Yadong Wang, 1, Steven Lu, 2, Peter Gabriele, 2, Jeremy J. Harris, 2
1Department of Bioengineering, University of Pittsburgh, Pittsburgh, PA 15261 USA, 2Research and Development, The Secant Group, LLC, Telford, PA 18969 USA
Material Matters, 2016, 11.3
The world of commercial biomaterials has stagnated over the past 30 years as few materials have successfully transitioned from the bench to clinical use. Synthetic aliphatic polyesters have continued to dominate the field of resorbable biomaterials due to their long history and track record of approval with the U.S. Food and Drug Administration (FDA). Despite a plethora of research to develop biocompatible, biodegradable polymers, new biomaterials have suffered from compliance mismatch, i.e., the failure to successfully mimic the mechanical properties of natural tissue. To address these concerns, poly(glycerol sebacate) (PGS) was developed in the lab of Professor Robert Langer as a tough, biodegradable elastomer.1 Since this discovery, the biomedical engineering community has used PGS in a multitude of implantable applications in the areas of cardiovascular, neurovascular, orthopedic, and soft tissue.2
PGS is a simple glycerol-ester polymer created from the basic mammalian metabolites of glycerol and sebacic acid, both of which have a regulatory background with the FDA.1 Originally designed as a biodegradable polymer with improved elastic mechanical properties and biocompatibility, research on PGS-based medical applications has uncovered a number of unique properties that have bolstered its utility as a biomaterial. In addition to its elasticity, PGS demonstrates minimal swelling, undergoes surface degradation and exhibits mild acute and chronic inflammatory responses in vivo. Although the majority of researchers use the thermoset elastomer form of PGS, the polymer is customizable through a continuum of resin forms. Depending on the degree of polymerization, PGS can be manufactured as a soft gel, a lubricious Vaseline®-like paste, a thermoplastic, or a thermoset depending on the applications. Through manipulation of the various morphologies, PGS can also be formulated as a coating for a range of medical implants and extruded into luminal structures, sheets, rods, and other geometric shapes. Furthermore, the polymer is compatible with a host of biologic materials such as collagen, bone minerals and extracellular matrix (ECM)-like compositions, further making it an ideal bioresorbable material for tissue engineering, regenerative medicine applications, and the biomedical device industry.
Current interest in bioelastomers for tissue engineering is largely driven by the need for soft-tissue repair. PGS was created specifically for engineering soft tissues in dynamic mechanical environments such as the cardiovascular system.1 Made from glycerol and sebacic acid via polycondensation, the ester bonds in the polymer backbone covalently crosslink to form a three-dimensional (3D) network of random coils, which resembles the structure of vulcanized rubber and gives rise to rubber-like elasticity (Figure 1). Hydrogen bonding interactions between hydroxyl groups further enhance the mechanical properties of PGS. Due to the ester linkages, both the polymer backbone and crosslinks can undergo hydrolytic degradation.
Figure 1.Reaction schematic for poly(glycerol sebacate). PGS is synthesized from glycerol and sebacic acid and further curing under heat and vacuum to produce a crosslinked PGS thermoset.
PGS is synthesized via a polycondensation reaction between glycerol and sebacic acid to first form a pre-polymer resin which is then converted into the thermoset elastomer. In particular, both starting materials are inexpensive and can be obtained from renewable resources; sebacic acid for example, is derived from castor oil.3 The synthesis of PGS also utilizes an environmentally-friendly chemistry without the need for toxic solvents or catalysts, resulting in a synthetic biomaterial through an overall sustainable synthesis. Several synthesis routes are reported in the literature but the most commonly used two-step method is described here. The reactor is first charged with the monomers and heated to 120 °C under a blanket of N2. Upon forming a homogenous solution, the mixture is further heated for 24 hours. The reactor is then placed under vacuum (40 mTorr–10 Torr) for an additional 24–48 hours depending on the desired degree of polymerization. At this point, the resin or pre-polymer is complete and ready for the second step—to create the thermoset elastomer. The resin can be used neat as a casted film, molded to a specific shape, or reduced in a solvent for casting or dip-coating applications. The resin is cured for 24–96 hours depending on the desired mechanical properties required for the elastomer.
One advantage of PGS is the ability to tune its mechanical properties by making small changes to the polymerization and curing procedures. Elastomer modulus values are directly related to the degree of crosslinking and range from 0.77–1.9 MPa for 48 and 96 hour cure times, respectively. A wider window of modulus values is achieved by manipulating the monomer stoichiometry where modulus values can be tuned from 0.01–5 MPa. Altering monomer stoichiometry also allows fine tuning of both molecular weight and free chemical functionality. Average molecular weight values range from 2,000 to >200,000 Da and can be tuned simply by changing the glycerol:sebacic acid ratio. Chemical functionality (as measured by acid number titration) ranges from 110–10 mg/g, resulting in PGS with varying hydrophilicity and reactivity of the elastomer.
PGS degrades primarily through hydrolysis of the ester linkage into smaller oligomers and ultimately to the starting monomers, glycerol and sebacic acid. PGS degradation is unique and differs from other resorbable polymers (e.g., polylactide, polyglycolide, and copolymers) in that PGS degrades via surface erosion as opposed to bulk erosion.4 The significance of this is demonstrated in a linear degradation profile over time with a controlled loss of mechanical properties, in contrast to bulk erosion where mechanical properties exhibit a catastrophic decrease. Many studies have evaluated the in vitro degradation hoping to model in vivo performance. However, there exists a poor correlation between in vitro and in vivo degradation behavior. In vitro studies performed under a range of conditions typically show upwards of a 20% mass loss at 30 days compared to a 70% mass loss observed in subcutaneous tissues.4b Despite the accelerated degradation kinetics of PGS in vivo relative to in vitro model conditions, the degradation rate can be tuned by modulating crosslinking density via cure time and temperature.5
The hydrolytic degradation of PGS into its component monomers, glycerol and sebacic acid, provides a resorbable material with high biocompatibility. Glycerol is a metabolic building block for lipids and has a long history of use in pharmaceuticals. Sebacic acid is the natural metabolite intermediate in ω-oxidation of medium and long-chain fatty acids. Further, co-polymers containing sebacic acid are used in chemotherapeutic drug delivery.6 Various studies have evaluated the biocompatibility of PGS with both in vitro assays and in vivo implantation studies. PGS has shown to be non-cytotoxic in vitro1 and induces a minimal inflammatory response with little fibrous capsule formation, likely due to the surface degradation behavior of PGS.7
Tissue engineering has advanced rapidly over the last three decades, and part of the driving force of innovation in the field comes from novel biodegradable elastomeric biomaterials.8 All bodily tissues are inherently elastic to some degree and many implants/grafts partly fail due to a mechanical property mismatch between native and engineered constructs.9 For materials interfacing with vascular tissue, substrate elasticity and mechanical stimulation significantly influence cell functions and tissue development.10 Thus elastomeric materials are recognized as an important class of scaffold materials for vascular tissue and other soft tissue regenerative applications.
Mechanical properties are a particularly important criteria for the selection of materials used in cardiovascular applications. In particular, PGS shows little plastic deformation, making it attractive for engineering cardiovascular tissues. Small-diameter arterial grafts are still a major challenge in tissue engineering, and highly porous PGS scaffolds can be particularly effective in the engineering of small arteries.11 Moreover, endothelial progenitor cells and smooth muscle cells (SMCs) adhere and proliferate well on PGS.11d SMCs cultured in PGS scaffolds co-express elastin and collagen, leading to highly compliant engineered blood vessels.11c Furthermore, while the amount of tropoelastin made by these cells is identical on PGS and PLGA scaffolds, the elastic PGS substrate allows for crosslinking of tropoelastin into desmosinecrosslinked elastin.11b In a rat abdominal aorta model, composite arterial grafts comprised of PGS tubes reinforced with a polycaprolactone nanofiber sheath demonstrated constructive remodeling of the graft into neoarteries within 3 months.12 The neoartery mimicked the native artery mechanically, biochemically, and anatomically, and the neoarteries were well integrated with host vasculature. Remarkably, the neoartery pulsed synchronously with host arteries. After one year post-implant, the neoarteries contained the same amount of elastin as their native counterpart and had regenerated in the adventitia of the neoarteries (Figure 2).13
Figure 2.Gross morphology and tissue architecture of neoarteries resemble native arteries. A) Top left: transformation of graft into neoartery in situ over the course of 1 year. Nondegradable sutures (black) mark the graft location. Top right: Transverse view of explanted neoarteries resembles that of native aortas. Bottom: Longitudinal view of explanted neoarteries resembles the adjacent native aorta. All ruler ticks are 1 mm. B) H&E stained transverse sections of the middle of neoarteries show similar tissue architecture with native aortas, with no visible graft material residues. Scale bar 100 mm. C) Neoartery sections immunostained for von Willebrand factor (vWF, red) and α-smooth muscle actin (α-SMA, green). The luminal surface of neoarteries is completely covered by vWF positive cells (red), suggesting a confluent endothelium. Neoarteries contain a media-like middle layer of the vascular wall rich in α-SMA positive cells with circumferentially elongated nuclei, similar to vascular smooth muscle found in native aortas. The outermost layer of neoarteries lacks α-SMA, resembling native adventitia. Some cells in the media-like layer are negative for α-SMA, and some cells adjacent to the endothelium are α-SMA positive but not circumferentially elongated. Scale bar 100 mm. L indicates vessel lumen. Nuclei stained with DAPI (blue). D) En face view of the luminal surface of neoarteries shows complete coverage by CD31 positive cells with cobblestone-like morphology and alignment parallel to the direction of blood flow, an arrangement similar to that found in native aortas. Neoarteries were cut open longitudinally and imaged as whole mounts using confocal microscopy and z-stack flattening. Arrow indicates the direction of blood flow. Scale bar 100 mm. Reprinted by permission from Reference 13. Copyright 2013, Elsevier Ltd.
PGS is also used extensively in cardiac tissue engineering14 due to the ease in modulating the mechanical properties of PGS to readily match those of myocardial tissues.5 In one application, PGS was used to fabricate highly porous scaffolds with parallel channels that mimic the capillary networks found in native myocardium.14c Co-cultures of cardiac fibroblasts and cardiomyocytes in a perfusion bioreactor with oxygen carriers yielded contractile constructs within 11 days.14d When placed in vivo, cell-free PGS scaffolds vascularize after implantation in an infarcted rat myocardium model within 2 weeks.14d Recently, a PGS scaffold with an accordion-like honeycomb microstructure was created (Figure 3).14a Its stiffness was controlled by curing time in order to match the mechanical properties of rat right ventricular myocardium. Additionally, PGS scaffolds have been precoated with ECM proteins to provide ligands for increased cell interaction which increased cellularity, enhanced ECM protein production, and modulated the differentiation of endothelial progenitor cells.15
PGS has also demonstrated promise as a scaffold material for nerve regeneration.7 The in vitro and in vivo neural biocompatibility of PGS has been systematically evaluated. Primary Schwann cells showed similar attachment rate and metabolic activity on both PGS and PLGA surfaces in vitro. The cells on PGS had a higher proliferation rate and lower apoptotic activity than those on PLGA. In vivo implantation juxtaposed to the sciatic nerve revealed PGS causes a significantly lower chronic inflammatory response than PLGA, likely due to the minimal swelling and surface eroding characteristics of PGS. A recent study investigated microfabricated PGS porous scaffolding for retinal progenitor cell (RPC) grafting. The scaffold had a Young’s modulus of 1.66 } 0.23 MPa and a maximal strain of 113 } 22%.16 These mechanical properties more closely resemble those of retinal tissue (Young’s modulus of 0.1 MPa and maximal strain of 83%) than the traditional PLA/PLGA blend (Young’s modulus of 9.0 } 1.7 MPa and maximal strain of 9%) used for RPC delivery. The in vitro study revealed that RPCs adhere to and proliferate well in the PGS scaffold, and shows a trend toward differentiation. Subretinal transplantations demonstrated long-term RPC survival and high levels of RPC migration into host retinal tissue.17
Figure 3.Accordion-like honeycomb scaffolds yield anisotropic mechanical properties similar to native myocardium. A,B) Schematic diagrams illustrating the accordion-like honeycomb design constructed by two overlapping 200 × 200 μm squares rotated 45° (diamonds). Preferred (PD) and orthogonal cross-preferred (XD) material directions, respectively, corresponding to circumferential and longitudinal axes of the heart, are indicated. Scale bars: 1 mm (A) and 200 μm (B). C) Scanning electron micrographs demonstrating the fidelity of excimer laser microablation in rendering an accordion-like honeycomb design in PGS. Scale bars: 200 μm. D) PGS curing time was systematically varied, yielding a linear dependence of PGS effective stiffness (EPGS) on curing time within the tested range. E) Representative uniaxial stress–strain plots for accordion-like honeycomb scaffolds with cultured neonatal rat heart cells (scaffolds were fabricated from PGS membranes cured for 7.5 h at 160 °C; neonatal rat heart cells were cultured for 1 week). Reprinted by permission from Reference 19. Copyright 2008, Nature Publishing Group.
Although bone is a hard tissue, it develops from soft collagenous tissue in the embryonic stage. Similarly, the natural bone healing process also starts from a soft provisional tissue called the callus. For this reason, scaffolds made from PGS elastomer were used to heal a non-union bony defect. To do this, a porous PGS tube was used to join two ends of a completely transected ulna in a rabbit model.18 Healing started with the formation of a cartilage tissue similar to a callus, which progressively mineralized and completely bridged the defect within 2 months as examined by micro-CT. Results revealed the lower stiffness PGS elastomer can create a load-transducing environment in which bone regeneration more effectively occurs. In contrast, metallic implants can cause stress shielding of the bone making healing more difficult.
Coating technology plays a significant role in medical device development as it provides a means of modifying the underlying substrate and enhancing the performance of the device. PGS has shown tremendous promise as a coating material; its resin is easily reducible in a wide range of solvents (e.g., ethyl acetate, THF, acetone, 1,3-dioxolane, and various alcohols) resulting in a solution that can be used in dip and spray coating applications. Figure 4 shows a range of textile substrates (PET, polypropylene, PGA, and nitinol) coated with a smooth conformable thin film of PGS. Enhanced mechanical properties, improved biocompatibility, and antimicrobial properties are all features imparted by a PGS coating, illustrating its utility in the medical device space.
Figure 4.SEM images of Regenerez coatings deposited on a range of commonly used medical device textile components: A) dip-coated poly(ethylene terephthalate) woven, B) dip-coated poly(glycolic acid) knit, C) dip-coated PEEK mesh, D) spray-coated nitinol braid, and E) dip-coated poly(propylene) mesh. SEM images provided by Carissa Smoot of the Secant Group.
Biomaterials will continue to play an important role in medical device and regenerative medicine as a need will always exist for materials that are able to mimic the properties of native tissue. PGS has a multitude of properties which makes it an ideal material to fill the many technological demands of device and tissue engineering applications. Over the past 15 years, PGS has been utilized in applications within the fields of cardiovascular, nervous, soft, and hard tissue and continues to find new uses today such as coatings for implantable devices. Over this timeframe, PGS has progressed from the research lab to commercialization with the introduction of Regenerez® Poly(glycerol Sebacate) Resin. Recent research advances, as outlined here, will certainly expand the utility and application of this versatile biomaterial.