Cardiac tissue engineering

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Cardiac tissue engineering Q Z C H E N , S E H A R D I N G , N N A L I , H J A W A D and A R B O C C A C C I N I , Imperial College London, UK

16.1 Introduction Heart disease remains the leading cause of death and disability in the industrialised nations. It afflicts 1.8 million Britons and 25 million people worldwide, with approximately 120 000 new cases diagnosed each year in the UK. Heart failure, which is characterised by a dilated or hypertrophied heart, is a syndrome of breathlessness, oedema and fatigue resulting from damaged or defective myocardium. The enlargement in ventricular volume leads to progressive structural and functional changes in ventricles (called ventricular remodelling), representing a predisposing factor towards the end stage of heart failure. The most frequent initiating cause of heart failure is myocardial infarction, also known as heart attack, which is the single most common cause of death in economically developed countries, including the US and Western Europe. Myocardial infarction typically results in fibrous (collagen) scar formation and permanently impaired cardiac function because, after a massive cell loss due to ischaemia, the myocardial tissue lacks significant intrinsic regenerative capability. Eventually, heart transplantation is the ultimate treatment option to end-stage heart failure. Owing to the lack of organ donors and complications associated with immune suppressive treatments, scientists and surgeons constantly look for new strategies to repair the injured heart. Around the mid-1990s studies veered to an intriguing strategy: cell therapy (Koh et al., 1993). A number of studies carried out so far (e.g. Li et al., 1996; Scorsin et al., 1997; Taylor et al., 1998; Reinecke et al., 1999) have indicated that cell implantation (Fig. 16.1a) in models of myocardial infarction can improve contractile (mostly diastolic) function. An alternative approach to deliver isolated cells into the heart is to use a synthetic biodegradable construct (Fig. 16.1b), which is manipulated in vitro with cells and implanted later in vivo. So far, many studies have been published using different cells and different synthetic materials (Cima et al., 1991; Zammaretti and Jaconi, 2004; Leor et al., 2005). In this chapter, we present a review on the achievements of myocardial tissue engineering, focusing on construct (scaffold)-based strategies. It must be

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16.1 Two strategies in myocardial tissue engineering: (a) direct transplantation of isolated cells, and (b) implantation of an in vitro engineered tissue construct (adapted from: http://www.umm.edu/news/releases/myoblast.htm).

mentioned that a number of excellent reviews focusing on different aspects of cardiac tissue engineering, including cell-based therapies for myocardial regeneration, have been published recently (Eschenhagen, 2005; Leor et al., 2005; Gerecht-Nir et al., 2006; Zimmermann et al., 2006).

16.2 Cell sources The selection of cell sources for myocardial tissue engineering should be based on the studies of related cell-based approaches. A variety of cell models have been under intensive investigation with the hope of improving myocardial function. They can be categorised into three groups: (1) somatic muscle cells, such as fetal or neonatal cardiomyocytes and skeletal myoblasts; (2) myocardium-regenerating stem cells, such as embryonic stem cells and (possibly) bone marrow-derived mesenchymal stem cells; and (3) angiogenesis-stimulating cells, including fibroblasts and endothelial progenitor cells. Owing to their relevance, the first two cell types are briefly reviewed in the following sections.

16.2.1 Somatic muscle cells Fetal or neonatal cardiomyocytes Early cell transplantation studies focused on using fetal or neonatal rodent (rat or mouse) cardiomyocytes, as these cells have the inherent electrophysiological, structural and contractile properties of cardiomyocytes and still retain some proliferative capacity. In their pioneering study, Soonpaa et al. (1994)

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established the principles of cardiac cell implantation in the heart. They demonstrated that fetal or neonatal cardiomyocytes could be transplanted and integrated within both the healthy and dystrophic myocardium of mice and dogs, and that the surviving donor cells were aligned with recipient cells and formed cell-to-cell contacts. Similar results have later been reported by other research groups (Li et al., 1996; Reinecke et al., 1999; Murry et al., 2002). All results showed that early-stage cardiomyocytes (fetal and neonatal) were better candidates than more mature cardiac cells due to their superior in vivo survival capability (Reinecke et al., 1999). Several mechanisms have been proposed for improved heart function following cardiac myocyte transplantation (Kessler and Byrne, 1999; Etzion et al., 2001a): · direct contribution of the transplanted myocytes to contractility; · attenuation of infarct expansion by virtue of the elastic properties of cardiomyocytes; and · angiogenesis induced by growth factors secreted from the cells resulting in improved collateral flow. Skeletal myoblasts Theoretically, skeletal muscle cells may be superior to cardiomyocytes for infarct repair, because skeletal myoblasts have almost all the properties of the ideal donor cell type except their non-cardiac origin: (1) autologous sources, which obviate the need for immune suppression; (2) rapid expandability in an undifferentiated state in vitro; and (3) capability to withstand ischemia (Caspi and Gepstein, 2006; Etzion et al., 2001b). Although it was originally hoped that skeletal myoblasts would adopt a cardiac phenotype, it is now clear that in the heart cardiac phenotype myoblasts remain committed to form only mature skeletal muscle cells that possess completely different electromechanical properties from those of heart cells. However, studies in small and large animal models of infarction demonstrated beneficial effects of grafting of these cells on ventricular performance (Scorsin et al., 2000). Clinical trials with these cells were initiated some years ago but have now been discontinued, largely because of the pro-arrhythmic effect of skeletal myoblast transplantation.

16.2.2 Stem cell-derived myocytes In principle, stem cells are the optimal cell source for tissue regeneration, including myocardium. Firstly, they are capable of self-replication throughout life such that an unlimited number of stem cells of similar properties can be produced via expansion in vitro. Secondly, the stem cells are clonogenic, and thus each cell can form a colony in which all the cells are derived from this single cell and have identical genetic constitution. Thirdly, they are able to

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16.2 Types of stem cells in terms of potency.

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differentiate into one or more specific cell types. Hence, after expansion stem cells can be directed to differentiate into cardiomyogenic lineage. For these reasons, stem cell-based therapies for cardiac muscle regeneration have been under intensive research since 1995. Figure 16.2 shows types of stem cells in terms of potency. In this section, we briefly describe the applications of three types of stem cells in myocardial tissue engineering: bone marrow-derived stem cells (multipotent somatic or cord blood stem cells), native cardiac progenitor cells (multipotent somatic stem cells) and embryonic stem cells (pluripotent). Bone marrow-derived stem cells Bone marrow stem cells are the most primitive cells in the marrow. These cells can be classified into: (1) bone marrow-derived mesenchymal stem cells (MSC) and (2) haematopoietic stem cells (HSC) or called haematopoietic progenitor cells (HPC). Bone marrow-derived mesenchymal stem cells Bone marrow-derived mesenchymal stem cells are a subset of bone marrow stromal cells. This potential multipotent stem cell is derived from the nonhaematopoietic, stromal compartment of the bone marrow which can grow into non-marrow cells, such as bone, cartilage, adipose, endothelial, and myogenic cells. A number of studies suggested that bone marrow-derived MSCs could differentiate into cardiomyocytes both in vitro and in vivo (Orlic et al., 2001; Pittenger and Martin, 2004; Amado et al., 2005). One possible advantage of mesenchymal stem cells is their ability to be either autotransplanted or allotransplanted, as some reports suggested that they may be relatively privileged in terms of immune compatibility (Le Blanc, 2003). Haematopoietic stem cell It has been hypothesised that bone marrow stem cells might be able to differentiate into myocardium in vivo. Studies in the animal models of ischaemia and clinical trials suggested that delivery of haematopoietic stem cells and circulating endothelial progenitor cells (they originate from bone marrow stem cells) may result in improvement in the ventricular function in ischaemic heart disease patients (Caspi and Gepstein, 2006). However, Balsam et al. (2004) and Murry et al. (2004) demonstrated that the haematopoietic stem cells continued to differentiate along the haematopoietic lineage, suggesting the functional improvement observed may not be related to transdifferentiation into the cardiac lineage, but rather from indirect mechanisms. Clinical trials have been carried out with autologous bone marrow-derived cells. These are often unfractionated, and therefore contain an unquantified

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amount of the HSCs and MSCs. At present, the results of three medium size clinical trials (100±200 patients) show a variable and modest improvement in cardiac function (Assmus et al., 2006; Lunde et al., 2006; Schachinger et al., 2006). Native cardiac progenitor cells It had long been accepted that the adult mammalian heart, a terminally differentiated organ, had no self-renewal potential. This notion about the adult heart, however, has been challenged by accumulated evidence that myocardium itself contains a resident progenitor cell population capable of giving rise to new cardiomyocytes (Beltrami et al., 2003; Oh et al., 2003; Messina et al., 2004; Pfister et al., 2005; Laugwitz et al., 2005). The existence of these progenitor cells will no doubt open new opportunities for myocardial repair, though many issues still need to be addressed. Embryonic stem cells Embryonic stem cells are thought to have much greater translational potential than other stem cells because of their several advantages over other stem cells, since they are: (1) pluripotent, which means they have a broader multilineage expressing profile, (2) robust, which means they have the long-term proliferation ability with a normal karyotype and (3) genetically manipulable (Thomson et al., 1998). The first study using embryonic stem cells as a source for cell transplantation into the myocardium was reported by Klug et al. in 1996. By using genetically selected mouse embryonic stem cell-derived cardiomyocytes, it was shown that the differentiated cells developed myofibrils and gap junctions between adjacent cells and performed synchronous contractile activity in vitro. This study proved the feasibility of guiding an unlimited number of embryonic stem cells into cardiomyogenic cell linage and to utilise them for myocardial regeneration. Later studies (Menard et al., 2005), utilising the infarcted rat heart model, demonstrated that transplantation of differentiated mouse embryonic stem cellderived cardiomyocytes can result in short- and long-term improvement of myocardial performance. The vast biomedical potential of human ESCs has stirred enthusiasm in the field of tissue engineering. An overview of the utilisation of human ESCs in cardiac tissue engineering is given in Table 16.1. These studies demonstrated that the human ESC-derived cardiomyocytes displayed structural and functional properties of early-stage cardiomyocytes (Snir et al., 2003; Laflamme et al., 2005).

Table 16.1 Human ES cells in myocardial tissue engineering Method of hES differentiation

State before transplantation

Major result

Group and reference

In vitro; via EBs in suspension

NA

ES cells differentiated into cardiomyocytes, even after long-term culture. Upon differentiation, beating cells were observed after 1 week, increased in numbers with time, and retained contractility for >70 days. The beating cells expressed markers of cardiomyocytes.

Xu et al. (2002), Laflamme et al. (2005)

In vitro; via EBs in suspension

NA

ES cells showed consistence in phenotype with earlystage cardiomyocytes, and expression of several cardiac-specific genes and transcription factors.

Kehat et al. (2001)

In vitro; via EBs in suspension

NA

ES cells showed a progressive ultrastructural development from an irregular myofibrillar distribution to an organised sarcomeric pattern at late stages.

Snir et al. (2003)

In vitro; via EBs in suspension

NA

ESC-derived myocytes at mid-stage development demonstrated the stable presences of functional receptors and signalling pathways, and the presence of cardiac-specific action potentials and ionic currents.

Satin et al. (2004)

In vitro; via EBs in suspension

NA

ES cells differentiated into cardiomyocytes. Upon differentiation, beating cells were observed after 9 days, and retained contractility for longer than 6 months.

Harding et al. (2007)

Table 16.1 Continued Method of hES differentiation

State before transplantation

Major result

Group and reference

In vitro; via EBs in suspension

In vitro: coculture of differentiated rat cardiomyocyte and hESCM In vivo transplantation of hESCM into swine

Tight electrophysiological coupling between the engrafted human ESCms and rat cardiomyocytes was observed.

Kehat et al. (2004)

In vitro (coculture of undifferentiated hESC with mouse endoderm-like cells)

Differentiated

ESCs differentiated to beating muscle. Sarcomeric marker proteins, chronotropic responses, and ion channel expression and function were typical of cardiomyocytes. Electrophysiology demonstrated that most cells resembled human fetal ventricular cells.

Mummery et al. (2003)

In vitro via Ebs in suspension

In vitro co-culture of hESCM with rat myocytes

Electrically active, hESC-derived cardiomyocytes are capable of actively pacing quiescent, recipient, ventricular cardiomyocytes in vitro and ventricular myocardium in vivo

Xue et al. (2005)

Differentiated hESCM transplanted into guinea pig NA ˆ Not applicable

The transplanted hES cell-derived cardiomyocytes paced the hearts of swine.

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16.3 Construct-based strategies in myocardial tissue engineering In tissue engineering strategies based on the use of scaffolds, the regenerative ability of the host body should be increased through a designed construct that is populated with isolated cells and signalling molecules, aiming at regenerating functional tissue as an alternative to conventional organ transplantation and tissue reconstruction. In the following paragraphs, we review and discuss the design criteria and potential biomaterials used in myocardial tissue engineering approaches involving the use of scaffolds.

16.3.1 Design criteria of myocardial tissue engineering constructs One of the major challenges in tissue engineering is the design and fabrication of tissue-like materials to provide a scaffold or template for cells. An ideal scaffold should mimic extracellular matrix of the tissue that is to be engineered, which means that it must meet several stringent criteria. The specific requirements for myocardial tissue engineering scaffolds are as follows: · Ability to deliver cells. The material should not only be biocompatible, but also foster cell attachment, survival, differentiation and proliferation. · Biodegradability. The composition of the material, combined with the porous structure of the scaffold, should lead to biodegradation in vivo at a rate that matches the tissue regeneration rate. In other words, a synthetic scaffold should remain in the body for as short a period as possible; at the same time it must maintain its viability long enough for the cells to make their own matrix. · Mechanical properties. The principle in the mechanical design of the scaffold is that the scaffold should not interrupt the normal beating process of cardiomyocytes, while providing mechanical support for the cells to attach and to secrete their own matrix. · Porous structure. The scaffold should have an interconnected porous structure with porosity > 90% and diameters between 300±500 m for cell penetration, tissue ingrowth and vascularisation, and nutrient and waste transportation. · Commercialisation. The synthesis of the material and fabrication of the scaffold should be cost-effective, being suitable for commercialisation. Among these criteria, biocompatibility and cell-supporting and fostering ability are of highest importance for tissue engineering. So far, numerous biocompatible and biodegradable materials, including polymers, ceramics and composites, have been developed to support and foster cells, as described in other chapters in this book. Based on these available biomaterials, the foremost

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criteria to be considered are the mechanical properties and the porous structure of scaffolds, which are discussed in the following two sections. Mechanical design Ideally, a construct material should display mechanical and functional properties of native myocardium, such as coherent contractions, low diastolic tension and syncytial propagation of action potentials. Since the extracellular matrix (ECM) of myocytes is collagen, it is reasonable to hypothesise that cardiomyocytes would be able to beat adequately in a scaffolding material with mechanical properties similar (if impossible to be the same) to those of myocardial collagen. While natural collagens are an obvious option, their variable physical properties, including mechanical properties, with different sources of the protein matrices have hampered their application. Concerns have also arisen regarding immunogenic problems associated with the introduction of foreign collagen. As such, it is not surprising that much attention has been paid to synthetic polymers which have reproducible properties and are considered highly reliable materials for tissue engineering. The myocardium collagen matrix mainly consists of type I and III collagens, which form a structural continuum. Synthesised by cardiac fibroblasts, type I and III collagens have different physical properties. Type I collagen mainly provides rigidity, whereas type III collagen contributes to elasticity. The two types of collagens together support and tether myocytes to maintain their alignment, whereas their tensile strength and resilience resist the deformation, maintain the shape and thickness of the construct, prevent the rupture and contribute to the passive and active stiffness of the myocardium. The ratio of collagen types (type I/type III) within a healthy heart is typically about 0.5 (Pauschinger et al., 1999) and the stiffness of the collagen matrix of heart muscle is of several tens to hundreds kPa. Design of porous structures Tissue regeneration can be induced to take place in vitro. However, in the absence of true vascularisation, in vitro tissue engineering approaches face the problem of critical thickness: mass transportation into tissue is difficult beyond a thin peripheral layer of a tissue construct even with artificial means to supply engineered tissue constructs with nutrients and oxygen (<100 m without the help of a bioreactor, and <200 m with the support of a bioreactor) (Radisic and Vunjak-Novakovic, 2005). Diffusion barriers that are present in vitro are most likely to become more deleterious in vivo owing to lack of vascularisation. Once the engineered tissue construct is placed in the body, vascularisation becomes a key issue for further remodelling in the in vivo environment. It was thought that a pore size in the range of 50±100 m was sufficient to allow the vascularisation

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of a scaffold following transplantation (Shachar and Cohen, 2003). Recently, Radisic and Vunjak-Novakovic (2005) suggested a larger (100 m) pore size for cardiac tissue constructs. There are other authors who have suggested that pore sizes larger than 300 m are necessary for long-term survival, i.e. vascularisation (Temenoff et al., 2000). It is suggested that pore size in the range of 300± 500 m will not obstruct either vascularisation or mass transportation of nutrients and waste products, while the scaffold can maintain adequate mechanical integrity during in vitro culture and in vivo transplantation. In summary, a potentially advantageous material for myocardial tissue engineering scaffold should be a soft elastomer with stiffness of the order of several tens to hundreds kPa, and the scaffold should be made to exhibit pore sizes in the rage of 300±500 m.

16.3.2 Potential scaffolding biomaterials So far, a number of polymeric biomaterials have been developed or are under development for myocardial tissue engineering. An overview of biomaterials applied in myocardial tissue engineering is given in Table 16.2. In this section, the studies conducted on each of these polymers are discussed. Collagen gel matrix In 1997, Eschenhagen et al. (1997) reported, for the first time, an artificial heart tissue, which was termed engineered heart tissue (EHT) (Fig. 16.3). In this work, embryonic chick cardiomyocytes were mixed with collagen solution and allowed to gel between two Velcro-coated glass tubes. By culturing the cardiomyocytes in the collagen matrix, they produced a spontaneously and coherently contracting 3D heart tissue in vitro. Immunohistochemistry and electron Table 16.2 Overview of biomaterials used in myocardium tissue engineering Biomaterial

Physical state

Reference

Natural Collagen Collagen mesh (or sponge) Gelatine mesh Alginate mesh

Liquid/gel Solid Solid Solid

Zimmermann et al. (2000) Kofidis et al. (2003) Li et al. (1999) Leor et al. (2000)

Synthetic PGA and copolymer with PLA PLLA PCL and copolymer with PLA PGS, PGA

Solid Solid Solid Solid

Radisic and Vunjak-Novakovic (2005) Radisic and Vunjak-Novakovic (2005) Shin et al. (2004) Vunjak-Novakovic et al. (2006)

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16.3 Engineered heart tissue for regeneration of diseased hearts. (Reprinted from: Zimmermann WH, Melnychenko I, Eschenhagen T., Biomaterials, 25(9) 1639±47 (2004), with permission from Elsevier.)

microscopy revealed a highly organised myocardium-like structure exhibiting typical cross-striation, sarcomeric myofilaments, intercalated discs, desmosomes and tight junctions. More recently, large (thickness 1±4 mm and diameter 15 mm), mechanically supportive EHTs were produced with neonatal rat heart cells, and were implanted on hearts with myocardial infarction in immune-suppressed rats (Zimmermann et al., 2006). When evaluated 28 days later, EHTs showed several beneficial effects: · developing electrical coupling to the native myocardium without evidence of arrhythmia induction; · preventing further dilation; · inducing systolic wall thickening of infarcted myocardial segments; and · faster fractional area shortening of infarcted hearts, compared with controls (sham operation and non-contractile constructs). In summary, the research has confirmed that EHTs have many structural, functional and physiological characteristics of cardiac tissue, and that EHTs can be implanted to both healthy and infarcted hearts and can survive in vivo in both situations. Although this is important proof-of-concept work, there is no realistic possibility of human or rat neonatal cardiomyocytes coming to clinical application. The potential for expansion of this work to the construction of large implants with embryonic stem cell-derived cardiomyocytes remains to be established. Collagen fibrous mesh (or collagen sponge) The application of collagen gel matrix is limited by insufficient mechanical strength. This has led researchers to look for new approaches based on solid scaffolds. Kofidis et al. (2003) seeded neonatal rat cardiomyocytes in vitro into a

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3D solid collagen mesh and called the product artificial myocardial tissue (AMT). The artificial myocardial tissue was shown to possess structural, mechanical, physiological and biological characteristics similar to native cardiac tissue (Kofidis et al., 2003). More recently, Kofidis et al. (2005) utilised undifferentiated embryonic stem cells as the substrate of artificial myocardial tissue. The bioartificial mixtures were implanted in the infarcted area of the rat hearts. Studies revealed that embryonic stem cells formed stable intramyocardial grafts that were incorporated into the surrounding area without distorting myocardial geometry, thereby preventing ventricular wall thinning. In contrast to collagen gel, solid collagen mesh is too stiff to match host heart muscle. This drawback is in fact shared by all existing solid scaffolds discussed below. Gelatine mesh Gelatine mesh is the second type of solid scaffold applied in cardiac muscle engineering. Li et al. (1999) seeded fetal rat ventricular muscle (not isolated cells) into a gelatine foam to form grafts. The grafts were cultured in vitro for 7 days, forming a beating cardiac graft. The grafts implanted into the subcutaneous tissue contracted regularly and spontaneously. When implanted onto myocardial scar tissue, the cells within the grafts survived and formed junctions with the recipient heart cells. Alginate mesh In addition to protein-based materials, there is intensive activity in the area of natural polysaccharides. Alginate, a negatively charged polysaccharide from seaweed that forms hydrogels in the presence of calcium ions, was initially developed for drug delivery and it is now under development for tissue engineering scaffolds. The group of Cohen (Zmora et al., 2002) produced an alginate sponge using a freeze-drying technique, with porosity being 90% and pore size 50±150 m. Moreover, Leor et al. (2000) seeded fetal rat myocardial cells into this sponge to form an engineered heart construct. After 4-day culture in vitro, the engineered constructs were implanted into the rat hearts with myocardium infarct. Hearts were harvested 9 weeks after implantation. A large number of blood vessels were found in the grafting area, indicating intensive neovascularisation. The specimens showed almost complete disappearance of the scaffold and good integration into the host. In contrast to the control animals which developed significant left ventricular dilatation accompanied by progressive deterioration in left ventricular contractility, the graft-treated rats showed attenuation of left ventricular dilatation and unchanged contractility in left ventricle.

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Poly(glycolic acid) (PGA) and its copolymer with poly(lactic acid) (PLA) Synthetic biopolymers are thought to have a future in tissue engineering due to not only their excellent processing characteristics, which can ensure the off-theshelf availability, but also their advantage of being biocompatible and biodegradable. The biodegradable synthetic polymers most often utilised for 3D scaffolds in tissue engineering are the poly(-hydroxy acids), including poly(lactic acid) (PLA) and poly(glycolic acid) (PGA), as well as poly(lactic-coglycolide) (PLGA) copolymers. The first published work on a synthetic polymer-based scaffold designed for cardiac muscle engineering was by Freed and Vunjak-Novakovic in 1997. In this pioneering work, it was demonstrated that cultivation of primary neonatal rat cardiomyocytes on highly porous (porosity being 97%) PGA scaffolds in bioreactors could result in contractile 3D cardiac-like tissues, which consisted of cardiomyocytes with cardiac-specific structural and electrophysiological properties, contracting spontaneously and synchronously. In the subsequent studies, the group invested great efforts to overcome the limitation on the thickness of engineered tissue through improving cell-seeding and tissue culturing conditions (bioreactor). A maximal construct thickness of 1± 2 mm has been reported (Carrier et al., 2002). Poly(-caprolactone) (PCL) and its copolymer with PLA PCL is another important member of the aliphatic polyester family. It has been used to effectively entrap antibiotic drugs, and a construct made with PCL has been considered as a drug delivery system. The degradation of PCL and its copolymers involves similar mechanisms to PLA, proceeding in two stages: random hydrolytic ester cleavage and weight loss through the diffusion of oligomeric species from the bulk. More recently, Vacanti's group (Shin et al., 2004) has demonstrated the formation of contractile cardiac grafts in vitro using a nanofibrous PCL mesh. The nanofibrous mesh, which was produced by the electrospin technique, had an extracellular matrix-like topography. The average fibre diameter of the scaffold was about 250 nm, well below the size of an individual cardiomyocyte. After neonatal rat cardiomyocytes were seeded in the nanofibrous mesh, the construct was cultured, while being suspended across a wire ring that acted as a passive load to contracting cardiomyocytes. The cardiomyocytes started beating after 3 days and were cultured in vitro for 14 days. The cardiomyocytes attached well on the PCL meshes and expressed cardiac-specific proteins such as alphamyosin heavy chain, connexin43 and cardiac troponin I. This work indicated that using this technique, cardiac grafts can be matured in vitro to obtain sufficient function prior to implantation.

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Elastomers: poly(glycerol-sebacate) (PGS) and PET-DLA To engineer the tissue of heart, which beats cyclically and constantly throughout life, the biomaterial should be as soft and elastic as heart muscle. These mechanical characteristics are impossible with polyester-based thermosetting polymers, such as PGA, PLA, PCL and their copolymers, because they undergo plastic deformation and they are prone to failure when exposed to long-term cyclic strains. The limitation of their use in engineering elastomeric and flexible tissues has made biomaterial scientists turn to elastomers for cardiac tissue engineering. PGS was recently developed for the field of soft tissue engineering (Wang et al., 2002). This polymer is a biodegradable, biocompatible and inexpensive elastomer, and has already shown potential in nerve (Sundback et al., 2005) and vascular tissue engineering (Fidkowski et al., 2005). PGS also has superior mechanical properties, being capable of sustaining and recovering from deformation due to its intrinsic elasticity, and is thus suited to work in a mechanically dynamic environment, such as heart. So far, no systematic in vitro and in vivo studies have been reported, regarding applications of PGS in cardiac tissue engineering. Another elastomer being investigated for myocardial tissue engineering is a multiblock poly(aliphatic/aromatic-ester) containing phthalic acid sequences such as poly(ethylene terephthalate) (PET) and a dimmer fatty acid (DFA) (El Fray and Boccaccini, 2005). An initial in vitro assessment by the present authors using mouse ES cell-derived cardiomyocytes has shown that this material can support healthy, functional heart muscle cells (Fig. 16.4).

16.4 SEM image showing differentiating mouse ES cells seeded on PET-DLA substrate (authors' own results, PET-DLA material supplied by Prof. M. El Fray, Szczecin University of Technology, Poland).

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16.4 Conclusions and future trends Heart muscle engineering aims to regenerate functional myocardium to repair diseased and injured heart. Huge efforts have been invested in the development of cell sources for myocardial regeneration, including fetal cardiomyocytes, skeletal myoblasts, bone marrow stem cells, endothelial progenitors, native cardiac progenitor cells and embryonic stem cells. A number of biocompatible polymeric materials have also been investigated for cardiac regeneration strategies involving artificial constructs (patches). No matter what types of cells and biomaterials they are built from, the utilisation of artificial matrices to engineer myocardium has not been fully successful yet, with no examples of human application. The limitations are not only set by cell-related issues (such as scale-up in a rather short period, efficiency of cell seeding or cell survival rate and immune rejection), but also caused by the properties of the engineered tissue construct. First of all, engineered heart muscle must develop systolic (contractive) force with appropriate compliance; at the same time it must withstand diastolic (expansive) load. The material used to build the construct has no ability to beat without cells. The contractile movement of the engineered construct is completely driven by the seeded myocardial cells that inherently have a beating ability. One can envisage that the transfer of mechanical signals from cells to the scaffold would be jeopardised if the scaffold material is too stiff. Most of the above reviewed biomaterials, including collagen fibres, are much stronger than myocardium (Table 16.3 and Fig. 16.5). This explains why solid engineered constructs lack a contractile function. On the other hand, collagen gels are too weak to sustain the required mechanical loads. Another equally critical issue is size limitation, which is still a barrier in the field. Perfusion in vitro may improve the nutrient supply to augment the size of Table 16.3 Properties of potential biomaterials for cardiac muscle engineering Polymer

Elastomer (E) or thermoplastic (T)

Young's modulus (or stiffness)

Tensile strength

Degradation (month)

PGA PLLA or PDLLA PGS Collagen fibre (tendon/cartilage/ ligament/bone) Collagen gel (calf skin)

T T E E

7±10 GPa 1±4 GPa 0.282 MPa 2±46 MPa

70 MPa 30±80 MPa 0.5 MPa 1±7 MPa

2±12 2±12 Degradable Degradable

E

0.002±0.022 MPa

1±9 kPa

Degradable

Myocardium of rat Myocardium of human

E E

Maximum stiffness 0.14 MPa 30±70 kPa 0.2±0.5 MPa 3±15 kPa

NA NA

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16.5 Typical stress±strain curves of synthetic polymer (PGS) and heart muscles from different species.

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the engineered constructs. However, once they are implanted in vivo, the portion of thick tissue will need to be vascularised so that the cells within it will receive the necessary nutrients and waste products will be removed. Hence, prevascularisation is highlighted as a crucial step in the development of an engineered tissue construct. The use of novel materials to enhance cell growth and properties, to encourage vascularisation and to support the damaged heart, will be needed to bring this new technology into clinical application.

16.5 Acknowledgement We acknowledge financial support from BBSRC/EPSRC (Grant number BB/ D011027/1).

16.6 References and further reading Amado L C, Saliaris A P, Schuleri K H, St John M, Xie J S, Cattaneo S, Durand D J, Fitton T, Kuang J Q, Stewart G, Lehrke S, Baumgartner W W, Martin B J, Heldman A W and Hare J M (2005), `Cardiac repair with intramyocardial injection of allogeneic mesenchymal stem cells after myocardial infarction', Proc Natl Acad Sci USA, 102 (32), 11474±11479. Assmus B, Honold J, Schachinger V, Britten M B, Fischer-Rasokat U, Lehmann R, Teupe C, Pistorius K, Martin H, Abolmaali N D, Tonn T, Dimmeler S and Zeiher A M (2006), `Transcoronary transplantation of progenitor cells after myocardial infarction', N Engl J Med, 355 (12), 1222±1232. Balsam L B, Wagers A J, Christensen J L, Kofidis T, Weissman I L and Robbins R C (2004), `Haematopoietic stem cells adopt mature haematopoietic fates in ischaemic myocardium', Nature, 428 (6983), 668±673. Beltrami A P, Barlucchi L, Torella D, Baker M, Limana F, Chimenti S, Kasahara H, Rota M, Musso E, Urbanek K, Leri A, Kajstura J, Nadal-Ginard B and Anversa P (2003), `Adult cardiac stem cells are multipotent and support myocardial regeneration', Cell, 114 (6), 763±776. Carrier R L, Rupnick M, Langer R, Schoen F J, Freed L E and Vunjak-Novakovic G (2002), `Perfusion improves tissue architecture of engineered cardiac muscle', Tiss Eng, 8 (2), 175±188. Caspi O and Gepstein L (2006), `Stem cells for myocardial repair', Eur Heart J Suppl, 8 (Supplement E), E43±E54. Chen M K and Beierle E A (2004), `Animal models for intestinal tissue engineering', Biomaterials, 25 (9), 1675±1681. Cima L G, Vacanti J P, Vacanti C, Ingber D, Mooney D and Langer R (1991), `Tissue engineering by cell transplantation using degradable polymer substrates', J Biomech Eng ± Trans ASME, 113 (2), 143±151. El Fray M and Boccaccini A R (2005), `Novel hybrid PET/DFA-TiO2 nanocomposites by in situ polycondensation', Mater Lett, 59 (18), 2300±2304. Eschenhagen T (2005), `Engineering myocardial tissue', Circ Res, 97 (12), 1220±1231. Eschenhagen T, Fink C, Remmers U, Scholz H, Wattchow J, Weil J, Zimmerman W, Dohmen H H, Schafer H, Bishopric N, Wakatsuki T and Elson E L (1997), `Threedimensional reconstitution of embryonic cardiomyocytes in a collagen matrix: a

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