1.2 INTRODUCTION TO BIOMATERIALS

There is a necessity for replacing bone substance which has been lost due to traumatic or nontraumatic events. The lost bone can be replaced by endogenous or exogenous bone tissues, which is connected with several problems. The use of endogenous bone substance involves additional surgery; moreover the endogenous bone is available only in limited quantities. In case of exogeneous bone implants, the major disadvantage is that they may be rejected by human body, disease may be transmitted together with the implant, and also the clinical performance of exogenous bone is considerably inferior to fresh endogenous graft material. For these reasons there is growing need for fabrication of artificial hard tissue replacement implants. Research into novel materials for biomedical applications is ever increasing as the medical community look to improve the way in which disorders and trauma are treated. Issues with current materials and the additional trauma associated with the use of bone grafts has pushed research towards new materials to aid the required repair and/or the regeneration of bone after fracture or the removal of bony defects or cancerous bone. Many new materials have been developed in an attempt to address these concerns but there are still some issues surrounding the appropriateness of their mechanical properties, the ability of degradable materials to retain their properties once implanted and the ability to form the material in situ to the requirements of the surgeon. A biomaterial is a synthetic material used to replace part of a living system or to function in intimate contact with living tissue. There are a number of definitions describing biomaterials, but the most widely accepted definition is as follows: “a biomaterial is a nonviable material used in a medical device, intended to interact with biological systems” (Williams, 1987). The biomaterials industry world wide has an annual turnover of $2.3 billion in the field of hard tissue repair and replacement (total of $12 billion); there is currently a projected growth rate of 7-12% per annum for biomaterials in clinical applications (Bonfield, 1995; Suchanek and Yoshimura, 1998).

1. CURRENT CRANIOFACIAL BIOMATERIALS

The wide diversity of materials currently used in medicine is testimony to the number of advances that have occurred over the past fifty years. Biomaterials can be classified in four catagories, namely metals, ceramics, polymers and natural materials (of both plant and animal origin). In addition to this, two classes of materials may be combined to form a composite, which is a fifth class of biomaterial.

1. Metals

Currently there are three metals used for fixation devices: stainless steel (316L), cobalt chromium alloy and titanium, which is either commercially pure or alloyed with vanadium, aluminium and/or niobium. Of these, titanium and its alloys are the most popular for craniofacial use because of its high strength and reduced artefacts on CT and MR imaging (Saxe et al., 1982; Fiala et al., 1994). Permanent metal fixation implants leave a residual implant within the body, which may cause secondary problems. Metal implants may cause stress shielding and subsequent bone weakening in load bearing applications; implants may be palpable and cause tissue irritation and pain, necessitating their removal. There are also concerns of implant migration and interference with diagnostic and therapeutic applications of radiation (Pietrzak and Eppley, 2000). Metallic implants are often darker than the surrounding tissue and may cause discolouration, also, due to the rigidity of metals, it is often difficult to match the implant to the contours of the underlying bone. This may result in implant edges that are visible through the skin (Ahn et al., 1997). The application of metallic implants close to the skin may result in temperature effects. During cold weather the implant may be cooled sufficiently to cause pain or discomfort (Daniels et al., 1990). Observations of malignancy in association with metal fixation devices are rare, with only 20 cases reported in the literature (Rubin et al., 1997), with latencies between 1-30 years. Only one was reported to occur above the neck. This implies that the local risk of cancer is low with metallic implants. However, it has been found that there is a significant increase in the risk of systematic tumours, e.g. lymphatic, in patients who have undergone total hip replacement (Gillespie et al., 1988). It should be noted that there are other factors associated with arthritic disease which may have contributed. The release of metal ions from implants via corrosion has raised concerns over system toxicity. The accumulation of ions can result in local fibroblastic tissue reaction which has been implicated in implant failure of maxillofacial plates (Byrne et al., 1973; Thomas et al., 1988).

1.2.1.2 Polymers

There are many different polymers that have been used in maxillofacial surgery, however, only the most popular will be considered here. Cancer associated with a polymer implant is very rare. Even a single case is not reported in association with a polymer craniofacial implant (Rubin and Yaremchuk, 1997).

1.2.1.2a Polyethylene This is normally used in its ultra high molecular weight form as the low friction acetabulum cup part of total hip replacement. It is also made as a porous implant, Medpor, which allows the in growth of both soft tissue and bone, which makes the implant very difficult to remove. It has been used in orbital fracture repair and in the management of cosmetic and post traumatic facial deformities (Rubin and Yaremchuk, 1997).

1.2.1.2b Polytetrafluoroethylene (PTFE) Proplast is a porous composite of PTFE and carbon fibres. It has been used since 1960 as filler in non-load bearing applications because of its porous surface and consequential ingrowth of fibrous tissue. Proplast has been used for malar, chin, nasal and orbital floor implants with acceptable complication rates. When it was used as an interpositional disc implant within the temporomandibular joint, complication rates rose to well above acceptable limits. Nearly all implants fractured under the load of the joint. Additionally, the Proplast particulates caused a vigorous foreign body reaction and subsequent erosion of the joint (Rubin and Yaremchuk, 1997).

1.2.1.2c Polymethylmethacrylate (PMMA) This is commonly used as cement, formed by mixing powdered polymer with liquid monomer. The resultant chemical reaction is exothermic, and toxicity has been associated with the unreacted monomer (Rubin and Yaremchuk, 1997). The hardened methylmethacrylate is brittle and prone to fracture. Hard tissue replacement (HTR), is a porous composite of PMMA and polyhydroxylmethacrylate. It has a calcium hydroxide coating which is thought to impart a negative surface charge to encourage bony in-growth (Eppley et al., 1990). Methylmethacrylate has been shown experimentally to cause vasodilatation (Peebles et al., 1972). After insertion of hip prostheses with freshly mixed MMA cement into the femur, patients have suffered acute hypotension that can progress to cardiovascular collapse and death (Kepes et al., 1972; Milne, 1973; Monteny et al., 1978).

1.2.1.3 Composite Materials

Composite materials allow the use of two materials which individually, would be unsuitable for the desired application. This allows the combination of polymers with insufficient mechanical properties with ceramics and glasses of low fracture toughness to produce a material which possesses the best of both materials. Composite used within medicine can be broadly cast into two classes: particulate reinforced or fibre reinforced. Particulate reinforcement is easier to produce as the composites are isotropic; hence may be prepared in-situ, however, this leads to lower mechanical properties. Fibre reinforced composites are anisotropic, as is bone, and possess superior mechanical properties to their particulate analogues.

1.2.1.4 Bioresorbable Materials

Due to their nature, resorbable materials are usually utilised in short term non load bearing applications such as sutures and drug delivery devices, however, applications for medium term such as fracture fixation are being considered as new, novel materials are researched. As discussed previously, metallic implants are the material of choice to stabilise fractures, but they may exhibit complications such as stress shielding, migration, tissue irritation and pain. These complications may require secondary surgery to remove the implant. It presents the patient with prolonged trauma and recovery time, whilst adding to the cost of surgery and requiring more theatre time. Resorbable materials offer a potential solution to these problems, by eliminating the need for follow up surgery to remove the implant. Stress shielding is eliminated as a gradual loss of material, and hence material properties, exposes the immature bone to gradually increasing levels of stress, aiding remodelling. This will clearly be beneficial to medicine as patient trauma, surgical time and related costs are all reduced. Ideally, a resorbable polymer should serve its temporary need, be metabolically eliminated completely by the body within an acceptable time period, without eliciting an adverse physiological response (Pietrzak et al., 2000). The current market for resorbable biomaterials is dominated by poly-lactic acid (PLA), and poly-glycolic acid (PGA), or co-polymers of these materials. PGA is not used on its own due to its fast resorbtion rates. The low mechanical properties of PLA and PLA/PGA copolymers have limited their use in low load bearing applications. The relatively fast degradation of these polymers may lead to the build up of degradation products, which leads to a decrease in pH and possible aseptic foreign body reactions (Rubin et al., 1997). Applications within craniomaxillofacial surgery include: plates, screws and mesh panels.

1.2.1.5 Ceramics and Glasses

1.2.1.5a Alumina Extensive research on Al2O3 was done during the 1950s and 1960s. It was used as a coating for the articulating surfaces in total hip replacement as it has very favourable wear and corrosion properties as well as good biocompatibility. However, its use was discontinued after complaints of post-operative pain.

1.2.1.5b Hydroxyapatite Hydroxyapatite (Ca10(PO4)6(OH)2) is a complex calcium phosphate mineral, the main structural element of vertebrate bone which can be formed synthetically in either dense or porous forms. Porous hydroxyapatite is too brittle to bear a load but will allow bony in-growth when used as an onlay.

1.2.1.5c Bioactive glasses Hench et al. (1971) discovered that bone can bond chemically to certain glass compositions. This group of glasses has become known as bioactive glasses based on the following definition (Hench and Andersson, 1993): “A bioactive materials is one that elicits a specific biological response at the interface of the material which results in formation of a bond between the tissues and material”. Bioactive glasses have numerous applications in the repair and reconstruction of diseased and damaged tissue, especially hard tissue (bone). One aspect that makes bioactive glasses different from other bioactive ceramics and glass ceramics is the possibility of controlling a range of chemical properties and rate of bonding to tissue. The most reactive glass compositions develop a stable, bonded interface with soft tissues as shown by Wilson et al. (1981). It is possible to design glasses with properties specific to a particular clinic application. This is also possible with some glass-ceramics, but their heterogeneous microstructure restricts their versatility.

1.2.2 TYPES OF BIOMATERIAL-TISSUE INTERFACE

The ideal goal of surgery is to restore the structure and function of the body to its natural state (Pietrzak et al., 2000). Bone is a unique tissue in that it can heal itself without scarring, a process referred to as regeneration (Clokie et al., 2000). Developers of new materials for bone regeneration applications are faced with the challenge of utilising this feature. Even a single material implanted in the living tissues is inert; all materials elicit a response from the host tissue. The response occurs at the tissue-implant interface and depends upon many factors (Hench and Ethridge, 1982), listed in Table 1.2. There are four general types of implant-tissue responses, as summarized in Table 1.3.

Table 1.2 Factors affecting implant-tissue interface response (Hench and Wilson, 1993).
|Tissue side |implant side |
|Type of Tissue |Composition of Implant |
|Health of Tissue |Phases in Implant |
|Age of Tissue |Phase Boundaries |
|Blood Circulation in Tissue |Surface Morphology |
|Blood Circulation at interface |Surface Porosity |
|Motion at Interface |Chemical Reaction |
|Closeness of Fit |Closeness of Fit |
|Mechanical Load |Mechanical Load |

It is critical that any implant material avoid a toxic response that kills cells in the surrounding tissues. Released chemicals can migrate within tissue fluids and cause systemic damage to the patient (Black, 1984). One of the interests in glass / ceramic implant is their lack of toxicity and corrosion resistiveness.

Table 1.3 Consequences of implant-tissue interaction (Hench and Wilson, 1993).

|Impalnt-Tissue reaction |Consequence |
|Toxic |Tissue dies |
|Biologically nearly inert |Tissue forms a non-adherent fibrous capsule aroundde implant |
|Bioactive |Tissue forms an interfacial bond with the implant |
|Dissolution of implant |Tissue replaces implant |

The most common response of tissues to an implant is formation of a non-adherent fibrous capsule. The fibrous tissue is formed in order to “wall off” or isolate the implant from the host. It is a protective mechanism and with time can lead to complete encapsulation of an implant within the fibrous layer. Metals and most polymers produce this type of interfacial response. Biologically inactive, chemical inertness of alumina or zirconia, results in a very thin fibrous layer under optimal conditions. More chemically reactive metallic implants elicit thicker interfacial layers. However, it is important to remember that the thickness of an interfacial fibrous layer also depends upon motion and fit at the interface, as well as the other factors indicated in Table 1.2. The third type of interfacial response, indicated in Table 1.3, is due to bond forms across the interface between implant and the tissue. This is termed as “bioactive” interface. The interfacial bond prevents motion between the two materials and mimics the type of interface that is formed when natural tissues repair themselves. This type of interface requires the material to have a controlled rate of chemical reactivity. An important characteristic of a bioactive interface is that it changes with time, as do natural tissues, which are in a state of dynamic equilibrium. When the rate of change of a bioactive interface is sufficiently rapid the material “dissolves” or “resorbs” and is replaced by the surrounding tissues. Thus, resorbable biomaterials must be of a composition that can be degraded chemically by body fluids or digested easily by macrophages. The degradation products must be chemical compounds that are not toxic and can be easily disposed of without damage to cells.

3. MECHANISM OF BIOACTIVE BONDING

When a bioactive material is implanted in the body, a series of biophysical and biochemical reactions occur at the implant-tissue interface. These reactions eventually result in mechanical strong chemical interfacial bonding (Hench and Ethridge 1982; Hench 1987). This attachment is called “Bioactive Fixation”. Hench (1971) reported that certain compositions of silicate glasses can form a bond with bone tissue. Research has shown that these bioactive glasses can also bond with certain types of connective tissue through attachment of collagen to the glass surface (Hench, 1991). The basis of the bone bonding property of bioactive glasses is chemical reactivity of the glass in body fluids. The surface chemical reactions result in the formation of a hydroxycarbonate apatite (HCA) layer to which bone can bond. Bonding occurs due to a sequence of reactions. On immersion of a bioactive glass in an aqueous solution, three general processes occur: (i) Leaching (ii) Dissolution (iii) Precipitation Leaching is characterized by release, usually by cation exchange with H+ or H3O+ ions, of alkali or alkaline earth elements. Ion exchange is easy because these cations are not part of the glass network, they modify the network by forming non-bridging oxygen bonds. The release of network-modifying ions is rapid for glasses in the bioactive compositional region. This ion exchange process leads to an increase in interfacial pH (> 7.4). Network dissolution occurs concurrently by breaking of -Si-O-Si-O-Si- bonds through the action of hydroxyl (OH) ions. Breakdown of the network occurs locally and release silica into solution in the form of silicic acid [Si(OH)4]. The rate of dissolution of silica depends very much on glass composition. The dissolution rate decreases greatly for the composition of > 60% SiO2 because of the larger number of bridging oxygen bonds in the glass structure. The hydrated silica (SiOH) formed on the glass surface by these reactions undergoes rearrangement by polycondensation of neighboring silanols, resulting in a silica-rich gel layer. In the precipitation reaction, calcium and phosphate ions released from the glass together with those from solution form a calcia-phosphate-rich (CaP) layer on the surface. When formed in vitro, the CaP layer is mainly located on the top of the silica gel, whereas in vivo it is formed within the gel layer. The calcium phosphate phase that accumulates in the gel surface is initially amorphous (a-CaP). It later crystallizes to a hydroxycarbonate apatite (HCA) structure by incorporating carbonate anions from the solution within the a-CaP phase. The mechanism of nucleation and growth of the HCA layer appears to be the same in vitro and in vivo and is accelerated by the presence of hydrated silica. All these processes can be summarized by five reaction stages: Stage 1: Leaching and formation of silanols (SiOH). Stage 2: Loss of soluble Silica and formation of silanols. Stage 3: Polycondensation of silanols to form hydrated silica gel. Stage 4: Formation of an amorphous calcium phosphate layer. Stage 5: Crysatllization of a hydroxycarbonate apatite layer.
Table 1.4 summarizes these five reaction stages in detail.

1.2.4 TYPES OF BIOACTIVE GLASSES

Hench broadly divided the bioactive glasses in two classes: Class A and Class B.

1. Class A bioglasses These are materials, which are osteoproductive in nature. They can bond to both soft and hard tissues. They release Si ions in the form of silicic acid, which provides a silica gel layer, which further enhances the precipitation of amorphous CaP layer and also rapidly crystallizes HCA layer. Generally HCA layer is formed within 1-10 hours for this type of materials. Bioglass 45S5® (45 wt.% SiO2, 24.5 wt.% CaO, 24.5 wt.% Na2O and 6.0 wt.% P2O5) is a class A bioactive glass.

2. Class B bioglasses These are the materials, which are osteoconductive in nature. They don’t produce silica gel layer and they only bond to hard tissues. In this type of glasses HCA layer is formed within 24 hours to several days.

Table 1.4 Reaction stages of a bioactive glass implant (Hench and Wilson, 1993).

|Stage |Reactions |
|1 |Rapid exchange of Na+ or K + with H + or H3O + from solution: |
| |Si – O – Na+ + OH ─ [pic] Si–OH+ Na+ (solution) + OH─ |
| |This stage is usually controlled by diffusion and exhibits a t-1/2 dependence. |
|2 |Loss of soluble silica in the form of Si(OH)4 to the solution , resulting from |
| |breaking of Si–O–Si bonds and formation of Si – OH (silanols) at the glass |
| |solution interface: |
| | |
| |Si – O – Si + H2O [pic] Si – OH + OH – Si |
| | |
| |This stage is usually controlled by interfacial reaction and exhibits a t |
| |dependence. |
|3 |Condensation and repolymerization of a SiO2- rich layer on the surface |
| |depleted in alkalis and alkaline-earth cations: |
| | |
| |[pic] |
|4 |Migration of Ca2+ and [pic]groups to the surface through the SiO2-rich layer |
| |forming a CaO-P2O5-rich film on top of the SiO2 rich layer, followed by |
| |growth of the amorphous CaO-P2O5-rich film by incorporation of soluble calcium |
| |and phosphates from solution |
|5 |Crystallization of the amorphous CaO-P2O5 film by incorporation of OH─, CO[pic] |
| |or F─ anions from solution to form a mixed hydroxyl, carbonate, fluorapatite layer. |

4. CLINIC APPLICATIONS FOR BIOGLASSES

Biomaterials science involves both biology and material science & engineering. It has developed into an important division of the biomedical field. Hench et al. (1971) reported that certain compositions of silicate glasses are bioactive. The main clinical use of bioactive glasses has been in the area of bone grafting in the oral cavity and as a glass ceramic for orthopedic use. There has also a significant amount of work on the development of bioactive glass bone cement. Successful tooth root implants were achieved in baboons by Stanley et al. (1976) using bulk Bioglass®. In 1973, a surface-active glass ceramic of lower bioactivity (Ceravital®, trademark of E. Leitz & Co.) was first tested in Germany (Bromer, 1976). Developed by Bromer (1976) and Blenke (1978), it was tested by Gross and Strunz (1982) in maxillofacial applications and as an ossicular chain prosthesis by Reck (1981). FDA approval for sales of middle ear prosthesis devices was given in 1982. Soft-tissue bonding to Bioglass® was demonstrated by Wilson et al. (1981). After establishment of biocompatibility in vivo and in vitro by Wilson et al. in 1981 and successful comparison with a wide range of alternative materials in the middle ear by Merwin et al. (1982), the FDA gave approval, which allowed sales of Bioglass® middle ear prostheses in 1984. The first successful use of 45S5 bioglass® was a replacement for the ossicles in the middle ear, as a treatment for the conductive hearing loss which develops when sound waves impinging on the tympanic membrane do not reach the oval window in the inner ear. Conduction loss in the middle ear can result from trauma, chronic infection or be due to congenital abnormality. Replacement of one or more of the ossicles can restore the continuity of the conducting system. Initially, polymers, both porous and solid, and metals of various types are used. These materials engender a fibrous tissue reaction which effectively holds the implant in place. Scar tissue around an implant however will dampen, rather than transfer sound waves and implants which promote scar tissue become gradually less efficient. The major mode of failure is extrusion through the tympanic membrane. When metals or plastic implants are in continuous contact with soft tissue of the ear drum they wear through and are lost through the hole. These two problems, immobilization by a means other than fibrous tissue and prevention of extrusion, are solved by the special properties of 45S5 Bioglass® which bonds to both hard and soft tissue (Wilson et al., 1993). Beginning in 1985, bioactive glasses have been used in a number of clinical applications, and their use has increased steadily over the recent years (Williams, 1985; Hulbert et al., 1987; LeGeros, 1988; Yamamuro et al., 1989; Cook et al., 1990; Yamamuro et al., 1990; De Groot, 1991). Research has shown that these bioactive glasses can also bond with certain types of connective tissue through attachment of collagen to the glass surface (Hench, 1991). These bioglasses have as bone graft substitutes, in dentistry, as drug delivery vehicle, macroporus scaffolds and in molecular tailoring. Clinical trials of Bioglass® tooth root implants for maintenance of the alveolar ridge of edentulous patients were completed by Stanley et al. (1986). Preclinical testing in monkeys of Bioglass® powder for periodontal testing in dogs of bulk Bioglass® for facial augmentation was completed by Wilson et al. (1985) and Merwin et al. (1982). In oral surgery (dentistry), two main uses are reported. Dental pins are devices which are drilled into the mandible and are used to hold prosthetic teeth in place. These are usually made from bioactive materials such as low silica glass. A large number of dental filers are also used, cone-shaped devices made from 45S5 Bioglass®, have been used to fill the defect in the jaw which is created when a tooth is removed. Removal of one or more teeth produces changes in the jaw bone which are followed by gradual bone loss so that the normal shape of the bone which supports healthy teeth changes to narrow “knife edge” ridge with reduced height which cannot comfortably support dentures. Without some means of preventing this bone loss, denture wearers are often destined to suffer increasing discomfort from ill-fitting dentures and in many case may eventually become unable to wear dentures at all. These devices have now been in use for almost a decade and those made from bioactive glass have proven to be more successful than others which have been tried (Wilson et al., 1993). There have been a number of recently published clinical studies on bone grafting in the oral environment (Low, 1997; Forum, 1998; Lovelace, 1998; Furusarwa and Mizunama, 1997). These clinical studies have compared the amount of bone fill of the bioactive glasses in periodontal defects to standard clinical practice of debridement and found significantly more bone formation In Finland an innovative use of solid, cast bioactive glass implants has been used in the treatment of facial injuries in which the bone which supports the eye is damaged. Another use, in ENT surgery at present, but with the potential for much wider applications, is that of providing a soft tissue seal for an implant which passes through the skin. Electrodes, which are an essential part of an extracochlear implant developed at the University of London to treat profound deafness, must be connected both to the cochlea (or inner ear) and to the complex electronics on the outside. Any material which passes through the skin and subcutaneous tissues without an effective seal can provide a channel along which bacteria, which are always present on the skin, can move to cause infection. This is in any case dangerous in any position close to brain because of potentilly fatal meningitis. The anchors which contain the electrodes in this implant are coated with 45S5 Bioglass® and are implanted in the cranium. The implant is placed so that part is in bone where the Bioglass provides a bond to immobilize it. The part which passes through the skin bonds to the soft tissues and provides the essential seal. The soft tissue bond is protected from damage due to movement, by the bone bond and relatively thin layer of soft tissue overlying it. Such implants have been in place for several years with no significant problems. Another application which is being developed using 45S5 Bioglass® particulate is for the treatment of patients with paralysis of one of the vocal cords. According to a report published in “The glass researcher” by Hench (2002) by inventor of Bioglass, trademarked 45S5 bioglass® is a miracle material, it has saved more than 2,000, 000 human teeth and repaired numerous bony defects in the jaw, head and cranium in the past 16 years. More than 800,000 units of Bioglass®, marketed as Perioglas® by US Biomaterials Corp., are estimated to have been sold since 1986. Approvals by regulatory bodies for use of bioactive glasses in orthopedics have recently been achieved and the bioactive glasses are used in a variety of applications where large bone defects are present, such as in revision surgery for total hip replacements and in spinal repair. Their clinical performance is an improvement over previously used bioinert implants or Class B, osteoconductive implants as shown in Figures 1.15 and 1.16. Several thousand successful cases have been documented. Bioceramics serve many different functions in repair of the body, which are summarized in Figure 1.17.

1.2.6 LATEST DEVELOPMENTS

From the advances in related medical technologies, bioactive materials appear to be entering a new stage of development in which they will have to be skillfully designed to exhibit a variety of new functions. Recently, research at Imperial College, London, conducted in Dr. Hench’lab in collaboration with Prof. Julia Polak of the Faculty of Medical Sciences, has shown that the unique biological behavior of bioactive glasses is due to activation of seven families of genes that are present in bone growing cells (osteoblasts).
[pic]
Fig. 1.15 Survivability of alveolar ridge maintenance implants (Stanley et al., 1995).

The genes are activated to produce various proteins that serve as growth factors and enhance the proliferation of new osteoblasts. An important aspect of this discovery is that the biological control is due to the controlled release of critical concentrations of soluble silica and calcia ions from the bioactive glasses. In other words, the bioactive glasses serve as a controlled release system as well as a biologically active substrate for anchoring and growth of new cells. This discovery provides a conceptual base for new cell-based medical therapies. Two alternative cellular-based methods can be used to repair diseased, damaged or aged tissues: (i) in situ stimulation of tissue regeneration (ii) in vitro growth of tissues followed by in vivo transplantation. Both methods are enhanced by use of bioactive materials that release biologically active molecules at controlled rates.

[pic]
Fig. 1.16 Extrusion-displacement rate vs. time of implantation for various middle ear ossicular replacement (Lobel, 1995).

Binary gel-glasses are being designed to stimulate specific responses at the molecular level. Molecular modifications of S70C30 are expected to elicit specific interactions with cell integrins and thereby direct cell proliferation, differential, and extracellular matrix production and organization. Macroporous foams of this composition are being designed to activate genes that stimulate regeneration of living tissues whilst reducing the possibility of infection after surgery. Bioactive glasses and newly developed bioactive gel-glasses made from the SiO2-CaO-P2O5 and SiO2-CaO systems can be used as controlled release systems and bioactive, resorbable scaffolds. Bioactive foams made from the gel-glass systems are especially attractive for these new applications because they are hierarchical materials that provide interconnected porosities of 100-500 micrometers to produce the 3-D architecture required for stable cell-cell interactions and maintenance of an intercellular blood supply.
[pic]
Figure 1.17 Clinical uses of bioceramics (Hench and Wilson, 1993).

Molecular tailoring of multifunctional scaffolds by incorporation of specific LaCourse’s group, US Biomaterials and Prof. Joseph Simmons’ group at the University ofArizona has shown that both bone and lung cells can be grown on special compositions of bioactive glass fibers. Current studies are underway to incorporate the bioactive fibers within the porous bioactive scaffolds as a multifunctional bio-photonics detector system. The US Defense Advanced Research Projects Agency (DARPA) is funding a multi-disciplinary, multi-institutional project to use integrated bioactive fibers and scaffolds as hosts for living lung cells (scaffolds) and as a means of optically monitoring the vitality of the cells in situ (fibers). Recent publications of Dr. Ion Notingher and Dr. Sophie Verrier, have shown the feasible of obtaining Raman signatures of mouse lung type cells grown on bioactive glasses without harming the cells. For these new functions (different mechanical properties, drug delivery capability, the capability to activate proteins and/or cells for tissue regeneration and tissue engineering, biomimetics, nanotechnology, etc. these materials have been successfully applied as solids and particulates and may be combined with other materials, both natural and synthetic, to provide treatment for many disparate clinic conditions. Combinations of bioactive glasses and apatites could give very satisfactory results in terms of acceleration of the bioactivity of the apatites, which, due both to their composition and structure, are the ceramics most similar to the mineral component of the bone. Use of glass as an active interface with biological systems has come a long way in just over 36 years. Now that we know that there is a genetic basis for the behavior of bioactive glasses, this understanding could be used to develop novel compositions that will prevent deterioration of tissues by enhancing the human body’s own regeneration capacity by stimulating genes which initiate the repair process, bringing us one step closer in facing the greatest challenge. It is thought that the economic advantages of these new approaches may aid in solving the problem of caring for an aging population.