Cellular Ceramics / 5

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5.8

Biomedical Applications: Tissue Engineering

Julian R. Jones and Aldo R. Boccaccini

5.8.1

Introduction

Life expectancy is increasing as healthcare and technology improve, but not all body parts can maintain their function with the ageing process. Bone and cartilage are needed to support the ageing body even though the cells that produce them become less active with age, while the heart, kidneys, and liver must operate for much longer than ever before.

The most common bone disease is osteoporosis, which causes a loss of bone density and affects everyone as they age. Bone is a natural composite of collagen (polymer) and bone mineral (ceramic). Collagen is a triple helix of protein chains, which has high tensile and flexural strength and provides a framework for the bone. Bone mineral is a crystalline calcium phosphate ceramic [hydroxyapaptite, Ca10(PO4)6(OH)2] that contributes the stiffness and compressive strength of bone. Cortical bone is a dense structure with high mechanical strength and is also known as compact bone. Trabecular bone is a network of struts (trabeculae) enclosing large voids (macropores) with 55–70 % interconnected porosity, which is a supporting structure [1]. The struts of the trabecular bone, which are present in the ends of long bones such as the femur or within the confines of the cortical bone in short bones, are most affected by osteoporosis. The loss of bone density is caused by a reduction in number of osteogenic cells (osteoblasts) with age. The only treatment for severe cases of osteoporotic joints is total joint replacement.

Surgical procedures for the repair of tissue lost as a result of trauma or by the excision of diseased or cancerous tissue involve graft implants (transplants). Grafts can be taken from a donor site in the same patient (autograft), from another human donor (homograft), or from other living or nonliving species (xenografts). Grafts can be used to attain the ultimate goal of restoration of a tissue to its original state and function, but there are many limitations [2]. Autografts and homografts are restricted by limited material availability and complicated multistage surgery to the detriment of the harvest site. Homografts and xenografts run the risk of disease transmission. Therefore, there is a great demand for synthetic substitutes specially designed and manufactured to act as a scaffold for the regeneration of tissues to their natural state and function, which constitutes the fundamentals of the tissue engineering discipline [3].

Cellular Ceramics: Structure, Manufacturing, Properties and Applications.

Michael Scheffler, Paolo Colombo (Eds.)

Copyright 2005 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim

ISBN: 3-527-31320-6

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This chapter reviews how cellular ceramics are being considered in bone reconstruction and other orthopaedic applications with focus on the development of highly porous scaffolds for bone tissue regeneration. Where appropriate, reference to the application of cellular ceramics in other areas of tissue engineering and implant technology are made. The rest of the chapter is organized in the following manner: Section 5.8.2 includes a general consideration of the new scientific field of regenerative medicine and the opportunities for biomaterials, Section 5.8.3 covers the field of bioceramics, that is, ceramics used in medical applications in general, and presents the classification of bioceramics accepted in the current literature, Section 5.8.4 covers the field of porous scaffolds for tissue engineering, while Section 5.8.5 is devoted specifically and extensively to cellular structures of hydroxyapatite, calcium phosphate, and bioactive glass. Reference to other bioceramics used in cellular form in tissue engineering applications is also made in Section 5.8.5. Finally, Section 5.8.6 contains a summary of the chapter, identifying progresses made and needs for future developments, particularly in the area of bioactive porous scaffolds.

5.8.2

Regenerative Medicine and Biomaterials

Regenerative medicine is a general term used to describe techniques that are being developed to regenerate diseased or damaged tissues to their original state and function. Tissue engineering and tissue regeneration are two branches of this broad field. In some tissue engineering applications synthetic or natural porous scaffolds are used as templates for tissue growth [1, 3]. There is the potential for stem cells to be extracted from a patient and seeded on a (natural or synthetic) scaffold of the desired architecture in vitro, where they will be given the biological signals to proliferate and differentiate, and new tissue will grow, ready for implantation [4, 5]. Tissue regeneration techniques involve the direct implantation of an engineered scaffold (seeded with cells or not) into a defect to guide and stimulate tissue repair in situ. In both cases the scaffold should be resorbed (dissolve) as the tissue grows, leaving no trace of damage or surgery [3, 6]. In any application of natural or artificial scaffolds, choice and design of the material are important.

Any material that is implanted into the body should be biocompatible, that is, noncytotoxic (not toxic to cells). There are three classes of noncytotoxic materials: bioinert, bioactive, and bioresorbable [7].

No material is completely inert on implantation, but the only response to the implantation of bioinert materials is encapsulation of the implant by fibrous tissue (scar tissue). Examples of bioinert materials are medical-grade alumina, zirconia, stainless steels, and high-density polyethylene, materials that are used, for example, in total hip replacements.

Millions of orthopaedic prostheses made of bioinert materials have been implanted, an example of which is the Charnley total hip replacement, which is heralded as one of the most successful surgical inventions [7, 8]. However, long-term monitoring of 20 000 Charnley joints revealed that after 25 years of implantation

5.8 Biomedical Applications: Tissue Engineering 549

24 % required revision surgery [8]. The most common reason for failure was aseptic loosening of the femoral stem, where bone resorption occurred due to a mismatch in the Young’s modulus of bone and the metal stem (stress shielding) [9].

Resorbable materials are those that dissolve in contact with body fluids and whose dissolution products can be secreted via the kidneys. The most common biomedical resorbable materials are polymers that degrade by chain scission such as poly(glycolic acid), PGA, poly(lactic acid), PLA, and their copolymers (PLGA). Some bioceramics are also resorbable in vivo, such as calcium phosphates [7], as discussed below.

Bioactive materials stimulate a biological response from the body such as bonding to living tissue [10]. There are two classes of bioactive materials intended for bone reconstruction and orthopaedic implants: class B bioactive materials bond to hard tissue (bone) and stimulate bone growth along the surface of the bioactive material (osteoconduction). Examples of class B bioactive materials are synthetic hydroxyapatite and tricalcium phosphate ceramics.

Class A bioactive materials not only bond to bone and are osteoconductive but they are also osteoproductive, that is, they stimulate the growth of new bone on the material away from the bone/implant interface and can bond to soft tissue such as gingiva (gum) and cartilage [10]. Examples of class A bioactive materials are bioactive glasses.

The mechanism of bone bonding to bioactive materials is thought to be due to the formation of a hydroxyapatite (HA) layer on the surface of the materials after immersion in body fluid. This layer is similar to the apatite layer in bone and therefore a strong bond can form. The layer forms quickest on the class A bioactive materials [10].

There is a wealth of specialized literature on bioceramics; some relevant books and review articles about the general use of porous bioceramics in biomedical applications are those in Refs. [7, 10–13].

5.8.3

Bioactive Ceramics for Tissue Engineering

Ceramic materials used for medical implants as well as in the repair and reconstruction of diseased or damaged parts of the muscoskeletal system and of other tissues, including dental applications, are called bioceramics [7, 11]. Many ceramic compositions have been tested for use in the body, but only a few have achieved human clinical application to date. Depending on the type of response in the body, bioceramics can be broadly classified as bioinert, bioactive, and resorbable, following the classification given above. Three factors influence the choice of ceramics and glasses as biomaterials: 1) physical and mechanical properties, 2) degradation of the material in the body, and 3) biocompatibility.

Bioactive ceramics are the bioceramics of choice for applications in regenerative medicine strategies and are the focus of the present chapter. The most widely used bioactive bioceramics are briefly described in this section, while the specific development of bioactive ceramics of cellular structure and their applications in bone tissue engineering and regeneration are considered in detail in subsequent sections.

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Hydroxyapatite [HA, Ca10(PO4)6(OH)2] is the main mineral constituent of teeth and bones. Synthetic hydroxyapatite has been used in several clinical applications such as a filler for bone defects, bone spacers and plates, bone-graft substitutes, and as a coating of the metal femoral stem in total hip replacement [11, 13–15]. The aim of the HA coating is to improve the interface between metal and bone [9]. Due to its excellent biocompatibility and bioactivity, HA has attracted major interest and research efforts in the last 20 years and, as discussed in detail below, is one of the most extensively considered materials for production of porous structures for bone tissue regeneration and tissue engineering scaffolds.

Tricalcium phosphate (e.g., b-TCP, Ca3(PO4)2, with Ca/P = 1.5) is an osteoconductive material that is also resorbable in the body. b-TCP is usually used in conjunction with synthetic HA to improve the resorbability of HA in applications such as the filling of bone defects left by cysts, sinus-floor augmentation, and bone cements. These bioresorbable ceramics became commercially available in the early 1980s for medical and dental applications [16]. Biphasic calcium phosphate consists of a mixture of HA and b-TCP. The bioactivity of this material can be controlled by manipulating the HA/b-TCP ratio. In this group of bioceramics, also calcium phosphate cements (CPCs) should be included, a term which encompasses a wide variety of formulations [17, 18]. Unlike sintered HA, CPCs can be actively remodelled in vivo [19].

Selected compositions of silicate glasses (e.g., 45S5 Bioglass) and glass ceramics (e.g., apatite-wollastonite) as well as some calcium phosphate glasses are bioactive glasses [7, 10]. The first bioactive silicate glass composition (46.1 SiO2, 24.4 Na2O, 26.9 CaO, 2.6 P2O5, in mol %), was reported in 1971 by Hench [20]. This bioactive glass, known as 45S5 Bioglass, is now used in the clinic as a treatment for periodontal disease (Perioglas) and as a bone filling material (Novabone) [10]. Bioglass implants have also been used to replace damaged middle ear bones, restoring hearing to thousands of patients [7, 10].

After immersion in body fluid bioactive glasses undergo a dissolution process that is instrumental in the formation of an apatite layer, which in the case of bioactive glasses is a carbonated hydroxyapatite layer (HCA). This not only means the glasses are resorbable in the body, but recent findings have shown that the dissolution products of bioactive glasses up-regulate seven families of genes that regulate osteogenesis and the production of growth factors [21, 22]. These findings may provide the reasons why certain compositions of bioactive glasses are specially effective in promoting bone growth and are class A bioactive materials (see Section 5.8.2) [10]. As discussed in detail below, cellular solids made from bioactive glasses (glass foams) are being increasingly considered as scaffolds in different tissue engineering strategies.

5.8.4

Scaffold Biomaterials for Tissue Engineering

Essentially, the above-mentioned materials and combinations thereof are the biomaterials available for scaffold design for tissue engineering applications. An ideal scaf-

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fold is one that mimics the extracellular matrix of the host tissue so that it can act as a template in three dimensions on which cells attach, multiply, migrate, and function. The criteria for an ideal scaffold for bone regeneration can be summarized as follows [3]:

The first criterion above refers to porosity and pore structure, which are key parameters determining the properties and the applicability of scaffolds for tissue engineering [1, 3, 26]. Figure 1 shows a summary of the different functions related to the pore structure in a tissue engineering scaffold [27]. In general, scaffold porosity, pore morphology, and pore orientation must be tailored to the particular tissue under consideration, and extensive evidence in the literature documents the critical influence of scaffold porosity and pore structure on the success of bone-tissue engineering approaches. For example, bone morphogenic protein (BMP)-induced osteogenesis has been shown to depend on pore size [28], porous structure [29], and overall scaffold geometry [30, 31] of hydroxyapatite-based scaffolds. The list above summarizes the ideal criteria for a versatile scaffold, but all the criteria may not have to be fulfilled for all applications. For example, for bone regeneration, where the scaffold is implanted directly into a bone defect, mechanical properties of the scaffold are critical, and the modulus of elasticity and mechanical strength of the porous material should match that of natural bone. If a scaffold with a modulus much lower than the host bone is implanted into a load-bearing site, the scaffold will fracture. If the modulus of the scaffold is much higher than that of bone the load will be transmitted through the scaffold instead of the bone (stress shielding), causing bone resorption rather than bone regeneration. On the other hand, for tissue engineering

5.8 Biomedical Applications: Tissue Engineering 553

ter < 10 mm). As mentioned above, macropore diameters should be in excess of about 100 mm to allow bone ingrowth, and if bone is required to penetrate within the ceramic implant, macropores should be large enough that they are connected by pore windows with diameters in excess of 100 mm to allow bone ingrowth. However, the presence of micropores is significant as they may act as attachment sites for cells such as osteoblasts (bone-growing cells). Moreover, interconnected micropores are important because they allow HA bodies to be machinable, a requirement to produce complex shapes. It has been discussed that optimum pore sizes required for bone ingrowth differ between in vitro and in vivo environments, and an excellent discussion about the design strategies of 3D porous structures for bone tissue engineering scaffolds is provided in the chapter of Baksh and Davies in Ref. [1].

The classical way to produce porous HA and TCP ceramics with pore sizes greater than about 100 mm is sintering of powders with suitable porogenic additives, such as paraffin, naphthalene, hydrogen peroxide [13, 51], and other organic particles [28, 48, 52–55], which are burnt out at elevated temperature with evolution of gases and leave mainly spherical pores in the HA or TCP structure. Traditional methods based on ceramic-slip foaming [56, 57], salt leaching [58,59], emulsion [60], and dual-phase mixing of polymer and ceramic slurries [61] have also been investigated, as well as the use of naturally occurring porous calcium-based structures, such as in the hydrothermal conversion of coral or bone [62, 63].

Different forms of porous HA substrates produced by the above methods are available commercially, such as Endobon (Merck, Darmstadt, Germany), Bio-Ostetic (Berkeley Advanced Biomaterials, Inc., Berkeley, CA, USA) and Pro Osteon (Interpore International, Irvine, CA, USA). Endobon is manufactured from bovine cancellous bone, whereby the organic portion of the bone is removed and the bone mineral (still maintaining the trabecular structure) is hydrothermally treated to obtain hydroxyapatite. The advantage of this product is that it mimics the structure of trabecular bone well and has high strength, but there is concern over the possibility of disease transfer. Pro Osteon is made by the hydrothermal conversion of marine coralline excavations to hydroxyapatite. The porous structure resembles trabecular bone, but of course is a less accurate mimic than HA produced from natural bone itself. A resorbable form of Pro Osteon has recently been developed [64]. Bio-Ostetic is an alternative resorbable porous implant based on TCP and HA.

There is little long-term data on the survivability of these implants, although what exists is favorable. After two years of implantation of Pro Osteon 200 with rigid anterior plating, as a bone-replacement implant in the cervical spine, there has been no record of plate breakage, screw breakage, resorption of the implant, or pseudarthrosis [65].

These implants are limited to bone repair and have generally been designed to repair bone as bone replacements, rather than by bone regeneration.

Traditional fabrication methods are in general not suitable for production of highly porous foams of sufficient strength, as required for tissue engineering scaffolds. In fact, most commercial suppliers of porous HA and calcium phosphate materials for bone reconstruction or substitution recognize the need for more research focusing on developing new products with improved strength. Further-

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more, pore interconnectivity obtained by traditional fabrication methods is low when compared to the total volume of the pores [52, 66]. As a consequence, in the last 15 years numerous novel fabrication procedures have been developed for production of improved cellular HA and TCP structures for scaffold applications, as described in subsequent paragraphs, and the subject is of continuous interest.

Template or Replication Techniques

Methods based on powder-filled polymer sponges, which are widely used for production of ceramic foams (see Chapter 2.1) have been used also for many years for production of porous HA and other calcium phosphate cellular structures. Recent reports dealing specifically with scaffolds for tissue engineering are available [67–72]. In these methods, sponges having the desired pore structure are infiltrated with a ceramic slurry and then dried and fired. The polymer substrate, usually polyurethane (PU), is removed by pyrolysis and the HA skeleton is sintered. Polyurethane systems free from silicones are favored to avoid contamination of the ceramic on pyrolysis [73]. Recently, HA foams with porosities in excess of 90 %, reticulated open-cell structure, and crushing strength of 0.2 MP have been produced by this method [73]. Sacrificial PU foams have been also used to produce glass-reinforced HA foams [74]. Addition of a phosphate glass to HA leads to increased mechanical strength of the foams, and methods have been developed to incorporate soluble bioactive glasses as reinforcement in HA [75–77]. The polymer-sponge method has been also proposed to manufacture macroporous calcium phosphate glass scaffolds of composition CaO–CaF2–P2O5–MgO–ZnO [70].

In related developments, HA foams with tailored pore structure have been produced from slurries of high solids content (60 wt %) by using commercially available HA powders and adding a dispersant [78]. Polyurethane foams and rapid-proto- typing models (see also below) were used as substrates, which could be eliminated by heat treatment at 650 and 700 C, respectively. Porosities in the range 50–96 % were produced by using different types of polymer substrates [78].

Several approaches are being investigated to improve the mechanical properties of HA foams produced by replication methods. It has been reported, for example, that combination of replicating techniques with gel casting of foams (see below) leads to HA cellular structures of improved fracture strength [67]. Compression strengths in the range 0.5–5 MPa were measured for these foams, depending on HA concentration in the starting suspension [67]. Another investigated approach to increase the mechanical strength of HA foams is coating of the porous structure with polymer or polymer/HA composite layers. In particular, the brittleness and low strength of HA foams obtained by the reticulate method have been improved by applying coatings based on poly(e-caprolactone) (PCL)/HA composites [35]. The PCL coating can be used also to effectively entrap antibiotic drugs and thus the construct can be considered a drug delivery system and used to enhance bone ingrowth and regeneration in the treatment of bone defects.

5.8 Biomedical Applications: Tissue Engineering 555

Foaming of Ceramic Suspensions and Gel-Casting Processes

Approaches to build foamlike structures into ceramic suspensions are being investigated as a means to obtain cellular HA structures. The methods are based on the addition of foaming agents to ceramic suspensions and forming foam structures by agitation. However, the consolidation of the formed foam, that is, its transformation into a rigid network, is a critical step that is difficult to accomplish by using traditional shaping methods involving simple liquid removal [79, 80]. Novel setting mechanisms are being developed with this aim, which can make use of nonporous molds offering advantages in terms of shaping capability and shape complexity. In these techniques, which may be grouped under the heading “direct consolidation techniques” [80], the structure of the ceramic slips is consolidated without powder compaction or removal of liquid, thus preserving the homogeneity achieved in the slurry state. Starch consolidation, protein consolidation, and direct coagulation casting (DCC) are some of the methods proposed to achieve porous structures that mimic that of cortical and/or trabecular bone [79, 81–83]. Well dispersed aqueous suspensions of HA powders with solids loadings as high as 60 vol% in the presence of suitable amount of ammonium polycarbonate as dispersant are used. The consolidator agents considered include etherified potato starch modified by hydroxypropylation and cross-linking, albumin, chicken egg white, and mixtures of two polysaccharide powders, namely, agar and locust bean gum [80, 84].

To obtain complex shapes and pore structures resembling the different parts of bone, well-known methods such as slip casting, foaming, and use of fugitive porogenic additives should be combined with starchand protein-consolidation techniques. It is claimed that smart combination of traditional and advanced techniques will enable the structure of the HA scaffold to be tailored according to the application requirements [79].

Gel-casting of foams is another recently developed method to produce macroporous ceramics [85–87]. This method yields compounds in various porosity fractions and controlled pore size that are noncytotoxic and have optimized strength and open spherical-cellular structure [86, 88]. Gel-casting involves dispersion of an aqueous suspension of HA powder with polyacrylate derivatives as dispersing agent. Acrylic monomers are also incorporated into the suspensions to promote gelation by in situ polymerization. Prior to this, the mixtures are foamed by agitation, aided by the addition of nonionic surfactants that reduce the surface tension of liquid–gas interfaces and serve as stabilizers of the foams. The gelation process of foamed suspensions is promoted by addition of initiator and catalyst for in situ polymerization of the added monomers by means of the redox system of ammonium persulfate and tetramethylethylenediamine. Before gelation, the mass is cast into molds of the required size and shape and dried. A subsequent sintering stage is required at 1350 C for at least 2 h for densification of the HA matrix [86, 88].

An improved method, termed “slurry expansion”, has been recently developed that leads to isotropic cellular HA structures with porosity close to 80 vol % and bimodal pore size distribution [89]. Figure 2 shows the obtained scaffold micro- (left) and macrostructures (right). Compressive strength and Young’s modulus were in the ranges 6–9 MPa and 0.5–2.5 GPa, respectively [89].

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Fig. 2 SEM images showing microstructure (left) and macrostructure (right) of HA foam scaffolds prepared by the slurryexpansion method of Martinetti et al. [89]. (Photos courtesy of Dr. R. Martinetti, FIN-Ceramica, Faenza, Italy.)

Alternative routes to produce calcium phosphate scaffolds at low temperature, to avoid the formation of high-temperature apatite phases, have been developed [90]. This method has the objective of increasing the in vitro and in vivo solubility of the scaffold by yielding a calcium-deficient hydroxyapatite. a-TCP powder is used as starting material and is mixed with hydrogen peroxide as foaming agent. Consolidation is not obtained by sintering but through low-temperature setting reactions [90]. Similarly, highly porous carbonated apatite bodies have been produced recently at low temperatures from mixtures of poly(vinyl alcohol) fibers, sodium chloride, and nanocrystalline carbonated apatite powders [91].

Use of Naturally Occurring Cellular Structures

Naturally occurring porous structures are considered to fabricate HA scaffolds [92]. A frequently used structure is coral. Hydrothermal and solvothermal methods are used to transform natural coral into HA after removal of the organic component by, for example, immersion in sodium hypochlorite [93]. Pore size in typical coral formations is in the range 200 to 300 mm. The porosity is interconnected and the structure resembles that of trabecular bone.

Wood has been proposed as a suitable template to fabricate porous HA scaffolds [94]. Heating wood in a nonoxidizing atmosphere at temperatures above 600 C results in decomposition of the polyaromatic constituents to form a carbon residue which can reproduce the original porous structure. A typical microstructure of a pyrolyzed specimen of Quercus alba is shown in Fig. 3 [94], which exhibits in principle adequate macropore size (> 100 mm) and pore interconnectivity required for bonetissue engineering applications. This porous carbon substrate can be subsequently transformed into a calcium phosphate skeleton after impregnation with a suitable precursor and calcination [94].