Journal of Cardiovascular Translational Research

, 4:592

Stem Cell-Based Cardiac Tissue Engineering


    • Institute of Biomaterials and Biomedical EngineeringUniversity of Toronto
    • Toronto General Research InstituteUniversity Health Network, MaRS Centre
  • Hannah Song
    • Institute of Biomaterials and Biomedical EngineeringUniversity of Toronto
  • C. Katherine Chiang
    • Institute of Biomaterials and Biomedical EngineeringUniversity of Toronto
    • Institute of Biomaterials and Biomedical EngineeringUniversity of Toronto
    • Department of Chemical Engineering and Applied ChemistryUniversity of Toronto
    • Heart and Stroke/Richard Lewar Centre of ExcellenceUniversity of Toronto

DOI: 10.1007/s12265-011-9307-x

Cite this article as:
Nunes, S.S., Song, H., Chiang, C.K. et al. J. of Cardiovasc. Trans. Res. (2011) 4: 592. doi:10.1007/s12265-011-9307-x


Cardiovascular diseases are the leading cause of death worldwide, and cell-based therapies represent a potential cure for patients with cardiac diseases such as myocardial infarction, heart failure, and congenital heart diseases. Towards this goal, cardiac tissue engineering is now being investigated as an approach to support cell-based therapies and enhance their efficacy. This review focuses on the latest research in cardiac tissue engineering based on the use of embryonic, induced pluripotent, or adult stem cells. We describe different strategies such as direct injection of cells and/or biomaterials as well as direct replacement therapies with tissue mimics. In this regard, the latest research has shown promising results demonstrating the improvement of cardiac function with different strategies. It is clear from recent studies that the most important consideration to be addressed by new therapeutic strategies is long-term functional improvement. For this goal to be realized, novel and efficient methods of cell delivery are required that enable high cell retention, followed by electrical integration and mechanical coupling of the injected cells or the engineered tissue to the host myocardium.


Cardiac tissue engineeringStem cellsCell therapyRegenerative medicineInduced pluripotent stem cellsEmbryonic stem cells


Myocardial infarction (MI) occurs when one or more of the blood vessels supplying the heart are occluded. The blockage of coronary arteries leads to a sudden decrease in the supply of nutrients and oxygen to the portion of heart muscle supplied by the artery. When blood supply is not reestablished in time, irreversible cell death within the affected part of the heart muscle will occur. Such injuries to the myocardium result in the formation of scar tissue that does not have contractile, mechanical, or electrical properties of normal myocardium. By consequence, cardiac output is reduced, decreasing the ability of the heart to supply blood efficiently. The deterioration of heart function will lead to heart failure, and at the end-stages, the only treatment options are heart transplants and ventricular assist devices [1, 2]. Therefore, new therapies are required to prevent the progression of pathological remodeling and cell death, as well as to induce tissue recovery in the affected areas.

Regenerative medicine aims to achieve this goal through the restoration of tissue structure and organ function, thus delaying or preventing disease progression [3]. In the case of therapies for heart disease, the desired outcome would be maintenance of normal ventricular function and anatomy [4], replacement of cardiomyocytes lost post-MI, prevention of left ventricular wall thinning, and improvement in overall cardiac output towards physiological levels (average of 5.6 L/min for human males and 4.9 L/min for human females) [5]. Therefore, cell-based therapies represent a potential cure for patients with cardiac diseases such as MI, heart failure, and congenital heart diseases. Towards this goal, cardiac tissue engineering is now being investigated as an approach to support cell-based therapies and enhance their efficacy. In a classical approach, cardiac tissue engineering involves the integrated use of cells, biomaterials, and bioreactors with the purpose of generating a contractile myocardium [6].

Cardiac Tissue-Engineering Strategies

Direct Cell Injection Strategies

Since MI may result in loss of up to one billion cardiomyocytes in the infarct zone, the idea of regenerating myocardium by cell injection has emerged [7] and was tested in a number of notable in vivo studies over the past 20 years. Although direct cell injection into the heart does not result in a piece of beating cardiac tissue such as in conventional in vitro cardiac tissue engineering approaches, we classify it here as tissue engineering since it enables cardiac regeneration through application of cells and occasionally biomaterials. Initially, it was thought that injection of beating cells is required for restoration of the function. However, a large body of evidence suggests functional improvements even with the injection of non-contractile cells such as bone marrow cells or endothelial progenitor cells [8]. Of note, while not the focus of this review, others have demonstrated efficacy of injecting biomaterials alone [9], such as decellularized matrices [10], in improving cardiac function.

In clinical trials, non-immunogenic autologous bone marrow-derived mononuclear cells or progenitors expressing particular surface markers such as CD34 or CD133 have been assessed [8]. In animal studies, more diverse cell populations have been examined including those derived from pluripotent stem cells. In this section, we review animal studies that satisfy the following criteria: (1) used animal MI model, (2) showed functional cardiac measurements at 4 weeks or longer post cell injection, and (3) injected cells with or without injectable biomaterials. To compare the outcomes from different studies, we investigated 12 parameters focusing on experimental conditions and methods. These parameters, summarized in Table 1, include: (1) types of animals used for the study, (2) methods of inducing MI, (3) time of cell injection post-MI, (4) type and number of cells injected, (5) characteristics of cells at the time of injection, (6) delivery vehicles used for cell injection, (7) time of cardiac functional measurements, (8) results of cardiac functional measurements, (9) integration between host and injected cells, (10) characteristics of injected cells at the end time point, and (11) changes in infarct size at the end time point. Also, the studies are divided into two groups depending on whether any biomaterial was injected or not.
Table 1

In vivo cell injection studies in myocardial infarction model








End time

Cardiac function



Infarc size


Cell injection

SD rat (M)

Comp. LAD ligation

7–10 days

hESC (H9.2)-derived CM

1.5 × 106

71% TnI + (ICC)


4 and 8

FS ↑, LVDd ↓ (echo)

CX-43 (IHC)

α-actinin + (IHC), no teratoma



Lewis rat (F)

Comp. LAD1 ligation


Amnion-derived cell

2 × 106


2 and 6

LVDA ↓ (echo)


Some TnI + (IHC)

Wall thickness ↑


C57BL/6 or AN mice (M)

Comp. LAD ligation


Fibroblast or iPS

0.2 × 106



EF↑ , LVDd ↓ (echo) wrt. fibroblast


α-Actinin + (IHC), no teratoma

Heart Size ↓


SVE129 mice (F)

Comp. LAD ligation


D3 mESC-derived Nkx2.5 + CPC

0.5–1 × 106




LVDd↓, EF ↑ (echo)

CX-43 (IHC)


Infarct size ↓ (IHC)


SCID mice

Comp. LAD ligation


Tet Notch mESC- derived CM

0.5 × 106

40% TnT+


2 and 8

EF↑ at 2 week (MRI)


TnT + (IHC)




Comp. LAD ligation


hES3-GFP-derived CM

3 × 106



No improvement (MRI)




Comp. LAD ligation


Human fetal sca-1+ and -derived CM

0.5 × 106

70% TnI+ (ICC)





5% Retention, 50% TnI + from both CP and CM (IHC)

Wall thickness ↑ (MRI)


Biomaterial with or without cells

SD rats (M)

Comp. LCx ligation

2 weeks


HA-based hydrogel


EF↑ (P-V catheter)


Infarct size ↓


Athymic SD rats (m)

Temp. LAD ligation

4 weeks

hESC (H7)-derived CM

1 × 107

50–65% β-MHC+

PSC + GFR-matrigel

8, 16, and 32

No difference

Grafts were separated by scar tissue (IHC)



AN rats

Temp. LAD ligation

4 days

hESC (H7)-derived CM

1 × 107


PSC + matrigel


LVEDD ↓, EF ↑(echo, MRI)

Cadherin (IHC)

TnI + (IHC)



SD rats (F)

Comp. LAD ligation

1 week


1 × 107




ES/D D↓, EF↑ (echo)


Mixed TnT + and TnT-cells

Infarct size ↓


MI myocardial infarction, TI time of injection post-MI, NUM injected cell number, DV delivery vehicle, END TIME time of cardiac functional measurement (weeks), PRE-CHAR cell characterization prior to injection, POST-CHAR cell characterization at the end of the study, SD Sprague–Dawley rats, Lewis Rat LEW/Crl rat, AN athymic nude, SCID severe combined immunodeficiency, M male, F female, Comp. LAD ligation complete left anterior descending artery ligation, Temp. LAD ligation temporary left anterior descending artery ligation, Comp. LCx ligation complete left circumferential artery ligation, CM cardiomyocyte, hESC human embryonic stem cell, mESC mouse embryonic stem cell, TnI troponin I, ICC immunocytochemistry, IHC immunohistochemistry, FS fractional shortening, LVDd left ventricular diastolic diameter, LVDA left ventricular diastolic area, Echo echocardiography, SVE1

There were profound differences in the type of animal MI model and the cell injection time point, which rendered direct comparisons of the results difficult. Among the studies reviewed here, another significant difference was that all the studies with biomaterial injections (with or without cells) were performed many days after an MI using rat model, as opposed to those with cell injections, where most report injection of cells immediately post-MI using mouse model. This might be due to difficulties in producing large amounts of cells for injection; however, the difference in cell injection time point between these two animal models makes comparison of the outcome more difficult.

From all of the studies that we examined, there was only one study that investigated all the parameters listed above [11]. When the studies in Table 1 are compared, the number of injected cells varied by as much as 100-fold. Furthermore, a broad variety of different cell types were tested, including human embryonic stem cells (hESC), mouse embryonic stem cell (mESC)-derived cardiac progenitors and cardiomyocytes [1116] (Fig. 1), undifferentiated cells including induced pluripotent stem (iPS) cells [17], mESC [18] or amnion-derived cells [19], and fetal cardiac resident stem cells [20]. Low retention rates are commonly reported, regardless of the cell type injected. The cell retention at 12 weeks was quantified to be only 5% of injected cells in one of the studies [20]. Therefore, the idea of injecting cells with biocompatible materials to improve cell retention and survival has been emphasized [14, 15, 18, 21]. Most of the studies reviewed here demonstrated cardiac functional improvement such as increased ejection fraction or decreased left ventricular systolic or diastolic diameter compared with MI-only controls at 4 weeks. However, one study described no functional improvement at any of the time points tested up to 3 months [14]. The authors attribute this discrepancy to the fact that they use a model of chronic infarction since in a previous study these same cells were beneficial in an acute model of MI [15].
Fig. 1

Histological evaluation of human myocardial grafts at 4 weeks. Directly differentiated hES cell-derived cardiomyocytes were heat-shocked and injected into 4-day-old infarcts in athymic rats, using the pro-survival cocktail consisting of growth factor-reduced Matrigel supplemented with 100 μM, benzyloxycarbonyl-Val-Ala-Asp(O-methyl)-fluoromethyl ketone, 50 nM of Bcl-XL BH4, 200 nM cyclosporin A, 100 ng/ml IGF-1, and 50 μM pinacidil to enhance survival. ac Depict brightfield microscopic images from recipient hearts 4 weeks post-transplantation, whereas those in di were acquired with a three-laser confocal microscope. a Combined human pan-centromeric in situ hybridization and β-myosin heavy-chain immunostain. The implanted cells have formed a large graft of β-myosin-positive cardiomyocytes (βMHC, red stain) within the infarct scar tissue, and their human origin was confirmed by nuclear staining with the human-specific pan-centromeric in situ probe (huCent, brown chromagen). The left-ventricular cavity is at the upper right corner. Spared host subendocardial myocardium is present but is negative for β-myosin. Fast green counterstain. Scale bar, 100 μm. b High magnification of boxed region in a. All of the β-myosin-positive cells (red stain) have nuclear staining with the pan-centromeric probe (huCent, brown-black stain), whereas rat nuclei in the surrounding scar or myocardial tissue are unstained. c Hematoxylin and eosin stain. A serial section to the boxed region in a is shown. The graft cells have a vacuolated appearance due to the presence of glycogen. Note the numerous nuclei in the rat scar tissue and spared myocardium, which were negative for the human pan-centromeric probe in the contiguous section presented in b. Scale bar, 50 μm. d Colocalization of human pan-centromeric in situ hybridization and β-myosin heavy chain. Co-localization of β-myosin heavy-chain immunofluorescence (red signal) and the human-specific pan-centromeric probe (green) in the human cardiac grafts was confirmed by confocal microscopy. Note that the immediately adjacent rodent host myocardium shows minimal expression of β-myosin heavy chain. Scale bar, 100 μm; inset shows the corresponding boxed area magnified twofold. e Graft and host myosin heavy chain expression pattern. This section was double-immunostained for sarcomeric myosin heavy chain (sMHC all striated muscle; red) and β-myosin heavy chain (human cardiac muscle; green). The human myocardial graft is clearly identified by the dual staining for sarcomeric and β-myosin, which appears yellow in this merged image. The surviving subendocardial rat myocardium is identified by the red staining for sarcomeric myosin. The surrounding infarct scar tissue is unstained. Scale bar, 100 μm. f Host–graft contact. Infrequent but unequivocal sites of contact between host and graft myocardium were observed, as is illustrated by this section double-immunostained for β-myosin heavy chain (red) and cadherins (green). Points of close apposition between engrafted human myocardium (strongly immunoreactive for β-myosin heavy chain) and host muscle (minimally reactive) show shared adherens junctions, a component of intercalated disks, as indicated by cadherin staining (asterisks). Arrowheads indicate adherens junctions among the adjacent host cardiomyocytes. Scale bar, 10 μm. g Cardiac troponin I expression and sarcomeric organization of graft cells. This section was double-immunostained for cardiac troponin I (green) and β-myosin heavy chain (red). Note that, in this merged image, the graft cells are positive for both markers and appear yellow, whereas the adjacent host cardiomyocytes only stain for cardiac troponin I (note green cells in lower right-hand corner). The graft muscle shows definite areas of sarcomeric organization (note boxed area, magnified twofold in inset). Scale bar, 20 μm. h Combined Nkx2.5 and β-myosin heavy chain immunostain. This section was double-immunostained for β-myosin heavy chain (human cardiac muscle; green) and the cardiac-specific transcription factor Nkx2.5 (human-specific antibody; red nuclear signal). The human myocardial graft in the middle of the field is readily distinguishable by its strong positive reaction for both markers. Scale bar, 50 μm. i Combined human pan-centromeric in situ hybridization and myosin light chain 2 V immunostain. As demonstrated by this typical field, both host and graft (human-specific pan-centromeric positive, green nuclei) myocardium showed immunoreactivity for the ventricular-specific marker myosin light chain 2 V (red). Scale bar, 20 μm (adapted from [15])

Some studies investigated functional integration between the injected cells and the host cardiomyocytes by immunostaining for cell junctional molecules such as connexin 43 or cadherin (Table 1). However, none of the studies showed any significant integration. The study conducted by Fernandes et al. [14] indicated that grafts were separated by scar tissue from the host myocardium, which may be a result of a natural process occurring with foreign cell injections. Current studies indicate that injecting biomaterials with or without cells has a distinct beneficial effect on decreasing infarction size and improving wall thickness. However, the mechanism responsible for functional improvements as a result of injection of biomaterials alone is yet to be elucidated. One possible explanation is that the injection of biomaterials results in mechanical stabilization of the ventricle irrespective of effects on cellular function or survival [22].

To facilitate further comparison between different studies, it appears necessary to standardize the experimental conditions and MI models. We believe that MI models where injections are performed at a later time point after the induction of MI are more clinically relevant since they likely reflect more closely the human MI scenario than models in which MI is induced and cells are injected right away. In addition, long-term studies maybe necessary to ensure that the functional improvement is not transient. Also, the generation of new in vitro systems which allows for the systematical testing of many important parameters such as functional integration between injected and host cells would be beneficial.

In vitro cardiac test systems have been developed previously [23, 24]. Such systems can account for some of the heterogeneities inherent in the in vivo system (Fig. 2) and can function as a preliminary analysis of cell integration and electrical coupling. Engineered heart tissue, which is cultivated in a cardiac mimetic environment using an electrical stimulation chamber, forms a basis of this model system (Fig. 2). This system has been shown to successfully detect integration between injected cells and the in vitro “host” myocardium using optical mapping (Fig. 2d). Improving these systems will help us to improve therapeutic cell injection strategies in regenerative medicine.
Fig. 2

Engineered heart tissue (EHT) for in vitro cell injection studies. a Schematic diagram of experimental set up and integration between injected and host cells. b Isochronal activation maps representing the electrical impulse propagation through the EHT, and a red arrow indicating the electrode for the point stimulation. c Electrical tracing of spontaneous action potentials. d Entrainment of action potential-induced changes in intracellular calcium using rhod-2 fluorescence in YFP-expressing and non-expressing cardiomyocytes within EHTs. a Frame (left) and line-scan mode image (middle) taken from an EHT injected with YFP-expressing neonatal cardiomyocytes (nCMs). The line-scan image was obtained by repeatedly scanning along the white line in left and stacking the lines vertically. Right shows the spatially averaged changes in rhod-2 fluorescence as a function of time for each cell along the white scan line. Changes in rhod-2 fluorescence occur synchronously in YFP+ (green) and YFP-myocytes (red), indicating that both cell types are functionally coupled. b Identical results were obtained when images were taken during spontaneous activity of an EHT injected with ESC-CPs. (modified from [23] and [76])

Direct Tissue Replacement Strategies

The fact that some improvement in function is observed regardless of the cell type injected and that cell retention is very small, indicates that improvement in cardiac function is likely due to the secretion of paracrine factors by the injected cells [25, 26] resulting in an increased preservation of affected myocardium in a transient manner as opposed to transdifferentiation and direct integration of contractile cells [26, 27]. In addition, the unlikely possibility that injected cells would, on their own, recapitulate the complex tissue organization present in cardiac muscle where cardiomyocytes, vascular, and stromal cells are positioned together in an intricate organization has increased the drive for new direct replacement tissue-engineering strategies. The objective of cardiac tissue engineering for direct replacement is to generate tissue constructs or mimics that can functionally replace damaged myocardium. This requires the use of biocompatible and/or biodegradable materials that serve as a scaffold for tissue mimics [28]. In order to generate cardiac tissue mimics, a large number of cells are needed to ensure proper tissue function and integration when implanted in the affected areas.

In adults, the ability to repair damaged cardiac tissue is hindered by the reduced proliferative capacity of mature cardiomyocytes [29]. For true myocardial regeneration, both the beating myocardium and the functional vasculature need to be regenerated in the infarct zone. Pluripotent stem cells, such as ESC and iPS cells, hold the potential to differentiate into all cell types in the body [30] and have the potential to be propagated in vitro for long periods of time. By consequence, these cells are a useful tool to generate large numbers of cells that hold the potential to not only exert functional benefits by secretion of paracrine factors but also to enhance cardiac function through differentiation and direct integration into cardiac tissue. However, safety issues involving differentiation into non-cardiac cell types or even formation of teratomas [26], hinder direct clinical translation at this time. Therefore, another possibly safer source of autologous cardiovascular cells is adult stem cell such as resident cardiac progenitors. However, such cells are usually present in very low numbers in adults.

Embryonic Stem Cells

Many different techniques have been used to assemble cardiac tissue mimics, from scaffold-free self assembled cells [31, 32] to the use of different biocompatible and/or biodegradable materials as scaffolds (for a review, see [33]). Grafting of mESCs cultured in polyglycolic-acid biodegradable scaffolds into infarcted mouse myocardium significantly improved animal survival, blood pressure, and ventricular function [34]. Authors also report the presence of implanted cells in the infarcted area suggesting cell retention and possible myocardium repair [34]. Other studies utilizing injection of undifferentiated ESC reported the formation of teratomas suggesting that this approach is not clinically relevant [35].

In order to generate biomechanical support and cell delivery to the heart, Chen et al. [36] generated hybrid cardiac patches made with poly(glycerol sebacate) and supplemented with hESC-derived cardiomyocytes. These patches sustained cell beating for long periods in culture and, when sutured over the left ventricle of normal rats, remained intact without deleterious effect on ventricular function, suggesting that these patches could function as support devices for cardiac repair. In addition, the delivery of stem cells in poly(glycerol sebacate) could potentially be more effective by utilizing an accordion-like honeycomb microstructure since Engelmayr et al. have demonstrated that these scaffolds, microfabricated to mimic cardiac muscle mechanical properties, promote seeded heart cell alignment [37].

Engineered cardiac tissue (ECT) has also been generated by seeding mouse embryonic stem cell-derived cardiomyocytes into collagen type I supplemented with Matrigel [38]. After in vitro stretching for 7 days, authors demonstrated that the ECT can beat synchronously and respond to physical and pharmaceutical stimulation. In addition, no signs of tumorigenesis were found after subcutaneous implantation for 4 weeks. Cultivation of mESC-derived cardiomyocytes on elastic poly(lactide-co-caprolactone) scaffolds exposed to cyclic stretch, followed by implantation into infarcted hearts showed reduced fibrotic tissue formation and upregulation of cardiac gene expression as compared with unstrained controls [39]. Cyclic stretch was also described to upregulate expression of sarcomeric cardiac genes and to improve cell alignment and distribution of connexin-43, a gap junction protein, highlighting the importance of mechanical strain transduction in cardiac tissue engineering [39].

A different, complementary approach consists in the preservation of cardiac tissue by rescuing the vascular network compromised during myocardial infarction. Many authors have now described the potential to generate vascular-related cells such as endothelial and perivascular cells from hESCs [4042]. Injection of hESC-derived vascular cells in a bioactive hydrogel as an in situ forming scaffold after myocardial infarction in rats has shown that the delivered cells formed capillaries in the infarct zone [43]. In addition, magnetic resonance imaging revealed that the microvascular grafts effectively preserved contractile performance, attenuated left ventricular dilation, and decreased infarct size [43].

In a more complex approach, Caspi et al. [44] have engineered vascularized cardiac muscle using hESC-derived cardiomyocytes and hESC-derived endothelial cells seeded in biodegradable, biocompatible, and Food and Drug Administration (FDA) approved materials. Analysis of the engineered tissues indicated that increased cardiomyocyte and endothelial cell proliferation as well as formation of vessel-like structures occurred when these cells were cultured with mouse embryonic fibroblasts. When implanted into an uninjured rat heart, these tri-culture scaffolds displayed the formation of viable grafts with both human and host-derived patent vasculature within the implants [45]. A similar tri-culture approach was utilized by Stevens et al. [31] who reported that patches containing only cardiomyocytes do not form substantial grafts in vivo demonstrating the importance of both vascular and stromal elements in increasing survival and integration of the engineered cardiac tissue. However, neither group assessed the presence of residual undifferentiated cell activity in the implanted engineered tissues.

A promising cell type for use in cardiac tissue-engineering strategies would be iPS cells due to their documented ability to give rise to functional cardiomyocytes [46]. The use of these cells would not only overcome the ethical concerns related to the use of hESCs, but it might also allow for the generation of an unlimited supply of functional, proliferative, and possibly autologous human cardiomyocytes and vascular cells thus overcoming any immunogenic concerns as well. However, iPS cells hold the similar safety issues involving differentiation into non-cardiac cell types or teratomas.

Adult Stem Cells

While there are still ethical and immunogenic concerns related to the use of hESCs in tissue engineering with therapeutic purposes, the use of autologous adult cells as a source overcomes those issues. Adult stem cells with the potential to self-renew and to differentiate into specific lineages exist in different organs. Some examples include hematopoietic and mesenchymal cells of bone marrow. Some reports have implicated adult bone marrow cells in myocardial regeneration [4749] as well as functional improvement in infarcted hearts [50]. The current data indicate that bone marrow cells improve cardiac function by a paracrine mechanism dependent on the secretion of soluble factors and not trans-differentiation [28]. Nonetheless, the contradictory results obtained with bone marrow transplantation in patients with infarcted myocardium [51] and issues regarding the possible formation of bone and cartilage from these cells still need to be addressed before these cells can be safely used for cardiac therapy.

Another possible source of cells for cardiac engineering is resident cardiac stem cells. Cardiac progenitor cells expressing stem cell antigen-1 (Sca-1) have been reported in adult mouse myocardium [52] and have been implicated in cardiac homing and differentiation after infarction. Others report the existence of Lin(−) c-kit(+) cells that, when injected into an ischemic heart, reconstitute well-differentiated myocardium [53] and resident Isl1+ cardiac progenitors that have been shown to hold the potential to differentiate into cardiomyocytes [54]. However, their real potential for cardiovascular tissue engineering is still unknown given the fact that these cells are present in very low numbers.

Towards this goal, Domian et al. [55] employed a two-colored fluorescent reporter system (eGFP driven by a cardiac-specific Nkx-2.5 enhancer and dsRed driven by an Isl1-specific enhancer of the Mef2c gene) to isolate first and second heart field progenitors from mouse embryos and embryonic stem cells and to generate beating two-dimensional cardiac tissue. Different populations were isolated according with their expression profile of the above-mentioned fluorescent reporters by fluorescence-activated cell sorting. Authors show that the eGFP+/dsRed + Isl1 population was most similar to the myogenic population. When seeded on glass surfaces coated with poly(dimethyl siloxane) stamped with micropatterned fibronectin lanes, the double positive cells elongated in the direction of the patterns, expressed sarcomeric alpha-actinin, and formed muscular thin films (MTF) which generated contractile force comparable with that of neonatal rat ventricular cardiomyocytes [55]. Importantly, the MTF displayed spontaneous beating and could also be paced by field stimulation.

Adult cardiac progenitors can also be isolated from explant cultures of human endomyocardial biopsies and expanded in vitro as self-adherent clusters or cardiospheres (CSps) [56, 57]. These cells displayed both paracrine and direct regeneration effects in infarcted mice [58]. They were shown to secrete vascular endothelial growth factor, hepatocyte growth factor and insulin-like growth factor-1 when transplanted into a mouse model of myocardial infarction accounting for a paracrine activity that decreased apoptotic rates and caspase 3 levels while increasing capillary density [58]. In addition, based on the number of human-specific cells relative to overall capillary density and myocardial viability, direct differentiation accounted for 20% to 50% of the observed effects. The same group also described that magnetic targeting of iron-labeled cardiosphere-derived cells enhanced engraftment and functional benefits [59]. In order to do that, CSps-derived cells were labeled with superparamagnetic microspheres and guided to the heart by imposing an external magnetic field on the heart during and immediately after cell injection. With this approach, injected cells accumulated around the ischemic zone while the non-targeted cells washed out immediately after injection. Quantitative polymerase chain reaction analysis confirmed cell retention (24 h post-delivery) and engraftment (3 weeks post-delivery) in the recipient hearts by threefold compared with non-targeted cells [59]. Maximal attenuation of left ventricular remodeling and greatest functional improvement occurred in the cell-targeted group without incremental inflammation [59].

Challenges and Future Studies

Embryonic and induced pluripotent cell-based therapies are still in their infancy with respect to clinical translation. To date, only one embryonic stem cell-based therapy for spinal cord injury has been approved for clinical trial [60]. The use of embryonic stem cells for cardiac regeneration is currently in preclinical animal trials [15], although there have been several other cell-based therapies in clinical trials involving adult-derived progenitor cells or stem cells for cardiac repair [6168]. The risks associated with embryonic stem cell-derived therapies come from the possibility that even one undifferentiated cell, if present, could potentially result in the formation of teratoma. While researchers continue to elucidate the mechanisms of novel therapies that have shown great promise in animal studies, clinicians are more concerned with the unmet clinical needs of their patients [64].

The success of a potential therapy for cardiac regeneration not only depends on the ability to provide the necessary physiological improvements but also on the ability to achieve clinical implementation. Seger and Lee [26] outlined a few considerations that should be carefully examined when investigating a potential stem cell-based therapy for cardiac regeneration. First and foremost come the issues of cell source and isolation. While ESC can reliably give rise to large numbers of cardiomyocytes, potential immunogenicity of differentiated progeny and safety issues related to the residual undifferentiated cell activity remain. iPSC represent a potential autologous cell source, however, the time required for reprogramming and differentiation into cardiomyocytes, coupled with the low cardiomyocyte yield of current protocols [46] represent significant challenges for potential clinical translation. Viral incorporation of reprogramming genes into the host genome and the complexity of epigenetic modifications involved in reprogramming represent another challenge [6971].

Most cell injection studies have reported significant drawbacks with respect to the cell retention, viability, and distribution [26, 72] of the injected cells. Recent advances in tissue-engineered constructs have shown great promise in providing the necessary support for cell retention and distribution [7375]. However, issues related to the survival of cells in thick constructs remain. Functional coupling with the host myocardium is yet to be demonstrated for cardiac tissues based on human pluripotent stem cell-derived cardiomyocytes. The work of Zimmermann and Eschenhagen in the rat system clearly demonstrated the functional integration of the engineered tissue with the host myocardium [75]. In addition, although direct cell/biomaterial injection can be performed in a minimally invasive catheter-based route, implantation of engineered cardiac tissues requires open heart surgery thus it would be limited to a small number of patients. Issues related to storage and transport of these human-engineered cardiac tissues would also emerge and affect clinical availability. The tissue engineering bioreactor would likely assume the roles of the cultivation, storage, and transport vessel, thus motivating further progress in the bioprocess and bioreactor field.

Safety is perhaps the most important consideration when translating this technology to the clinics. If pluripotent stem cells are to be used in future therapies, rigorous purification of desired cell populations are required. New in vitro assays that test the residual undifferentiated cell activity and the stability of the differentiated progeny are required. Towards this goal, we have demonstrated that the in vitro engineered heart tissue can be a useful model for cell injection studies [76]. We have evaluated the regenerative potential of injected pluripotent stem cell-derived cardiomyocytes (ESC-CMs) as well as pluripotent stem-cell-derived Flk1+/PDGFα + cardiac progenitors (ESC-CPs). We used EHTs as surrogate heart tissues and studied their ability to integrate with injected ESC-CMs, ESC-CPs, neonatal cardiomyocytes, or neonatal cardiac fibroblasts. Functional and phenotypic analyses (Fig. 2) revealed that ESC-CPs improved excitation threshold, maximum capture rate, impulse propagation, and expression of cardiac markers such as cardiac troponin T and connexin while the other cell types exhibited limited to no functional integration with the surrogate EHT [76]. Additionally, EHT was instrumental in enabling identification of residual undifferentiated cell activity in mouse ESC-derived populations [23]. Injection of undifferentiated R1 mouse ESC into EHT leads to teratoma formation and formation of structures characteristics of all three germ layers consistent with the in vivo studies [35]. The long-term cultivation (4 weeks) of Flk1+/PDGFα + cardiac progenitors in the absence of cardiac inductive cytokines resulted in the formation of non-cardiac mesodermal structures such as bone and cartilage, reinforcing the importance of high purity selection of differentiated progeny and maintenance of cardiac inductive cytokines. By consequence, studies using mouse ESCs also represent tools to investigate the importance of biomaterials and the influence of different scaffolds in the assembly of cardiac tissue in vitro.

We have also developed a cardiac tissue based on neonatal rat heart cells that mimics some aspects of diabetic myocardium, by modulation of glucose and insulin concentrations during cultivation. Our results indicate that the diabetic rat heart and high glucose cultivation conditions exhibited diminishing electrophysiological properties and increased ratio of myosin heavy chain isoforms β to α, indicative of diseased states [77]. Our current studies involve miniaturization of the EHT for high-throughput studies and development of disease specific models.


While it is clear that progress has been made in cardiac tissue engineering and regenerative medicine, additional investigation is required in order to reach a consensus on the best biomaterials and cell types to be used as well as the optimal time point for cell injection or tissue implantation. This would be facilitated by the standardization of experimental conditions and would require the systematic analysis of many cell types and biomaterials. From the studies described above, we can conclude that the most important consideration to be addressed by new therapeutic strategies is long-term benefit that includes stable functional improvements and attenuation of pathological remodeling. For these goals to be achieved, efficient delivery and survival of the implanted cells and engineered tissues is required. Since human post-natal cardiomyocytes have limited ability to proliferate, the current lack of an autologous non-immunogenic source of cardiomyocyte has limited the progress of cardiac tissue engineering towards clinical applications. Functional cardiac tissues have been engineered based on hESC derived cardiomyocytes and the lessons learned from these studies pave the way to tissue engineering of cardiac patches based on human iPSC.


Financial support for our work is provided by a Natural Sciences and Engineering Research Council of Canada (NSERC) Discovery Grant (RGPIN 326982–10), Discovery Accelerator Supplement (RGPAS 396125–10), NSERC Strategic Grant (STPGP 381002–09), NSERC-Canadian Institutes of Health Research Collaborative Health Research Grant (CHRPJ 385981–10), and Heart and Stroke Foundation of Ontario Grant-in-Aid (T6946).

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