Bioabsorbable polymer scaffolds for tissue engineering capable of sustained growth factor delivery
Introduction
The loss and failure of tissues and organs necessitates millions of surgical procedures and costs the economy in excess of $400 billion per year, in the United States alone [1]. A significant shortage of tissues and organs suitable for transplantation seriously limits the number of patients suffering from these problems who can be treated with these therapies, and this problem continues to worsen each year [2]. Tissue engineering has been proposed as a therapeutic approach to create new tissues and treat patients suffering from the loss or failure of organs and tissues [1]. A number of tissue engineering strategies have been developed, and many involve the transplantation of cells on or within polymeric matrices [1], [3]. The success of these approaches is dependent on mass transport between the engineered and surrounding host tissues being sufficient to meet the metabolic requirements of the engineered tissue. Diffusion is sufficient for this purpose when small numbers of cells (e.g., islet or neural cells) are transplanted on or within thin (thickness typically less than 0.5 mm) polymer devices [4], or if the metabolic needs of the transplanted cells are very low (e.g., chondrocytes) [5]. However, these requirements are not met for the majority of tissues one may wish to engineer (e.g., bone, liver).
A rapid and high level of vascularization is required for the survival of most cell types following transplantation. Oxygen is typically the limiting factor for the survival of transplanted cells, and simple calculations of oxygen diffusion in tissues indicates that cells cannot be more than several hundred micrometers from a capillary if they are to survive [4]. Studies with hepatocyte transplantation confirm this prediction, as the majority of cells transplanted on 3 mm thick polymer matrices died within a short time [6]. The surviving cells were largely found at the edges of the matrices, near capillaries in the surrounding host tissue. Capillary ingrowth is often noted in the cell-polymer implants over time [7], [8]. However, the vascularization is either too slow or limited to provide sufficient nutrient transport for the majority of transplanted cells. Histological studies with a variety of other cell types (e.g., bone) also confirm this prediction [9]. Quantification of the mechanical properties of new tissues formed by transplanted smooth muscle cells reveal the functional impact this limitation has on the engineered tissues. The outer layer of the engineered tissue exhibited the elasticity desired from smooth muscle tissue, while the inner core of the tissue had mechanical properties similar to that of the scar tissue formed by implantation of the polymer matrix alone [10].
We hypothesize that vascularization of engineered tissues can be enhanced by the local, controlled delivery of molecules which induce capillary formation. A variety of growth factors have been identified that promote angiogenesis both in vitro and in vivo assays, and recombinant versions of many of these proteins are now available [11]. These substances act at various steps of the angiogenic process, which involves activation of the endothelial cells in capillaries to digest their basement membrane, migrate, proliferate, and form new capillary branches. A number of factors have been identified, and most act on a number of cells types in addition to endothelial cells [11]. We are focusing on the delivery of vascular endothelial growth factor (VEGF), as it is highly proangiogenic. In addition, VEGF is the only factor known to act solely on endothelial cells [11]. This may be a critical issue, as the goal in tissue engineering is to increase the net mass transport into the forming tissues. Stimulation of cell proliferation with non-endothelial cell specific factors may lead to an increase in fibroblast proliferation and no net increase in transport to the cells of interest in the engineered tissue.
We have previously investigated the delivery of VEGF from alginate beads and poly (lactide-co-glycolide) (PLG) microspheres. These alginate beads or PLG microspheres could be transplanted, as an auxiliary component, along with the cells in the three dimensional polymer matrix utilized as a scaffold to enhance the vascularization of engineered tissues [12]. PLG are attractive polymers for this application, as they degrade by hydrolysis to lactic and glycolic acid [13]. These polymers have been used for over twenty years in a variety of medical devices, including sutures [13]. PLG microspheres can be readily formed using standard double emulsion techniques, and incorporated growth factors can be released at constant rates for extended times from these polymers [12], [14], [15], [16]. Alginate, a naturally derived polysaccharide typically purified from seaweed, is a biocompatible polymer widely utilized in the food and chemical industry. In addition, it has more recently been developed as a cell transplantation material for a variety of cell types, including chondrocytes and islets [17], [18]. Alginate, when ionically cross-linked with divalent cations (e.g., Ca+2), readily forms hydrogels at low concentrations (e.g., 1–2%). VEGF, and other growth factors, can be readily incorporated into alginate beads, and subsequently released for time periods up to several weeks [19], [20], [21], [22]. Importantly, VEGF released from alginate beads not only maintains its ability to stimulate endothelial cell growth, but the activity of the released VEGF is greatly enhanced over control VEGF which is added directly to cultured cells [22].
Ideally, one would like to deliver angiogenic factors directly from the polymer matrices which are utilized to both transplant cells and provide a scaffold for new tissue formation. However, the processing techniques typically utilized to form these matrices typically involve high temperatures or organic solvents [23], [24]. Both conditions would be expected to denature proteins present during the process. We have recently developed a novel approach to fabricate highly porous structural matrices from PLG, and this process avoids the use of high temperatures and organic solvents [25], [26]. In this method, polymer particles are combined with NaCl particles and compression molded into the desired shape. The polymer/NaCl is then exposed to high pressure carbon dioxide gas. After equilibration at high pressure, the gas pressure is returned to ambient conditions. This results in a thermodynamic instability which causes the polymer and gas to phase separate, creating gas pores in the matrix. The polymer particles expand as pores form, and fuse to form a continuous structure. The NaCl particles are subsequently removed by leaching in distilled water to create an interconnected pore structure. This approach may be ideal for the incorporation and delivery of angiogenic factors due to the gentle conditions utilized in the fabrication process.
Matrices utilized for simultaneous cell transplantation and growth factor delivery must exhibit appropriate pore structure, mechanical properties, and growth factor release kinetics. The delivery of growth factors from PLG is typically coupled to the degradation rate of the polymer [27], and both the mechanical properties and degradation rate of PLG are typically regulated by the ratio of lactide to glycolide and degree of polymerization [13], [28]. In this study, the affect of various gas foaming processing parameters (gas type, polymer type and molecular weight, equilibration time) on the formation of porous PLG matrices were investigated, and preliminary studies which demonstrate sustained delivery of functional angiogenic factors from these matrices were performed.
Section snippets
Materials and methods
Pellets of poly l-lactic acid [PLLA], a 50:50 copolymer of d,l-lactide and glycolide (50:50 PLG) with intrinsic viscosity (i.v. of 0.2 dL/g), a 75:25 PLG copolymer (i.v.=0.2, 0.7), and an 85:15 PLG copolymer (i.v.=1.4) were obtained from Boehringer Ingelheim (Henley, Montvale, NJ, USA). PGA, 50:50 PLG (i.v.=.8) and 85:15 PLG (i.v.=.63) were purchased from Medisorb (Cincinnati, OH, USA). 85:15 PLG (i.v.=3.63) was obtained from Purasorb (Lincolnshire, IL, USA). High mannuronic acid sodium
Foaming solid polymer disks
The effects of a number of variables on the porosity and mechanical properties of matrices formed with the gas foaming process were studied to determine which conditions allowed formation of suitable matrices. In the first series of experiments, solid polymer disks (no NaCl) were foamed to investigate the role of the gas type, pressure release rate, and polymer composition and molecular weight on the expansion of solid polymer particles subjected to the foaming process. In the first experiment,
Discussion
To promote the vascularization of tissues engineered from transplanted cells via localized angiogenic factor delivery it will be necessary to develop mechanically robust matrices in which growth factors can be incorporated in a stable manner, and released in a controlled manner with desirable kinetics. A gas foaming process has now been developed to fabricate mechanically stable three-dimensional tissue engineering matrices using mild processing conditions, and the sustained delivery of
Acknowledgements
This research was funded by the NIH (R29 DK 50715-04) an NIDR summer research fellowship (T35-DE07101) to MHS, an NIDR training grant (DE07057) to LDS, a Whitaker Foundation graduate fellowship (MP), and by Reprogenesis.
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Note that M.H.S. and L.D.S. contributed equally to this publication and should be considered co-first authors.