Abstract

Introduction

Regeneration of the infarcted myocardium after a heart attack is one of the most challenging aspects in tissue engineering. Suitable cell sources and optimized biocompatible materials must be identified.

Sources of data

In this review, we briefly discuss the current therapeutic options available to patients with heart failure post-myocardial infarction. We describe the various strategies currently proposed to encourage myocardial regeneration, with focus on the achievements in myocardial tissue engineering (MTE). We report on the current cell types, materials and methods being investigated for developing a tissue-engineered myocardial construct.

Areas of agreement

Generally, there is agreement that a ‘vehicle’ is required to transport cells to the infarcted heart to help myocardial repair and regeneration.

Areas of controversy

Suitable cell source, biomaterials, cell environment and implantation time post-infarction remain obstacles in the field of MTE.

Growing points

Research is being focused on optimizing natural and synthetic biomaterials for tissue engineering. The type of cell and its origin (autologous or derived from embryonic stem cells), cell density and method of cell delivery are also being explored.

Areas timely for developing research

The possibility is being explored that materials may not only act as a support for the delivered cell implants, but may also add value by changing cell survival, maturation or integration, or by prevention of mechanical and electrical remodelling of the failing heart.

Heart disease

Cardiovascular disease is the leading cause of death in the UK, comprising 39% of all deaths per annum (www.bhf.org). Myocardial infarction (MI), commonly known as a heart attack, is caused by the abrupt occlusion of one or more of the blood vessels (coronary arteries) supplying blood to heart. This reduces the supply of nutrients and oxygen to the heart muscle (myocardium). If blood flow is not restored rapidly, irreversible cell death occurs within the blood-deprived myocardium, eventually impairing cardiac performance. Patients who survive the acute event may eventually develop heart failure, which is defined as the clinical state resulting from the inability of the heart to pump enough blood to meet the body's metabolic requirements. The adult human heart is unable to self-regenerate to a significant degree, and a fibrous non-contractile scar tissue is formed in the infarcted myocardial territory (Fig. 1). The replacement of the healthy myocardium after infarction with a non-contractile fibrous scar tissue, which also does not effectively conduct the electrical wavefront, reduces the heart contractile efficiency. Compensatory mechanisms are activated to assist the heart to maintain cardiac output. This ultimately places an extra burden on the weakened myocardium, eventually leading to end-stage heart failure, with a progressive fall in cardiac output, development of multi-organ failure from reduced perfusion and ultimately death. Pharmacological therapies can slow the progression to end-stage disease, but rarely prevent or reverse progression of the failing state. Currently, the only therapeutic options available to treat patients with terminal end-stage heart failure are heart transplantation or left ventricular assist devices (VADs). However these two options are not widely available, and have significant limitations of cost. While heart transplantation involves the replacement of the whole organ, VADs aim to prevent remodelling and dilation of the left ventricle via unloading the dilated chamber. Alternatives for advanced, but not terminal, heart failure are reverse remodelling strategies using either biventricular pacemakers where electrical dyssynchrony is present or external constraining mesh ‘jackets’ wrapped around the failing heart to prevent ongoing ventricular dilatation. Owing to mixed results achieved1 with some reporting improved cardiac function while preserving ventricular geometry and others reporting marginal improvements, alternative treatments are required. Below, we discuss the current suggested potential options to treat infarcted hearts, with focus on tissue engineering.

Fig. 1

Schematic diagram illustrating the damage caused by an MI in human heart. Source: http://www.uptodate.com (accessed 20 June 2008).

Fig. 1

Schematic diagram illustrating the damage caused by an MI in human heart. Source: http://www.uptodate.com (accessed 20 June 2008).

Suggested options for myocardial repair and regeneration

‘Resident Cardiac Progenitor Cells’

Traditionally the heart has always been thought to have no regenerative capability, with cardiomyocytes viewed as terminally differentiated cells that were unable to self-renew. Recently, this paradigm has been challenged. Anversa and co-workers have reported the presence of a group of ‘cardiac progenitor cells’ (CPCs) resident in the heart, with specific stem cell markers, Lin-, c-kit+ and Ki-67 and early cardiac marker such as transcription factors GATA-4 and Nkx2.5. Other groups have identified other potential resident CPCs using a variety of cell surface markers and demonstrated in vitro cell division and in vivo myocardial regenerative capacity.2 Their origin is still unclear, and it is not known whether these cells home to the heart from the bone marrow or may reside in the heart from foetal life. Limited evidence on their regenerative capacity and their low population in the elderly and diseased human heart have tempered the enthusiasm over CPCs contributing significantly to myocardial regeneration. More recently, Marban and co-workers3 expanded resident cardiac stem cells ex vivo and reported myocardial regeneration and functional improvements post-injection into infarcted mouse hearts.

Cell transplantation

Repair and regeneration of the infarcted myocardium has been studied with cellular cardiomyoplasty, which involves the injection of cells either directly into infarct or intravenously. A number of cell sources have been explored, including bone marrow,4 skeletal muscle5 and embryonic stem cells (ESCs).6 Table 1 summarizes the cell sources explored for cell transplantation, which have been extensively reviewed elsewhere.7–9 In general, all cell sources explored have had mixed results. No candidate cell has been uniformly successful in the clinical trials performed to date, though the modest reported improvements provide encouragement. However, many hurdles still face the clinical application of the autologous cell therapy strategies, including the ideal cell type, cell dose, optimal timing for cell transplantation and mode of delivery to optimize cell retention within the myocardium. Long-term survival and terminal differentiation post-implantation are yet to be determined.10 Whether cells are injected directly into the infarct or adjacent myocardium by thoracotomy, or delivered via coronary circulation, the need for surgery or the inevitable cell loss, respectively, raises doubts about cell transplantation.

Table 1

Cell sources used for cell transplantation.

Cell source General advantages General disadvantages 
Skeletal myoblasts5 Autologous, contractile, easily cultured and expanded in vitro Inability to transdifferentiate into CM, preventing electrical coupling causing arrhythmias 
Progenitor cells mobilized from bone marrow11 Autologous Results are not reproducible, cell route from bone marrow to infarct and bone marrow cell transdifferentiation to CM is questionable 
Bone marrow stem cells4 Easily accessible and autologous Mixed results achieved. Limited evidence on transdifferentiation. Possible formation of other tissue types 
Bone-marrow-derived cardiomyocytes12 Autologous Results of differentiating bone marrow into CM in vitro is difficult to reproduce 
ESC-derived cardiomyocytes6 Remain in culture for unlimited periods Immunological and ethical constraints
Teratoma formation 
Spermatogonial-derived cardiomyocytes13 Authentic cardiomyocyte phenotype Evidence of expandability and reproducibility at an early stage 
Primary cardiomyocytes14 True morphological representation and preserve cardiac function in vitro Do not survive in culture longer than 48 h (adult ventricular cells) and 1 week (neonatal myocytes), dedifferentiate in culture, unable to proliferate and self-renew, human source problematic 
CPCs2 Potential autologous source Evidence of expandability and reproducibility at an early stage 
Cell source General advantages General disadvantages 
Skeletal myoblasts5 Autologous, contractile, easily cultured and expanded in vitro Inability to transdifferentiate into CM, preventing electrical coupling causing arrhythmias 
Progenitor cells mobilized from bone marrow11 Autologous Results are not reproducible, cell route from bone marrow to infarct and bone marrow cell transdifferentiation to CM is questionable 
Bone marrow stem cells4 Easily accessible and autologous Mixed results achieved. Limited evidence on transdifferentiation. Possible formation of other tissue types 
Bone-marrow-derived cardiomyocytes12 Autologous Results of differentiating bone marrow into CM in vitro is difficult to reproduce 
ESC-derived cardiomyocytes6 Remain in culture for unlimited periods Immunological and ethical constraints
Teratoma formation 
Spermatogonial-derived cardiomyocytes13 Authentic cardiomyocyte phenotype Evidence of expandability and reproducibility at an early stage 
Primary cardiomyocytes14 True morphological representation and preserve cardiac function in vitro Do not survive in culture longer than 48 h (adult ventricular cells) and 1 week (neonatal myocytes), dedifferentiate in culture, unable to proliferate and self-renew, human source problematic 
CPCs2 Potential autologous source Evidence of expandability and reproducibility at an early stage 

CM, cardiomyocytes.

Tissue engineering

Tissue engineering comprises the combination of biomaterials, cells and specific growth factors. It is a interdisciplinary field where material scientists and cell biologists form a construct (ex vivo or in situ), eventually implanted onto the injured site by clinicians (Fig. 2). Tissue engineering is currently receiving much attention in the field of regenerative medicine as an alternative to current treatments.15 Myocardial tissue engineering (MTE) has been proposed as an alternative option to replace the scarred non-contractile fibrous tissue caused post-infarction. The basic MTE paradigm is to seed cells capable of forming cardiomyocytes onto a biocompatible material in vitro, followed by implantation of the construct on or in the infarcted region of the failing heart. The grafted tissue will in turn direct new tissue formation as the cells integrate with the native tissue while the scaffold degrades over time.16

Fig. 2

Schematic diagram illustrating the principle of tissue engineering for myocardial regeneration.

Fig. 2

Schematic diagram illustrating the principle of tissue engineering for myocardial regeneration.

Myocardial tissue engineering

Initially, a scaffold made from a biomaterial is designed. A selection of cells (cardiac or non-cardiac) is expanded in vitro with additional growth factors. Cells, with or without growth factors, are seeded onto the scaffold to form the MTE construct and mechanical, electrical and morphological properties optimized. This is eventually sutured onto or into the infarcted region.

Requirements for MTE constructs

Although many biomaterials have been suggested for various tissue-engineering purposes, all constructs share the following basic requirements. Additional requirements for MTE constructs include the following. In terms of the cardiac function itself, the physical characteristics of the material should lie in a range of values that are close enough to the natural myocardium that they do not massively enhance diastolic stiffness (and hence impede relaxation). It might be argued that a certain degree of support of the scar would be beneficial to prevent scar expansion and the consequent deleterious remodelling of the myocardium, and therefore stiffness slightly in excess of the scar itself would be acceptable. For the systolic function, the material alone will confer no benefit other than, possibly, contributing to elastic recoil: the important point is that the shape, attachment method and stiffness should not physically hinder the full contraction of the ventricle. If the material is combined with beating myocytes, then it should allow these cells to exert their contractile effect. One desirable potential strategy is for the material to support nascent new cardiomyocytes as they develop and integrate into the host and biodegrade when assimilation is complete.

  • Biocompatible. Biomaterial must not be rejected or induce an inflammatory response in vivo.

  • Mechanical integrity. Biomaterial must enable handling during transplantation. More importantly, mechanical properties should match the host tissue it intends to replace and provide mechanical support during regeneration.

  • Biodegradable. The degradation rate of the biomaterial should match the regeneration rate of the host tissue, and the degradation by-products must be non-toxic and readily removed from the body.

  • Cell ‘friendly’. Enhance cell adhesion and survival both in vivo and in vitro.

  • Biomimetic. Reflect the extracellular matrix (ECM) of the tissue it intends to replace.

  • Fabrication. Biomaterial must be easily accessible and designed with acceptable cost.

  • It is important that the biomaterial is able to withstand, or even contribute to the continuous stretching/relaxing motion of the myocardium that occurs at each heartbeat.

  • Biomaterial must encourage cell proliferation and differentiation into cardiomyocytes as well as supporting vascular cells.

  • Ideally, biomaterials could encourage cardiomyocyte alignment and maturation in vitro before implantation or in vivo, improving the contractile properties of the graft.

  • Biomaterial must enable electrical integration of engineered graft with the native tissue to allow synchronized beating between the artificial construct and the heart. This requires matched excitability of host and grafted tissue and support electrical of wavefront propagation.

  • Strategies to encourage vascularization of the construct to support the survival of grafted cells.

  • Suitable cell source that will not provoke arrhythmia once introduced into the body.

Methods adopted in MTE

Electrospinning

Collagen is the main constituent of the ECM, and therefore tissue-engineered constructs require an ECM-like structure and topography to mimic the in vivo size and scale of the collagen fibrils in the ECM. Electrospinning, a process patented in 1934 by Formhals, is used to develop scaffolds, made from synthetic, natural or a combination of both biomaterials, with sub-micron pores and nanotopography surfaces.17 Electrospinning involves an electrically charged jet of a polymer solution produced by a high voltage. A constant pressure generated by a metering pump causes the polymer mixture to flow from the pipette, at a constant rate, onto a collector screen which eventually dries/solidifies leaving a polymer fibre (fibre diameter can range from 3 nm to 5 µm). This method is common in MTE and has been suggested for various biomaterials (Table 2).

Table 2

Examples of the biopolymers, the processing used to produce the engineered construct and cell types used in MTE.

Scaffold material Method used for construct processing Cell source 
Natural materials 
 Collagen Commercially available 3D scaffold Acellular 
 Collagen gel Bioreactor Neonatal rat CM 
 Gelatin mesh (Gelafoam) Commercially available patches Foetal rat CM
Foetal human CM
Rat stomach smooth muscle cells
Rat skin fibroblasts 
Biostretch/bioreactor Foetal human CM 
 Collagen + glycosaminoglycan Two cross-linking methods Bone-marrow-derived mesenchymal cells 
 Collagen type I matrix Bioreactor Neonatal rat CM 
 Collagen type I sponge + matrigel Bioreactor Undifferentiated mESCs mESCMs 
 Sodium alginate Freeze-drying technique Neonatal CM
Foetal rat cardiac CM
hESC-CM
Neonatal rat CM 
Synthetic materials 
 Polyerethane Solvent cast and spin coating Neonatal rat CM 
 Elastomeric 1,3-trimethylene carbonate and d,l-lactide and copolymers Salt-leaching Rat cardiomyocytes cell (CRL-1446) line and human umbilical vein endothelial cells 
 Vicryl mesh (Dermagraft, Smith and Nephew) Commercially available Human dermal fibroblasts 
 Poly(N-isopropylacrylamide) Cell sheeting Neonatal rat CM 
 Poly(ε-caprolactone) Electrospinning Neonatal rat mesenchymal cells 
Neonatal rat CM 
 Poly-glycolic acid (PGA) Commercially available mESC-CM 
 Non-woven poly(lactide)- and poly(glycolide)-based (PLGA) Electrospinning Primary rat CM 
 Poly-co-caprolactone (PGCL) Solvent casting and particle leaching Rat bone-marrow-derived nuclear cells 
Combination of natural and synthetic materials 
 Poly(ester urethane) urea and collagen type I Electrospinning Aortic rat smooth muscle cells 
 Polycaprolactone (PCL) mesh coated with collagen type I Electrospinning Isolated rat cardiomyocytes 
 PGA, PLA and PCL polymers with commercially available hydrophilic collagen sponge (Ultrafoam®Three polymers mixed and pores achieved using gas-forming method, followed by immersion in collagen Neonatal heart cells 
 Collagen (types I and IV) and Matrigel matrix mixed with cells and seeded onto non-woven polymer mesh (poly-(l-lactide acid) reinforced with PTFE) Materials mixed in circular moulds Multipotent bone-marrow-derived mesenchymal progenitor cells 
Scaffold material Method used for construct processing Cell source 
Natural materials 
 Collagen Commercially available 3D scaffold Acellular 
 Collagen gel Bioreactor Neonatal rat CM 
 Gelatin mesh (Gelafoam) Commercially available patches Foetal rat CM
Foetal human CM
Rat stomach smooth muscle cells
Rat skin fibroblasts 
Biostretch/bioreactor Foetal human CM 
 Collagen + glycosaminoglycan Two cross-linking methods Bone-marrow-derived mesenchymal cells 
 Collagen type I matrix Bioreactor Neonatal rat CM 
 Collagen type I sponge + matrigel Bioreactor Undifferentiated mESCs mESCMs 
 Sodium alginate Freeze-drying technique Neonatal CM
Foetal rat cardiac CM
hESC-CM
Neonatal rat CM 
Synthetic materials 
 Polyerethane Solvent cast and spin coating Neonatal rat CM 
 Elastomeric 1,3-trimethylene carbonate and d,l-lactide and copolymers Salt-leaching Rat cardiomyocytes cell (CRL-1446) line and human umbilical vein endothelial cells 
 Vicryl mesh (Dermagraft, Smith and Nephew) Commercially available Human dermal fibroblasts 
 Poly(N-isopropylacrylamide) Cell sheeting Neonatal rat CM 
 Poly(ε-caprolactone) Electrospinning Neonatal rat mesenchymal cells 
Neonatal rat CM 
 Poly-glycolic acid (PGA) Commercially available mESC-CM 
 Non-woven poly(lactide)- and poly(glycolide)-based (PLGA) Electrospinning Primary rat CM 
 Poly-co-caprolactone (PGCL) Solvent casting and particle leaching Rat bone-marrow-derived nuclear cells 
Combination of natural and synthetic materials 
 Poly(ester urethane) urea and collagen type I Electrospinning Aortic rat smooth muscle cells 
 Polycaprolactone (PCL) mesh coated with collagen type I Electrospinning Isolated rat cardiomyocytes 
 PGA, PLA and PCL polymers with commercially available hydrophilic collagen sponge (Ultrafoam®Three polymers mixed and pores achieved using gas-forming method, followed by immersion in collagen Neonatal heart cells 
 Collagen (types I and IV) and Matrigel matrix mixed with cells and seeded onto non-woven polymer mesh (poly-(l-lactide acid) reinforced with PTFE) Materials mixed in circular moulds Multipotent bone-marrow-derived mesenchymal progenitor cells 

3D, three-dimensional; CM, cardiomyocytes; mESC-CM, mouse embryonic stem cell-derived cardiomyocytes; hESC-CM, human embryonic stem cell-derived cardiomyocytes.

Source: Jawad et al.17

Bioreactors

The general concept of a bioreactor is to encourage growth and development of biological cells or tissue on biomaterials as if under in vivo conditions. The ability of producing a 3D myocardial tissue comprising more than a few layers of muscle is the main advantage of using a bioreactor. Although the sole purpose of the construct developed with bioreactors is for scientific research, further improvements in in vitro design and quality control may eventually allow their application for MTE. Several different bioreactors have been suggested for cardiac constructs: static or mixed flask bioreactors, where constructs are suspended in a cultivation medium; rotating vessel bioreactors, where constructs are suspended in a medium that has a constant rotational flow; and finally perfusion cartridge bioreactors, where constructs are perfused at interstitial velocities, comparable to blood flow in native tissue.18

Cell sheeting (temperature-sensitive)

This method was initially reported by Shimizu et al.19 in 2002 and has been suggested for MTE. The concept of this method involves temperature-responsive dishes made from a specific polymer, poly(N-isopropylacrylamide), which is temperature sensitive. At 37°C, the polymer is hydrophobic and cell adhesive; however, a 5°C reduction in temperature can cause the polymer to become non-cell adhesive as it hydrates and swells because its hydrophobic nature is now hydrophilic. Cardiomyocytes seeded onto the polymer will produce individual spontaneously beating myocardial tissue sheets. More recently, Miyahara et al.20 reported the success of implanting a 100 μm thick cardiac tissue, made from six monolayered mesenchymal stem cell sheets layered together and implanted onto the infarcted region, improving the infarcted wall thickness. The caveats about the potential for the bone-marrow-derived mesenchymal stem cells becoming cardiomyocytes, indicated in Table 1, also hold for these cell sheets.

In situ engineering

Like tissue engineering, in situ engineering involves both biomaterials and cells. However, this is a ‘scaffold-free’ approach in MTE. It involves the direct injection of the biomaterial and cell mixture into the infarcted region. Unlike prefabricated scaffolds used in MTE, the injectable natural polymers will readily bond to the native tissue, as they can be easily shaped or cast to the heart's complex dimension, simultaneously providing a good support for the cells. Different polymers such as alginate,21 fibrin glue,22 collagen23 and matrigel24 have been suggested for in situ engineering. Acellular alginate with bioactive molecules has also been suggested, with the hope that CPCs will home to the infarcted region.25 More recently, biomaterials with peptides and growth factors,26 as well as ‘self-assembling’ peptide nanofibres,27 have been suggested. None of these strategies have in fact produced new myocardial tissue, and experience with injection of cells directly into the heart has suggested that there is a low limit on the amount of new material that can be introduced into either the dense and continually compressing myocardium or the stiff and avascular scar.

Biomaterials used in MTE

This section aims to highlight the achievements, to date, in producing MTE scaffolds (porous structure) or patches (dense structure). It must be noted that either a porous structure or a dense patch maybe suitable, depending on the purpose of the construct. If the engineered biomaterial is to support and possibly remould the infarcted area over a period of time, then it is vital the construct is a scaffold that consists of interconnected pores (>90% porosity) with diameters ranging between 300 and 500 µm for cell survival. This will allow cells to exchange nutrients and remove cellular secretions, enhance cell penetration and tissue vascularization.28 On the other hand, if the biomaterial will serve solely as a means of cell transport, to deliver cells to the desired region only and degrade over a given period of time (e.g. within 3 months), a dense patch will be adequate for this purpose.17 Wide ranges of biomaterials, mainly polymers, have been suggested for MTE, being synthetic and natural. This section will give an insight into the types of biomaterials reported specifically for myocardial regeneration.

Synthetic polymeric materials

Polyesters

The biodegradable poly(α-hydroxy acid) aliphatic polyesters such as poly ε-caprolactone, polylactic acid (PLA), polyglycolic acid (PGA) and their copolymers are the major classes of polymers suggested for MTE. The ability to tailor their mechanical properties, define their morphology and control their degradation kinetics by altering the copolymer ratio, to better suit the tissue it intends to replace, are attractive features of these synthetic polymers.15 Furthermore, the natural metabolites of PGA and PLA and their copolymers, glycolic and lactic acid, make these polymers more appealing as the human body is naturally able to completely remove monomeric compounds of lactic and glycolic acids, although it has been reported that these by-products increase acidic concentration in the body, causing inflammatory response which in turn damages the local tissue. A further disadvantage of these polymers is their bulk degradation kinetics which causes a sudden loss of mechanical properties of the construct.15 Freed and Vunjak-Novakovic in 1997 were the first to report the use of polyesters (PGA) as a scaffold for cardiac tissue engineering. A spontaneously contractile 3D engineered cardiac tissue with specific cardiac structural and electrophysiological properties was achieved using a bioreactor.28 However, there is increasing evidence that successful scaffold materials must have elastic properties that are similar to that of the native heart. This will allow the construct to withstand and possibly move in synchrony with each contraction/relaxation motion that occurs during each heartbeat, as well as prevent cells from detaching from the bioengineered construct.29 This means that the group of polyesters, which are less flexible than the heart tissue, will not fulfil all the requirements for an ideal MTE scaffold or patch.

Elastomeric polymers

The key advantage of these polymers is their elastic behaviour, in that they are able to withstand strong deformation forces and return to their original size upon removal of stress. This can circumvent the problems related to material stiffness. Elastomeric polyurethane (PU) has been suggested for MTE, where the 3D cardiac construct was achieved, reporting good cell adhesion.29 Furthermore, in vitro and in vivo studies have reported no tissue inflammation.15 However, a major setback of PU in MTE is its toxic by-product diisocyanate, which might be released upon degradation and known to be harmful to living tissue. The incorporation of diisocyanate is necessary in the synthesis of PU. Another elastomeric polymer used in MTE is 1,3-trimethylene carbonate (TMC),30 a copolymer containing TMC, and polyester d,l-lactide had the ability to sustain the heart's cyclic strains under physiological condition with no severe tissue reaction in vivo. Further example demonstrating the success of elastomeric polymer in MTE is polyethylene glycol (PEG), in which the spontaneous beating of PEG discs was achieved when cardiomyocytes were cultured on the surfaces.31 More recently, Chen et al.32 characterized a soft elastomer for MTE, made from poly(glycerol sebacate), which has already shown promising results in soft tissue engineering for nerves and vascular tissue engineering.32 A further group of materials being considered is that of poly(ethylenetephatalate)/dimer fatty acid block copolymer.33

Natural polymeric materials

Extracellular derivatives

Unlike natural polymers, synthetic polymers have the key advantage of allowing precise control of their hydrophilic/hydrophobic ratio, degradation rate and mechanical properties. However, they do not possess the biological specifics of natural polymers.15 ECM proteins and derivatives such as collagen type I and fibronectin are examples of natural polymers suggested for MTE. Although they are renowned to facilitate cell adhesion and proliferation and maintain cells in their differentiated states because of their particular adhesive properties, rapid degradation kinetics and weak mechanical properties hinder their success in MTE. In addition, immunogenicity and inconsistent material properties between various batches of natural polymers are still of major concern. Collagen has been widely investigated for both MTE 34 and in situ engineering,23 however, with many conflicting results. Eschenhagen et al.35 developed the all natural engineered heart tissue (EHT) made from a combination of neonatal cardiomyocytes and artificial matrix (collagen type I and Matrigel) in artificial moulds under mechanical strain. Recently, the EHT has been reported to improve the cardiac function, since 28 days post-implantation electrical coupling to native heart and systolic wall thickening of infarction were observed in vivo.14 Suitable cell source and graft size are examples of areas that need further investigation for the potential clinical use of the EHT.17

A complex protein mixture secreted by mouse tumour cells, Matrigel (trade name given by BD Biosciences), resembles the ECM environment. It is commonly used in laboratories as a substrate to enhance cell adhesion on to material surfaces. Interestingly, this has been suggested for in situ engineering as an acellular matrix and as a mixture with endothelial ESCs.24 In both cases, Matrigel was found to be effective in that the cardiac function improved attenuated the left ventricular function reported. Glycosaminoglycan (GAG) is also an abundant protein found in the ECM of the body; many studies have reported the success of incorporating GAG with collagen to form a nanofibrous scaffold.36 In 2005, Koifids et al.37 combined undifferentiated mESCs with collagen type I in a bioreactor to form an artificial myocardial tissue. Despite the success of the artificial myocardial tissue and other studies forming constructs from collagen, Matrigel and GAG, immunogenicity and weak mechanical properties are still issues that need to be addressed.

Gelatin is largely composed of denatured collagen, which is mainly obtained from connective tissue found in bones, ligaments, tendons and cartilage. A spontaneously contractile 3D cardiac tissue was obtained in 1999 by Li and co-workers by seeding foetal ventricular muscle cells onto a gelatin mesh in vitro. Post-implantation results showed cell survival and interconnection with native tissue in vivo, although no cardiac functional improvement was reported. More recently, they reported increased cellular proliferation and better performance upon mechanical stress on the gelatin foam.38

A more extreme and more imaginative use of natural ECM material has recently been described. Ott et al.39 decellularized rat hearts with detergents by coronary perfusion and obtained a perfusable acellular vascular construct with intact chambers and valves by preserving the myocardium matrix. This construct was then recellularized with cardiac or endothelial cells, and showed modest but incontrovertible contractile abilities. This approach has the potential to solve the problem of blood supply to the graft by retaining the natural structures guiding the formation of vessels. It also holds the vision of reconstruction of the whole organ, or large parts of it, which will be necessary if regeneration strategies are to address complex congenital conditions in which parts of the heart are absent or malformed.

Alginate

Although certain bacteria produce the natural negatively charged polysaccharide alginate, alginate is mainly derived from brown seaweed. Its unique property of forming hydrogels in the presence of calcium ions has attracted its use in MTE. Leor et al.21 produced alginate scaffolds with 90% porosity and pore sizes ranging between 50 and 150 µm by freeze drying. Intensive neovascularization was revealed post-implantation of the alginate scaffold with foetal rat cardiac cells. hESC have also been seeded onto alginate scaffolds, where no regeneration was observed.25 Although other studies have investigated alginate for in situ engineering where acellular alginate with bioactive molecules was injected into the infarcted area, with hope that the reported CPCs will home to the diseased region.25 However, due to the limited studies on the effect of alginate, one cannot conclude whether it was the cells alone, the alginate alone or the combination of both that contributed to the healing of the infarction.

Fibrin glue

Fibrin glue is a biopolymer formed by the polymerization of fibrinogen monomers. This natural polymer has only been suggested for in situ engineering. Different cell types have been investigated with this polymer, skeletal myoblasts, bone marrow cells and endothelial cells,17 all reporting enhanced neovascularization, improved left ventricular function and reduced cardiac remodelling. Although results are promising, this material is at infancy and requires further investigations.

Combination of synthetic and natural polymers

Owing to the poor mechanical properties of collagen, researchers have investigated the combination of collagen and synthetic polymeric materials. Both materials would contribute to the scaffold with equal importance. The synthetic polymer would provide a suitable mechanical support, whereas the natural collagen polymer would confer to the cells a more in vivo-like environment. The combination of collagen with various types of synthetic polymers is also summarized in Table 1. Both electrospinning and bioreactors have been used to combine the two materials.17 Krupnick et al.40 mixed bone marrow progenitor cells with a collagen and Matrigel matrix and then went on to seed the mixture onto the synthetic non-woven polymeric mesh. In vivo implantation onto injured hearts was successful, in that minimal intra-cardiac inflammation occurred as well as the cardiac function appeared normal, with no arrhythmias reported. However, the immunogenicity of collagen is still of great concern.

Closing remarks/limitations

A great effort is being made towards the development of an ‘optimal’ MTE construct; however, limitations to take this further into clinical trials are inevitable. The choice of biomaterial, cell source and suitable environment for cells to proliferate and differentiate in vitro before implantation remain obstacles in the field. The support and enhancement of cardiac contractile and electrical properties need careful consideration when designing the construct. Although many biomaterials, with varying compositions and properties, are continually being suggested for MTE, the current challenge is to focus investigations on the already available materials of proved biocompatibility in order to improve their performance towards clinical trials. The proposed materials, which use animal-derived products such as Matrigel, add another layer of complication to the testing and approval process for clinical application. There is little doubt about which type of cells needs to be introduced into the infarcted region; cardiomyocytes have shown to help regeneration and improve the cardiac function. However, the source of cardiomyocytes, whether those resident in the heart or derived from the bone marrow, embryonic or spermatogonial stem cells, to be used as the cell source, needs to be determined before constructs can be introduced into human hearts. Furthermore, at what stage of development cells should be used remains to be determined: whether immature cells which could further proliferate in vivo or mature cardiomyocytes with the properties of adult myocytes. In addition, the question remains as to whether cardiomyocytes alone should be introduced into the region or rather a multi-type culture consisting of cardiomyocytes, endothelial and fibroblastic cells to further enhance neovascularization. Extreme purification strategies for cardiomyocytes may be counterproductive for this reason. Moreover, whether the cellular constructs need to be cultured in vitro over a period before introducing into the myocardium or implantation should occur immediately after cell seeding remains questionable. A further possibility is the initial implantation of an acellular construct into the diseased myocardium to allow impregnation of newly formed vessels and ECM on the scaffold to provide a suitable environment for the implanted cells that will follow. Attachment of the material is another challenge, given the strong and repetitive forces generated during myocardial contraction. Compressive forces upon constructs placed within the muscle, or stretching of suture sites, will be experienced in each of the ∼100 000 daily heartbeats. Although many groups have provided evidence of the beneficial effects of MTE in vitro and in vivo (animal models), the mechanism behind this functional improvement is yet to be elucidated. In the meantime, it is important to overcome these obstacles and further improve our understanding of myocardial regeneration for the design and development of the most suitable, reliable and affordable construct.

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