissue engineering Solvent casting

BBG 6Na2O–8K2O–8MgO–22CaO–54B2O3–2P2O5 (mol%) HAp


0.15–0.2 (C)

4–6·10−3

[88]

[72]

[141] 5.5–9·10−3

[103–105]

Bone tissue engineering CS

45S5 Bioglasss® 46SiO2-23Na2O–27CaO–4P2O5 (wt.%)

Film (thickness ~ Particles/commercially 50 μm) available/5 μm/30 wt.% Particles/flame spray synthesis/30–50 nm/30 wt.% Particles/melting/b50 μm/~70 wt.% Pellet

Solvent casting

23–33 (C)

17–20 (Storage modulus)

[73]


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apatite layer on the surface of a composite with mWS was observed after 7-days soaking in the SBF medium. In the case of a composite with the same amount of classic WS it took 14 days [129,131]. Higher specific surface area provided by mWS particles present in a polymer matrix promotes cells interaction (osteoblast, chondrocyte) [76,132–133]. Further studies on the application of mWS showed that even 0.5% concentration of mWS in chitosan/carboxymethylcellulose (CS/CMC) scaffolds significantly improves their physicochemical and biological properties [131]. Inclusion of mWS particles had no effect on altering the porous architecture of the CS/CMC scaffolds but influenced protein adsorption and biomineralization ability. The CS/CMC/mWS scaffolds were biocompatible and enhanced differentiation of MSCs to osteoblasts [132]. A small amount of additives present in a degradable polymer matrix characterizes nanocomposite materials. When PCL was used as a matrix with nanometric WS (0.5–1 wt.%.) as a filler a homogenous dispersion was obtained. A result of a good distribution of the nanofiller within the polymer matrix was better mechanical properties (Young's modulus, tensile strength and work-of-break) [134]. It suggests that there is some interaction between WS nanoparticles and the polymer matrix (hydrogen bonding, van der Waals interaction) indicating that the ceramic filler contributes to the overall elastic properties of the composite samples. Enhancement of the modulus of nanocomposites at such a low concentration of ceramic particles cannot be attributed to the introduction of a ceramic filler with higher modulus, which is observed in traditional composite materials i.e. with 40 wt.% of WS. The same interaction between a polymer chain and nanometric WS could be confirmed by spectroscopic studies such as FTIR in the transmission mode, the FTIRATR technique and the proton-induced X-ray emission (PIXE) method [135]. Dispersion of WS nanoadditives is well visible in PIXE where analytical elements for WS (i.e. Ca and Si) could be observed (Fig. 4). To conclude, the biodegradable polymer/WS composite possess many interesting properties (biocompatibility, bioactivity, biodegradability and combining with them carrying therapeutic agents) which were guaranteed by WS particles. Some authors pointed out that these materials can be used as an alternative for conventional bioactive ceramic fillers such as HAp, TCP, BCP and BG used to modify the polymer matrix. 6. Calcium phosphate ceramics based composites Over the past few decades, calcium phosphate (CaP) ceramics alone have been widely used as bone graft substitutes, especially because of similarity of their chemical composition to the mineral phase of bone [136–137]. Due to their nature, calcium phosphate ceramics show also high biocompatibility and ability to bond with bone tissue under certain conditions, however, because of their brittleness, their clinical applications have been limited to the non- or low-load bearing parts of the skeleton. Many types of calcium phosphates have been considered as biomaterials for bone reconstruction in dental, orthopaedics and maxillofacial application due to different behavior in the living organism, including bioactivity, biodegradability and biological response. Bioactivity, degradation behavior and osteoconductivity/osteoinductivity of CaP ceramics generally depend on the Ca/P ratio, crystallinity and phase composition [138–139]. Synthetic HAp (Ca10(PO4)6(OH)2) shows good stability in the body, while tricalcium phosphates (α-TCP, β-TCP, Ca3(PO4)2) are more soluble, whereas biphasic calcium phosphate (BCP; an intimate mixture of HAp and β-TCP) exhibits intermediate properties depending on the weight ratio of stable/degradable phases. Thus, the dissolution rate decreases in the following order: α-TCP N β-TCP N BCP N HAp [1, 138]. Because the natural bone hydroxyapatite is nonstoichiometric and contains, beside the main components i.e.; Ca, (PO4)3 −, (OH)−, − 2+ some other groups and trace elements (e.g. CO2– , Na+, K+, 3 , F , Mg 2+ 2+ Sr , Zn ), new trends in calcium phosphates preparation, especially HAp, consist in obtaining substituted CaP ceramics. They do not only

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have a different solubility and bioactivity than the parent materials, but also cause modified biological response due to the release of biologically active ions during dissolution [136]. CaP ceramics exhibit different biological effects in vivo, while most of them are osteoconductive, only certain types are osteoinductive. Samavedi et al. proposed that a range of osteoinductive potentials of CaPs decrease in the following order BCP N TCP N HAp [138]. Currently HAp, TCP and BCP ceramics are quite common types of materials used for various biomedical applications and they are available on the market. In the past few years, many efforts have been directed to develop calcium phosphate-containing composite materials, especially polymer-matrix composites. 6.1. Hydroxyapatite based composites Stoichiometric HAp (Ca/P molar ratio of 1.67) is considered to be osteoconductive but not osteoinductive. Furthermore, because HAp is the most stable among CaP ceramics, its surfaces provide highly effective nucleating sites for the precipitation of an apatite crystal in contact with culture medium and body fluids [138]. Solubility, bioactivity and biological response of HAp can be modified by anionic and cationic substitution [136]. Mainly for these reasons, HAp is widely used to prepare polymer-ceramic composite materials usually with the aim to impart the bioactivity and osteoconductivity and also improve mechanical properties [140–147]. It is believed that one of the most promising bone graft material is collagen/nano-sized HAp (nHAp) composite because of its ability to mimic structure and composition of natural bone. New technics such as self-assembly and biomineralization have been recently used to obtain collagen/nHAp composites with oriented and hierarchical structure considering the spatial relationship between organic and inorganic phases [148–149]. Using the self-assembling method, a highly oriented and with morphology similar to compact bones, collagen/nHAp composite was obtained [148]. A pre-crystallization method allowed obtaining a composite in which HAp nanocrystals were situated among the collagen fibers creating a suitable structure for bone defect repair [149]. Several studies were conducted on the use another natural polymer, namely CS, as a component of HAp-containing composites. The presence of nHAp in a CS matrix increased compressive strength [140] and Young's modulus of composites [141]. It can be related to hydrogenbonding interactions between NH2 and OH groups of nHAp and chelation between NH2 and Ca2+ when the co-precipitation method of composite preparation was used [140]. The incorporation of nHAp affected a degradation rate of composites [140,141], induced bioactivity [140] and favored murine osteoblast-like MC3T3-E1 cells attachment and proliferation [141]. In the literature, HAp fillers have also been shown to modify numerous properties of synthetic polymer matrices, especially poly(α-hydroxy esters). The comparison of hydroxyapatite with different morphologies, namely nanoparticles (nHAp) and whiskers (wHAp), as PLGA matrix modifiers, showed that composites with nHAp had higher bending strength in comparison with wHAp-containing materials. It can be attributed to a more homogeneous distribution of nHAp in the polymer matrix and also enhanced crystallization of the PLGA matrix. It was also shown that morphology of HAp fillers influenced in vitro degradation behavior of the composites [150]. In turn, the use of various contents (0–30 wt.%) and also sizes (0–50 μm, 5 μm and N 200 nm) of HAp particles led to changes in thermal properties and/or crystallinity, as well as the mechanical strength of the PLLA and PLGA-based composites [151]. In order to improve the adhesive strength between HAp nanoparticles and PLLA and PLGA matrix, the hydroxyl groups on the surface of the ceramic nanoparticles were grafted with PLLA by chemical bonding [142–143]. It was shown, that composites containing grafted HAp (gHAp) exhibited significantly improved tensile strength, bending

strength, and impact energy. Furthermore, a composite with a low content (4 wt.%) of gHAp possessed higher mechanical properties than the pure PLLA. The implantation study of repairing critical-sized defects in the radius of a rabbit forelimb showed that the gHAp/PLGA scaffold exhibited rapid and strong mineralization and osteoconductivity [143]. Degradation studies showed that the incorporation of HAp [144] and nHAp particles [145] into PCL and PLGA matrix, respectively, accelerated degradation of the composites. Furthermore, a degradation rate of composites can be adjusted by varying the HAp content [144–145]. Kim et al. obtained nHAp/PLGA composite scaffolds using the gas forming particulate leaching method. The scaffolds exhibited significantly higher rat calvarial osteoblast growth, alkaline phosphatase activity, and extracellular matrix (ECM) mineralization in vitro compared to the pure PLGA material. Furthermore, the presence of nHAp fillers enhanced hydrophilicity and in vivo osteoconductivity (higher bone formation area and more extensive calcium deposition) of the material [146–147]. Another study showed that incorporation of Ag-doped HAp into PHBV nanofibers imparted antibacterial activity against E. coli and S. aureus bacteria and enhanced in vitro bioactivity when compared to the material containing unmodified HAp, while did not show any cytotoxic effect [152]. 6.2. Tricalcium phosphate based composites Tricalcium phosphate (β-TCP, TCP) is a well-known CaP-based bioceramics. β-tricalcium phosphate has been widely investigated and successfully used in clinical application as a biomaterial for bone repair applications due to its remarkable biocompatibility, in vivo resorbability, bioactivity and good osteoconductivity [153–183]. Moreover, some results indicated that TCP is considered to be osteoinductive [138]. For many years TCP has been combined with other ceramic materials such as; HAp, Al2O3, ZrO2 to produce small dental or orthopedic implants [154–156] in the form of composites or composite layers on the implants [156–159]. Last decades showed that the most perspective form of TCP is small particles or nanoparticles which modify a biodegradable polymer matrix [157,160–162]. Incorporation of the ceramic particles into polymer-based composite materials presented significant osteogenic benefits and encouraged formation of a new bone on implant surfaces [163–164]. In addition, during degradation process of composite implants/scaffolds based on biodegradable aliphatic polyesters, particles of TCP which modified the polymer matrix, improved not only osteoconductivity but also neutralized acidic pH of the environment typical for degradation of a polymer matrix [163–165]. These phenomena decrease the risk of local complications (inflammatory reactions) and increase the degradation rate comparing to pure polymer implants [166]. Moreover, employment of a room temperature preparation technique of such composites may allow incorporating of thermally unstable antibiotics, as well as growth factors e.g.; BMP-2 as stimulants of bone formation. Osteoconductive composite materials such as; PLLA/β-TCP, PCL/β-TCP comprising a large volume fraction of β-TCP (≥ 60 vol.%) and a smaller amount of polymer (PLLA, PCL) were studied as degradable antibiotic carrier materials for the orthopedic treatment [166–167]. A study on PCL/TCP beads loaded with vancomycin (1–4 wt.%) showed that a gradual drug release observed over the period of 4–11 weeks had diffusional character and depended on the composite matrix homogeneity and porosity [167]. Methods of fabrication of PLA/β-TCP and PCL/β-TCP composites i.e.; cold sintering and/or salt leaching allowed obtaining materials suitable for load-bearing bone scaffolds [167–169]. These materials were characterized by high strength and Young's modulus required as biomechanical properties of such implants. It is, therefore, believed that the first resorbable and a strong bone graft substitute will consist of a polymer-calcium phosphate composite and will contain a large ceramic fraction [170]. A small fraction (≤40 vol.%) of the particulate ceramic phase decreased mechanical properties of composite implants and scaffolds


but they were still osteoconductive and promoted osteoblasts adhesion, proliferation, penetration, and ECM deposition and influenced their degradation rate [171]. Kobayashi et al. showed that small amount of TCP (contents of 5–15 wt.%) in PLLA agglomerated and these regions acted as water diffusion paths and started hydrolytic degradation of the polymer matrix near the TCP-PLLA interfaces [172]. Results of some investigations pointed out that nanometric particles of β-TCP with high specific surface area and chemical groups capable of interaction with a polymer chain can affect both bioactivity/osteoconductivity and mechanical properties [173]. The fibrous PLA/TCP (1 wt.%) nanocomposites produced by the electrospinning method belong to a group of biomimetic materials because they chemically and structurally mimic a natural bone tissue. The nanocomposite fibers with nanometric or submicrometric diameters have particle size and composition similar to the natural collagen fibers present in a bone. A three dimensional (3D) form of the material facilitates settling and adhering of bone cells and inducting apatite formation on the surface of nanofibers (Fig. 5) [184]. Despite of this, implants/scaffolds made of PLGA and nanometric particles of β-TCP would not induce stronger osteointegration than implants made of PLGA and micrometric-size β-TCP particles e.g. a push out test showed similar results in both materials. To overcome these limitations, the reinforcement of a polymer matrix by TCP nanoparticles was proved when crosslinking CS was used as the matrix polymer [174–176]. As a crosslinking agent could be used; genipine, gluatre aldehyd or tripolyphosphate. It was shown that a composite scaffold obtained with genipine-crosslinked CS and nanometric β-TCP improved mechanical, bioactivity and cell supportive properties such as viability and osteogenic differentiation of human MSCs [177–178]. Such scaffolds had slightly higher extent of mineralization after 21 days of incubation in an osteogenic medium. By changing the crosslinking agent e.g. to tripolyphosphate it was possible to reduce its degradation rate without harmful side effects [179–180], pores size and wettability, swelling and compressive strength of scaffolds [176– 177]. In order to enhance osteoconductive process supported by a polymer composite with TCP a new strategy was proposed. To improve and accelerate healing of bone grafts for orthopedic and dental use a platelet-rich plasma (PRP) was applied on 3D scaffolds made of PCL/TCP composite materials. The PRP is an autologous source of concentrated platelets that contains several prepackaged growth factors including transforming growth factor (TGF), platelet-derived growth factor (PDGF), and vascular endothelial growth factor (VEGF). Immersion of a PCL/TCP scaffold in PRP fluid and its implanting into a critical-size femoral defects of a rat led after 3 months of the implantation to neovascularization and bone bridging [181–182]. In another study, when PCL/TCP scaffolds in combination with PRP were used in critical-sized defects healing of a canine mandible the results were confirmed in long-term observations: a bone volume fraction after both control times i.e. 6 and 9 months was better than in the control group. The regular deposition of osteoid with alveolar bone trabeculae was shown for all treated defects [183]. 6.3. Biphasic calcium phosphate based composites Biphasic calcium phosphate ceramics are a family of two-phase materials that combine the low solubility and osteoconductive character of HAp with the osteoinductivity of a more soluble phase, namely β-TCP. The main advantage is the possibility of controlling bioactivity, resorption rate and also mechanical properties of CaP ceramic by manipulating the HAp/β-TCP ratio [138,185–186]. BCP properties, especially biological effects, can be modified also by ion substitution. Kim et al. showed that that Sr incorporation into CaP ceramics increased a proliferation rate and differentiation (ALP activity) of the MG63 and HOS osteoblast-like cells. Such behaviours can be attributed to the combined effects, that is, not only to the Sr ions release but also to higher

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concentration of β-TCP phase in the BCP [187]. In turn, Mg substitution of BCP ceramics enhanced their biodegradation and bioactivity, and also improved biocompatibility, as well as in vitro and in vivo osteoconductivity tested with human adipose tissue-derived mesenchymal stem cells (hAT-MSCs) [188]. Above-mentioned features make BCP ceramics useful as a modifier of biodegradable synthetic polymers and copolymers; PCL [189], PLLA [190], PLGA [191–192] as well as natural polymers and their blends; collagen [193–194], gelatin-pectin [195]. Ebrahimian-Hosseinabadi et al. demonstrated that the modification of PLGA with various amounts (10–50 wt.%) of BCP nanoparticles affects in vitro degradation behavior of the composite scaffolds. On one hand, with increasing amount of BCP nanoparticles in the composites, reduction of Young's modulus increased as a result of higher weight loss of the scaffolds. It could be related to an increase in hydrophilicity and the water uptake of the composite scaffolds. On the other hand, a buffering effect of basic degradation products of BCP on a degradation process of PLGA was observed. Namely, more pronounced decrease in the molecular weight of the polymer with reducing amount of BCP in the matrix occurred during the degradation process [192]. Another work showed that BCP particles, homogeneously dispersed in PLGA nanofibers, enhanced expression of osteogenic differentiation markers (ALP and BSP) and also ECM mineralisation as a function of increasing content of the ceramic particles in the composites. These findings demonstrated a beneficial effect of BCP inclusion in PLGA nanofibers on the osteogenic differentiation of osteoblast-like MC3T3E1 cells and confirmed osteoinductive character of the composites [191]. The incorporation of BCP nanoparticles in another polymer matrix, namely PCL, resulted in an increase in ALP activity of human MSCs and also significantly improved compressive strength and Young's modulus of the materials [189]. In order to enhance mechanical properties of PLLA-based composites, the surface of BCP microparticles was modified by direct grafting with L-lactide. The results showed improved interfacial interaction between the BCP filler and the polymer matrix, and thus higher compressive strength in comparison to composites containing the unmodified particles [190]. It was also shown that when TIPS [192] and SCPL [189] methods were used to obtain PLGA and PCL-based composite scaffolds the porosity of the materials decreased with increasing BCP nanoparticles content. A commercially available collagen/BCP composite (Osteon™ II collagen, Genoss. Co. Ltd) was studied using a rabbit calvarial defect model by Lee et al. [193–194]. The composite material showed good osteoconductive properties and also relatively slow resorption and thus, in contrast to collagen sponge, it maintained the space needed for bone tissue regeneration [193]. It was also shown, that the collagen/BCP composite is a promising candidate for a carrier of recombinant human BMP-2 providing a constant release profile [194]. In turn, a gelatin-pectin/BCP nanocomposite scaffold was fabricated for delivering growth factors (BMP-2 and VEGF). The release studies exhibited that the scaffold had constant release properties. Furthermore, in vitro and in vivo tests conducted with osteoblast-like MC3T3-E1 and rat models, respectively, showed that the composites enhanced cell proliferation and new bone formation [195]. To conclude, calcium phosphates, due to their wide range of solubility, bioactivity, osteoconductivity/osteoinductivity and various mechanical behaviours, can be successfully applied as fillers of biodegradable polymer matrices to modulate composite properties. 7. Summary As pointed out above, ceramic materials discussed in this work, namely silica, bioactive and resorbable glasses, wollastonite and calcium phosphate ceramics have a number of unique and beneficial properties for medical applications, especially for tissue engineering and regenerative medicine. Depending on the type of ceramics, these properties include; bioactivity, osteoconductivity, osteoinductivity, resorbability,

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antibacterial and anti-inflammatory activity, ability to induce vascularisation and many others. As it has been shown in this review, the current biomaterials age belongs to composite materials which are classified as the fourth generation of biomaterials. They combine not less than two different phases i.e.; polymer and ceramic particles or nanoparticles. These new materials are mostly characterized by many advantageous physicochemical, mechanical and biological properties and what is even more important, they give scientists the opportunity of a better control of such properties as; mechanical behavior, degradation rate, surface topography, surface charge and wettability and interaction with cells. Many of the new composite materials modified with the ceramic fillers gain unique futures as antibacterial properties or induct specific cells reactions e.g.; differentiation or activation of secretion of growth factors. Apart from the ceramic filler amount, other parameters such as its size (nanometric or micrometric), its shape (particles, whiskers, nanotubes, fibers), distribution and a type of surface functional groups influence final properties of the composites. Type of matrix polymer and its parameters, such as the molecular weight, polydispersity, crystallinity, chain orientation, functional groups and overall hydrophilicity also strongly affect the composites properties. Of a great importance and impact for applications in the medicine is that natural and synthetic polymers have good formability. As presented in Table 3, numerous methods of fabrication of polymer-based composites combined with many types of ceramic fillers allow obtaining the composites in various forms (porous scaffolds, films, membranes, fibers) and with different morphology and mechanical properties. Moreover, as described in the subsequent sections of this work, these composites also vary in terms of biological effects and other relevant properties. The present work shows that through the selection of materials of the ceramic filler and the matrix, their properties and relationships, it is possible to design and obtain materials with a wide range of properties for various applications in the medicine, especially in the bone defects treatment. To conclude, the composite materials provide a multistage design approach and therefore greater possibilities to control their material and biological properties than the ceramics and the polymers alone. Acknowledgments The authors would like to acknowledge the financial support from the National Science Centre, Poland Grant Nos. 2015/17/N/ST8/00226 (MD), 2014/13/B/ST8/02973 (KCK), 2012/07/B/ST8/03378 (ESZ) and from the Polish Ministry of Science and Higher Education Grant No. N N507 401 939 (ESZ).

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