Mythili Prakasam
email: mythili.prakasam@icmcb.cnrs.fr

Ali Chirazi
email: ali.chirazi@fei.com

Grzegorz Pyka
email: grzegorzpyka@fei.com

Anna Prokhodtseva
email: anna.prokhodtseva@fei.com

Daniel Lichau
email: daniellichau@fei.com

Alain Largeteau
email: alain.largeteau@u-bordeaux.fr

Fabrication and multiscale structural properties of interconnected porous biomaterial for tissue engineering by Freeze Isostatic Pressure (FIP)

Keywords: tissue engineering, porous materials, microstructure, biomaterials, bone regeneration

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Introduction

Repairing of human injured tissues with the help of biomaterials was evidenced even in the pre-historic period [START_REF] Ratner | Biomaterials Science-An Introduction to Materials in Medicine[END_REF]. Usage of metals such as gold for dental applications was observed even in the Roman era [START_REF] Anderson | Smart biomaterials[END_REF]. At present, huge research progress is observed in the case of bone tissue engineering. Bone tissue engineering is a field with continuous evolution and the demand of this field is continuously increasing owing to the growing population. In addition, bone tissue engineering is also in demand due to orthopedic reconstructions required in the case of trauma, birth defects, and degeneration. Various types of implants are employed for orthopedic applications such as autografts, allografts, and xenografts. Autografts have the upper hand in terms of superior bone fusion whereas allografts and xenografts have high probability in transmitting pathogens and causing inflammatory reactions in the human body. The aforesaid issues have led to the development of synthetic biomaterials for musculoskeletal tissue engineering to satisfy the requirements of orthopedics.

Human bone can be defined as a composite of hydroxyapatite [START_REF] Ratner | Biomaterials Science-An Introduction to Materials in Medicine[END_REF] (HAp) Ca 10 (OH) 2 (PO 4 ) 6 , type-I collagen, water, cells, and lipids. Bones [START_REF] Anderson | Smart biomaterials[END_REF][START_REF] Alvarez | Metallic scaffolds for bone regeneration[END_REF][START_REF] Li | The development of binary Mg-Ca alloys for use as biodegradable materials within bone[END_REF][START_REF] Wu | Hydrogen release from titanium hydride in foaming of orthopedic NiTi scaffolds[END_REF][START_REF] He | Collagen-infiltrated porous hydroxyapatite coating and its osteogenic properties: In vitro and in vivo study[END_REF][START_REF] Weiner | The material bone: Structure-mechanical function relations[END_REF][START_REF] Legeros | Properties of osteoconductive biomaterials: Calcium Phosphates[END_REF][START_REF] Gauthier | Macroporous biphasic Calcium Phosphate ceramics: Influence of macropore diameter and macroporosity percentage on bone ingrowth[END_REF][START_REF] Davies | In vitro modeling of the bone/implant interface[END_REF] are formed in the body as a result of the osteoblast matrix formed by HAp crystals. Bone is comprised of two distinct forms: one is porous (cancellous bone) and the other dense (cortical bone). Cancellous bone contains hemocytoblasts, proerythroblasts, and bone marrow. Cancellous bone has a lower Young's modulus and is more elastic compared to cortical bone. The porous structure consists of pore sizes in the range of 200-500 µm, and cancellous bone constitutes 30%-90% of the porosity. The porosity content alters depending on the load, age, and health state of the bone. The constituents of human bone are 60%-70% bone mineral (Ca 10 (PO 4 ) 6 (OH) 2 ), 10%-20% collagen, and 10%-20% water. In the 1920s, stainless steel metallic implants were developed for bone repair. Later, artificial bone implants were based on ceramics, composites, metals, and polymers for bone engineering [START_REF] Alvarez | Metallic scaffolds for bone regeneration[END_REF][START_REF] Li | The development of binary Mg-Ca alloys for use as biodegradable materials within bone[END_REF][START_REF] Wu | Hydrogen release from titanium hydride in foaming of orthopedic NiTi scaffolds[END_REF][START_REF] He | Collagen-infiltrated porous hydroxyapatite coating and its osteogenic properties: In vitro and in vivo study[END_REF].

Synthetic scaffolds developed should be capable of withstanding the mechanical loading, structural integrity, and possibility to support or host tissue and allow cell proliferation. Development of cancellous bone implants primarily depends on the structural design of the scaffold such as porosity, pore shape and size, interconnectivity of pores, orientation, and control of microstructure. Well defined mechanical modulus and porosity of the scaffold will help in accelerating the bone tissue regeneration and support growth of new bone tissue cells. Mechanical properties of natural bone are found to differ based on the asymmetry and anisotropy. Further, the highest modulus is observed in the bone when the loading axis is perpendicular to the axis of the long bone [START_REF] Weiner | The material bone: Structure-mechanical function relations[END_REF]. Research data has shown an estimation of minimum pore size of ~50 µm for blood vessel formation and a ~200 µm for osteonal in-growth [START_REF] Legeros | Properties of osteoconductive biomaterials: Calcium Phosphates[END_REF]. Both porosity and their architecture are critical in gauging the fluid transport rate through porous bioceramics, which determines the rate and the degree of the bone growth in vivo [START_REF] Gauthier | Macroporous biphasic Calcium Phosphate ceramics: Influence of macropore diameter and macroporosity percentage on bone ingrowth[END_REF].

The fabrication of porous scaffolds with complex architectures represents a challenge in tissue engineering. Recent studies have shown that it is possible to construct tissue-engineered bone repair scaffolds with tight pore size distributions and controlled geometries using rapid prototyping techniques. Despite the wide use of porous scaffolds, its design, geometry optimization, and mechanical assessment require further developments and studies for successful integration in tissue engineering. Calcium phosphate scaffolds are regarded as an interesting material for scaffold application. Calcium phosphate (CP) based materials aid in osteoblast adhesion and proliferation [START_REF] Davies | In vitro modeling of the bone/implant interface[END_REF][START_REF] Anselme | Osteoblast adhesion on biomaterials[END_REF]. Various processing routes to fabricate CP-compounds include uniaxial compaction, cold or hot isostatic pressing, granulation, loose packing, slip casting, gel casting, pressure mold forming, injection molding, polymer replication, extrusion, slurry dipping, and spraying. In addition, formation of ceramic sheets by tape casting is also widely employed. Sintering of bioceramics is also well-known [START_REF] Prakasam | Properties and Applications of Dense Hydroxyapatite: A Review[END_REF].

Porous biomaterials consolidated at low temperature with increased mechanical strength can enable inclusion of therapeutic molecules and is one of the domains of investigation at ICMCB-CNRS (Institut de Chimie de la Matière Condensée de Bordeaux-CNRS). The present article discusses the innovative experimental methodology used in the fabrication of porous scaffolds by freeze isostatic pressing (FIP). To demonstrate the feasibility of FIP, TiO 2 anatase phase was selected. TiO 2 has various applications such as in dye-sensitized solar cells, hydrogen production and storage, sensors, rechargeable batteries, electro catalysis, self-cleaning and antibacterial surfaces, and photocatalytic cancer treatment. No influence of the fabrication technology on the chemical composition of the starting compounds should be altered. To show the proof of concept on the non-influence of the fabrication technology on the composition, we have used TiO 2 in the present work. The interconnected porosity and the pore volume were studied by employing 3D-correlative and multimodal imaging techniques to give an insight about the pore structure characteristics, homogeneity, and interconnectivity. This type of an integrated approach will throw light on pore size range, their size and form by employing a wide variety of imaging techniques such as multiscale X-ray tomography, FIB-SEM Slice & View stacking, and high-resolution STEM-EDS electronic tomography. This combined approach allows quantification of morphological and geometrical spatial distributions of the multiscale porous network through length scales spanning from tens of microns to less than a nanometer. The spatial distribution of the wall thickness was investigated as well and its possible relationship with pore connectivity and size distribution were studied and is discussed in detail in this article. The observations of the microstructure through these imaging techniques enabled us to quantify and validate the efficiency of our fabrication technique for porous biomaterials used in bone tissue engineering applications.

Results

Simultaneous study of the nano-and micro structural properties of materials and quantitative assessment of their relationship with the micro-or macroscopic functional properties of the materials are crucial prerequisites for development of imaging for porous biomaterials. The 3D imaging methodology will also assist in quantitative and qualitative analyses when used in conjunction with other experiments requiring nanotoxicity and cell interaction studies. Scaffolds, being an integral part of bone tissue engineering, require their three-dimensional area to be quantified. Three-dimensional information on scaffolds will help to analyze properties such as mechanical support, protein production through biochemical and mechanical interactions to aid in cell attachment, assisting in bone tissue formation. Chemical composition and fabrication methodology are the major parameters to judge their suitability to be used as scaffold. Fabrication methodology governs the pore size, pore volume, and mechanical strength of the scaffold and their performance. To date, conventionally used techniques for fabrication of porous scaffolds involve foaming, solvent casting, leaching, freeze casting, thermally induced phase separation, and additive manufacturing techniques. Apart from pore size and pore volume, the interconnectivity of the pores is essential to judge their role in their functioning. Open and interconnected pores allow nutrients and molecules into the inner part of the scaffold to enable vascularization. High porosity presents high surface area per unit volume. Biodegradation process is highly dependent on pore parameters and cell-mediated process through chemical dissolution to replace the scaffold with new bone. Pore sizes >300 µm of scaffolds assists in vascularization and enhanced bone formation. Reports on poly-(D, L-Lactic acid) scaffolds with pore sizes in the range of 320-420 µm proved to produce a well-organized collagen network. However, when the pore sizes were smaller, <270 µm prevented osteosarcoma derived osteoblasts to proliferate and produce extracellular matrix. High pore volume also aids in compressive strength.

The scaffold fabricated should initially support and resist the host tissue for optimum bone healing. The degradation and mechanical properties depend on the pore size, geometry, and load application direction. The pore size and morphology are tailored and controlled based on the fabrication methodologies. Interconnectivity of the scaffolds is also essential to analyze but is difficult to control and analyze during fabrication methodology. Additive manufacturing and rapid prototyping allows complex scaffolds shapes. Interconnectivity in porous resorbable bioceramics are of utmost importance in comparison to their pore and interconnectivity density. Microstructural studies have enabled the quantifying and qualifying of the scaffolds for tissue engineering. µ-CT scanning is used to quantify the pore size and porosity of the scaffolds. µ-CT imaging and analysis is used to determine porosity and pore sizes in 3D biomaterial scaffolds. To identify the interconnectivity further, microstructural analyses and software were developed.

Discussion

Freeze Isostatic Pressure: FIP Application of isostatic pressure in ambient temperatures on dry metallic powders such as Fe and Al led to a full densification by plastic deformations to achieve cold welding due to the ductility of metallic materials. In the case of inorganic compounds such as Alumina (Al 2 O 3 ), no plastic deformation was observed resulting in less densification due to the repacking of grains as the applied pressure was below the yield strength of grains [START_REF] Eksi | Effects of Powder Hardness and Particle Size on the Densification of Cold Isostatically Pressed Powders[END_REF]. Other low temperature sintering techniques also exist for the fabrication of biocomposites such as hydrothermal sintering. Hydrothermal reaction can be defined as "any heterogenous chemical reaction in the presence of a solvent (whether aqueous or non-aqueous) above room temperature and at pressure greater than 1 atm in a closed system" [START_REF] Byrappa | Handbook of Hydrothermal Technology; A Technology for Crystal Growth and Materials Processing[END_REF]. Sintering is achieved through thermal treatment of the green bodies or powders creating grain bonding into a predominantly solid structure via mass transport events occurring on the atomic scale. The bonding by formation of necks leads to improved strength and a lower system energy. The mechanisms for sintering depend on the conditions within the system, whether there is liquid present as well as solid, and whether there is an additional external applied pressure. Diffusion in liquid state is much more rapid than solid state diffusion and can exist at lower temperature. Bouville and Studart [START_REF] Bouville | Geologically-inspired strong bulk ceramics made with water at room temperature[END_REF] demonstrated that cold isostatic pressure (CIP) on powder in wet conditions of powder for ceramics (CaCO 3 ) leads to sintering in presence of water at room temperature of an inorganic compound moderately soluble in water. Ion dissolution and liquid induced plasticity is not favored either in oil or in air, however, a dissolution-precipitation and/or plastic deformation caused by solvent (water in this case) could lead to the densification of the powder [START_REF] Bouville | Geologically-inspired strong bulk ceramics made with water at room temperature[END_REF]. When the grains are in contact under external force, there is a high stress concentration induced by pressure solution creep caused during the movement of ions. Dissolution of the contact point takes place in the high stress region in the presence of the liquid interface layer. The dissoluted particles precipitate in the less stressed zone. This dissolution at the grain contact point is akin to dissolution-precipitation processes that are observed in the cold sintering process and in the hydrothermal sintering process [START_REF] Ndayishimiye | Low temperature hydrothermal sintering process for the quasi-complete densification of nanometric α-quartz[END_REF]. Plastic deformation will also be favored in the metallic powders in addition to the dissolution-precipitation process leading to consolidation under high pressure in presence of a liquid medium. In the case of the ceramic powders, the dissolution and precipitation process plays a major role in consolidation in the presence of high pressure and a liquid interface layer. The role of liquid is indispensable for consolidation with high pressure and low temperatures.

In nature, only the crystals of ice I exist; pressure sintering of ice crystals is relevant for the densification of glaciers. The necks form between grains of ice crystals by diffusion and plastic flow due to the pressure [START_REF] Blackford | Sintering and microstructure of ice: A review[END_REF]. In the phase diagram of water, it can be inferred that there is a possibility to retain the liquid form of ice even at -20 • C in the presence of high pressure (around 200 MPa). The pressure-temperature diagram of water is already well discussed by many authors in the literature [START_REF] Bail | High pressure freezing and thawing of foods: A review[END_REF]. From the phase diagram of water, the possibility to variate the formation and shape of ice crystals and to control the size by different ways by selecting the coupled pressure/temperature were already explored in pressure-assisted freezing and pressure-assisted thawing applied in foods [START_REF] Bail | High pressure freezing and thawing of foods: A review[END_REF]. The control of formation of ice crystals favors fine nucleation without growing or, in contrast, to permit the growth of a massive number of small ice crystals and also to obtain large crystals of ice I, ice III, or Ice V regarding the level of pressure, as shown in the Figure 1, adapted from Le Bail et al. [START_REF] Bail | High pressure freezing and thawing of foods: A review[END_REF]. Ice VI exists at ambient temperature under very high pressure, higher than 600 MPa.

Sintering is achieved through thermal treatment of the green bodies or powders creating grain bonding into a predominantly solid structure via mass transport events occurring on the atomic scale. The bonding by formation of necks leads to improved strength and a lower system energy. The mechanisms for sintering depend on the conditions within the system, whether there is liquid present as well as solid, and whether there is an additional external applied pressure. Diffusion in liquid state is much more rapid than solid state diffusion and can exist at lower temperature. Bouville and Studart [START_REF] Bouville | Geologically-inspired strong bulk ceramics made with water at room temperature[END_REF] demonstrated that cold isostatic pressure (CIP) on powder in wet conditions of powder for ceramics (CaCO3) leads to sintering in presence of water at room temperature of an inorganic compound moderately soluble in water. Ion dissolution and liquid induced plasticity is not favored either in oil or in air, however, a dissolution-precipitation and/or plastic deformation caused by solvent (water in this case) could lead to the densification of the powder [START_REF] Bouville | Geologically-inspired strong bulk ceramics made with water at room temperature[END_REF]. When the grains are in contact under external force, there is a high stress concentration induced by pressure solution creep caused during the movement of ions. Dissolution of the contact point takes place in the high stress region in the presence of the liquid interface layer. The dissoluted particles precipitate in the less stressed zone. This dissolution at the grain contact point is akin to dissolution-precipitation processes that are observed in the cold sintering process and in the hydrothermal sintering process [START_REF] Ndayishimiye | Low temperature hydrothermal sintering process for the quasi-complete densification of nanometric α-quartz[END_REF]. Plastic deformation will also be favored in the metallic powders in addition to the dissolution-precipitation process leading to consolidation under high pressure in presence of a liquid medium. In the case of the ceramic powders, the dissolution and precipitation process plays a major role in consolidation in the presence of high pressure and a liquid interface layer. The role of liquid is indispensable for consolidation with high pressure and low temperatures.

In nature, only the crystals of ice I exist; pressure sintering of ice crystals is relevant for the densification of glaciers. The necks form between grains of ice crystals by diffusion and plastic flow due to the pressure [START_REF] Blackford | Sintering and microstructure of ice: A review[END_REF]. In the phase diagram of water, it can be inferred that there is a possibility to retain the liquid form of ice even at -20 °C in the presence of high pressure (around 200 MPa). The pressure-temperature diagram of water is already well discussed by many authors in the literature [START_REF] Bail | High pressure freezing and thawing of foods: A review[END_REF]. From the phase diagram of water, the possibility to variate the formation and shape of ice crystals and to control the size by different ways by selecting the coupled pressure/temperature were already explored in pressure-assisted freezing and pressure-assisted thawing applied in foods [START_REF] Bail | High pressure freezing and thawing of foods: A review[END_REF]. The control of formation of ice crystals favors fine nucleation without growing or, in contrast, to permit the growth of a massive number of small ice crystals and also to obtain large crystals of ice I, ice III, or Ice V regarding the level of pressure, as shown in the Figure 1, adapted from Le Bail et al. [START_REF] Bail | High pressure freezing and thawing of foods: A review[END_REF]. Ice VI exists at ambient temperature under very high pressure, higher than 600 MPa. Pressure is easier to modify than temperature, which is subjected to the mass inertia of the pressure-vessel. The pressure shift on the frozen mass allows smaller and more homogeneous ice crystals during the starting of level of pressure from depressurization at constant temperature, and also the rate of depressurization. ABEFG: pressure shift freezing (PSF), freezing process increases the ice-nucleation rate. d/Phase transition under a pressure change continued at a constant pressure. GFEBA: pressure-induced thawing (PIT), reverses process of PSF. e/Phase transition under a pressure change continued at a constant pressure. ABEFHI: freezing to ice III and IHFEBA: thawing to ice III, ice III forms with higher density than liquid water and ice I.

Crystallization and liquefaction are controlled by level of pressure and temperature, but also by the pressurization and depressurization rate which could be very rapid; the temperature rate is more complicated to control quickly in a pressured vessel.

In the present work, we have worked on using ice as a template similar to a porogen and apply the pressure at the same time to consolidate the materials. The possibility to freeze the water at different temperatures and to create different morphologies of ice in the presence of high pressure is explored. This innovative technique called freeze isostatic pressure (FIP) consists on the application of pressure at minus temperature on a mixture of sample made of powder and water. Water is used as a template under solid state which is eliminated after return to ambient condition of pressure and temperature. Powder made of inorganic compound is compacted on such template which gives the porosity [START_REF] Largeteau | Trends in high pressure developments for new perspectives[END_REF]. FIP allows the control of the ice crystals nucleation/growing of water. Water is easily eliminated by return at ambient conditions through sublimation. The objective by using the parameters P and T, is to select the path for selecting the crystals I, III, and V. The characteristics of such ice are as follows: Ice I (0.92 g•cm -3 ), Ice III (1.14 g•cm -3 ), and Ice V (1.23 g•cm -3 ). Both ice III and V are denser than ice I and liquid phase. The rate of compression and decompression, and freezing and thawing, determine the final size of porosities initiated by crystallization of ice crystals. At the contact point between grains of powder inorganic compound, the dissolution and precipitation into the interfacial water film occurs and could be also activated by a solid-state welding processes (cold-welding) process to obtain a well consolidated monolith. The dissolution localized at grains contact favored by high stress contact point acts like the dissolution and precipitation processes expected for HyS process, explained previously where temperature enhances this phenomenon. The transportation of species along the grain boundary in the less stress contact point leads to a precipitation on the grain surface and initiates the neck formation in the free spaces between grains, acting like osmotic pressure effect.

Then, we could assume that hydrothermal reaction as "any heterogenous chemical reaction in the presence of a solvent (whether aqueous or no aqueous) under liquid state at pressure greater than 1 atm in a closed system" because the solvent under liquid phase-even if temperature is negative-could enhance dissolution and precipitation by pressure at the contact point of the grains. This consolidation between grains could be also enhanced by adding binder such as collagen, gelatin, polymer, and so on. Figure 2 shows the TiO 2 anatase porous ceramics obtained by applying various pressures and temperatures. An example of pressure and temperature condition which yielded porous TiO 2 ceramics was the pathway FG following a return to ambient temperature with the extraction of the sample from the pressure vessel. Interconnected pores were obtained at the aforesaid experimental conditions.

Materials and Methods

As noted above, an innovative fabrication methodology of the porous monolith was presented in this article. ICMCB, France has developed a very low temperature processing method that could help in obtaining high strength materials with the help of parameters such as pressure, temperature, and dwell time. The aforesaid method of consolidation yielded the monolith with porosity in the range of 80-93%. We have used Titanium oxide (TiO2) in anatase phase (grain size ~15 nm). The porous monolith consisted of high porous body with fine thin pore walls and interconnected spaces. The presence of the interconnected pores will assist in cell proliferation. With the currently existing imaging techniques, it is difficult to differentiate the thin pore walls in the materials fabricated at low temperatures with little mechanical strength. Therefore, the aim of the current study was to demonstrate the feasibility and the necessity of a multiscale and multi-modal correlative approach to the characterization and structural quantification of such a material system as detailed below and shown in Figures 34567. Figure 8 shows a summarized and schematic correlative workflow used for this study.

Low resolution μCT on HeliScan™

•

The biocomposite sample with dimensions of 4.3 mm × 4.6 mm × 6.4 mm was examined to select smaller region of interest (ROI) for further high-resolution imaging. The sample was scanned with double-helix trajectory with projections per revolution set to 2880. Total amount of the acquired projections was 10,180. The volume had an isotropic voxel size of 2.35 μm. AvizoTM software was applied for building 3D model of the sample and selection of the ROI for further investigation. Sample block of 1.3 mm × 1.3 mm × 2 mm was cut off using a scalpel according to the 3D model based defined ROI.

High resolution μCT on HeliScan™

•

The sample ROI was examined afterwards, using HeliScan™ system. The voxel size of the reconstructed images was 0.84 μm.

High resolution Automated Slice & View stacking on Scios DualBeam™ microscope

• Simultaneous slicing with FIB and SEM imaging was performed on FEI's Scios DualBeam™ microscope (Merignac, France). Auto Slice and View 4.0 software was used to automatically collect 3D data from a user-defined volume of the sample by milling serial sections (slices) and then acquiring high resolution images of each slice. This software enabled study of the 3D structure and composition of samples at the nanometer scale.

Materials and Methods

As noted above, an innovative fabrication methodology of the porous monolith was presented in this article. ICMCB, France has developed a very low temperature processing method that could help in obtaining high strength materials with the help of parameters such as pressure, temperature, and dwell time. The aforesaid method of consolidation yielded the monolith with porosity in the range of 80%-93%. We have used Titanium oxide (TiO 2 ) in anatase phase (grain size ~15 nm). The porous monolith consisted of high porous body with fine thin pore walls and interconnected spaces. The presence of the interconnected pores will assist in cell proliferation. With the currently existing imaging techniques, it is difficult to differentiate the thin pore walls in the materials fabricated at low temperatures with little mechanical strength. Therefore, the aim of the current study was to demonstrate the feasibility and the necessity of a multiscale and multi-modal correlative approach to the characterization and structural quantification of such a material system as detailed below and shown in Figures 34567. Figure 8 shows a summarized and schematic correlative workflow used for this study.

Low resolution µCT on HeliScan™

•

The biocomposite sample with dimensions of 4.3 mm × 4.6 mm × 6.4 mm was examined to select smaller region of interest (ROI) for further high-resolution imaging. The sample was scanned with double-helix trajectory with projections per revolution set to 2880. Total amount of the acquired projections was 10,180. The volume had an isotropic voxel size of 2.35 µm. AvizoTM software was applied for building 3D model of the sample and selection of the ROI for further investigation. Sample block of 1.3 mm × 1.3 mm × 2 mm was cut off using a scalpel according to the 3D model based defined ROI.

High resolution µCT on HeliScan™

•

The sample ROI was examined afterwards, using HeliScan™ system. The voxel size of the reconstructed images was 0.84 µm.

High resolution Automated Slice & View stacking on Scios DualBeam™ microscope

• Simultaneous slicing with FIB and SEM imaging was performed on FEI's Scios DualBeam™ microscope (Merignac, France). Auto Slice and View 4.0 software was used to automatically collect 3D data from a user-defined volume of the sample by milling serial sections (slices) and then acquiring high resolution images of each slice. This software enabled study of the 3D structure and composition of samples at the nanometer scale.

•

Area of interest of 23 µm × 20 µm × 14 µm was analyzed, thickness of a unique slice removed by FIB being 15 nm, total of 933 slices were milled. SEM imaging was done with SE and BSE in-lens detectors simultaneously, image resolution set to 1536 × 1024 pixels. In this way, voxel size is 16 nm × 20 nm × 15 nm.

High resolution TEM

• A TEM thin foil sample was lifted out near the previously done Slice and View site. Due to the delicate nature of the sample, the lifted out chunk was thinned to 300 nm thickness with 30 kV ion beam. 5 kV and 2 kV ion beam cleaning was then performed to make the sample electron transparent.

detectors simultaneously, image resolution set to 1536 × 1024 pixels. In this way, voxel size is 16 nm × 20 nm × 15 nm.

High resolution TEM

• A TEM thin foil sample was lifted out near the previously done Slice and View site. Due to the delicate nature of the sample, the lifted out chunk was thinned to 300 nm thickness with 30 kV ion beam. Five kV and 2 kV ion beam cleaning was then performed to make the sample electron transparent. Porous media morphological changes, from more elongated to purely spherical pores, while decreasing in pore size, is also observed and quantified as shown in Figures 9101112.

To summarize quantifications, the three-dimensional interconnectivity of the porous media through different scales for the innovative manufactured biomaterial was investigated using newly developed 3D correlative and multi-modal imaging techniques. Porous media morphological changes, from more elongated to purely spherical pores, while decreasing in pore size, is also observed and quantified as shown in Figures 9101112.

To summarize quantifications, the three-dimensional interconnectivity of the porous media through different scales for the innovative manufactured biomaterial was investigated using newly developed 3D correlative and multi-modal imaging techniques.

I

Multiscale X-ray tomography, FIB-SEM Slice & View stacking and high-resolution STEM-EDS electronic tomography observations were combined allowing quantification of morphological and geometrical spatial distributions of the multiscale porous network through length scales spanning from hundreds of microns to nanometer. II 3D macro-micro-nano pore and defect network connectivity, size distribution, anisotropy, and form factors were quantified. III 3D wall thickness distribution and localized internal wall's densification were investigated. Porous media morphological changes, from more elongated to purely spherical pores, while decreasing in pore size, is also observed and quantified as shown in Figures 9101112.

To summarize quantifications, the three-dimensional interconnectivity of the porous media through different scales for the innovative manufactured biomaterial was investigated using newly developed 3D correlative and multi-modal imaging techniques.

Conclusions

Hydrothermal reaction could be defined as "any heterogenous chemical reaction in the presence of a solvent (whether aqueous or no aqueous) under liquid state at pressure greater than 1 atm in a closed system" because the solvent under liquid phase even if temperature is negative (like -20 • C/200 MPa for water) could enhance dissolution and precipitation by pressure at the contact point of the grains.

The innovative process freeze isostatic pressure (FIP) based on this hydrothermal phenomenon allows the reaching of consolidated monoliths at negative temperature by using a clean template, such as water.

To validate the fabrication process, microstructural studies were mandatory and the interconnectivity of the pores. Microstructural studies carried out in the present work enabled the quantifying and qualifying of the scaffolds for tissue engineering. To identify the interconnectivity further microstructural analyses and software were developed. Three-dimensional-correlative and multimodal imaging techniques and the software developed by Thermofisher Bordeaux was highly helpful in quantifying the highly porous biomaterial for tissue engineering applications. Hence, it is foreseen that both the fabrication process and the 3D-correlative and multimodal imaging alongside the software developed can be successfully employed for other porous materials for energy or biomedical applications.

Figure 1 .

 1 Figure 1. Le Bail [18]-Possibilities of high pressure processing based on the phase diagram of water.

Figure 1

 1 Figure 1 shows various pathways possible for the formation of ice crystals: a/Freezing in liquid phase

Figure 1 .

 1 Figure 1. Le Bail [18]-Possibilities of high pressure processing based on the phase diagram of water.

Figure 1

 1 Figure1shows various pathways possible for the formation of ice crystals: a/Freezing in liquid phase Water is in liquid state up to -21 • C under 210 MPa; this state could be the starting state for crystallization of ice crystals. This liquid phase could be reached by different pathways and could be like an initial state for the processes described below. ABEF: freezing without ice crystal formation, water stays in liquid phase b/Phase transition under constant pressure

J

  . Funct. Biomater. 2018, 9, x FOR PEER REVIEW 6 of 12 porous TiO2 ceramics was the pathway FG following a return to ambient temperature with the extraction of the sample from the pressure vessel. Interconnected pores were obtained at the aforesaid experimental conditions.

Figure 2 .

 2 Figure 2. Porous TiO2 ceramics obtained by FIP processing.

Figure 2 .

 2 Figure 2. Porous TiO 2 ceramics obtained by FIP processing.

Figure 3 .

 3 Figure 3. Correlating low resolution (LR) μCT to high resolution (HR) μCT. ROI = region of interest.

Figure 4 .

 4 Figure 4. Correlating high resolution μCT to FIB-SEM Stacking.

Figure 3 .

 3 Figure 3. Correlating low resolution (LR) µCT to high resolution (HR) µCT. ROI = region of interest.

Figure 3 .

 3 Figure 3. Correlating low resolution (LR) μCT to high resolution (HR) μCT. ROI = region of interest.

Figure 4 .

 4 Figure 4. Correlating high resolution μCT to FIB-SEM Stacking.

Figure 4 .

 4 Figure 4. Correlating high resolution µCT to FIB-SEM Stacking.

Figure 5 .

 5 Figure 5. Example of porous wall's thickness variation, topology, and tortuosity.

Figure 6 .

 6 Figure 6. Defining and extracting ROI for TEM lamella.

Figure 7 .

 7 Figure 7. TEM Lamella correlation.

Figure 5 .

 5 Figure 5. Example of porous wall's thickness variation, topology, and tortuosity.

Figure 5 .

 5 Figure 5. Example of porous wall's thickness variation, topology, and tortuosity.

Figure 6 .

 6 Figure 6. Defining and extracting ROI for TEM lamella.

Figure 7 .

 7 Figure 7. TEM Lamella correlation.

Figure 6 .

 6 Figure 6. Defining and extracting ROI for TEM lamella.

Figure 5 .

 5 Figure 5. Example of porous wall's thickness variation, topology, and tortuosity.

Figure 6 .

 6 Figure 6. Defining and extracting ROI for TEM lamella.

Figure 7 .

 7 Figure 7. TEM Lamella correlation.Figure 7. TEM Lamella correlation.

Figure 7 .

 7 Figure 7. TEM Lamella correlation.Figure 7. TEM Lamella correlation.

Figure 8 .

 8 Figure 8. Schematics of full correlation workflow.

Figures 9 -

 9  show quite clearly the scale overlap between porosities characterized and quantified at different scales, where the interval of size and volume variations makes multiscale correlation mandatory. Without such an approach, the complete characterization and understanding of the interconnected porous media is impossible.Porous media morphological changes, from more elongated to purely spherical pores, while decreasing in pore size, is also observed and quantified as shown in Figures9-12.To summarize quantifications, the three-dimensional interconnectivity of the porous media through different scales for the innovative manufactured biomaterial was investigated using newly developed 3D correlative and multi-modal imaging techniques.

I

  . Multiscale X-ray tomography, FIB-SEM Slice & View stacking and high-resolution STEM-EDS electronic tomography observations were combined allowing quantification of morphological and geometrical spatial distributions of the multiscale porous network through length scales spanning from hundreds of microns to nanometer. II. 3D macro-micro-nano pore and defect network connectivity, size distribution, anisotropy, and form factors were quantified. III. 3D wall thickness distribution and localized internal wall's densification were investigated.

Figure 9 .

 9 Figure 9. Porous media quantification throughout scales.

Figure 8 .

 8 Figure 8. Schematics of full correlation workflow.

Figures 9 -

 9  show quite clearly the scale overlap between porosities characterized and quantified at different scales, where the interval of size and volume variations makes multiscale correlation mandatory. Without such an approach, the complete characterization and understanding of the interconnected porous media is impossible.Porous media morphological changes, from more elongated to purely spherical pores, while decreasing in pore size, is also observed and quantified as shown in Figures9-12.To summarize quantifications, the three-dimensional interconnectivity of the porous media through different scales for the innovative manufactured biomaterial was investigated using newly developed 3D correlative and multi-modal imaging techniques.

Figure 8 .

 8 Figure 8. Schematics of full correlation workflow.

Figures 9 -

 9  show quite clearly the scale overlap between porosities characterized and quantified at different scales, where the interval of size and volume variations makes multiscale correlation mandatory. Without such an approach, the complete characterization and understanding of the interconnected porous media is impossible.Porous media morphological changes, from more elongated to purely spherical pores, while decreasing in pore size, is also observed and quantified as shown in Figures9-12.To summarize quantifications, the three-dimensional interconnectivity of the porous media through different scales for the innovative manufactured biomaterial was investigated using newly developed 3D correlative and multi-modal imaging techniques.

I

  . Multiscale X-ray tomography, FIB-SEM Slice & View stacking and high-resolution STEM-EDS electronic tomography observations were combined allowing quantification of morphological and geometrical spatial distributions of the multiscale porous network through length scales spanning from hundreds of microns to nanometer. II. 3D macro-micro-nano pore and defect network connectivity, size distribution, anisotropy, and form factors were quantified. III. 3D wall thickness distribution and localized internal wall's densification were investigated.

Figure 9 .

 9 Figure 9. Porous media quantification throughout scales.Figure 9. Porous media quantification throughout scales.

Figure 9 .

 9 Figure 9. Porous media quantification throughout scales.Figure 9. Porous media quantification throughout scales.

Figure 10 .

 10 Figure 10. Porous media quantification throughout scales.

Figure 11 .

 11 Figure 11. Porous media quantification throughout scales.

Figure 12 .

 12 Figure 12. Porous media quantification throughout scales.

Figure 10 . 12 Figure 10 .

 101210 Figure 10. Porous media quantification throughout scales.

Figure 11 .

 11 Figure 11. Porous media quantification throughout scales.

Figure 12 .

 12 Figure 12. Porous media quantification throughout scales.

Figure 11 .

 11 Figure 11. Porous media quantification throughout scales.

Figure 10 .

 10 Figure 10. Porous media quantification throughout scales.

Figure 11 .

 11 Figure 11. Porous media quantification throughout scales.

Figure 12 .

 12 Figure 12. Porous media quantification throughout scales.Figure 12. Porous media quantification throughout scales.

Figure 12 .

 12 Figure 12. Porous media quantification throughout scales.Figure 12. Porous media quantification throughout scales.

J. Funct. Biomater. 2018, 9, 51; doi:10.3390/jfb9030051 www.mdpi.com/journal/jfb

J. Funct. Biomater. 2018, 9, x FOR PEER REVIEW 4 of 12

J. Funct. Biomater. 2018, 9, x FOR PEER REVIEW 7 of 12

Acknowledgments:

The authors thank Stéphane Toulin for providing the scientific articles for references in the article.

Author Contributions: A.L. and M.P. were involved in Conceptualization and Methodology of FIP processing. A.C., D.L., G.P. and A.P. were equally involved in software and validation related to microstructures. M.P. involved in writing, reviewing and editing. A.L. and M.P. were involved in characterization analyses.

Funding: This research received no external funding.

Conflicts of Interest:

The authors declare no conflict of interest.


Mythili Prakasam
email: mythili.prakasam@icmcb.cnrs.fr

Janis Locs
email: janis.locs@rtu.lv

Kristine Salma-Ancane
email: kristine.salma-ancane@rtu.lv

Dagnija Loca
email: dagnija.loca@rtu.lv

Alain Largeteau
email: alain.largeteau@icmcb.cnrs.fr

Liga Berzina-Cimdina

Biodegradable Materials and Metallic Implants-A Review

Keywords: bioceramics, biocomposites, biodegradable metal alloys, bioimplants

Recent progress made in biomaterials and their clinical applications is well known. In the last five decades, great advances have been made in the field of biomaterials, including ceramics, glasses, polymers, composites, glass-ceramics and metal alloys. A variety of bioimplants are currently used in either one of the aforesaid forms. Some of these materials are designed to degrade or to be resorbed inside the body rather than removing the implant after its function is served. Many properties such as mechanical properties, non-toxicity, surface modification, degradation rate, biocompatibility, and corrosion rate and scaffold design are taken into consideration. The current review focuses on state-of-the-art biodegradable bioceramics, polymers, metal alloys and a few implants that employ bioresorbable/biodegradable materials. The essential functions, properties and their critical factors are discussed in detail, in addition to their challenges to be overcome.

Introduction

Significant advances in science and technology have contributed to various domains, including the medical field. Advances in the field of medical sciences and engineering have created the possibility to employ implants in the human body. Research and development as early as the 1890s, led to the development of pacemakers and defibrillators [START_REF] Elmqvist | Cardiac Pacemakers and Defibrillators[END_REF]. Implants appeasing the requirements for replacing/functioning/assisting the natural organs in the human body are called bioimplants. Biocompatibility and implants' various roles to match the natural bodily functions are studied. Bioimplants [START_REF] Soler-Botija | Engineered 3D bioimplants using elastomeric scaffold, self-assembling peptide hydrogel, and adipose tissue-derived progenitor cells for cardiac regeneration[END_REF] in today's medical scenario are used as brain/neural implants [START_REF] Takmakov | Rapid evaluation of the durability of cortical neural implants using accelerated aging with reactive oxygen species[END_REF], sensory implants [START_REF] Huang | Sensory innervation around immediately vs. delayed loaded implants: A pilot study[END_REF], spinal implants [START_REF] Bagga | Bioactive Spinal Implants and Method of Manufacture Thereof[END_REF], organ stimulation implants [START_REF] Oka | Thirst driving and suppressing signals encoded by distinct neural populations in the brain[END_REF], subcutaneous implants [START_REF] Guo | Substrate modulus of 3D-printed scaffolds regulates the regenerative response in subcutaneous implants through the macrophage phenotype and Wnt signaling[END_REF], dental implants [START_REF] Klein | Systematic review on success of narrow-diameter dental implants[END_REF], cosmetic implants [START_REF] Yahyavi-Firouz-Abadi | Imaging of cosmetic plastic procedures and implants in the body and their potential complications[END_REF], convenience implants and other structural implants [START_REF] Silva | Microstructural and mechanical study of zirconia-hydroxyapatite (ZH) composite ceramics for biomedical applications[END_REF], such as stents, braces, rods, heart valves, bones, pins, hip prosthesis, eye, ear, skull implants and knee replacement implants. An upsurge in bioimplants markets is predicted to reach US$116 billion by 2020, primarily because of the rise in road accidents and trauma injuries. Similar demands are also analogous in the domain of cardiovascular, spinal and ophthalmic disorders, leading to a rise in the demand for bioimplants.

Bioimplants are protheses made for regularizing physiological functions, and bioceramics represent one classification of bioimplants based on material type. The classification of biomaterials [START_REF] Shackelford | Bioceramics-Current status and future trends[END_REF] includes single crystals, polycrystals, glass, glass-ceramics, polymers and composites. The type of interaction and the mechanism of tissue attachment at the implant interface is based on its inherent nature. The type of tissue implant is considered (1) biotoxic if the surrounding tissue dies; (2) it is bioinert-nontoxic and biologically inactive; (3) it is bioactive-nontoxic and biologically active; and (4) bioresorbable/biodegradable-nontoxic and dissolves itself to replace the surrounding tissues. Forms and shapes of bioceramics [START_REF] Hench | An Introduction to Bioceramics[END_REF] are available as dense, porous, bimodal and amorphous states. To avoid the complications of ceramics which are bioinert, the bioactive coating of hydroxyapatite (HAp) is used. Ceramic implants and resorbable/biodegradable bioceramics are the focus of this review. Resorbable bioceramics [START_REF] Shikinami | Bioresorbable devices made of forged composites of hydroxyapatite (HA) particles and poly-L-lactide (PLLA): Part I. Basic characteristics[END_REF] are designed to gradually degrade in a time frame followed by a replacement of natural host tissue. Some of the complications related to resorbable bioceramics are maintaining strength and stability for its whole lifetime (until complete degradation, rate of resorption and the compositional changes in the body). Ideal bioactive [START_REF] Cao | Bioactive materials[END_REF] ceramics should be biocompatible, osteoconductive, osteoinductive, enable angiogenesis, biodegradable, nontoxic and have the flexibility to form different shapes. Bioceramics based on calcium phosphates (CP) with the mineral component similar to natural bone are a very well-known class of materials used for tissue engineering applications such as scaffolds.

Reconstructive surgery using interference screw fixations is prevalently employed in the femur and tibia [START_REF] Harilainen | Randomized prospective study of ACL reconstruction with interference screw fixation in patellar tendon autografts versus femoral metal plate suspension and tibial post fixation in hamstring tendon autografts: 5-year clinical and radiological follow-up results[END_REF]. Prior to this, bioresorbable implants and metallic biomaterials [START_REF] Okazaki | Comparison of metal release from various metallic biomaterials in vitro[END_REF] were used. The disadvantages of metallic biomaterials are varied; for example, the requirement of postoperative complications, distortion of post-operative metallic screw and inflammatory reactions. Recent progress in development of bioresorbable implants for traumatology and sports injuries [START_REF] Weiler | Biodegradable implants in sports medicine: The biological base[END_REF] is continuously evolving. However, there are still reports about the disadvantages of bioresorbable biomaterials, such as inflammation, osteolysis and the decrease in mechanical strength. Polymers [START_REF] Vert | Degradable and bioresorbable polymers in surgery and in pharmacology: Beliefs and facts[END_REF] are used as bioresorbable materials for clinical applications for bone implants, sutures, and as drug delivery devices.

Biodegradable materials, irrespective of their constituent form, are expected to degrade progressively over a period of time to assist as scaffolds or for the healing process. Metals are stiff and have high mechanical strength in comparison to their polymer counterparts; hence, the former is usually chosen for medical applications. Magnesium [START_REF] Chen | Recent advances on the development of magnesium alloys for biodegradable implants[END_REF] and iron [START_REF] Schinhammer | Design strategy for biodegradable Fe-based alloys for medical applications[END_REF] based biodegradable alloys are expected to degrade with the metabolism at an appropriate degradation rate inside the human body. Porous Mg alloy degrades rapidly in vivo. The main characteristics that are considered for porous metallic implants are porosity, pore size and pore interconnectivity, to cater for the replication of natural human-like implants. Higher porosity aids in angiogenesis. Bone is comprised of two distinct forms: one is porous (cancellous bone) and the other dense (cortical bone) [START_REF] Zioupos | Some basic relationships between density values in cancellous and cortical bone[END_REF]. Cancellous bone contains hemocytoblasts, proerythroblasts and bone marrow. Cancellous bone has a lower Youngs modulus and is more elastic compared to cortical bone. The porous structure consists of pore sizes in the range of 200-500 µm, and cancellous bone constitutes 30-90% of the porosity. The porosity content alters depending on the load, age and health state of the bone. Cortical bone is the outer layer of the bone that aids in providing the shape and form of the bone. Eighty percent of the skeleton is composed of cortical bone. Cortical bone stacks osteons or Harversian systems in the form of interstitial lamellae. In the case of loss of bone, bone grafting is used, and solutions are chosen based on the required biomechanical properties, chemical composition, bone mass and size of the defect site. Different types of bone grafting methods are employed [START_REF] Finkemeier | Bone-grafting and bone-graft substitutes[END_REF]. A few of the bone grafting methods are autografting (cancellous/cortical bones), allograft (cancellous/demineralized bone matrix (DBM)) and bone graft substitutes (HAp/tricalcium phosphates Ca 3 (PO 4 ) 2 (TCP)/biphasic calcium phosphates (BCP)/bioactive composites, growth factors).

The objective of this paper is to review the progress and development of biodegradable/bioresorbable biomaterials for various applications and their current challenges and future developments.

Bioresorbable Composites and Implants

Bioceramics

Scaffold design factors [START_REF] Hutmacher | Scaffold design and fabrication technologies for engineering tissues-State of the art and future perspectives[END_REF] in bone tissue engineering are given in Table 1. Biomaterials are considered as the alternative solution for natural bone with similar osteoinduction, osteoconduction, inflammation and mechanical integrity as native bone. Usage of bioceramics implants avoids the risk of disease transmission and immunogenicity. Bioceramics will experience humid conditions upon implantation. Currently, bioceramics are used for applications [START_REF] Legeros | Biphasic calcium phosphate bioceramics: Preparation, properties and applications[END_REF] related to dental, periodontal, cochlear, maxillofacial, spinal discs and otolaryngology. Initially, developed bioinert bioceramics were principally from zirconia (ZrO 2 ) [START_REF] Lohbauer | Hydroxylation of dental zirconia surfaces: Characterization and bonding potential[END_REF] and alumina (Al 2 O 3 ) [START_REF] Wittenbrink | Low-aspect ratio nanopatterns on bioinert alumina influence the response and morphology of osteoblast-like cells[END_REF] to fabricate femoral heads of total hip arthroplasty prostheses. The aforesaid ceramics are well known for their good mechanical properties. Later, bioactive ceramics were developed and are currently widely known for their ability to form the bone bonding without causing inflammation. Calcium phosphates (CPs) [START_REF] Bohner | Calcium orthophosphates in medicine: From ceramics to calcium phosphate cements[END_REF] have high biocompatibility which led to studies on a variety of CPs. Some of the CPs based on hydroxyapatite (HAp, Ca 10 (PO 4 ) 6 (OH) 2 ), dicalcium phosphate dihydrate (DCPD, brushite, CaHPO 4 •2H 2 O) and tricalcium phosphate (TCP, Ca 3 (PO 4 ) 2 ) are extensively studied as bone replacement materials. β-TCP exhibits higher resorbability than HAp under physiological conditions. Bioceramics are commonly used for various applications such as fracture repair, bone defect filling and as bone replacement tissues. Degradable bioceramics [START_REF] De Groot | Clinical applications of calcium phosphate biomaterials: A review[END_REF] (Table 2) have a significantly lower mechanical strength in comparison to their non-resorbable counterparts. The physical and chemical composition of bioceramics determines their biological response. Bioactive ceramics also assist in the production of proteins and cell adhesion (osteoconduction) because of their presence of inorganic phase (CP). Bioceramics, with their capability of stimulating bone regeneration tissues, have led to a new class of biomaterials. Tricalcium phosphate (TCP) [START_REF] Knaack | Resorbable calcium phosphate bone substitute[END_REF] is known to have a close chemical composition similar to bone tissue mineral. TCP has two different crystalline forms: α-TCP and β-TCP. TCP has a good resorbability and bioactivity with a higher rate of biodegradation than hydroxyapatite under in vivo conditions [START_REF] Teramoto | Resorption of apatite-wollastonite containing glass-ceramic and beta-tricalcium phosphate in vivo[END_REF]. Due to its osteoconductivity and resorption properties, it is widely used for bone graft remodeling. TCP is also used for the applications in orthopedic, dental and maxillofacial applications. Complete resorption of TCP is reported in the case of rat tibia and for cancellous bone reformation in canine models. Bone grafts made of TCP show a slow resorbability by osteoclasts over a span of 10 months to 2 years.

Calcium phosphate is the major inorganic component of bone tissue, with hydroxyapatite (HAp) having an atomic ratio of 1.67 for calcium and phosphate. HAp is bioactive and resorbable (when it is either in amorphous or low crystalline state). Natural and synthetic HAp reactivity varies based on the preparation method. Synthetic HAp prepared at high temperature yields high crystallinity. Biodegradation and resorbability of HAp are very slow. HAp bioceramics are commonly used for the small defects in the case of bone loss or in the case of fractures of the tibia, calcaneus and vertebra. HAp is not employed for load-bearing bone applications because of the poor mechanical properties. Currently, there is huge progress being made towards the development of HAp with improved mechanical and biological properties for tissue engineering. Metal ions such as strontium, magnesium and silicon are substituted in HAp, and have resulted in the improvement of biological and mechanical properties [START_REF] Boanini | Ionic substitutions in calcium phosphates synthesized at low temperature[END_REF]. Zinc- [START_REF] Thian | Zinc-substituted hydroxyapatite: A biomaterial with enhanced bioactivity and antibacterial properties[END_REF] and manganese- [START_REF] Mayer | Manganese in precipitated hydroxyapatites[END_REF] doped HAp has a quick resorption mechanism. Continuous research on the improvement of properties of HAp has shown a good response in terms of osteoconductivity and osteogenicity with future prospects to be employed as a bioactive delivery vehicle. Dicalcium phosphates (DCP) [START_REF] Gbureck | Resorbable dicalcium phosphate bone substitutes prepared by 3D powder printing[END_REF] consist mainly of calcium phosphates and water, and are commonly used to alter the physical properties of the composition. A dehydrated form of DCP is known by the mineral name brushite, with the Ca/P ratio 1. DCP-based bioceramics are biodegradable. Brushite is known to convert into HAp in vivo or under physiological conditions that restricts its resorbability and degradation rate. Brushite is used for the treatment of fractures in tibial plates, distal metaphysis bone. Monetite (CaHPO 4 ) bioceramics [START_REF] Tamimi | The effect of autoclaving on the physical and biological properties of dicalcium phosphate dihydrate bioceramics: Brushite vs. monetite[END_REF] are also reported for their applications in bone augmentation and regeneration for dental and orthopedic applications. The bioceramics [START_REF] Arcos | Bioceramics for drug delivery[END_REF][START_REF] Lind | Transforming growth factor-6 stimulates bone ongrowth: Hydroxyapatite-coated implants studied in dogs[END_REF][START_REF] Meseguer-Olmo | A bioactive sol-gel glass implant for in vivo gentamicin release. Experimental model in Rabbit[END_REF][START_REF] Fan | Hard-templating synthesis of a novel rod-like nanoporous calcium phosphate bioceramics and their capacity as antibiotic carriers[END_REF][START_REF] Solberg | Efficacy of gentamycin-impregnated resorbable hydroxyapatite cement in treating osteomyelitis in a rat model[END_REF][START_REF] Ferna | Calcium phosphate bone cements for clinical applications. Part I: Solution chemistry[END_REF] which are employed for the applications in hard tissue replacement are zirconia and alumina (and are also commonly employed in the fabrication of femoral heads).

Biodegradable Polymers

Natural-based [START_REF] Chiellini | Biodegradable polymeric materials[END_REF] and synthetic based [START_REF] Gunatillake | Biodegradable synthetic polymers for tissue engineering[END_REF] polymer materials are some of the preferred choices for biodegradable materials. Polymers can be defined as macromolecules composed of covalently bonded monomers. Depending on the number of repeating monomers belonging to the same or different molecules, they can be classified as homopolymers and copolymers, respectively. Some of the known nature-based polymers are starch, chitosan, derivatives of hyaluronic acid, collagen, fibrin gels and silk. Practical applications of the aforesaid natural-based polymers are impeded by their low mechanical strength, unknown rate of degradation, repellency and high physiological activity. To overcome the challenges posed by natural polymers and in order to replicate them, synthetic polymers (Table 3) were studied with catered design of properties according to the requirements. Depending on the type of polymer chain arrangement-linear, branched, or cross-linked [START_REF] He | Injectable biodegradable polymer composites based on poly (propylene fumarate) crosslinked with poly (ethylene glycol)-dimethacrylate[END_REF]-and its crystallinity/amorphous nature, the degradation rate can be altered. The glass transition temperature of polymers makes them flexible [START_REF] Lodge | Self-concentrations and effective glass transition temperatures in polymer blends[END_REF]; hence, it is necessary to make polymers biodegradable above normal body temperature. Some of the commonly used biodegradable polymers [START_REF] Hollinger | Biodegradable bone repair materials: Synthetic polymers and ceramics[END_REF] are polyglycolic acid (PGA), polylactic acid (PLA), poly-β-hydroxybutyrate (PHB), poly (lactic acid-co-glycolic acid) (PLGA) and poly-ε-caprolactone (PCL). Among these polymers, PLA, PGA and PLGA are used for degradable sutures. PLA was the earliest reported polymer used for sutures [START_REF] Middleton | Synthetic biodegradable polymers as orthopedic devices[END_REF]. PLA was synthesized from natural sources such as starch and corns. Crystalline PLA (L-PLA) is less resistant to hydrolysis than amorphous PLA (DL-PLA) [START_REF] Karst | Molecular modeling study of the resistance of PLA to hydrolysis based on the blending of PLLA and PDLA[END_REF]. PGA, a synthetic polymer, has high crystallinity, low solubility and high degradation rate (because of hydrophilism). The high degradation rate of PGA results in the decrease of its mechanical strength after implantation [START_REF] Sung | The effect of scaffold degradation rate on three-dimensional cell growth and angiogenesis[END_REF]. PCL, an aliphatic polyester, although difficult to control its degradation rate, is widely used in medical implants. PCL is usually employed in the case of long-term implants and as a drug delivery system owing to its crystallinity and permeability, respectively. PCL also has high flexibility and shaping options to form different shapes. PHB is another type of biodegradable polymer with high crystallinity and is used as a model for physical properties of polymers [START_REF] Piskin | Biodegradable polymers as biomaterials[END_REF]. However, due to the brittleness of PHB, it is not employed for practical applications, but is blended with other polymers to study their miscibility and crystallization properties. Poly-para-dioxanone (PPD) [START_REF] Pezzin | Poly (para-dioxanone) and poly (L-lactic acid) blends: Thermal, mechanical, and morphological properties[END_REF] is another biodegradable polymer which is commonly used for its biodegradability, resorbability, compatibility, and its flexibility as a medical implant. PPD is used as fracture internal fixation material and in the form of films, foams, molded products and coatings. Studies have shown the possibility of complete disintegration of PPD implants within 6-7 months after implantation by varying the molecular weight, crystallinity and their melting temperature [START_REF] Tan | Biodegradable materials for bone repairs: A review[END_REF].

The choice of biodegradable polymers is chosen based on the required properties, such as their chemical, physical and mechanical properties. PLA, classified on the basis of its stereoisomers such as poly (L-lactide) acid (PLLA) and poly (D-lactide) acid (PDLA) [START_REF] Maurus | Bioabsorbable Implant Material Review[END_REF], is influenced by their constituent polymer chains which determine PLA's characteristics. The degradation of the polymers is through hydrolysis of ester bonds resulting in alteration of their chemical structure. When the surface of the polymer is attracted by the microorganism, the enzymes secreted by the microorganism converted the macromolecules into tiny molecular debris and eventually converted them into CO 2 and H 2 O. A physical biodegradation process occurs as the microorganisms interact with the surface of the polymers, resulting in hydrolysis (which is a chemical process) to ionize and convert them into oligomer debris. However, chemical biodegradation occurs due to the enzymatic interaction of the microorganism with polymers leading to conversion into CO 2 and H 2 O. However, it has to be noted that all the biodegradation processes of all the polymers do not follow the same methodology, and the complex process of biodegradation is studied extensively by various research groups [START_REF] David | The biodegradation of polymers: Recent results[END_REF][START_REF] Premraj | Biodegradation of polymers[END_REF][START_REF] Shtilman | Biodegradation of Polymers[END_REF].

Implants for bone repair and other load bearing implants require high mechanical strength; hence, ceramics and metals are preferred. However, in some cases, counter-effects such as inflammation [START_REF] Broggini | Peri-implant inflammation defined by the implant-abutment interface[END_REF], corrosion [START_REF] Witte | Degradable biomaterials based on magnesium corrosion[END_REF] and bone loss [START_REF] Cappiello | Evaluation of peri-implant bone loss around platform-switched implants[END_REF] are observed. Hence, polymers are considered for bone repair applications [START_REF] Sheikh | Biodegradable materials for bone repair and tissue engineering applications[END_REF] owing to their tensile strength, elastic modulus and non-corrosion properties. It has to be noted that polymers also experience a loss of mechanical strength with the increase in implantation time as a result of degradation/resorbability. Currently, polymers such as PGA, PLA and PLGA are widely employed for bone implants. Other fields of applications include screws, bone fixation objects, anklebones, patella and other fixing objects. Some disadvantages of these polymers involve X-ray transparency, unidentified influences of foreign body interactions and decreased mechanical strength with time. Other types of polymers such as collagen, hyaluronan and fibrin glue are also degraded enzymatically through collagenase, hyaluronidase and plasmin, respectively. Cyanoacrylates [START_REF] Almeida | First human use of cyanoacrylate adhesive for treatment of saphenous vein incompetence[END_REF] are known to be super glues or instant adhesives in surgery. The mechanism of degradation appears to involve base-catalyzed unzipping from the chain terminus, followed by re-equilibrium of the chains. The degradation decreases adhesive bond strength on heating and poor resistance to prolonged contact with water [START_REF] Robello | Degradation and Stabilization of Polycyanoacrylates[END_REF]. More intensive research is currently being carried out on polymers.

Magnesium Alloys

From the biomedical point of view, magnesium alloys with appropriate mechanical properties and biodegradation are necessary, as well as their biosafety in terms of toxicity. Magnesium alloys are used for tissue engineering, orthopedic and cardiovascular applications. Nan Li et al. [START_REF] Li | Novel Magnesium alloys developed for biomedical application: A review[END_REF] reported on the novel magnesium alloys developed for biomedical applications. Witte et al. [START_REF] Witte | In vitro and in vivo corrosion measurements of magnesium alloys[END_REF] investigated in vivo corrosion of different types of magnesium alloys and reported that an accumulation of biological calcium phosphates and all related alloys displayed a better osteointegration. The mechanical properties of magnesium alloys are increased by alloying with aluminium [START_REF] Tang | Mechanical properties and microstructure of magnesium-aluminum based alloys containing strontium[END_REF] and rare earth elements [START_REF] Tekumalla | Mechanical Properties of Magnesium-Rare Earth Alloy Systems: A Review[END_REF]. However, owing to the neurotoxicity due to the accumulation of aluminium [START_REF] Taïr | Aluminium-induced acute neurotoxicity in rats: Treatment with aqueous extract of Arthrophytum (Hammada scoparia)[END_REF] and hepatotoxicity of rare earth materials [START_REF] Wang | In vivo evaluation of Mg-6Zn and titanium alloys on collagen metabolism in the healing of intestinal anastomosis[END_REF], the advantage of increasing mechanical properties of Magnesium alloys cannot be used for biomedical applications. There is extensive research pointed in this direction to identify the new magnesium alloys with nontoxicity or low toxicity [START_REF] Gu | A review on magnesium alloys as biodegradable materials[END_REF]. Various biomedical magnesium alloys such as Mg-Zn-based [START_REF] Gao | Characterization of strengthening precipitate phases in a Mg-Zn alloy[END_REF], Mg-Ca-based [START_REF] Li | The development of binary Mg-Ca alloys for use as biodegradable materials within bone[END_REF], Mg-Si-based, Mg-Sr-based and Mg-rare earth alloy-based are studied extensively [START_REF] Zhang | Microstructure, mechanical properties and bio-corrosion properties of Mg-Si(-Ca, Zn) alloy for biomedical application[END_REF] by various research groups for the development of biodegradable magnesium alloys. Rapid corrosion [START_REF] Song | Understanding magnesium corrosion-A framework for improved alloy performance[END_REF] of pure magnesium is well-known, but with the increase of purity by purification, the corrosion rate is reduced to a large extent. The common impurities found in magnesium are Fe, Cu and Ni [START_REF] Song | Corrosion mechanisms of magnesium alloys[END_REF]. The corrosion rate of magnesium is highly dependent on the ratio of impurities present in it. Pure magnesium cannot be an apt material for applications in orthopedics because of their mechanical properties. Coarsening of magnesium grains [START_REF] Ali | Current research progress in grain refinement of cast magnesium alloys: A review article[END_REF] is caused by their heat treatments such as forging or rolling. Calcium (Ca) is used for grain refining in magnesium alloys [START_REF] Lee | Grain refinement of Magnesium[END_REF]. The solubility limit of Ca in Mg matric is only 1.34 wt %. There are also reports on the increase of coarse grains in Mg-Ca alloys in the case of increase in Ca [START_REF] Wang | Novel biocompatible magnesium alloys design with nutrient alloying elements Si, Ca and Sr: Structure and properties characterization[END_REF]. Rapid solidification [START_REF] He | Microstructure Evolution and Biodegradation Behavior of Laser Rapid Solidified Mg-Al-Zn Alloy[END_REF] of Mg-Ca alloys yielded fine grains that had superior corrosion resistance and improved cell activity. Degradation [START_REF] Lee | Long-term clinical study and multiscale analysis of in vivo biodegradation mechanism of Mg alloy[END_REF] of Mg-Ca alloys was reported over a period of 3 months following the bone formation. Zinc (Zn) also forms alloys with Mg through which the improvement of mechanical properties of Mg-Zn alloys can be achieved. Other than binary Mg-Zn alloys [START_REF] Zhang | Research on an Mg-Zn alloy as a degradable biomaterial[END_REF], ternary and quaternary alloys of Mg-Zn with Zr, Y, Mn and Ca are formed [START_REF] Wang | Effects of Ca on the formation of LPSO phase and mechanical properties of Mg-Zn-Y-Mn alloy[END_REF]. A binary alloy of Mg-6Zn [START_REF] Wang | Surface modification of magnesium alloys developed for bioabsorbable orthopedic implants: A general review[END_REF] is resistant to galvanic corrosion due to its single phase formation; in vitro cytotoxicity tests indicated that it has good biocompatibility properties; in vivo tests proved that it is not harmful to the internal organs [START_REF] He | Biocompatibility of bio-Mg-Zn alloy within bone with heart, liver, kidney and spleen[END_REF]. However, the bio corrosion was slightly higher in the case of in vivo conditions than in the in vitro conditions. The biodegradation of Mg-Zn alloy was rather average with around 2 mm/year while used as alloy rods in rabbit femur shafts; adding Y in the binary alloy of Mg-Zn aids to diminish the corrosion rate. On the other hand, the addition of Zr in the binary alloy of Mg-Zn helps in grain refinement, and at the same time helps to increase the osteointegration properties superior to titanium alloys. Adding the intermetallic group of metals such as manganese (Mn) to Mg-Zn can contribute to the corrosion-resistive properties and subsequently remove heavy metals such as Fe. The degradation of Mg-Zn-Mn alloy is highly dependent on the individual constituent of each component in the alloy. The alloy compound of Mg-Zn-Ca [START_REF] Hofstetter | High strength low alloy (HSLA) Mg-Zn-Ca alloys with excellent biodegradation performance[END_REF][START_REF] Hofstetter | Assessing the degradation performance of ultrahigh-purity magnesium in vitro ad in vivo[END_REF] has superior mechanical and corrosion properties when Zn content is <4 wt %, even under in vivo conditions. Other alloys of Mg with Zn, Ca and Si and Sr are also prevalent. The aforesaid were found to contribute primarily to the refinement of grains and change of grain morphology, which assists in corrosion resistance. Although rare earth elements were added as dopants, the toxicity of the rare earth elements is well known. Hence, rare earth elements are not yet employed as implant materials. In the current state-of-the-art [START_REF] Agarwal | Biodegradable magnesium alloys for orthopaedic applications: A review on corrosion, biocompatibility and surface modifications[END_REF], magnesium-based alloys cannot be used for orthopedic applications due to their high degradation rate. Concerning the corrosion rate, magnesium alloys based on Mg-Mn-Zn and Mg-Ca have a low in vivo corrosion rate. The mechanical strength of these Mg-based alloys is increased through the change in their structural properties because of their biocompatible properties. The fabrication methodology [START_REF] Xu | Development and microstructural characterizations of Mg-Zn-Ca alloys for biomedical applications[END_REF] through powder metallurgy, melt-casting and spin melting, is used for the fabrication of magnesium alloys and magnesium-based metal composites. Another strategy employed for controlling the degradation rate of magnesium alloys is through coating their surfaces with calcium phosphate [100], fluorinated coatings, polymer coatings, etc. Magnesium-based implants are currently used as micro clips for laryngeal microsurgery, orthopedic and cardiovascular systems. There are further developments in this domain of developing magnesium alloys. The common challenge in magnesium alloys is their formation of hydrogen. Magnesium glasses are considered as an alternative for crystalline alloys to avoid the formation of the hydrogen bubble effect [101]. During the manufacturing process of magnesium alloys, their influence on the implant surface with controllable properties is required for controlling the degradation and corrosion process. The microstructure and properties of Ti alloys vary depending on the type of interstitial/substitutional element chosen [108]. Many substitutional metals such as, V; Mo, Nb, Cr, Fe, Si, Ta, Cu, Ni, Pd, Co, M, and W are used as additives to stabilize the β phase. Titanium alloys have considerably superior biocompatibility compared to their other counterparts due to their excellent corrosion resistance. In comparison to Ti-6Al-4V, β-titanium alloys with stabilizing elements (Mo, Ta and Zr) are considered to be safe. B-titanium alloys have a modulus closer to that of bone. However, the fatigue strength is poor compared to other α-β alloys. The fatigue strength of these alloys was enhanced by the addition of Y 2 O 3 , SiO 2 and ZrO 2 which causes dispersion; hence, strengthening the alloys. Ti alloys require surface modifications to increase their wear resistance. Titanium alloys form easy bonding with bone, demonstrating good integration with bone tissue. Currently, pure Ti alloys are used in pacemaker cases, ventricular devices, implantable drug pumps, screws and staples in spinal surgery, dental implants and in craniofacial implants. Ti-6Al-4V alloys are used for dental implants, hip, knee, wrist, spine, shoulder, elbow and orthodontics. However, there is a long-term risk associated because of the release of Al and V that causes Alzheimer's, osteomalacia and other neurological conditions. Currently, Vanadium free Ti-6Al-7Nb and Ti-5Al-2.5 Fe are being developed [109] and used in femoral prosthesis stems, fracture plates, nails, rods, screws and wires. Titanium alloys cause allergic tissue reactions because of fretting, and have relatively less wear resistance in comparison to other alloys used. Because of the poor bending properties of α-β titanium alloys, Ti-Mo-Zr-Fe [110] was used instead for total hip replacement prostheses. Titanium alloy's wear resistance is improved by addition of refractory metal elements and by surface modification techniques [111]. The application of Ti alloys for implant materials is impeded by their poor mechanical strength.

Orthopedic Implants

Orthopedic implants is the term representing implants that replace a missing joint or any damaged bone. Orthopedic implants represent a large part of the implants made in the field of medical implants. Orthopedic implants are mostly composed of chromium, cobalt, molybdenum, nickel, titanium and zirconium alloys. The disadvantages of metallic orthopedic implants such as metal allergy, metal implant failure, and complications related to joint replacement are reported. Other serious complications of orthopedic implants include device failure, pain, metallic staining of the surrounding tissue, muscular necrosis, periprosthetic fibrosis and loosening of prosthesis. Orthopedic implants used for total hip arthroplasty (THA) and total knee arthroplasty (TKA) metal allergy and device failure are reported. Metallic orthopedic implants failure results in pain, instability, malrotation, and local pain leading to infection. Wendy et al. reported on the metal hypersensitivity reactions to orthopedic implants [112]. Other orthopedic implants also include bone screw with self-tapping threads used as fastening elements in prosthetics. Steel and titanium screws are used prevalently for fabrication of bone screws. Recently, there has been progress made in the direction of better performance of bone screws leading to the study of magnesium (Mg) and Mg-coated screws. Recently, Ullrich et al. [113] reported on the bone screw with minimum risk of fracture with miniaturization, by reducing the friction on the surface of the screw implant. This was achieved by means of surface functionalization, self-tapping and self-lubricating thread profile by incorporating some bioactive materials to accelerate bone growth/healing. By improving the tapping efficiency of the abrasives, friction is minimized during the screwing-in process. The mass and size of the implant is reduced which will reduce patient rehabilitation time; this also opens avenues for resorbable implants such as in the skull and small limb field.

Metals Used in Knee Replacement

Many different types of metallic systems are used for knee replacement. Steel has higher stiffness than the cortical bone with an elastic modulus of 200 GPa [114]. Stiff steel implants bare all the load, and thus weaken the normal bone. Hence, bone cement is used as an interface to decrease the stress shielding effect [115]. Fixing of stiff steel implants is usually carried out with the help of bone cement. Most commonly used steels are cheaper than other types of metals. In medical fields, the commonly used iron alloys are of low carbon alloyed, chromium (17-20%), molybdenum (2-4%) and nickel (12-14%), with the addition of other small amounts of other elements. The corrosion resistance of these metals is under intensive investigations [116] as release of these metal ions inside the body could cause allergy, toxicity and other symptoms. The release of iron from stainless steel favors bacteria infections by acting as an iron source for their proliferation [117]. To avoid these inconveniences, stainless steel implants are used in the body as short-term implants or as coated implants. Recently, cobalt-and titanium-based materials were used as joint replacement materials. The fabrication methodology and the microstructural defects lead to fatigue-based fracture. Good fatigue resistance of approximately 400 MPa is considered necessary for total hip implants to last a long time [118]. Fatigue resistance is diminished by wear, corrosion and material defects. Steel has an elastic modulus of approximately 200 GPa and titanium has an elastic modulus of approximately 110 GPa. Hence, while using the titanium-based alloys, cement used as an intermediate layer is not considered important. To ensure fixation, porous-coated implants are used. Hydroxyapatite is commonly used as a coating material. In total joint arthroplasty, aseptic loosening is the prime cause of failure. Stainless steel knee implants were reported to have an 80-90% success rate, even after 25 years of implantation [119]. Radiographic loosening was inferred in few cases of cobalt-chrome alloys. As a cementless implant, the surface finish governs the osseointegration potential.

Orthodontic Wire

Orthodontic brackets and wires are widely used in cases of a long span of treatments. Researchers are investigating new types of archwires to minimize the discomfort caused [120]. Various materials such as nickel titanium, beta titanium, copper and stainless steel are currently used. Different shapes and forms (such as square, round, rectangular and beveled surfaces) of orthodontic archwires [121] are used. There was continuous evolution between the archwire and the bracket to reduce the friction significantly. Heat-activated multimodulus wire, due to its increased flexibility, minimizes angular friction. Nickel titanium (NiTi) wires [122] with added resiliency and varying heat activation variation is considered mostly by dentists, whereas the overstressing of the NiTi wire causes permanent distortion. To reduce the wire surface friction, surface coatings are employed on NiTi wires. Beta titanium wire appears to be the intermediate wire between NiTi wires and stainless steel, with twice the flexibility and half the force of stainless steel. Another material, copper NiTi, is reported to be more resistant to permanent distortion than other nickel titanium wires [123]. Various materials in the following classifications such as noble metal alloys, stainless archwires, cobalt-chromium archwire, nickel titantium archwire, copper nickel titanium alloy and beta titanium archwire are used for archwires. However, detailed information on these materials is not given here as they are out of the scope of the present review article. Biodegradable metals are also used as metallic stents. Biodegradable metals are used for treating coronary artery diseases. Mg, Fe and Zn are widely used for biomedical applications. Mg-6Zn alloys, Zn, Fe and Co-Cr stents are biocompatible, but the latter two have less degradation time.

Conclusions

Bioresorbable/biodegradable composites and ceramic implants are of growing interest because of their ability to provide a necessary mechanical function for the tissue reconstruction process after the designated function in the body. These bioresorbable materials should degrade at a tolerable range for the body and an optimized degradation rate is required based on the function that the material serves in the body. There are not enough human in vivo data to validate the biodegradability of bioresorbable materials. This inhibits the valorization of the materials for marketing purposes. There are many challenges, such as faster than expected degeneration of the mechanical properties of Mg-based alloys during the tissue remodeling process. Surface coatings and modifications are suggested as an alternative to achieve the extended mechanical integrity of these materials. More research has to be focused on materials such as Mg-based and Fe-based alloys to determine their mechanical and degradation kinetics. The biosafety of the degrading material should be ensured with the host tissue and implant interface. Various other parameters such as age, physical condition, risk of infection, and type of fracture are primordial for the choice of the required bioresorbable composites and ceramic implants. Currently, biodegradable/bioresorbable polymers are successfully used for biomedical and pharmaceutical applications. Polymers such as PCL, PGA and PLA are used for craniomaxillofacial fixation, sutures, interference screws, fixation plates and pins and for meniscal repair. However, the degradation rates of these polymers are not well controlled; they lack mechanical strength and induce inflammation reactions in the acidic medium. To date, few metallic polymer composites such as Mg/PCL have demonstrated compressive moduli similar to human cancellous bone, in addition to their excellent cytocompatibility and osteoblastic differentiation properties. In the current state-of-the-art, biodegradable/bioresorbable composites still require much fundamental research and development for use in bioimplants. Metals have to be selected carefully to avoid metal-related toxicity in the body and to avoid corrosion in the body. There are many other factors to be considered such as the influence of the pH change on adjacent tissue and the metabolic activity in supplementary phases. In the present scenario, magnesium-based implants have good future prospects for development.

2. 4 .

 4 Titanium Alloys Titanium, being one of the low-density elements, is one of the materials used for biomedical applications [102]. The strengthening process of a titanium alloy is done either by alloying or deformation processing [103]. Ti alloys were favored as biomaterials because of their good biocompatibility [104], corrosion resistance [105] and lower modulus [106]. Ti alloys such as α + β Ti-6Al-4V and metastable β alloys are currently being investigated [107]. Pure titanium undergoes a transformation temperature from α phase to β phase at ~885 • C. The transition temperature is influenced by the interstitial elements such as, O; N, C, B and H and various other substantial elements.

Table 1 .

 1 Scaffold characteristics and requirements.

	Scaffold Characteristics	Requirements
		Non-toxicity to the host tissue, support normal cellular activity,
	Biocompatibility Biodegradation/bioresorbability	osteoconductive, osteoinductive and osteogenic, angiogenesis A controlled resorption rate to host new bone tissue. Possibility to vary the degradation rate, controlled rate of drug delivery and
		incorporate biomolecules
		Minimum pore size of 100 µm for diffusion of nutrients, cell
		survival and proliferation. To enable bone tissue in-growth, pore
	Pore size and shape	sizes in the range of 200-350 µm are required. Multiscale porosity
		with a combination of micro and macro pores allows cell growth,
		but may be detrimental in terms of mechanical strength
		Capability to withstand mechanical stress and loading. Scaffold
	Mechanical properties	possibility to have good mechanical properties and to mimic as a
		natural body component

Table 2 .

 2 Influencing parameters of biodegradation.

		Fabrication methodology
		Size/shape
	Scaffold	Pore size and porosity surface roughness
		surface area to volume ratio
		Additives or impurities
		pH/ionic strength
		cell type and density
	In vitro factors	Mechanical loads
		Incubation temperature
		Biological medium's composition
		Tissue modelling and remodeling
	In vivo factors	Mechanical loads Enzyme concentrations
		Implantation site

Table 3 .

 3 Synthetic degradable polymers and applications.

	Synthetic Degradable Polymers	Applications
	Polycyanoacrylates [61]	Adhesives, drug delivery
	Polyanhydrides [62]	Drug delivery
	Poly(amino acids) [63]	Drug delivery, tissue engineering, orthopedic applications
	Poly(ortho ester) [64]	Drug delivery, Stents
	Polyphosphazenes [65]	Blood contacting devices, drug delivery, skeletal reconstruction
	Poly(propylene fumarate) [66]	Orthopedic applications
		Barrier membranes, drug delivery, guided tissue
	Polylactic acid (PLA) [67], poly glycolic (PGA) [68]	regeneration (in dental applications), orthopedic
	and copolymers	applications, stents, staples, sutures, tissue
		engineering
	Polyhydroxybutyrate (PHB) [69],	Long-term drug delivery, orthopedic applications,
	polyhydroxyvalerate (PHV) [70], and copolymers	stapes stents
	Polycaprolactone [71]	Long-term drug delivery, orthopedic applications, staples, stents
	Polydioxanone [72]	Fracture fixation in non-load-bearing bones, sutures, wound clip

J. Funct. Biomater. 2017, 8, 44; doi:10.3390/jfb8040044 www.mdpi.com/journal/jfb

Acknowledgments: This work has been supported by the European Social Fund within the project "Involvement of new scientist group for synergistic investigation to development of nanostructured composite materials for bone tissue regeneration", No. 2013/0007/1DP/1.1.1.2.0/13/APIA/VIAA/024.

Author Contributions: All the authors contributed equally to the present review article.

Conflicts of Interest:

The authors declare no conflict of interest.