Porous bioactive glass scaffolds have been of interest for biomaterial design for some time. A microstructure that can stimulate the healing process and promote tissue growth while being made from a material that will eventually degrade is attractive. Unfortunately, highly porous materials inherently have low strength which can be a major detriment. Another major hurdle to be overcome is the poor handleability of porous scaffolds as they easily break when handled, can easily get tangled together, and ultimately become ineffective for the intended purpose. The majority of the orthopedic industry uses dense particles of cadaver bone, calcium phosphate based ceramics, or bioactive glass in their devices as they have the required strength and handleability required during surgery.
The use of silicate based bioactive glasses such as 45S5 and S53P4, are currently used in products approved by the US Food and Drug Administration and other world safety organizations for use in implantable devices such as orthopedic implants and are known to be capable of making three dimensional porous scaffolds. Glasses with compositions similar to 45S5 and S53P4 crystallize rapidly when heated above each glasses respective glass transition temperature (Tg), making viscous sintering by traditional heat treatments difficult without crystallization. Therefore, currently there are no amorphous, rigid, and porous scaffolds composed of silicate glasses are commercially available. Glasses with wider working ranges with multiple alkali and alkaline earth elements in combination with additional silica have been developed to fill the need, but these glasses convert to hydroxyl apatite (HA) slower than 45S5 and S53P4, and none of these scaffolds are in currently available products approved by the FDA for any clinical market.
Silicate glass particles that are larger than >500 μm can take years to fully react with body fluids and be converted to the inorganic component of bone known as hydroxyapatite (HA) because, large particles of glass (>500 μm) leave voids in healing tissue that take years to remodel into natural tissue, and the large particles have a relatively small surface area mass ratio and don't allow for tissue penetration by bone or blood vessels.
Porous scaffolds allow for tissue penetration, and the surface area available for reaction with body fluids is relatively large and penetrates throughout the entire scaffold, which significantly reduces conversion time to HA and final remodeling. The conversion kinetics of silicate glass 45S5 to HA has been shown to slow from the contracting volume model to the diffusion model once the silica gel layer achieves a thickness great enough to become the diffusion barrier. This is the reason why a porous scaffold composed of 45S5 will convert to HA in a matter of weeks as opposed to a solid glass 45S5 bead of comparable size that could take years.
The crystallization properties and glass transition properties of a particular glass are important when treating glass and making it into a porous scaffold. Glass compositions composed of relatively low concentrations of glass forming oxides and relatively high concentrations of alkaline and alkaline earth oxides tend to crystallize rapidly when heated above the glass transition temperature, making bonding by viscous flow difficult. Therefore the ability to make porous materials or scaffolds rigid from these glasses by traditional thermal treatments has not been possible.
Typically, glass is bonded by heating above the glass transition temperature to a viscosity appropriate for viscous flow. Depending on the amount of time allowable for the process and the amount of flow required, the viscosity used for processing can vary with the application. Glasses used for such applications are designed to resist crystallization at the processing temperatures because the mobility of ions is high. The more fluid the glass, typically the easier it may be for a glass to crystallize, but this is dependent on how close the glass composition is to a crystalline phase and how great the activation energy is to allow the glass to crystallize.
The glasses that will benefit from this method of bonding require relatively low energy input to allow the atoms to rearrange and start forming crystals. The formation of crystals, especially at the particle surface, is what inhibits the viscous flow of these particles. Crystalline phases that crystallize from glass typically melt at significantly higher temperatures than the temperatures required for crystallization; therefore the crystals do not form a viscous flow that would aid in the sintering process. In addition, the tendency towards crystallization increases as the surface area to volume ratio of the glass component increases, so the smaller the component, the greater the tendency toward surface crystallization and inhibited bonding.
Scaffolds, particularly those designed for use as bone grafts, should be highly porous (>50%) and are often formed by infusing a slurry composed of glass particles and other organic and inorganic components into a preform (foam or sponge or other porous polymer) that must be slowly burned out prior to sintering. To keep the desired microstructure of the preform, the heating rate is typically kept low, a few ° C./min, to the sintering temperature, and then the sintered part is slow cooled as to eliminate thermal shock of the glass/ceramic scaffold. For glasses that crystallize quickly (45S5 and S53P4), these methods are not effective in making rigid glass scaffolds.
The graph shown in FIG. 1 describes the usable regions of interest for the possible heat treatments of the following invention. Sintering bioactive glass is a time-temperature-crystallization dependent process, and no single component can be ignored.
Typically, scaffolds discussed in the literature are rough and sharp, which is not an issue for bench scale testing, but can be a significant problem for clinicians such as orthopedic surgeons that want to press on the implant material without puncturing a glove. A punctured glove opens the clinician and the patient up to possible disease transmission, and the clinician could be injured by the implant material if it were to penetrate the clinician's skin. Therefore, no fully amorphous scaffolds composed of silicate based 45S5 or S53P4 bioactive glass are currently approved for by the FDA for clinical use.
The surface roughness of a scaffold is certainly a disadvantage from a handling point of view, but in a product such as bone putty or even just a single phase implant material such as loose granules, the rough edges make each scaffold catch on one another decreasing the flowability of the particles. This decrease in flowability decreases the overall scaffold loading the putty can incorporate as the putty itself acts as a lubricant to improve the handleability. The debris from broken edges may also increase the overall immune response as macrophages try to remove/engulf the small particles.
Therefore a need exists for a method of bonding of silicate based bioactive glass with a high affinity for crystallization.