The need to treat bone defects resulting from age-related degenerative disease, trauma, and reconstructive surgery continues to grow at a significant rate. An estimated 2.2 million bone grafting procedures are performed annually worldwide at a cost of $2.5 billion. Considering the “graying” of the U.S. population, in which the percentage of persons 55 years and older is expected to nearly double over the next 30 years, tremendous societal impact is inevitable. Consequently, the market for orthopedic biomaterials will continue to expand.
Autologous (i.e., autografts) and cadaveric (i.e., allografts) bone are the most commonly used grafting materials to treat bone defects. Each, however, has drawbacks related to morbidity associated with a second surgical procedure for autografts and the potential for disease transmission with allografts. These observations have led to development of synthetic materials and tissue engineering approaches for use in bone regeneration. Bone graft substitutes have used materials of natural and synthetic origin. Ongoing developments focus on enhancing biological activity, such as by incorporating stem cells (e.g., mesenchymal stem cells) and growth factors (e.g., bone morphogenetic protein (BMP) 2) into bone graft substitute materials.
Synthetic biodegradable polymers are also used for drug delivery and to aid in tissue regeneration. Candidate materials include polyanhydrides, polyamides, polycarbonates, and polyorthoesters. The majority of resorbable synthetic polymers utilized for drug delivery and tissue engineering belong to the polyester family, such as poly(glycolic acid), poly(lactic acid), and poly(lactic-co-glycolic acid) (PLGA). These materials are relatively biocompatible, can degrade by the hydrolytic cleavage of ester bonds, and have degradation and mechanical properties that can be tailored by changing monomer ratio. Their degradation products of glycolic and lactic acid are metabolically removed from the body by conversion to carbon dioxide and water in the Krebs cycle. Other common polyesters include poly(ε-caprolactone), polyvalerolactone, polydioxanone, and their blends and copolymers.
Furthermore, although certain drugs are intended for systemic therapy, many are most effective if targeted to or placed within a specific site. To this end, drugs are routinely encapsulated in polymers. Entrapment in a solid matrix protects the molecules from environmental effects, and controlled release can be achieved. Persistent challenges, however, include instability of encapsulated drugs, incomplete release, and initial burst.
Therefore, since cells and tissues require exposure to bioactive agents at particular concentrations and doses for certain durations, limited release kinetics is a shortcoming of many drug delivery systems in regenerative medicine. Consequently, some have attempted to develop drugs conjugated to polymers to extend release duration. Among others, water soluble polymers, such as poly(ethylene glycol), polylysine, polyglutamic acid, and N-(2-hydroxypropyl)methacrylamide (HPMA) have been used for this purpose. With these systems, drugs are attached as pendants linked to the polymeric backbone via ester, amide, and hydrazone bonds. Depending on the spacer molecule chosen, drug release can be prolonged until cleavage in a desired environment, such as pH-sensitive release in a lysosome. Other applications include targeting to specific cells and prolonging circulation time by shielding the drug from degradative enzymes and preventing opsonization. The number of molecules (i.e., payload) that can be attached to the backbone, however, is limited by the number of functional groups required for conjugation.
Hence, there remains a need for degradable compositions and compounds for treating tissue wounds, including bone tissue wounds, that can release drugs at a wound site in a controlled manner. There also remains a need for such compositions and compounds whose payload is not limited by the number of functional groups present on a polymer backbone.