Amalgam

For many years, amalgam was the material of choice for restoration of teeth. It consists of liquid mercury and silver–tin alloy powder. Reaction between these components gives a hard silvery-grey material. Nowadays, with increasing concern about restoration aesthetics and mercury toxicity, alternative polymer-based, tooth-coloured materials such as composites, or glass ionomer cements, and hybrids of the two, have been brought into the market. These, unlike classical amalgam restorations, are bonded to the excavated tissue. The ability to bond to dentine has meant that less tissue needs to be removed. Additionally, larger cavities can be successfully restored as the bonded material provides support for the remaining tooth structure. A major problem, however, if the bond fails is that, unlike silver and mercury, polymers are not generally antibacterial.

Composites

Dental composites consist of a resin matrix (mostly high molecular weight, hydrophobic, dimethacrylate monomers), silane-treated inorganic fillers (often silica based) and an initiator/activator (typically camphorquinone and dimethylparatoluidine) system. Upon exposure to a blue light (e.g. from an LED source), the highly filled composite pastes placed within a tooth structure set via free radical initiated polymerisation. This produces a cross-linked three-dimensional network and a material with suitable strength and wear characteristics to replace lost tissues. Gaining a strong, resilient bond between the composite and dentine can, however, be difficult. It generally requires a number of complex dentine pre-treatment steps and use of additional more hydrophilic methacrylate resin adhesives (see below).

A major problem with dental composites is polymerisation shrinkage during set. This causes stress on the bond between the material and tooth structure leading to gap generation at their interface. Additionally, differences in thermal expansion coefficients of the material and tooth can mean changes in temperature widen these gaps. This problem is then further exacerbated by cyclic loading. Water and enzyme penetration (‘nanoleakage’) at the material/dentine interface may weaken the adhesive and break down collagen, respectively. Ultimately ‘bacterial microleakage’ through widened gaps of micron dimension can occur. This can lead to staining around the restoration margins, breakdown of adhesion, recurrence of caries, inflammation of the tooth pulp and also postoperative pain (Hersek et al., 2002; Murrey et al., 2002).

Recurrent caries is considered the major contributing factor responsible for tooth restoration failure. Attempts made to reduce the polymerisation shrinkage by modifying the composite composition (e.g. increasing the filler loading, using a high molecular weight methacrylate monomer or pre-cured resin as filler) and stress on the tooth by placement of the material in thin consecutive layers have been only partially successful. Additionally, in modern dentistry, there is a drive to remove less of the disease damaged tissue (Splieth et al., 2001). In this situation, there is greater possibility of residual bacteria within the tooth cavity and so development of materials with some antibacterial action is becoming ever more crucial.

Glass ionomer cements (GICs)

Glass ionomer cements are self-adhesive, aesthetic filling materials. They consist of fluoroalumino silicate glass and an aqueous solution of polyacid, mostly polyacrylic. GICs set by acid-base reaction between the glass filler and the polyacid in the presence of water. GIC restorations have significantly reduced bacterial microleakage compared with those of composites for a number of reasons. Firstly, provided they do not dehydrate, GICs are dimensionally more stable during setting. Additionally, the polyacid in GICs associates with calcium ions of hydroxyapatite in enamel and dentine providing a mechanism for chemical bonding to tooth structure. There is therefore no need for the complex dentine pre-treatment procedures required with composite adhesion. Instead, simple conditioning of the excavated tooth surface with a polyacrylic acid solution is generally all that is required prior to material application. Although initial bond strengths of GICs are lower than those for composites, experimentally and clinically they have been found to be more durable.

Furthermore, GICs exhibit an initial burst fluoride release upon set. The acid/base setting process additionally leads to fluoride within the polymer matrix phase of the set cement which can subsequently be gradually released over a prolonged period (Forsten, 1990; De Moor et al., 1996). GICs also take up fluoride, for example from toothpaste, and re-release it over time. Fluoride can react with hydroxyapatite and form a more acid-resistant fluroapatite (De Bruyne and De Moor, 2004; Wiegand et al., 2007). At high concentration, it will inhibit bacterial metabolism and growth (Hamilton, 1990; Forss et al., 1991). Initial moisture sensitivity, poorer wear resistance and in particular low flexural strength of GICs, however, limit their application to low-load-bearing clinical situations. This has lead therefore to the development, by many companies, of GIC/composite hybrid materials.

Hybrids of composites and GICs

Polyacid-modified composite resins (compomers) contain all the components of composites but in addition methacrylate monomers with acidic functional groups. They also have some reactive glass powder of a similar type to that used in GICs. After setting by polymerisation, these fillers attract water (2–3.5% by mass) from their surroundings. The water can elicit a surface acid-base reaction between the acidic monomer and the reactive glass filler providing some fluoride release (Nicholson, 2007). Fluoride release is, however, often limited (Meyer et al., 1998) and water uptake can cause decline in mechanical properties. Additionally, polymerisation shrinkage and the need for composite adhesives remain a major concern of compomers.

Resin modified glass ionomers (RMGICs) are a further hybrid of dental composites and GICs but unlike compomers they have significant water content in the original formulation. This enables them to set via both polymerisation and acid–base reaction. They often contain the hydrophilic monomer hydroxyl ethyl methacrylate (HEMA) in addition to all GIC components (Smith, 1998; Xie et al., 2004). The hydrophilicity of HEMA can aid spreading and penetration of these formulations into dentine. Additionally, the polyacid gives a chemical bonding mechanism to tooth structure. Furthermore, RMGICs show a comparable level of early fluoride release and recharging capability to GICs (Xu and Burgess, 2003). Modern RMGICs can have much improved flexural strength compared with conventional GICs (Xie et al., 2000). Low surface hardness and wear resistance, however, remain potential problems (Peutzfeldt, 1997; Xie et al., 2000). Clinically, to overcome these limitations, GICs and RMGIC are often used as a base or liner beneath stronger composites but there is still room for further dental material improvement.

Resin adhesives

Resin adhesives are used to enhance retention of both composites and compomers to tooth structure and hence prevent bacterial microleakage (Van Meerbeek et al., 2001). There are a wide range of different formulations but they generally consist of the components of RMGICs and composites but with much lower levels of filler to improve fluidity. Preparation of enamel for bonding is relatively straightforward requiring just partial etching of its hydroxyapatite crystals with 35% phosphoric acid gel for ~30 s followed by rinsing and drying. This produces a microscopically roughened surface into which methacrylate monomers can penetrate and bond by micromechanical retention (Van Landuyt et al., 2007).

The heterogeneous structure and high water content of dentine (~22% by volume) makes it less compatible with hydrophobic methacrylate resin adhesives than enamel. Currently the most reproducible dentine bonding systems comprise an etchant, primer and adhesive resin. The etchant (commonly phosphoric acid gel) is required to remove any cutting debris left by excavation of carious tissue and additionally a surface layer of hydroxyapatite. This leaves a mat of collagen fibrils into which the primer can subsequently penetrate. The primer is often an aqueous or ethanolic solution of largely hydrophilic monomers (e.g. HEMA) which can aid penetration of a low-viscosity adhesive methacrylate resin into demineralised collagen. Upon drying and cure a hybrid layer of resin modified collagen is formed within which the resin and collagen are micromechanically interlocked. Subsequent bonding between the hybrid layer and composite restoration is chemical in nature and achieved by copolymerisation of methacrylate groups in the primer, adhesive and composite. A major drive in modern dental adhesive technology has been to simply the procedure but this has often been at the expense of bonding reliability (Yoshida et al., 2004).

Calcium phosphate-releasing materials

Calcium and phosphate ions released into collagen from filler particles embedded within methacrylate composites can reprecipitate as hydroxyapatite. Release kinetics of these ions from the set composite depends strongly upon both the methacrylate hydrophilicity and the calcium-phosphate filler chemistry/aqueous solubility. Calcium phosphates vary in solubility according to the Ca/P ratio as well as the surrounding pH. The solubility of calcium phosphates generally increases as the Ca/P ratio decreases. Under neutral conditions, the solubility increases in the following order: fluroapatite [FA,Ca₁₀(PO₄)₆F₂], <hydroxyapatite [HA,Ca₁₀(PO₄)₆OH₂], <tricalcium phosphate [TCP), Ca₃(PO₄)₂] ~ amorphous calcium phosphate [ACP, Ca_xH_y(PO₄)_6−x·nH₂O] <dicalcium phosphate [DCP, CaHPO₄], <monocalcium phosphate [MCP, Ca(H₂PO₄)₂] (Bohner, 2000a; Dorozhkin, 2008). Calcium-phosphate solubility also tends to increase as the pH is reduced. It is potentially possible therefore to produce ‘responsive’ composite materials that release higher levels of calcium and phosphate when surrounding bacteria produce acid.

Amorphous calcium phosphate (ACP) has been included in various methacrylate monomers (Regnault et al., 2008), and proven to give levels of calcium and phosphate release sufficient to promote mineral re-deposition (O’Donnell et al., 2006). ACP composites have recently been commercialised as a cement, adhesive (Aegis Ortho) and sealant. The ACP adhesive formulation has been observed to have shear bond strength comparable with that of a commercial RMGIC adhesive (Foster et al., 2008). Further work is still required, however, to assess if continuing release of calcium phosphate can make the ACP adhesive bond more durable under cyclic stress, thermal loading and ultimately in vivo as observed with GIC bonding.

Nanoparticles of dicalcium phosphate (DCP) have also been combined with silicon-carbide whiskers and added to dental resins. Chemical curing formulations were developed and proven to have calcium and phosphate release comparable with that of ACP composites (Xu et al., 2006, 2007). The small size of the filler particles may have limited the filler loading capacity (60 wt%) but this problem may be overcome through use of reactive fillers such as MCP and TCP that produce DCP as below. Although calcium phosphate-releasing materials have proved to be successful in re-mineralising defective tooth structure, they have no antibacterial action.

Antibacterial monomer-containing materials

The antibacterial monomer, 12-methacryloyloxydodecyl-pyridinium bromide, MDPB has been included in some methacrylate composites and adhesives. This can be achieved without adverse effect on mechanical strengths because after polymerisation the MDPB is copolymerised with the resin phase and immobilised. Incorporation of this monomer can increase the size of bacterial inhibition zones around composite samples in agar diffusion tests. Moreover, it can reduce surface bacterial growth. The antibacterial action of this monomer, however, is not considered as effective as might be achieved with a free leaching antibacterial agent as it can only affect bacteria in close proximity to the composite surface (Imazato et al., 1994, 1999; Imazato and McCabe, 1994; Imazato, 2003). Despite this limitation a product containing MDPB (Clearfil Protect Bond) has been commercialised.

Antibacterial filler-containing materials

An alternative antibacterial approach has been to introduce silver into dental composite fillers. Implantation of silver ions into conventional composite silica fillers has been shown to produce antibacterial action (Yamamoto et al., 1996). In more recent studies, colourless silver containing silica fillers more suitable for aesthetic dental composites were produced. In short-term studies, dental composites produced using these fillers were found to be highly effective against bacterial suspensions (Kawashita et al., 2000). Antibacterial silver and zinc have also been included in some calcium-phosphate fillers. When included in composite restorations these fillers were found to reduce bacterial microleakage in vitro. This action could have been due to formation of antibacterial complexes that render the resin surface less prone to bacterial adhesion or adsorption of the leached chemicals onto the bacterial surface hence reducing its adhesion tendency, or reducing the metabolism of bacteria (Syafiuddin et al., 1997). The next stage for these materials would be an in vivo animal model before human trial and commercialisation.

Chlorhexidine-releasing materials

Chlorhexidine is a highly effective, broad spectrum antimicrobial agent that also has anti-fungal activity. At low concentration, it affects the integrity of the bacterial membrane and reduces their growth, but at high concentration, it kills bacteria (McDonnell and Russell, 1999). Chlorhexidine is used in mouth washes (e.g. Corsodyl™) (Wirthlin et al., 2005) and a wide range of devices to control oral infection and recurrence of caries (Patel et al., 2001). It additionally binds to enamel and reduces bacterial adherence to teeth (Dols-Lafargue et al., 2008). Furthermore this agent can inhibit proteins that break down demineralised collagen.

Chlorhexidine can be easily incorporated into materials as different salt forms. It may also be present as a powder or fully dissolved in the fluid phase of composite pastes (Jedrychowski et al., 1983; Leung et al., 2005). The level of chlorhexidine release can be varied by changing its form and concentration. For example, in one study, the less water soluble acetate form enabled more sustained release than the highly water soluble gluconate salt (Anusavice et al., 2006). The drug-release kinetics is also, however, strongly affected by the material type.

Early chlorhexidine-containing dental composites, developed by simple addition of chlorhexidine diacetate, showed good antibacterial action but only for a few days (Takemura et al., 1983). More recently, upon addition of hydrophilic HEMA to chlorhexidine-containing hydrophobic dental composites, increased water sorption was observed and a concomitant enhanced release of drug (Leung et al., 2005). The level of early chlorhexidine release was sufficient to have much greater antibacterial action against oral bacteria within a biofilm grown on the material surface than fluoride release from GICs. Additionally, the chlorhexidine-releasing composites were highly successful at preventing bacterial microleakage in an in vitro model designed to quantify this problem. Associated with enhanced water sorption and release of chlorhexidine, however, could be a substantial decline in material strength.

In a further study (Mehdawi et al., 2009), chlorhexidine diacetate was added to dental adhesive methacrylate monomers filled with a mixture of water-soluble monocalcium phosphate monohydrate (MCPM) and less soluble β-tricalcium phosphate (β-TCP). As with hydrophilic monomers, the MCPM could increase water sorption to promote both chlorhexidine and calcium-phosphate release. The β-TCP, however, can subsequently react with remaining MCPM within the polymer and any water absorbed and then re-precipitate within the polymer as dicalcium phosphate dihydrate (DCPD also known as brushite). This provides a mechanism for limiting the reduction in resin strength that can occur upon water sorption or component release. At the present time, these systems are only experimental as further work is required to simultaneously optimise both mechanical and antibacterial properties.

Bone grafts

Bone grafting may include surgical transfer of relevant tissue from healthy parts of a patient (autogenous grafts) or from a donor of same species (allografts) but material supply is a major problem. Grafts may additionally be from another species (xenografts, e.g. hydroxyapatite extracted from bovine bone) or synthetic (alloplasts). Synthetic hydroxyapatite and tricalcium phosphate (TCP) particles have both been used to fill bone defects (Komlev et al., 2002). Bone may re-grow in the spaces between particles as they are remodelled but if greater mechanical support is required then calcium-phosphate cements (see below) may be more suitable.

Guided tissue-regeneration (GTR) membranes

A guided tissue-regeneration membrane acts as a barrier preventing fast-growing soft tissue from invading space required to be filled with new bone. Degradable polymeric (e.g. polylactide or glycolide) or natural (e.g. collagen) membranes are often used. Many of these membranes have been formulated with antibiotics, the most common of which have been from the tetracycline family (Zarkesh et al., 1999). In addition to its antibiotic action, tetracycline reduces in vivo degradation of collagen within bone. Antibacterial chlorhexidine (Chen et al., 2003) and zinc (Giertsen et al., 1989; Chou et al., 2007), however, have also been included. Generally, drug-release kinetics is dependent on drug concentration and membrane porosity and hydrophilicity. Provided drug-release kinetics is well controlled, antibacterial agents can have significant beneficial clinical effects.

In various studies drug has been incorporated into GTR membranes by simple dropping of drug solution onto the material or by membrane immersion in the drug solution for a controlled period of time. The first method has been shown to provide good impregnation of GTR membranes with tetracyclines (Hung et al., 2005). Upon placement of the membrane back into water, however, most of the drug was released within 24h. To improve both the loading efficiency and provide more sustained drug release, various authors have modified the surfaces of GTR membranes for example with carbohydrate derivatives (Lepretre et al., 2007; Tabary et al., 2007), tridode-cylmethylammonium (Hung et al., 2005), or with a dense polymeric layer (Lee et al., 2008). Significant drug release for 80 days has been made possible in some cases through such modifications (Lepretre et al., 2007).

Of concern is that antibacterial drugs can affect other cell growth in addition to the bacteria. In one study, for example, viability of periodontal ligament (PDL) cells grown on tissue culture plastic and treated with various chlorhexidine concentrations (0–50 µg/mL of cell growth media) was observed to decline by 50% when using 15 µg/mL of chlorhexidine (Chen et al., 2003). Although cells grown on tissue culture plastic may be less protected and therefore more sensitive to drugs than those in vivo, it should be appreciated that 15 µg/mL of chlorhexidine are orders of magnitude lower than currently used in mouthwashes. In order to understand the full implication of this result for chlorhexidine-releasing devices in vivo, fluid flow around any specific site must be fully appreciated so that length of time of tissue exposure to the drug can additionally be estimated.

It should be noted that bone can only grow if provided with space to do so. On the other hand, thin polymeric membranes, beyond holding back soft tissue, provide no significant mechanical support during bone healing. For greater mechanical strength, a polymeric adhesive or calcium-phosphate cement is required. In this case bone can only replace the material if/when they degrade. Gaining bone cement degradation at a rate commensurate with possible bone repair is a significant challenge.

Poly(methyl) methacrylate bone cements

Poly(methyl) methacrylates (PMMAs) were first introduced as bone cements in the 1960s by Charnley and Smith (Dalby et al., 2002). PMMA cements consist of a resin matrix (mostly methyl methacrylate monomer), contrast agent (often barium sulfate) and initiator/activator (typically benzoyl peroxide/dimethylparatoluidine) system. Upon mixing PMMA components and injection into the defect, they set chemically via free radical initiated polymerisation. This produces long entangled polymer chains and a material with high-strength characteristics. PMMAs mainly depend on micromechanical not chemical interaction with the surrounding bone for adhesion (Clarkin et al., 2010).

Modern commercially available PMMA formulations are usually loaded with antibiotics, commonly gentamicin (a broad spectrum antibiotic that can withstand the high temperature produced during PMMA setting). Drug loading is typically 1.25–2.5 wt% to prevent infection (Ensing et al., 2008) and 2.5–10 wt% to treat an existing infection (Dunne et al., 2007). Gentamicin release, however, is generally not high. A maximum of 15% of the total amount of the drug may be released within the first few days. The release can be enhanced by increasing the actual amount of the drug (Dunne et al., 2007), or adding a release modulator such as lactose (Virtoa et al., 2003). If the drug release is too fast, however, the antibiotic may be washed away from the wound site and at effective concentrations for too short a period of time (Wroblewski et al., 1986; Dunne et al., 2008).

A major problem with PMMA cements is polymerisation shrinkage during set. This causes stress on the bond between PMMA and surrounding bone and may lead to gap generation at their interface (as observed with dental composites bonded to dentine). Tissue fluid, enzyme penetration and ultimately ‘bacterial microleakage’ at this interface may subsequently weaken any adhesive bond and break down collagen. Additionally, the mismatch in rigidity of PMMA and bone may also lead to bond deterioration upon loading (Clarkin et al., 2010). Often PMMA undergoes wear that may elicit long-term inflammatory response upon formation of particulates (Peter et al., 1999). The use of PMMA is also complicated by the possibility of both thermal and chemical necrosis of the surrounding tissues and nerve ending as a result of the release of heat and residual monomer respectively (Donkerwolcke et al., 1998). Furthermore, PMMA cements are non-degradable and cannot remodel or integrate into the surrounding tissues (Lieberman et al., 2005).

Calcium-phosphate cements

The first calcium-phosphate cements (CPCs) were developed in the late 1980s to help in remineralisation of tooth structure destroyed by dental caries (Ishikawa et al., 1995). These consisted of equimolar tetracalcium phosphate [TTCP, Ca₄(PO₄)₂O] and dicalcium phosphate anhydrous [DCPA, CaHPO₄] or dicalcium phosphate dihydrate [DCPD, CaHPO₄ 2H₂O]. They were subsequently, in 1996, FDA approved for cranial defect repair (David et al., 2004). Nowadays, there are several commercially available formulations employed for maxillofacial and orthopaedic applications (Cuisinier et al., 2004, Tamimi et al., 2009). Generally, these formulations consist of a powder and liquid phase. The powder usually contains two or more calcium-phosphate compounds {tetracalcium phosphate [TTCP, Ca₄(PO₄)₂O]/dicalcium phosphate [DCP, CaHPO4] or monocalcium phosphate [MCP, Ca(H2PO₄)₂]/tricalcium phosphate [TCP), Ca₃(PO₄)₂]/calcium carbonate [CaCO₃] or amorphous calcium phosphate [ACP, Ca_xH_y(PO₄)_6−x nH₂O]/MCP/DCP}, while the liquid can be water, saline or sodium-phosphate solution.

Upon mixing the powder and liquid, a paste that can be easily injected and set in situ is formed. The setting reaction is a dissolution–precipitation reaction, that is, dissolution of soluble calcium phosphates from the powder and then precipitation of less soluble forms. Dependent upon the initial composition and pH, the final precipitated forms can be brushite [DCPD, CaHPO₄2H₂O] or apatite {hydroxyapatite [HA, Ca₁₀(PO₄)₆OH₂] or carbonated apatite [CA, Ca₁₀(PO₄)₆OH₂CO₃] (Gerhart et al., 1988; Miyamoto et al., 1995). For instance, TCP tends to form brushite under acidic conditions, but hydroxyapatite at neutral pH. Brushite cements can be faster setting and more resorbable than hydroxyapatite forming formulations but also of lower compressive strength.

Although, CPCs have been proven to be successful for bone repair, they have no antibacterial action. They are therefore usually loaded with antibiotics (Stallmann et al., 2006; Tamimi et al., 2008; Hofmann et al., 2009), particularly gentamicin sulfate. Antibiotic addition has been proven effective in reducing the number of Staphylococcus aureus and methicillin-resistant Staphylococcus aureus (MRSA) in experimental chronic osteomyelitis models (Stallmann et al., 2004; Joosten et al., 2005) and clinically been used to both prevent and treat infection.

Gentamicin release kinetics is commonly diffusion controlled, that is, cumulative release is proportional to the square root of time. From diffusion theory, release percentage from flat discs is expected additionally to be inversely proportional to specimen thickness. For samples of the order 1 mm thick, total release is typically a few days. Gentamicin release is also dependent on the cement type (brushite or hydroxyapatite), porosity and drug form (e.g. Bohner et al., 1997).

The relatively fast diffusion-controlled drug release from CPCs compared with that from polymers such as PMMA is primarily due to the high cement porosity. Porosity can be reduced and cement mechanical properties increased by reduction in the cement liquid component but only to a limited extent. Various authors have therefore attempted to add polymers into the aqueous phase to bind drugs and slow their release. For example, it has been proven that polyacrylic acid can hinder release of positively charged antibacterial agents such as gentamicin and chlorhexi-dine presumably due to its dissociation in water and production of COO⁻ groups (Bohner et al., 1997, 2000b; Young, unpublished data). Polyacrylic acid has also the potential to enhance the durability of the adhesive bond between the cement and the surrounding bone due to its known ability to chelate with calcium in hydroxyapatite. In a further study, prior mixing of gentamicin with polylactide powder enabled effective gentamicin release from HA cement over a period of weeks instead of days (Sasaki and Ishii, 1999).

Upon incorporation of some antibiotics into CPCs, the handling characteristics (consistency, setting time and injectability) and mechanical properties can both be affected. Tetracylcine, for example, has tended to increase the setting time and reduce the compressive strength of the apatite forming CPCs (Ratier et al., 2001) but gentamicin sulfate can have the opposite effect on brushite forming cements (Bohner et al., 1997; Joosten et al., 2005). Major problems with CPCs, however, include sensitivity of the setting process to the presence of blood and tissue fluid, difficulty in degradation rate control and low mechanical strength. Their application is therefore limited to small defects and non-weight bearing situations.

Injectable degradable polymers

The following will focus upon injectable poly(anhydrides), poly(ester) dimethacrylates and poly(propylene) fumarates. With these types of polymer, degradation rates can be more readily controlled than with CPCs. The polymer degradation products can be eliminated safely in the body but might be considered less biocompatible than calcium phosphate as they are not required for tissue repair. The anhydrides and poly(ester) dimethacrylates are relatively small fluid molecules with methacrylate groups on both ends. They can be rapidly cross-linked at body temperature using either light or chemical cure means. Cross-linking of poly(propylene) fumarate, however, is achieved through the use of dimethacrylate diluents. Cross-linking/polymerisation reactions provide a micromechanical means for the polymer to bond rapidly to roughened structure such as bone. When drugs are incorporated into these polymers, they can be slowly released over several weeks upon polymer degradation. In some cases water sorption may enhance drug release. Water sorption and acid production in the bulk of materials may additionally, however, cause the polymer to degrade catastrophically instead of via a surface erosion mechanism.

Poly (anhydrides)

Poly(anhydrides) are generally surface eroding polymers with a hydrophobic backbone and hydrolytic anhydride linkage (Muggli et al., 1999). Surface erosion is potentially advantageous for drug-delivery applications as then both polymer loss and drug release may be linear with time. This assumes diffusion-controlled drug release to be much slower than the polymer degradation. With diffusion control the level of drug release per second declines with time. More seriously, however, with polymers undergoing bulk degradation a sudden dose dumping effect may be observed. The anhydride links within the cross-linked polymer can degrade rapidly in the presence of water potentially providing materials for short-term drug-release applications (Gopferich and Tessmar, 2002). The degradation time of these polymers may, however, be increased from days to years by increasing their hydrophobicity.

Photo-polymerisable polyanhydrides prepared from monomers of sebacic acid (SA) alone, or copolymers of SA and 1,3-bis(p-carboxyphenoxy) propane (CPP), or 1,6-bis(p-carboxyphenoxy) hexane (CPH) have been formulated for bone repair (Chiu et al., 2002). These polymers have also had Food and Drug Administration (FDA) approval as delivery devices of antibiotic for the treatment of chronic bone infections (Septacin™ implant) (Jain et al., 2005). More recently, an injectable degradable poly(sebacic-co-ricinoleic-ester-anhydride) loaded with gentamicin (10–20 wt%) was also developed. This was shown to provide constant drug release (up to 20–80% depending on the molecular weight of the polymer and the drug concentration) over 28 days and to be effective when injected subcutaneously (Krasko et al., 2007) or in an experimental bone model (Brin et al., 2008).

A potential problem associated with the high reactivity of the anhydride group, however, may be difficulties in synthesis and purification.

Poly (propylene fumarates) (PPFs)

Injectable, biodegradable, poly(propylene fumarate) (PPF) formulations have been produced using unsaturated polyesters synthesised by reaction between fumaric acid and propylene glycol. The repeating units in PPF contain one double bond that can be cross-linked by any other C = C containing monomer. The PPF units also contain two ester groups that enable subsequent hydrolytic degradation of cross-linked polymers into fumaric acid and propylene glycol (He et al., 2001). PPF formulations largely undergo bulk degradation; the degradation time is dependent on polymer structure but has generally been of the order of several months (Gunatillake and Adhikari, 2003).

PPF has been extensively studied as potential bone cement (Peter et al., 1999; Jayabalan et al., 2001; Choll et al., 2007). One problem with these polymers, however, may be their low modulus. Although β-TCP addition can increase the material modulus there is concern that other mechanical properties may also be too low and with some formulations in vivo degradation was too fast for many orthopaedic applications (Mano et al., 2004; Ginebra et al., 2006). A possible partial explanation for low mechanical properties and fast degradation could be steric hindrance around the double bond in the fumarate systems limiting polymerisability.

Poly (ester) methacrylates

With methacrylates in non-degradable PMMA bone cements incomplete monomer conversion is a common observation. Greater and faster polymerisation of injectable triblock poly(propylene glycol-co-lactide) dimethacrylate monomers at body temperature, however, has been observed. In one study, for example, complete monomer conversion was achieved with some formulations using light cure initiators and less than 60s of light exposure. Upon decreasing molecular weight of the monomer, polymerisation rate was raised. The complete polymerisation is made possible by the high flexibility of the final cross-linked polymer chains. Whereas high conversion in dental composites and PMMA bone cement causes significant polymerisation shrinkage and heat generation, this is not the case with the studied formulations due to their high monomer molecular weight and therefore low reactive group (C=C) density.

The degradation rate of the cross-linked poly(propylene glycol-co-lactide) dimethacrylates can increase substantially upon reducing the length of the propylene glycol blocks (from months to weeks) providing a means to match degradation with bone re-growth (typically four–six months) (Ho and Young, 2006). Of additional importance, however, is the mechanism of polymer degradation (surface or water sorption and acid catalysed bulk). By reducing the molecular weight of these monomers there is a concomitant increase in cross-linking density of the resultant polymers and reduction in the probability of high water sorption and bulk sudden potentially catastrophic degradation.

When chlorhexidine diacetate particles (2.5–10wt %) were incorporated into the above polymers, the release was primarily via a diffusion mechanism. At ten weeks the level released from 2-mm thick discs, varied between 15% and 82%, being greater with lower polymer cross-link density and higher initial drug concentration or particle size. The results were consistent with negatively charged degradation products within the polymer associating with the positively charged chlorhexidine delaying both polymer erosion and drug release (Young and Ho, 2008).

The modulus (rigidity) of the above cross-linked polymers was proportional to the inverse monomer molecular weight, but as a result of the high flexibility of poly(propylene glycol), they have much lower modulus than bone. A further problem was the acidic products associated with the polymer degradation. The section below will demonstrate how these problems might be solved.

Injectable degradable composites

Recently, it has been shown that using an equimolar mixture of monocalcium phosphate (MCPM)/β-tricalcium phosphate (β-TCP) as filler for poly(propylene glycol-co-lactide) dimethacrylate and upon placement in water, the modulus of the fully polymerised composite increased 10-fold in 24h. After 1–10 weeks in water, the composite modulus was between 40 and 100 times that of the equivalent unfilled polymer. This increase in modulus was attributed to water sorption catalysed filler conversion to finer dispersed brushite crystals. Furthermore, the polymer degradation rate increased upon addition of this reactive filler mixture. Upon composite degradation both organic and inorganic components were released simultaneously. The acidic organic degradation products, however, were neutralised by the inorganic ions thereby reducing the possibility of adverse in vivo reactions as well as bulk catalysed polymer degradation (Young et al., 2009).

When loaded with 10 wt% chlorhexidine diacetate, these composites released 50–80% of drug via diffusion over a period of 10 weeks. Early release levels were proven sufficient to inhibit the growth of both S. aureus and MRSA in vitro. The level of release could also be tailored by adjusting the mass fraction of calcium-phosphate fillers; the higher the filler fraction, the faster the release (Young, unpublished data). Further work is now required to optimise these formulations for specific bone-repair applications.