The significance of pore microarchitecture in a multi-layered elastomeric scaffold for contractile cardiac muscle constructs
Introduction
Cardiovascular disease is the leading cause of death in developed countries [1] and congenital heart disease, which affects approximately one percent of newborns world-wide, is associated with high morbidity [2]. The functional consequences of myocardial infarction (MI) and other heart defects in which muscle fibers and collagen networks are disrupted are loss of myocardial elasticity, compliance and pumping action [3]. Current myocardial regeneration strategies, while promising [4], are unable to recreate the robust mechanical and contractile properties of normal heart muscle. In particular, an effective graft for myocardial repair is a critical unmet need, where combining elasticity and strength without compromising heart cell viability and contractility have proved challenging [5], [6], [7].
In the prototypical tissue engineering approach, three dimensional (3D) scaffolds provide the delivery vehicle for transplanting large numbers of viable cells toward a goal of tissue regeneration [8], [9]. Numerous 3D biomaterials have been explored as cardiac tissue engineering scaffolds, including non-woven poly(glycolic acid) (PGA) mesh [10], [11], [12], collagen gel [13], [14], collagen foam [15], [16], [17], [18], [19], alginate foam [20], [21], chitosan foam [22], knitted poly(lactic acid) [23], knitted hyaluronan ester [24], poly-4-hydroxybutyrate foam [25], poly(lactic acid)/poly(glycolic-co-lactic acid) (PLLA/PLGA) foam [26], and composites of natural and synthetic polymers [27]. However, these scaffolds are either thermoplastic polymers, which tend to be stiffer than normal soft tissues, degrade by bulk hydrolysis, and fail under long-term cyclic loading [28], or naturally occurring materials with intrinsic variability, immunogenicity, and mechanical strength concerns [29]. Langer and colleagues [30] developed a tough bioresorbable elastomer, poly(glycerol-sebacate) (PGS), that degraded predominately by surface hydrolysis [31] and has been tested in various tissue engineering applications [32], [33], [34] including myocardial repair. The mechanical properties of the PGS elastomer, both in the context of non-porous membranes [7], [35], [36] and porous scaffolds [37], [38], could be tailored to match those of normal heart muscle. Recently, one-layered (1L) PGS scaffolds with in-plane pore anisotropy, i.e., rectangular and accordion-like honeycomb pores produced by laser microablation of ∼250 μm thick PGS membranes [37], were shown to guide the alignment of cultured neonatal rat heart cells [37] and C2C12 myoblasts [39].
Alternatives to the cell-scaffold paradigm include “scaffold-free” approaches based on transplanting cell–cell or cell-extracellular matrix (ECM) grafts. As examples, engraftment and vascularization were demonstrated for heart cell patches comprised of human embryonic stem cell-derived cardiomyocytes, endothelial cells, and fibroblasts [40] and electrical and vascular integration were demonstrated after implantation of thin (∼100 μm) heart cell sheets comprised of interconnected cardiomyocytes [41]. However, scalability remains a major limitation of scaffold-free approaches [9], [13], [42], [43]. Other alternative approaches include “cell-free” biomaterials for myocardial repair. However, biomaterials used for congenital heart defect repair in pediatric patients are limited by lack of potential for growth and remodeling [44], [45], and although cell-free, non-porous PGS membranes were recently shown to reduce post-infarction myocardial hypertrophy in rodents, these implants could not assist contractile function, suggesting a role for cell-PGS implants in future approaches [35].
In the present study, multi-layered elastomeric PGS scaffolds with controlled pore microarchitectures were fabricated and combined with heart cells to engineer contractile cardiac muscle constructs in vitro. Excitation threshold, gene expression, and cardiac specific marker proteins were assessed under different conditions of cell seeding and cultivation, in particular scaffold coating with laminin (LN) to promote heart cell attachment [11], [38], [46] and interstitial perfusion to promote heart cell viability [12], [18], [19], [20], [47].
Section snippets
Methods
Fig. 1 provides an overview of methods used to microfabricate and demonstrate the multi-layered PGS scaffold.
Scaffold microfabrication
To produce 1L scaffolds with accordion-like honeycomb pores in PGS membranes, we adapted our previously described method [37] for use with a frequency quintupled 213 nm Nd:YAG laser. A program was written in Visual Basic for Applications to generate sequences of coordinates and laser parameters, and suitably formatted for uploading into the software controlling this laser (DigiLaz II, v.3.1.0; CETAC Technologies) such that a specified in-plane pore microarchitecture could be automatically
Discussion
Tissue engineered cardiac muscle remains challenged by cell sourcing, mass transport, and scaffold limitations [4], [7]. Recent advances in PGS microfabrication have permitted the design of biodegradable, elastomeric scaffolds with precisely defined pore microarchitectures amenable to both cardiomyogenesis and predictive mathematical modeling. Toward scaling-up our previous 1L accordion-like honeycomb PGS scaffolds for cardiac tissue engineering [37], laser-microablated PGS membranes were
Conclusion
Multi-layered elastomeric PGS scaffolds with controlled pore microarchitectures were fabricated by laser ablation and oxygen plasma-mediated lamination, seeded with heart cells, and cultured with interstitial perfusion. The laser-microablated PGS exhibited UTS and ɛf higher than normal rat left ventricular myocardium and stiffnesses ranging from 220 to 290 kPa. Heart cell culture on these scaffolds yielded cardiac muscle constructs. Excitation thresholds were unaffected by scaffold scale up
Acknowledgments
This work was funded by the American Recovery and Reinvestment Act (ARRA), Award 1-R01-HL086521-01A2 (to LEF) from the National Heart, Lung and Blood Institute (NHLBI). The content is solely the responsibility of the authors and does not necessarily represent the official views of the NHLBI or NIH. We are indebted to R. Langer for general advice, J. Wang and J. Hsiao for help with polymer synthesis, processing, and SEM, N. Watson for help with microscopy, M.G. Moretti and G. Talo and E. Kim for
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