Cellular Ceramics / 5

5.8 Biomedical Applications: Tissue Engineering 559

Fig. 5 Perspective view of a typical HA scaffold fabricated by the RP technique developed by Wilson et al. [96]. (Micrograph courtesy of Dr. C.E. Wilson, Twente University, The Netherlands.)

Fig. 6 Typical microstructure of a sintered latticework of HA produced by extrusion free-forming by the method developed by Gomes de Sousa and Evans [99] (Courtesy Ms. C. Gomes de Sousa, Queen Mary University of London, UK.)

However, the number of HA-based systems fabricated to date is relatively small, and data on scaffold performance are scarce, so that a conclusive comparison of the different RP and SFF methods in terms of optimal 3D design of porous structures, attained properties, and in vitro or in vivo behavior is not yet possible.

There is continuous research focused on the improvement of computational optimization procedures to be coupled with RP and SFF techniques. This should aid in creating scaffold structures such that both the porous ceramic scaffold and the eventual regenerated tissue will match host tissue stiffness, while at the same time meeting constraints on scaffold porosity, material, and fabrication method [97, 98].

560Part 5 Applications

5.8.5.2

Melt-Derived Bioactive Glasses

Although melt-derived bioactive glasses have had clinical success in particulate form, such as the commercially available 45S5 Bioglass composition [10], it has proven difficult to produce scaffolds with interconnected pore networks from these materials. Livingston et al. [33] produced a simple sintered scaffold by mixing 45S5 Bioglass powders, with particle size range of 38–75 mm, with 20.2 wt % of camphor (C10H16O) particles with a size range of 210–350 mm. The mixture was dry-pressed at 350 MPa and heat treated at 640 C for 30 min. The camphor decomposed to leave porous Bioglass blocks. The pores were in the range of 200–300 mm in diameter, but total porosity was just 21 %, hence distances between pores were large, and interconnectivity was low.

Yuan et al. [34] produced similar scaffolds by foaming Bioglass 45S5 powder with

a dilute H2O2 solution and sintering at 1000 C for 2 h to produce a porous glass ceramic. Although pore diameters were in the desired range of 100–600 mm, the

pores were irregular in shape with orientated channels running through the glass and large distances between channels. Interconnectivity was therefore low. The samples were implanted into the muscle tissue of dogs and were found to induce bone growth, containing osteoblasts and osteocytes, in soft tissue. Pathological calcification, that is the formation of nonhealthy bone that does not contain osteocytes, was also observed in pores near the surface of the material. Bone formed directly on the solid surface and on the surface of crystal layers that formed in the inner pores and in pores containing pathological calcification. Osteogenic cells were observed to aggregate near the material surface and secrete bone matrix, which then calcified to form bone. However, although the implants had a porosity of about 30 % only 3 % bone was formed.

The few, rather unsatisfactory, results available indicate that optimized highly porous bioactive glass scaffolds cannot be produced from melt-derived glass by conventional powder-technology methods. Thus the most favored bioactive glasses for produce porous scaffolds are those derived by sol–gel processing.

5.8.5.3

Sol–Gel-derived Bioactive Glasses

The sol–gel process is an alternative method to produce bioactive glasses [25]. An advantage of gel-derived bioactive glasses over their melt-derived counterparts is that they exhibit a mesoporous texture (pores diameters in the range 2–50 nm). This texture enhances bioactivity and resorbability of the glasses [105, 106]. A disadvantage of gel-derived glasses is that there is no clinical experience with them as yet; therefore, the route to clinical use would be longer than if melt-derived glasses were used.

The sol–gel process involves the production of a colloidal solution (sol) of Si-O groups by the hydrolysis of alkoxide precursors such as tetraethyl orthosilicate (TEOS) in excess water under acidic catalysis. Simultaneous polycondensation of

5.8 Biomedical Applications: Tissue Engineering 561

Si-O groups continues after hydrolysis is complete, beginning the formation of the silicate network. As the network connectivity increases, viscosity increases, and eventually a gel forms. The gel is then subjected to carefully controlled thermal processes of ageing (60 C) to strengthen it, drying (130 C) to remove the liquid byproduct of the polycondensation reaction, and thermal stabilization/sintering (600–800 C) to remove organics from the surface of the gel and to densify the structure. A chemically stable glass is then produced [107].

Sol–gel-derived bioactive glass scaffolds with cellular structure have been produced in several ways.

Many authors have cast silica-based sol–gel glasses around removable templates such as close-packed latex or polystyrene spheres [108–110] with similar techniques as used for HA slurries, except that the sol gels around the organic spheres, and the polymer is burnt out during stabilization of the gel. The pore network produced is very homogeneous and interconnected due to the close packing of the spheres. However pore size is limited by the difficulty in producing organic spheres with diameters on the order of 500 mm required to produce an ideal scaffold for tissue engineering applications.

Foaming agents are chemical additives that produce gas bubbles in a liquid. Gun et al. [111] added 6–10 vol % hydrogen peroxide to the base-catalyzed sol–gel process for methyl silicate materials. Decomposition of the hydrogen peroxide formed bubbles which remained throughout polymerization. Crack-free rods (20 cm length 0.5 cm diameter) were prepared that did not shrink during ageing. This technique could be applied to bioactive gel glasses, but there were few pores greater than 10 mm in diameter.

Polymer foaming techniques have been applied to sol–gel systems. Viscosity was controlled to stabilize bubbles created by Freon 11 (CCl3F) droplets dispersed in a sol to form porous silica with porosities of 55–90 % and corresponding mean cell diameters of 30–1000 mm [112]. Bubbles formed because Freon 11 has a boiling point of 23.8 C, and incubation was carried out at temperatures above this boiling temperature. The foam structure was stabilized by addition of an anionic surfactant and a rapid polymerization reaction, which caused fast gelation. The viscosity was controlled by adjustment of pH with H2SO4. A surfactant is generally used to stabilize any bubbles formed in the liquid phase by reducing the surface tension of the gas–liquid interfaces [113]. The bending strength of the silica foams made by these methods range from 2 MPa at a porosity of 86 % to approximately 8 MPa at a porosity of 66 %.

Direct foaming of sol–gel-derived bioactive glasses has been performed by vigorous agitation, with the aid of surfactants, to produce scaffolds that fulfill many of the criteria of the ideal scaffold [114]. The surfactant lowers the surface tension of the sol and stabilizes bubbles created by air entrapment. A gelling agent (hydrofluoric acid, HF) is added to the sol to increase the gelation time to a few minutes. As gelation occurs the bubbles are stabilized permanently. The highly porous foam is then heat-treated in the same way as a normal sol–gel glass. Figure 7 shows an SEM image of a typical foam of 70S30C bioactive glass (70 mol % SiO2, 30 mol % CaO).

562 Part 5 Applications

Fig. 7 SEM image of a typical sol–gel-derived bioactive glass foam of composition 70 mol % SiO2, 30 mol % CaO.

The properties of the pore network can be controlled at each stage of the process, for example, by the type and concentration of catalyst and surfactant, the glass composition, and the process temperature [115, 116]. The scaffolds can be produced with interconnected macropores, porosities in the range of 60–90 %, and a modal interconnected pore diameter of up to 150 mm [115] with many interconnections in excess of 300 mm in diameter.

The composition and textural pore size can be used to tailor the rate of resorbability of the scaffold and therefore the rate of release of elements that affect the gene expression of osteoblasts [115]. Due to the nature of the sol–gel process the scaffolds can be produced in many shapes, which are determined simply by the shape of the casting mold, or scaffolds can be cut to a required shape.

Primary human osteoblast cells, harvested from the tops of femurs removed during total hip replacements, have been seeded on the foamed glasses in vitro. The cells attached, proliferated, and secreted bone extracellular matrix, which mineralized after 10 d of culture. Figure 8 shows an SEM image of primary human osteoblasts cultured on a 58S bioactive glass foam for 10 d [117]. The micrograph shows a mineralized bone nodule inside a macropore. A bone nodule is a group of cells that have laid down some extracellular bone matrix. This bone nodule has mineralized without the addition of mineralization supplements to the culture media, which indicates the great potential of the bioactive glass foam as osseous tissue scaffold.

The only ideal scaffold criterion not addressed is the matching of mechanical properties of the scaffolds to bone for in situ bone-regeneration applications. Foams optimized by sintering have compressive strength of about 2.5 MPa, which is similar to that of trabecular bone (2–10 MPa) [118], but the fracture toughness and tensile strength of the foams are much lower than the values measured for bone. The mechanical properties of these foams should be sufficient for tissue engineering

5.8 Biomedical Applications: Tissue Engineering 563

Fig. 8 SEM image of primary human osteoblasts cultured on a bioactive glass foam (58S) for 10 d. The micrograph shows a mineralized bone nodule inside a macropore [117]. (Micrograph courtesy of Dr. Julie Gough, University of Manchester, UK.)

applications, where bone would be grown on a scaffold in the laboratory before implantation. To improve the toughness of the scaffolds, however, the complete structure of bone should be mimicked more closely. The foamed glasses and ceramics mimick bone mineral, but bone is a natural composite of bone mineral and collagen. An organic phase should therefore be introduced into the scaffold to provide toughness for in situ bone-regeneration applications. In recent developments, antimicrobial macroporous scaffolds based on sol–gel glasses doped with silver ions have been developed [119]. However, these foams, which combine bioactive and antimicrobial properties, are yet to be investigated further in terms of in vitro and in vivo behavior.

The interactions between cells and surfaces play a major biological role in cellular behavior. Cellular interactions with artificial surfaces are mediated through adsorbed proteins. A common strategy in tissue engineering is to modify the biomaterial surface selectively to interact with a cell through biomolecular recognition events. Adsorbed bioactive peptides can allow cell attachment on biomaterials, and permit three-dimensional structures modified with these peptides to preferentially induce tissue formation consistent with the cell type seeded, either on or within the device [120, 121]. To achieve full functionality, peptides must be adsorbed specifically. The surfaces of bioactive glass foam scaffolds can be modified with amine and mercaptan groups [122]. The modified scaffolds have been used as carriers for laminin, which was selected as a molecular model for the eventual tailoring of scaffold chemistry to engineer growth of arterial and lung tissues. The covalent bonds between the binding sites of the protein and the ligands on the scaffold surface did not

564 Part 5 Applications

denaturate the protein. The 3D architecture of the foams mimics the scale and interconnectivity of structures needed to grow a pulmonary artery and lung lobe in vitro. In vitro studies show that foams modified with chemical groups and coated with laminin maintained bioactivity and improved proliferation of mouse lung epithelial (MLE-12) cells compared to uncoated foams [123]. Sustained and controlled release from the scaffolds over a 30-day period was achieved as laminin release from the bioactive foams followed the dissolution rate of the material network [123].

Sol–gel-derived bioactive glasses can be processed to produce similar structures to the HA bone grafts used clinically. Although they have controllable resorbability and the potential to stimulate bone cells at the genetic level to encourage bone regeneration, sol–gel bioactive glass foams have not yet been submitted for FDA or EU (CE mark) approval. However, melt-derived bioactive glasses have FDA and EU approval and are commercially available as regenerative bone filler, so it is a matter of time before the sol–gel glass scaffolds are approved for implantation.

5.8.5.4

Other Bioceramics Exhibiting Cellular Structure

Research is also being carried out on the development of cellular structures from bioceramics other than calcium phosphates, hydroxyapatite, or bioactive glasses. For example, zirconia porous scaffolds coated with biphasic calcium phosphate (HA + TCP) have been developed [124]. Zirconia foams (90 % porosity, pore size 500–600 mm) exhibiting highly interconnected porosity were fabricated by a replication method using reticulated PU sponges. Slurry dipping was used to obtain calcium phosphate coatings of homogeneous thickness (ca. 30 mm) uniformly covering the pore walls of the zirconia foams. The adhesive strength of the coating layer was as high as 20–25 MPa. The biphasic calcium phosphate layer improved osteoconduction and stimulated the proliferation of bone cells [124].

Controlled porosity, bioinert alumina scaffolds with pore sizes in the range of 300–500 mm have been processed using a RP process based on fused deposition [95]. The technique allows production of scaffolds with designed volume fraction and pore size, which were used to evaluate mechanical properties and biological response and their interrelationship with porosity parameters [95]. The coating of porous alumina substrates with HA layers has been proposed to impart bioactive behavior to the scaffolds [125].

Alumina foams coated with HA or TCP have been prepared also by a PU-sponge method, followed by multiple slurry-dipping in aqueous HA or TCP suspensions and final heat treatment [126]. In these foams, the alumina structure provides the mechanical strength (compressive strength of 12 MPa for a porosity of 75 %), which is much higher than that of equivalent foams made totally of HA [127–129]. Good osteoconduction of HA-coated alumina scaffolds, equivalent to that of pure HA scaffolds, was proved by in vitro and in vivo studies [126].

Titania foams are gaining increasing attention for applications as tissue engineering scaffolds due to the high biocompatibility of TiO2 [130, 131]. Although traditionally classified as a bioinert material [132], it has been recently shown that surface-

5.8 Biomedical Applications: Tissue Engineering 565

modified porous titania structures can exhibit bioactive behavior [133–135]. Titania cellular structures have been produced by the replication method using fully reticulated polyester-based PU foams [130]. The aqueous ceramic suspension was prepared from titania powder with mean particle size of 0.3 mm in a concentration of 57 vol %. The heating schedule for burning out the polymers and sintering the foams comprised a first stage at 450 C for 1 h and subsequent heating at 1150 C. Pore sizes were 445 or 380 mm, depending on the PU foam used. Initial cell culture work with fibroblasts showed high adhesion of the cells to the foam structure [130], and further in vitro and in vivo are expected on these novel materials.

As a new generation of light, tough, and high-strength materials for medical implants and bone substitution, biomorphic SiC porous ceramics coated with bioactive glass have been developed [136]. Biomorphic SiC is fabricated by molten-Si infiltration of carbon templates obtained by controlled pyrolysis of wood. The structure exhibits high porosity (up to 70 %), high anisotropy, and resembles the structure of bone. The uniform and adherent bioactive glass coating is then applied by pulsed laser ablation. In vitro tests in simulated body fluid demonstrated the formation of HA layers after 72 h [136]. It is also claimed that biomaterials with high silicon content, such as SiC-based structures, do not present adverse physiological effects in the body, but little is known on the long term in vivo behavior of biomorphic SiC scaffolds, in particular considering the presence of impurities in the starting wood.

5.8.6

Properties of Some Selected Bioactive Ceramic Foams

Many techniques are being employed with different degree of success to produce cellular bioceramics tailored for applications as tissue engineering scaffolds, particularly for bone tissue and to a lesser extent for other (soft) tissues. Table 1 lists different characteristics of selected bioactive ceramic foams developed for bone-tissue engineering, including porosity characteristics, methods of fabrication, and typical mechanical properties. Mechanical properties of cortical bone are also shown for comparison. For HA-based scaffolds, innovative foaming of ceramic suspensions coupled with gel casting of foams have led to very satisfactory results in terms of porous structure and subsequent in vitro and in vivo behavior. Similarly, scaffolds produced by rapid prototyping and solid free-form fabrication methods exhibit highly ordered microstructures and can be readily manufactured to desired, complex shapes. It is intriguing that these methods have been developed so far mainly for HA scaffolds but not for bioactive glasses. Gel casting of foams produces bioceramic scaffolds with a structure very similar to that of trabecular bone. Both the gelcast foamed HA and sol–gel-derived bioactive glass foams showed favorable results in both in vitro and in vivo tests for bone regeneration. In general, optimized bioactive glass or HA scaffolds have the potential to fulfil eight out of nine of the criteria for an ideal scaffold for tissue engineering applications (as presented in Section 5.8.2). The only criterion not fulfilled is related to mechanical properties. Although the scaffolds can exhibit compressive strengths similar to those of trabecular bone,

566 Part 5 Applications

while maintaining a pore network suitable for tissue engineering, they have low flexural strength. This may not affect tissue engineering applications, where tissue is grown in bioreactors in vitro, but the scaffolds are not yet suitable for direct implantation in load-bearing applications.

Table 1 Selected bioactive ceramic foams for bone tissue engineering: porosity characteristics, methods of fabrication and typical mechanical properties. Data for cortical bone are also shown for comparison.