Biomaterials in myocardial tissue engineering

Abstract

Cardiovascular disease is the leading cause of death in the developed world, and as such there is a pressing need for treatment options. Cardiac tissue engineering emerged from the need to develop alternative sources and methods of replacing tissue damaged by cardiovascular diseases, as the ultimate treatment option for many who suffer from end-stage heart failure is a heart transplant. In this review we focus on biomaterial approaches to augmenting injured or impaired myocardium, with specific emphasis on: the design criteria for these biomaterials; the types of scaffolds - composed of natural or synthetic biomaterials or decellularized extracellular matrix - that have been used to develop cardiac patches and tissue models; methods to vascularize scaffolds and engineered tissue; and finally, injectable biomaterials (hydrogels) designed for endogenous repair, exogenous repair or as bulking agents to maintain ventricular geometry post-infarct. The challenges facing the field and obstacles that must be overcome to develop truly clinically viable cardiac therapies are also discussed.

Keywords: biomaterials; cardiac regeneration; cardiac scaffolds; cardiac tissue engineering; cardiac tissue models; injectable hydrogels.

Figure 1. Natural and synthetic scaffolds for cardiac tissue engineering

(A) Cardiac cell organization in peptide-attached alginate scaffolds. Cells were grown on scaffolds with both G4RGDY (RGD) and G4SPPRRARVTY (HBP) peptides; scaffolds with a single peptide attached, or unmodified alginate scaffolds. On day 14 post-seeding, cells were double stained with antibodies against α-sarcomeric actinin (green) for cardiomyocytes and F-actin (red) for both myocytes and non-myocytes (nuclei, blue; scale bar = 20 μm)(Sapir et al., 2011). (B) Isotonic myofiber arrangement of cardiomyocytes in EHT. Collagen scaffold with RGD were seeded with 5-day embryoid body-derived cardiomyocytes. On day 10 post-seeding, the ultrastructure of the EHT was assessed by transmission electron microscopy. Sarcomeric components, including organized bundles of myofibrils (MF) and Z bands (Z), are visible (scale bar = 500 nm)(Dawson et al., 2011). (C–F) Multi-loop EHT construct as cardiac patch. Individual loop-shaped EHTs were stacked onto a custom device designed to promote the fusion of the individual loops into a single multi-loop EHT and to permit their contraction under auxotonic load (scale bar = 10 mm) (C–E). The resulting synchronously contracting multi-loop EHT was engrafted onto the recipient heart using six single-knot sutures (F) (Zimmermann et al., 2006). (G–I) Four weeks after engraftment, EHTs form thick, organized, vascularized muscle. H&E staining of paraffin section through infarcted area shows engrafted EHTs form compact and oriented heart muscle (scale bar = 500 μm) (G). Sarcomeric organization of engrafted EHTs is visible by laser scanning microscopy (actin, green; nuclei, blue) (scale bar = 50 μm) (H). Newly formed vessels containing smooth muscle cells, with macrophages in close proximity are visible by confocal microscopy (actin, green; nuclei, blue; ED2, red and arrows) and erythrocytes in the vessels are visible by differential interference contrast imaging (asterisks; scale bar = 50 μm) (I)(Zimmermann et al., 2006).

Figure 2. Scaffolds made from decellularized extracellular matrix

(A–C) Decellularized whole adult rat hearts have patent vasculature and can be perfused. Coronary corrosion casts of cadaveric and decellularized rat hearts demonstrate vessel patency at both the macroscopic (upper row; scale bar = 1000 μm) and microscopic (lower row; scale bar = 250 μm) level (A). Functional perfusion was demonstrated by heterotypic transplantation of the decellularized heart with visualization before (B) and after (C) the host aorta was unclamped(Ott et al., 2008). (D) Recellularized heart constructs perform stroke work. Decellularized rat hearts were mounted into a working-heart bioreactor, seeded with neonatal rat cardiac cells, and function was assessed by real-time ECG, aortic pressure (after load) and left ventricular pressure (LVP) on day 0, 4 and 8. Decellularized heart construct on day 0 (left). Right lateral (top panel) and anterior view (bottom panel) of regions of movement and corresponding tracings in paced recellularized heart constructs on day 4 (centre). On day 8, electrical stimulation induced an increase in LVP, which was followed by repolarization (right) (Ott et al., 2008). (E) Composite scaffolds for cell delivery were assembled from thin sheets of decellularized human myocardium, fibrin hydrogel and immune selected MPCs (cultured ±TGF-β conditioning), and implanted into nude rat left anterior descending artery (LAD) ligation models of acute and chronic myocardial infarction. After 4 weeks, the hearts were explanted and function was evaluated by echocardiography (Godier-Furnemont et al., 2011). (F) Decellularized matrix maintains nearly the same extracellular matrix composition as native tissue, excepting fibronectin content (scale bar = 250μm)(Godier-Furnemont et al., 2011). (G) Composite scaffolds with MPCs promote angiogenesis and arteriogenesis. Four weeks post-implantation in nude rat models of acute cardiac infarction, the composite constructs with MPCs, with or without TGF-β preconditioning (MPC/scaffold group and TGF-β/MPC/scaffold group, respectively), had increased size and frequency of factor VIII- and SMA-positive blood vessels, relative to the scaffold only group (asterisk, patch; interface, dotted line; scale bar = 250μm)(Godier-Furnemont et al., 2011).

Figure 3. Scaffolds for engineered cardiac patches and tissue models

(A–C) SEM images of collagen sponges for (A) PBS only treated scaffold (PBS), (B) scaffold with lower dosed of co-immobilized VEGF and Ang1 (1/2VEGF+1/2Ang), (C) scaffold with higher dose of co-immobilized VEGF and Ang1 (VEGF+Ang1). SEM images suggested there were no significant differences in the pore structure and porosity of growth factor immobilized and control scaffolds(Chiu and Radisic, 2010). (D) Tensile modulus of scaffolds treated as listed. The tensile moduli of scaffolds with immobilized growth factors were significantly higher than that of PBS control sponges (P=0. 0353)(Chiu and Radisic, 2010). (E–I) Representative images of Factor VIII staining in Chicken CAM assay (brown represents positive staining; blue represents counter stain)(Chiu and Radisic, 2010). (J) Scaffolds with co-immobilized VEGF and Ang-1 enhanced angiogenesis in the chicken chorioallantoic membrane assay compared to scaffolds with immobilized VEGF or Ang-1 alone(Miyagi et al., 2011). (K–M) Representative images of Masson’s trichrome staining at 28 days after patch implantation. Arrows indicate thickness of the patch. At 28 days, wall thicknesses in both VEGF-treated patches were significantly greater than those in the control patches (p<0. 05 for Low VEGF vs. control; p < 0. 001 for High VEGF vs. control)(Miyagi et al., 2011). (N) Patch thickness at 28 days after implantation (Normal RV = normal right ventricular tissue; red dotted line indicates patch thickness at the time of implantation). Scaffolds with immobilized VEGF increased angiogenesis as compared to unmodified scaffolds(Miyagi et al., 2011). (O) Blood vessel density within the patches at 28 days after implantation. Scaffolds with immobilized VEGF increased angiogenesis as compared to unmodified scaffolds(Miyagi et al., 2011). (P–R) Representative images of CD31 expression (arrows identify CD31-positive vascular structures) at 28 days after patch implantation. Blood vessel density within the patch was significantly increased in the High VEGF group at 28 days (p < 0. 01 vs. control; p < 0. 05 vs. Low VEGF) after patch implantation(Miyagi et al., 2011). (S–W) The engineering of a vascular bed by induction and organization of capillary outgrowths from parent vascular explants. (S) The mouse or human artery and vein explants were placed at the two ends of a polydimethylsiloxane (PDMS) substrate with microgrooves of 25μm, 50μm or 100μm in width and coated with Tβ4-encapsulated collagen-chitosan hydrogel. The samples were cultivated for 3 weeks to achieve capillary outgrowths connected between the parent explants, or for 2 weeks with hepatocyte growth factor (HGF) or VEGF-supplementation in the culture medium. (T–U) The oriented capillary outgrowths connected from the arterial explant to the venous explant at Day 14 with HGF supplementation. Fluorescent microscopy images showing capillary outgrowths extending between mouse artery to vein at (T) Day 7 and (U) Day 14 of in vitro cultivation on PDMS substrate with 100μm grooves and 1500ng encapsulated thymosin β4. Arrows indicate locations of artery and vein explants. Scale bars, 100μm. Confocal microscopy images of capillary outgrowths indicate (V) cross-sectional and (W) longitudinal lumens(Chiu et al., 2012a).

Figure 4. Various injectable hydrogels, injection strategies, and outcomes

(A) Representative trichrome stains of transverse heart sections shows the effects of different delivery strategies (of modified PEG hydrogel with/without Tβ4 and cells) on cardiac structure 6 weeks after injection. Collagen in the infarct areas is shown in blue, whereas myocytes are in red. Scale bar corresponds to 2. 5 mm(Kraehenbuehl et al., 2011). (B–J) Fabrication of the myocardial matrix ECM hydrogel is done by (B) slicing porcine ventricular myocardium and then (C) decellularizing using sodium dodecyl sulfate. (D) Hematoxylin and eosin staining of a histological section reveals cellular removal. The decellularized extracellular matrix is then milled into a fine powder (E) and then solubilized through enzymatic digestion (F), which allows for injection via syringe and a 27-gauge needle (G) through percutaneous transendocardial delivery. (H) Image of myocardial matrix being injected through a MyoStar 27-gauge catheter, using a 1-ml Luer lock syringe, attached to the catheter. (I&J) NOGA maps for healthy animal representing final injection locations, indicated by orange dots(Singelyn et al., 2012). (K–R) Gelation properties of the (NIPAAm-co-AAc-co-HEMAPTMC) (86/4/10) hydrogel shows the cooled hydrogel solution (K) transitions to gel after incubation in 37 °C water bath for 30 s (L). Extracted hydrogel formed after 10 min (M) is very pliable and can be stretched to many times its original size (N). Representative images at 8 weeks following the injection procedure of the anterior view of PBS injected (O), and poly (NIPAAm-co-AAc-co-HEMAPTMC) (86/4/10) injected hearts (P). White arrow shows an aneurysm formation in the apex area (O), blue arrows indicate the injected hydrogel area (P). The composite histological sections of PBS injected (Q) and hydrogel injected (R) myocardial walls 8 weeks after injection stained with H&E. Black arrow shows an aneurysm formation (Q). Blue box indicates infarct site and injection area of hydrogel (R). Scale bar: 5 mm in (O,P), 500 μm in (Q,R)(Fujimoto et al., 2009).