Myocardial tissue engineering: a review

www.elsevier.com/locate/mser

Available online at www.sciencedirect.com

Materials Science and Engineering R 59 (2008) 1–37

Biomaterials in cardiac tissue engineering:

Ten years of research survey

Qi-Zhi Chen a,b,*, Sian E. Harding b, Nadire N. Ali b,Alexander R. Lyon b, Aldo R. Boccaccini a,*

a Department of Materials, Imperial College London, Prince Consort Road, London SW7 2AZ, UKb National Heart and Lung Institute, Imperial College London, Dovehouse Street, London SW3 6LY, UK

Received 24 June 2007; received in revised form 14 August 2007; accepted 14 August 2007

Available online 10 January 2008

Abstract

Driven by enormous clinical need, myocardial tissue engineering has become a prime focus of research within the field of tissue engineering.

Myocardial tissue engineering combines isolated functional cardiomyocytes and a biodegradable or nondegradable biomaterial to repair diseased

heart muscle. The challenges in heart muscle engineering include cell related issues (such as scale up in a short timeframe, efficiency of cell seeding

or cell survival rate, and immune rejection), the design and fabrication of myocardial tissue engineering substrates, and the engineering of tissue

constructs in vitro and in vivo. Several approaches have been put forward, and a number of models combining various polymeric biomaterials, cell

sources and bioreactors have been developed in the last 10 years for myocardial tissue engineering. This review provides a comprehensive update

on the biomaterials, as well as cells and biomimetic systems, used in the engineering of the cardiac muscle. The article is organized as follows. A

historic perspective of the evolution of cardiac medicine and emergence of cardiac tissue engineering is presented in the first section. Following a

review on the cells used in myocardial tissue engineering (second section), the third section presents a review on biomaterials used in myocardial

tissue engineering. This section starts with an overview of the development of tissue engineering substrates and goes on to discuss the selection of

biomaterials and design of solid and porous substrates. Then the applications of a variety of biomaterials used in different approaches of myocardial

tissue engineering are reviewed in great detail, and related issues and topics that remain challenges for the future progress of the field are identified

at the end of each subsection. This is followed by a brief review on the development of bioreactors (fourth section), which is an important

achievement in the field of myocardial tissue engineering, and which is also related to the biomaterials developed. At the end of this article, the

major achievements and remaining challenges are summarized, and the most promising paradigm for the future of heart muscle tissue engineering

is proposed (fifth section).

# 2007 Elsevier B.V. All rights reserved.

Keywords: Tissue engineering; Cardiac muscle; Biomaterials; Polymers; Cell therapy; Bioreactors

Contents

1. Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2

2. Cells applied in myocardial tissue engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5

2.1. Somatic muscle cells . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6

2.1.1. Foetal or neonatal cardiomyocyte . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6

2.1.2. Skeletal myoblast . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7

2.2. Angiogenic cells. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7

2.2.1. Fibroblasts. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7

2.2.2. Endothelial progenitor cells . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7

2.3. Stem cell-derived myocytes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8

2.3.1. Basics of stem cells . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8

* Corresponding authors. Tel.: +44 20 75946723; fax: +44 20 75946757.

E-mail addresses: [email protected] (Q.-Z. Chen), [email protected] (A.R. Boccaccini).

0927-796X/$ – see front matter # 2007 Elsevier B.V. All rights reserved.

doi:10.1016/j.mser.2007.08.001

mailto:[email protected]

mailto:[email protected]

http://dx.doi.org/10.1016/j.mser.2007.08.001

Q.-Z. Chen et al. / Materials Science and Engineering R 59 (2008) 1–372

2.3.2. Bone marrow-derived stem cells. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 9

2.3.3. Adipose-derived stem cells . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 9

2.3.4. Native cardiac progenitor cells . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 9

2.3.5. Embryonic stem cells . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 10

2.3.6. Human embryonic stem cells . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 11

2.4. Strategies to address immune rejection in cells . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 11

2.5. Strategies of cell delivery . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 12

2.6. Summary of cell-based therapy and their limitations. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 12

3. Biomaterials for myocardial tissue engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 13

3.1. Overview of substrate development for tissue engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 13

3.1.1. Criteria on tissue engineering substrates . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 13

3.1.2. Polymers used in soft tissue engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 13

3.1.3. Fabrication of tissue engineering substrates . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 16

3.2. Biomaterials for myocardial tissue engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 18

3.2.1. Selection of biomaterials and design of substrates . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 18

3.2.2. Biomaterials used in myocardial tissue engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 21

4. Biomimetic tissue engineering—bioreactors . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 28

4.1. Brief history of bioreactors . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 28

4.2. Comparison of bioreactors . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 29

4.3. Application of bioreactors in myocardial tissue engineering. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 29

4.4. Key issues . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 30

5. Summary . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 30

Acknowledgements . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 31

References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 31

1. Introduction

A historic perspective is presented in this section on the

evolution of cardiac medicine and the emergence of cardiac

tissue engineering as a major branch in the field of tissue

engineering. It is aimed to understand the rationale of cardiac

tissue engineering and related strategies. A general retrospect

on the history of tissue engineering has been given recently by

Vacanti [1].

Heart disease is the leading cause of death and disability in

both industrialised nations and the developing world, account-

ing for approximately 40% of all human mortality [2]. It is

estimated that 5 million Americans, 1.8 million Britons, and 25

million people worldwide suffer from heart failure, with

approximately 550,000 and 120,000 new cases diagnosed each

year in the United States (US) and the United Kingdom (UK),

respectively [3]. Prognosis is poor with 40% mortality within

12 months of diagnosis, and a 10% annual mortality rate

thereafter [4]. The economic burden imposed by this disease

has reached more than $33 billion in the US and more than £700

million in the UK annually [3].

Heart failure is a condition reflecting impairment of the

pumping efficiency of the heart, and it is caused by a variety of

underlying diseases, including ischemic heart disease with or

without an episode of acute myocardial infarction, hypertensive

heart disease, valvular heart disease, and primary myocardial

disease. The single most common cause of left-sided cardiac

failure is ischemic heart disease (also called coronary artery

disease) with an episode of acute myocardial infarction.

Myocardial infarction typically results in myocyte slippage.

The weakening of the collagen extracellular matrix results in

heart wall thinning and ventricular dilation. The impairment of

the heart wall muscle is permanent because, after a massive cell

loss due to infarction, the myocardial tissue lacks significant

intrinsic regenerative capability to replace the lost cells [5]. The

enlargement in ventricular volume leads to progressive

structural and functional changes in ventricles (called

ventricular remodelling) [5]. Ventricular remodelling is

initially compensatory, but adds further inefficiency to the

mechanical pumping of the ventricular muscle, predisposing

towards the end stage of congestive heart failure (CHF) (or just

heart failure) [5], a condition in which the heart cannot pump a

sufficient amount of blood to the meet the metabolic

requirements of the body [6].

Pharmacological therapy focuses on reduction of work load

(utilising diuretics, nitrates) and protection from the toxic

humoral factors which are overactivated in heart failure [7].

These include catecholamines (b-blockers), angiotensin-con-

verting enzyme (ACE inhibitors), and aldosterone (spirono-

lactone). Blockade of these humoral factors represents the

current standard conservative treatment for patients with mild

symptoms of heart failure and slight limitation during ordinary

activity [7]. Interventional therapy, such as surgery or

implantation of pacing devices to control electrical/mechanical

asynchrony, are now receiving more widespread application, in

particular for patients with marked symptoms and marked

limitation in activity [7–12]. However, both drug and

interventional therapies cannot adequately control disease

progression to the end stage [13]. Eventually, heart transplanta-

tion is the ultimate treatment option to end-stage heart failure.

Owing to the lack of organ donors and complications associated

with immune suppressive treatments, however, scientists and

surgeons constantly look for new strategies to repair the injured

heart [14].

Fig. 1. Heart–lung machine (http://heartonline.org/postpump.htm).

Q.-Z. Chen et al. / Materials Science and Engineering R 59 (2008) 1–37 3

Historically, all these surgical strategies started with the

development of the heart–lung machine (Fig. 1), which takes

over the functions of the heart and lungs during an open-heart

surgery [15], such as coronary bypass surgery and valve

replacement [16,17]. Cardiomyoplasty was an alternative

surgical approach for treating heart failure. Pre-prepared

skeletal muscle, which is able to function at power levels

analogous to those of the heart, was wrapped around the heart,

and paced to contract with the heart, thereby improving cardiac

pumping power [18,19]. Clinical studies reported that this

dynamic cardiomyoplasty could improve left ventricular

performance, reduce cardiac dilation, and interrupt disease

progression [20,21]. However, quantitative heamodynamic

analyses were not consistent, regarding the benefits of active

systolic assist, and mortality from the operation was

unacceptably high [22,23]. This prompted the suggestion that

passive mechanical constraint by the muscle wrap might halt or

even reverse the negative remodelling of the dilated ailing

heart. Inspired by this hypothesis, many studies have examined

Fig. 2. (a) Cardiac support device by Acorn CorCapTM, a typical approach of left ven

structure of the mesh [18]. Published with kind permission from Springer Science

the use of biomaterial supports to restrain the left ventricle [24].

Marlex mesh (polypropylene) [25], Merselene mesh (knitted

polyester) (26), Acorn CorCapTM heart mesh (knitted

polyester) (18) and MyocorTM Myosplint1 [18] are four

representative cardiac support devices that have been under

investigation. A typical approach of this strategy is illustrated in

Fig. 2. Although animal cardiomyoplasty showed distinct

benefits of the devices [27], these have not been translated to the

clinical setting, and currently evidence for their clinical benefit

is absent. None of these devices has received approval of the

Food and Drug Administration (FDA) [28].

Studies around the mid-1990s veered to an intriguing

strategy: the application of cell transplantation. Initially, it was

confirmed that diseased myocardium could be restored by the

transplantation of functional cardiac myocytes [29–31]. Since

then a number of research groups reproduced and refined

these pioneering experiments [32–45]. Most studies support the

conclusion that cell implantation in models of myocardial

infarction can improve contractile function. Clinical studies are

currently under way to investigate the safety and feasibility of

cell implantation in patients [39]. In the cell-based therapy,

isolated cells are injected to the infarct region via the

pericardium, coronary arteries, or endocardium.

In order to improve the site accuracy of cell delivery, an

alternative approach to deliver cells to the infarct region is to

rebuild 3D cell networks in vitro and to implant the cell bandage

onto the infarct heart, as shown in Fig. 3 [46].

The third approach involves the usage a man-made heart

patch (Fig. 4) [47–49], which is populated in vitro with cells and

implanted later in vivo. In this approach, the ring- or sheet-

shaped heart patch serves two functions: cell delivery and

mechanical support. According to theoretical simulations on

the effects of injected materials (Fig. 4b), the addition of a sheet

material to a damaged left ventricular wall could have

important effects on cardiac mechanics, with potentially

beneficial reduction of elevated myofibril stresses, as well as

tricle restraint. http://www.sciencedaily.com/releases/2004/11/. (b) The knitted

and Business Media.

http://www.cbmm.lodz.pl/res/sppb_big2.jpg


Fig. 3. Schematic graph showing the transplantation of a myocardial cell-sheet graft, also called heart bandage [46]. Reproduced by permission of the MRS Bulletin.


histological and functional changes to clinical left ventricular

metrics [50]. No simulation work has been reported yet, as far

as the authors are aware of, regarding the effects of epicardial

patches (Fig. 4a). The results may be similar in both cases.

A more ambitious strategy is the implantation of myocardial

tissue generated ex vivo, i.e. a 3D tissue regeneration strategy.

This classic strategy of tissue engineering is established on the

fact that living bodies have the potential of regeneration, and on

the supposition that the employment of natural biology (e.g.,

cells and biomolecules) will maximise the capacity for

regeneration and allow for greater success in developing

therapeutic strategies aimed at the replacement and repair of

tissue and the maintenance and enhancement of its function

[51–53]. The successful development of tissue engineering

constructs will have a profound impact in both the scientific

community and the public sphere. They could be used to

produce in vitro healthy cells for cell-based therapy. They may

also be used for many biomedical studies, such as cell biology,

organ development, functional cell differentiation from stem

cells, environment–cell interaction, cancer biology, new drug

treatment, and could ultimately be used for the repair of injured

or diseased tissues.

Fig. 4. Schematic illustrations of (a) epicardial and (b) endoventricular heart

patch approaches to deliver isolated cells to the infarct regions.

In essence, tissue engineering is a technique of imitating

nature. Natural tissues consist of three components: cells,

extracellular matrix (ECM), and signalling systems. The ECM

is made up of a complex of cell secretions immobilised in

spaces and thus forming a scaffold for its cells. Hence, it is

natural that the engineered tissue construct is a triad [54], the

three constitutes of which correspond to the above-mentioned

three basic components of natural tissues. Fig. 5 illustrates the

triad, i.e. a scaffold, living cells and signal molecules (such as

growth factors and cytokines).

The European Commission on Health and Consumer

Protection defined tissue engineering as ‘‘the persuasion of

the body to heal itself through the delivery, to the appropriate

site, independently or in synergy, of cells, biomolecules and

supporting structures’’ [53]. According to this definition, the

surgical approaches described above could be all classified into

the tissue engineering category, as listed in Table 1. It must be

mentioned that that these approaches are not independent of

one another. The heart patch approach, for example, is a

combinatory paradigm of passive diastolic constraint and cell

therapy. The classic 3D tissue engineering construction could

be optimised by the incorporation of drug and gene therapies.

The application of biomaterials has been mainly related with

the last four approaches listed in Table 1, i.e. left ventricular

constraint, scaffold-free cell sheet implantation, heart patch

implantation and 3D tissue engineering construction.

Cardiac tissue engineering covers heart valve, cardiovas-

cular and myocardial tissue engineering. This review focuses

on the myocardial tissue engineering. Recent reviews are

available on heart valve [55–58] and cardiovascular tissue

engineering [56,59–62].

In the past 10 years, many studies have been published using

different cells and different biomaterials for heart muscle

engineering [14,63–65], and excellent reviews focusing on

different aspects of cardiac muscle engineering are also

available [14,24,52,63–79]. In this article, we present a

Fig. 5. Triad of a classic tissue engineering construct.

Table 1

Currently applied or potential strategies for the treatment of heart failure

patients

1. Pharmaceutical therapy

2. Interventional therapy

(1) Reduction of the heart volume

(2) Implantation of a pace-maker

3. Heart transplantation

4. Tissue engineering strategy

(1) Cardiomyoplasty (active systolic assist)

(2) Cell-based therapy (isolated cell-delivery)

(3) Left ventricular restraint (passive diastolic constraint)

(4) Scaffold-free cell-sheet implantation

(5) Heart patch implantation (passive diastolic

constraint and cell delivery)

(6) 3D tissue engineering construction

(a scaffold + cells + macromolecules)


comprehensive review on the achievements of myocardial

tissue engineering, including cells, biomaterials, and biomi-

metic approach (i.e. bioreactors). Our intention is to identify

what we consider significant challenges and the most promising

approaches in the future of heart muscle tissue engineering. At

the same time, we are aware of the possibility that our own

biases might be embedded in the opinions expressed in the

review. We also apologise in advance to any individuals whose

significant effort in the field of heart tissue engineering was

neglected due to our misunderstanding or oversight.

2. Cells applied in myocardial tissue engineering

This section is devoted to provide a concise review on cells

for heart tissue engineering, including key studies. Although it

is not biomaterial specific, this knowledge is essential to

biomaterials scientists in this field, as cell implantation plays a

pivotal role in the engineering of the heart muscle that lacks

significant intrinsic regenerative capability. After reading this

section, material scientists hopefully will have a guide to

browse the huge amounts of reports available on a variety of

cell types applied in cardiac muscle regeneration and

engineering.

As early as in 1978, Bader and Oberpriller demonstrated the

regenerative capacity of amphibian hearts after autologous

implantation of minced ventricular tissue samples (i.e.

nonisolated cells) into injured newt hearts [80]. In this study,

a partial regeneration of injured newt ventricles was observed.

However, grafted tissue fragments remained morphologically


and functionally separated from the native myocardium. True

tissue engineering approaches emerged in early 1990s when

efforts to regenerate functional myocardial tissue were invested

in grafting of isolated cell [30,31]. Since then, numerous basic

studies on cells and several early-stage clinical trials have been

carried out using a variety of cell types with the hope of

improving myocardial function. So far, a variety of cell models

have been under intensive investigation. They can be

categorised into three groups (Table 2): (1) somatic muscle

cells, such as foetal or neonatal cardiomyocytes [42,44,45,81–

87] and skeletal myoblasts [88–92], (2) myocardium-generat-

ing cells, such as embryonic stem cells [93–97] (possibly) bone

marrow-derived mesenchymal stem cells [98–106] and adipose

stem cells [107]; and (3) angiogenesis-stimulating cells,

including fibroblasts [108] and endothelial progenitor cells.

Each of these cell types and cell delivery approaches are

reviewed in more detail in the following sections.

2.1. Somatic muscle cells

2.1.1. Foetal or neonatal cardiomyocyte

Early cell transplantation studies focused on using foetal or

neonatal rodent (rat or mouse) cardiomyocytes, as these cells

have the inherent electrophysiological, structural and contrac-

tile properties of cardiomyocytes and still retain some

proliferate capacity [31,81,87]. In their pioneering study,

Soonpaa et al. established the principles of cardiac cell

implantation in the heart. They demonstrated that foetal

cardiomyocyte could be transplanted and integrated within the

healthy myocardium of mice, and that the surviving donor cells

were aligned with recipient cells and formed cell-to-cell

contacts [31]. This group also reported the foetal cardiomyo-

cyte graft in the myocardium of dystrophic mice and dogs [87].

More studies have demonstrated that cardiac myocytes from

neonatal, embryonic or adult models can also be engrafted into

diseased (infarcted, cryoinjured or cardiomyopathic) hearts

Table 2

Potential cell sources for myocardial regeneration in human and their advantages

Cell source Autologous Easily

obtainable

Highly

expan

Somatic cells

Foetal cardiomyocytes No No No

Skeletal myoblasts Yes Yes Depen

Smooth muscle cells Yes Yes Yes

Fibroblasts Yes Yes Yes

Stem cells

Somatic stem cells

Mesenchymal stem cells Yes No Depen

Endothelial progenitor cells Yes Yes Depen

Crude bone marrow Yes Yes Depen

Umbilical cord cells No Yes Yes

(Hemaetopoietic stem cells)

Adipose stem cells

Yes Yes Yes

Embryonic stem cells

Human embryonic stem cells No No Yes

[30,42,44,45,81–83,85,86,109]. These results also indicate that

early-stage cardiomyocytes (foetal and neonatal) were better

candidates than more mature cardiac cells due to their superior

in vivo survival [42].

Time course studies for the survival of grafted cardiomyo-

cytes in the healthy heart were carried out by Muller-Ehmsen

et al. [34]. They isolated and injected male donor neonatal rat

cardiomyocytes into the left ventricular (LV) wall of adult

female inbred rats. They demonstrated that these cells could

survive and improve cardiac function for up to 6 months in a rat

model of chronic myocardial infarction [33]. Murry’s group

[110] showed, using syngeneic rat with cryoinjury, that cell

graft survival after 7 days could be up to 33%.

Cardiomyocyte transplantation, which was applied to

smaller infarcts [83], has been proved effective in the

prevention of cardiac dilation and remodelling following

infarction [36] and the improvement of the ventricular function

[44,85]. Several mechanisms have been proposed for improved

heart function following cardiac myocyte transplantation

[36,70,111,112]:

(1) d

and

dable

d on

d on

d on

d on

irect contribution of the transplanted myocytes to

contractility;

(2) a
ttenuation of infarct expansion by virtue of the elastic
properties of cardiomyocytes;

(3) a
ngiogenesis induced by growth factors secreted from the
foetal cells resulting in improved collateral flow;

(4) p
aracrine effects via the release of beneficial growth factors
from the transplanted cells, which support the cardiomyo-

cytes under strain in the failing heart, and may possibly

recruit residential cardiac progenitor cells.

However, the transplanted tissue decreased in size several

months after transplantation [42,83]. An electron microscopy

study revealed that dead cells had features of both necrosis and

apoptosis. Based on the most recent experiments, it is apparent

disadvantages for myocardial repair [63]

Cardiac

myogenesis

Clinical

trial

Safety

Yes No No

age Debated Yes Yes, arrhythmias

No No No

No No No

age Yes No Yes, fibrosis calcification

age Debated No Yes, calcification

age Debated Yes Yes, calcification

Debated No No

Yes No Yes

Yes No Yes, potential teratoma

if cells escape differentiation


that cell death is rapid and extensive after cardiomyocyte

grafting, with most cell deaths occurring during the first 2 days.

However, after 1 week the graft is relatively stable [34,110].

Although this is important proof-of-concept work, there is no

realistic possibility of human or rat neonatal cardiomyocytes

coming to clinical application [110,113,114]. As a result,

several alternative approaches have been developed to over-

come the limitations of foetal cardiomyocyte transplantation

and to obviate the need for immunosuppressants.

2.1.2. Skeletal myoblast

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. Skeletal myoblast satellite cells can be

harvested from autologous sources, which obviate the need for

immune suppression. Satellite cells are mononuclear progenitor

cells found in mature muscle. In undamaged muscle, the

majority of satellite cells are quiescent. Upon muscle damage,

satellite cells become activated and are able to differentiate and

fuse to augment existing muscle fibres and to form new fibres.

They can be rapidly expanded in an undifferentiated state in

vitro to clinically applicable numbers of myoblasts without a

risk for tumourgenecity, and they have the capabilities to

withstand ischemia better than many other cell types.

Continued proliferation in vivo may be an advantage when

engrafting into an injured heart, since the input of a smaller

number of cells might give rise to a large graft [70,115].

Although it was originally hoped that skeletal myoblasts

would adapt a cardiac phenotype, it is now clear that within

heart tissue the skeletal myoblasts remain committed to form

only mature skeletal muscle cells that possess completely

different electromechanical properties than those of heart cells.

Moreover, given the inability of myoblasts to form electro-

mechanical connections with host cardiomyocytes (due to lack

of expression of adhesion and gap junction proteins), it is not

surprising that physiological studies failed to demonstrate

synchronous beating of the grafted cells within the host tissue

[116].

However, studies in small and large animal models of

infarction demonstrated beneficial effects of grafting of these

cells on ventricular performance [117,118]. The mechanisms

underlying the beneficial effects of skeletal myoblasts remain to

be elucidated. The improvement in heart wall motion could be

achieved by contraction of the transplanted cells, a local effect

on scar remodelling by mechanical support and/or paracrine

influences on the remodelling process.

Nevertheless, given their autologous origin, the capacity to

amplify primary myoblasts from human muscle biopsies, and

the encouraging preclinical results, skeletal myoblasts were the

first cell type to reach clinical application [92,119–126]. The

pioneering Phase I clinical trials were performed either using a

direct surgery approach (during coronary artery bypass graft

surgery) or using a percutaneous endocardial catheter-delivery

approach and have demonstrated both the feasibility of the

procedure and the ability of the cells to engraft in the infarcted

myocardium [92,119–121,123–128].

The clinical application of autologous skeletal myoblasts is

currently limited by several concerns [115,129,130].

(1) L
ack of myocardial phenotype. This has been blamed for the
disturbingly high incidence of life-threatening ventricular

arrhythmias noted in the initial post transplantation phase in

these trials [119,120].

(2) L
ow recovery of satellite cells. The recovery of satellite
cells from muscle biopsies of elderly patients is low.

(3) E
fficiency. Grafting methods need be developed to improve
the efficiency of cell engraftment and survival.

(4) Q
uestionable beneficial effects. The efficacy of skeletal
myoblast therapy is still uncertain because of the following

facts. Firstly, the Phase I clinical trials were not randomised,

and might be placebo controlled. Secondly, widely different

cell transplantation protocols were used in a relatively small

number of patients. Finally, the trials were associated with

other confounding factors such as concomitant left ventricle

assistant device (LVAD) implantation or revascularisation.

Ongoing Phase II clinical trials will hopefully address these

concerns and thoroughly evaluate the safety and efficacy of

myoblast transplantation.

2.2. Angiogenic cells

2.2.1. Fibroblasts

Vascularisation is a key step in tissue repair. At sites of

injury, pheno-transformed fibroblast-like cells are responsible

for fibrous tissue formation. These cells are termed myofibro-

blasts because they contain alpha-smooth muscle actin

microfilaments and are contractile. In vivo studies of injured

rat cardiac tissues and in vitro cell culture studies [131] have

shown that such fibroblast-like cells contain requisite compo-

nents for angiotensin peptide generation and angiotensin II

receptors. Such locally generated angiotensin II acts in an

autocrine/paracrine manner to regulate collagen turnover and

thereby tissue homeostasis in injured tissue.

Human dermal fibroblasts have been applied for myocardial

regeneration to stimulate revascularisation and preserve left

ventricular (LV) function of the infarcted LV in mice [132]. It

has been shown that dermal fibroblasts functioned to attenuate

further loss of LV function accompanying acute myocardial

infarct and that this might be related in part to myocardial

revascularisation.

2.2.2. Endothelial progenitor cells

Endothelial progenitor cells (EPCs) are present in the bone

marrow and the peripheral blood and exhibit phenotypical

markers of mature endothelial cells [133]. It has been found that

rats with inflammatory-mediated cardiomyopathy exhibited a

significant mobilization of EPCs from the bone marrow to the

periphery and their ability to adhere to fibronectin, mature

endothelial cells and cultured cardiomyocytes was significantly

reduced when compared to healthy rats [134]. This result

prompted studies in the application of EPCs in attenuating

remodelling followed by acute myocardial infarction. Transfer

of EPCs resulted in a functional improvement in cardiac


performance. EPC transfer is effective in attenuating myo-

cardial damage in a model of non-ischemic dilated cardiomyo-

pathy [134], and probably exert their beneficial effects via new

vessel growth and improved blood supply to the failing heart.

2.3. Stem cell-derived myocytes

2.3.1. Basics of stem cells

Stem cells are primal undifferentiated, and thus unspecia-

lized, cells that retain the ability to differentiate into multiple

cell lineages [135]. 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 differentiate into

one or more specialised cell types. Hence, after expansion stem

cells can be directed to differentiate into cardiomyogenic

lineage [94,136–138]. For these reasons, stem cell-based

therapy for cardiac muscle regeneration has been under

intensive research during the last decade.

Stem cells can be categorised according to their potency

(totipotent, pluripotent, multipotent, and unipotent) (Fig. 6),

anatomic source (adult, embryonic, foetus, cord blood, or

cancer), or by cell surface markers, transcription factors, and

proteins they express. Four basic stem cell types classified

according to their potency are briefly introduced as follows [135].

(1) T
otipotent stem cells are produced from the fusion of an egg
and sperm cell. These cells can differentiate into embryonic

and extra-embryonic cell types.

(2) P
luripotent stem cells are the descendants of totipotent cells
and can differentiate into cells derived from the three germ

layers.

(3) M
ultipotent stem cells can produce only cells of a closely
related family of cells (e.g. haematopoietic stem cells

differentiate into red blood cells, white blood cells,

platelets, etc.).

(4) U
nipotent cells can produce only one cell type, but have the
property of self-renewal which distinguishes them from

non-stem cells.

There are five sources for stem cells, as described below

[139].

� A
dult stem cells are undifferentiated cells found among
differentiated cells of a specific tissue and are mostly multi-

potent cells. They are more accurately called somatic stem

cells because foetal and umbilical cord stem cells also fall into

this category. They are present in all tissues and seem to survive

long time periods and harsh conditions. Important sources of

somatic stem cells include bone marrow-derived stem cells,

such as mesenchymal stem cells, and adipose stem cells.

� E

Fig. 6. Categories of stem cells in terms of potency [140].

mbryonic stem cells include early embryonic stem cells

(totipotent) and blastocyst embryonic stem cells (pluripotent)

(Fig. 6). Blastocyst embryonic stem cells are cultured cells

obtained from the undifferentiated inner cell mass of an early

stage pre-implantation embryo. Unlike somatic stem cells,

embryonic stem cells can differentiate into any one of the

body’s more than 200 cell types.

� F
oetal stem cells. After eighth week of development, the
human embryo is referred to as a foetus. By this time it has

developed a human-like form. Stem cells in the foetus are

responsible for the initial development of all tissues before

birth. Like embryonic stem cells, foetal stem cells are

pluripotent.

� C
ord blood stem cells are derived from the blood of the
placenta and umbilical cord after birth. Umbilical cord stem

cells are multipotent.

� C
ancer stem cells arising through malignant transformation
of adult stem cells are proposed to be the source of some or all

tumours and cause metastasis and relapse of the disease.

The common features of all stem cells include:

(1) A
bility to self renew, which means they can divide and
produce stem cell progeny with similar properties.


(2) C
lonogenic, which means that each cell can form a colony
in which all the cells are derived from this single cell and

have identical genetic constitution.

(3) B
road differentiation profile, which means they can grow
into one or more mature cell types.

In this section, we focus on the applications of four types of

stem cells in myocardial tissue engineering: bone marrow-

derived stem cells, adipose stem cells, native cardiac progenitor

cells, and embryonic stem cells.

2.3.2. 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 cell (MSC), and (2) haematopoietic

stem cell (HSC).

2.3.2.1. Bone marrow-derived mesenchymal stem cells. Bone

marrow-derived mesenchymal stem cells are a subset of bone

marrow stromal cells (the term ‘‘mesenchymal stem cell’’ is

now used to include multipotent cells that are derived from

either bone marrow or other tissues, such as adult muscle or the

Wharton’s jelly present in the umbilical cord). This potential

multipotent stem cell is derived from the non-haematopoietic,

stromal compartment of the bone marrow, which can grow into

non-marrow cells, such as bone, cartilage, tendon, adipose, and

endothelial cells [141].

A number of studies suggested that bone marrow-derived

MSCs could differentiate into cardiomyocytes both in vitro and

in vivo [38,40,136,137,142–145]. Makino et al. [137] treated

murine mesenchymal stem cells with 5-azacytidine and isolated

a cardiomyogenic cell line after repeated screening of

spontaneous beating cells. This result was confirmed by

Tomita et al. [146]. Later Orlic et al. [38] reported that a

subpopulation of bone marrow stem cells were capable of

generating myocardium in vivo in mice. More recently, it was

reported that transplantation of mesenchymal stem cells into

the infarcted myocardium of rats and pigs resulted in improved

myocardial performance [144,147].

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 [148].

2.3.2.2. Haematopoietic stem cells. In addition to the initial

hypothesis that bone marrow stem cells might be able to

differentiate into myocardium in vivo, another major rationale

behind the research of bone marrow stem cells for cardiac

muscle regeneration was the essential roles of vascularisation

and angiogenesis in tissue regeneration. Studies in the animal

models of ischemia and Phases I and II clinical trials suggested

that delivery of haematopoietic stem cells and circulating

endothelial progenitor cells, both originating from bone

marrow stem cells, may result in improvement in the ventricular

function in ischemic heart disease patients [115]. Furthermore,

since bone marrow stem cells reside in the bone marrow of all

patients, they can be obtained by a relatively simple procedure

of bone marrow aspiration, expanded in vitro with or without

differentiation, and re-transplanted into the patient, thus

eliminating the need for immunosuppressants [70].

The initial assumption, regarding the capability of bone

marrow-derived stem cells to regenerate the heart by

transdifferentiation into cardiomyocytes, has been challenged

by a number of recent studied. Balsam et al. [149,150] and

Murry et al. [151] 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. A considerable

body of data indicates that a specific subset of bone marrow-

derived angioblasts, expressing endothelial precursor markers,

is responsible for neovascularisation and angiogenesis [99,152–

157]. Kocher et al. [99] for example, demonstrated that an

intravenous injection of human bone marrow donor cells to the

infarcted myocardium of rats resulted in a significant increase

in neovascularisation of post-infarction myocardial tissue,

attenuation of cardiomyocyte apoptosis and left ventricular

remodelling. The potential of bone marrow stem cells to heal a

damaged heart by inducing vasculogenesis in the injured

myocardium, thereby increasing heart viability and restoring

cardiac function has promoted the studies on bone marrow stem

cells quickly from small animals to clinical trials [115,158].

At present, the results of three medium size clinical trials

(100–200 patients) show a variable and modest healing function

of autologous bone marrow stem cells in cardiac function [159–

161]. The application of bone marrow cells for cardiac disease

is still in its preliminary phase, as optimal cell type, delivery

route, dose and timing require further optimisation. The

application of bone marrow-derived stem cells is also limited

by a safety issue: obtaining adequate autologous cells from a

patient with myocardium infarction in time to prevent post-

infraction remodelling may be difficult. In addition, the

presence of stem cells for cardiomyocytes in other parts of

the body, including bone marrow, has not been widely accepted

yet. Caspi and Gepstein [115] have given an excellently

tabulated overview on the clinical trial results of using bone

marrow stem cells in the treatment of acute and chronic heart

diseases.

2.3.3. Adipose-derived stem cells

Human adipose tissue provides a uniquely abundant and

accessible source of adult stem cells for applications in tissue

engineering and regenerative medicine [162–165]. Adipose-

derived stem cells have the ability to differentiate along

multiple lineage pathways. The cardiomyocyte phenotype from

adipose-derived cells has been reported [107,166]. Animal trial

with rats showed that adipose tissue-derived regenerative cells

improved heart function following myocardial infarction [167].

2.3.4. Native cardiac progenitor cells

It had long been believed 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 [168–172]. There are scientists who have

hypothesised that cardiac progenitor stem cells reside in the

hearts of neonatal animals, and that these progenitor stem cells

(if any) could eventually serve as the basis for cardiac cell

lineage formation and thus its application in the treatment of

cardiac disease in humans. Recently, cardiac progenitor cells

were found in the hearts of neonate humans, rats, and mice by a

multi-institution group of the United States and German

researchers [172].

Nonetheless, given the limited regeneration ability of the

adult heart, it is apparent that the existence of the above

mentioned cells within the adult heart do not translate to a

functionally significant cardiac differentiation following

myocardial infarction [115]. The role of these cells in the

normal adult heart is still to be elucidated. They may represent

an intrinsic repair system capable of replacing cells lost in the

normal process of ageing, or may simply reflect remnant cells

from organ development in early life. The existence of these

progenitor cells will no doubt open new opportunities for

myocardial repair, though many issues still need to be

addressed.

2.3.5. 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, in addition to the

common features shared by all stem cells as mentioned above.

First, they are pluripotent, which means they have a broader

multilineage expressing profile. Unlike adult stem cells, which

can differentiate to a relatively limited number of cell types,

embryonic stem cells have the potential to contribute to all adult

tissues. Second, they are robust. They have the long-term

proliferation ability with a normal karyotype, and can be

cryopreserved. Third, they can be genetically manipulated

[173]. Hence, research using embryonic stem cells remains at

the zenith of stem cell science.

The embryos, from which embryonic stem cells are derived,

are a hollow microscopic ball of cells called blastocyst. At this

blastocyst stage, a group of cells begins to separate from the

outer cell mass (or trophoblast) and forms the inner cell mass

(ICM) (also called embryoblast) (Fig. 6). While the outside

layer of cells of the blastocyst goes on to form the placenta, the

inner cell mass forms the embryo and will ultimately develop

into all the tissues in the body. Embryoblasts are, therefore,

truly pluripotent. In 1981, the inner embryoblasts were isolated

from mouse blastocysts and were successfully used to generate

pluripotent stem cell lines, which were termed embryonic stem

cells (ESC) [174,175]. However, the documentary world had

not seen a human embryonic stem cell line until 1998 when two

independent teams, Thomson et al. [94] and Shamblott et al.

[176], described the generation of human ESC lines.

The embryonic stem cell is capable of continuous

proliferation and self-renewal in vitro but also retains the

ability to differentiate into derivatives of all three germ layers

both in vitro and in vivo. Thus, following cultivation in

suspension, the ESCs tend to spontaneously create 3D

aggregates of differentiating tissue known as embryoid bodies

(EBs) [177]. Upon aggregation, differentiation is initiated and

the cells begin to a limited extent to recapitulate embryonic

development. Though they cannot form trophectodermal tissue

(which includes the placenta), cells of virtually every other type

present in the organism can develop. The aggregate at first

appears as a simple ball of cells, and then grow into an

increasingly more complex appearance. After a few days a

hollow ball (cystic embryoid body) forms, followed by the

appearance of internal structures, such as a yolk sac and heart

muscle cells (i.e. cardiomyocytes) which beat in a rhythmic

pattern to circulate nutrients within the increasingly larger

embryoid body.

The availability of the embryonic stem cell system has

boosted the hope of heart regeneration. The first study using

embryonic stem cell as a source for cell transplantation into the

myocardium was reported by Klug et al. [93]. By using

genetically selected mouse embryonic stem cell-derived

cardiomyocytes, Klug et al. showed that the differentiated

cells developed myofibrils and gap junctions between adjacent

cells and performed synchronous contractile activity in vitro for

up to 7 weeks. This study proved the feasibility to guide an

unlimited number of embryonic stem cells into cardiomyogenic

cell linage and to utilise them for myocardial regeneration.

Later studies [178–181], utilising the infarcted rat heart model,

demonstrated that transplantation of differentiated mouse

embryonic stem cell-derived cardiomyocytes can result in

short- and long-term improvement of myocardial performance.

Theoretically, the undifferentiated ESCs, which might be

able to differentiate in vivo into cardiomyocytes in the host

microenvironment containing cardiac-specific differentiation

signalling, could improve the heart function. However,

controversial results have been reported, regarding the in vivo

differentiation and outcome of the transplantation of undiffer-

entiated ESCs. Puceat’s group [182,183] showed that in

infarcted myocardium, grafted stem cells differentiated into

functional cardiomyocytes integrated with surrounding tissue,

improving contractile performance. However, Nussbaum et al.

[184] discovered that undifferentiated mouse ESCs consistently

formed cardiac teratomas in nude or immunocompetent

syngeneic mice, and that cardiac teratomas contained no more

cardiomyocytes than hind-limb teratomas, suggesting lack of

guided differentiation. Hence the authors concluded that

undifferentiated ESCs did not differentiate toward a cardio-

myocyte fate in either normal or infarcted hearts [184].

As regards immunogenicity of the transplantation of

undifferentiated ESCs, Menard et al. [185,186] reported that

cardiac committed mouse ESCs, which were transplanted to the

infarcted sheep heart following incubation with BMP-2,

differentiated to mature cardiomyocytes, and that cell

transplantation resulted in a significant improvement in cardiac

function independent of whether the sheep were immunosup-

pressed or not. However, another group [187] reported the

increased immunogenicity (i.e. rejection) of mouse ESCs upon

in vivo differentiation after transplantation into ischemic

myocardium of allogeneic animals, implying that clinical

transplantation of allogeneic ESCs or ESC derivatives for


treatment of cardiac failure might require immunosuppressive

therapy. This result was confirmed by Nussbaum’s group [184],

who found no evidence for allogeneic immune tolerance of cell

derivatives. Hence, successful cardiac repair strategies invol-

ving ESCs will need to control cardiac differentiation, avoid

introducing undifferentiated cells, and will likely require

immune modulation to avoid rejection.

2.3.6. Human embryonic stem cells

The vast biomedical potential of human ESCs has stirred

enthusiasm in the field of tissue engineering. From human ESCs,

scientists hope to grow replacement tissues for people with

various diseases, including bone marrow for cancer patients,

neurons for people with Alzheimer’s disease, pancreatic cells for

people with diabetes, and cardiomyocytes for patients with heart

damage. Furthermore, establishment of a tissue-specific differ-

entiation system may have significant impact on the study of

early human tissue differentiation, functional genomics,

pharmacological testing, and cell therapy. Given these reasons,

we think that the utilisation of human ESCs in cardiac tissue

engineering is worthy of a separate section. An overview of the

relevant reports is given in Table 3. After the establishment of

human ESC lines in 1998 [94,176], other research groups have

been able to develop a reproducible cardiomyocyte differentia-

tion system from the human ESCs [138,188–191]. The detailed

protocols can be found in these reports.

It has been demonstrated that the human ESC-derived

cardiomyocytes displayed structural properties of early-stage

cardiomyocytes [138,188,189,192]. The presence cardiac-

specific proteins and the absence of skeletal muscle markers

have also been confirmed [192,193]. These studies also

demonstrated the progressive maturation from an irregular

myofilament distribution to a more mature (i.e. organized)

sarcomeric pattern [138,192,193]. The human ESC-derived

cardiomyocytes were also shown to display functional

properties, consistent with an early-stage cardiac phenotype

[138,189,190,194,195]. Functional improvement of heart

following cell transplantation would require structural,

electrophysiological, and mechanical coupling of donor cells

to the existing network of host cardiomyocytes [115]. Hence, it

is important to investigate whether cells derived from human

ESCs can restore myocardial electromechanical properties. A

study of Gepstein group [196,197] demonstrated tight

electrophysiological coupling between the engrafted human

ESCs and host rat (or swine) cardiomyocytes both in vitro and

in vivo.

In spite of the exciting potential of human ESCs, the cells are

also giving rise to daunting legal and ethical concerns. ESCs are

controversial because they are obtained from the destruction of

a potential human embryo. Technologies for therapeutic and,

especially, reproductive cloning add further ethical problems.

In addition, a number of technical issues need to be addressed

prior to clinical application [115], as discussed below.

(1) S
cale up. One of drawbacks of cell therapy is the difficulty
in scaling up to meet larger production needs in clinical

applications.

(2) P
urification of the differentiating cardiomyocyte popula-
tion is necessary.

(3) I
n vivo delivery and efficiency of grafting. Cell-delivering
techniques should be developed to enable proper alignment

of the grafted tissue, high seeding rate of the transplanted

cells, and minimal damage to the host tissue.

(4) I
mmune rejection. Human ESCs (hESCs) are derived from
cell lines genetically distinct to the potential human

recipient. Therefore the potential for immune rejection

following hESC transplantation into an immunocompetent

adult recipient exists, as is observed with conventional solid

organ transplantation. As hESCs are therefore allogenic,

there is the potential requirement for co-administration of

immunosuppression [198–203]. Encouragingly there is

some evidence that hESCs have an immunopriviledged

status. Under certain conditions both in vitro and in vivo it

they display limited immunogenicity and are tolerated in

cross species (xeno-) transplantation without immunosup-

pression [204,205]. Reduced cell surface expression of both

major histocompatibility complex and accessory proteins is

believed to underlie this immunopriviledged status. How-

ever this is mainly observed in undifferentiated hESCs, and

with differentiation into specialised tissues such as

cardiomyocytes hESCs may acquire a greater immunogenic

phenotype. Strategies aimed at preventing immunological

rejection of the cells, such as genetic modification or graft-

recipient tissue type matching, should be explored.

2.4. Strategies to address immune rejection in cells

As mention above, the major limitation of cell therapy is

immune rejection associated with most used cell types.

Successful application of tissue engineering in man will

depend on the utilisation of an autologous or nonimmunogeneic

cell source, as well as synthetic scaffold materials, to avoid life

long immunosuppression. Several strategies aimed at achieving

immunological tolerance are being developed. These strategies

include [115]:

(1) E
stablishing banks of major histocompatibility complex
(MHC) antigen-typed human ESC lines.

(2) G
enetically altering the human ESC to suppress the
immune response (e.g., by knocking out the major

histocompatibility complexes).

(3) T
he concept of haematopoietic chimerism.
(4) G
enerating immune compatible ESC-derived cardiomyo-
cytes with the patient’s own genetic information, known as

somatic cell nuclear transplantation or therapeutic cloning.

The successful application of nuclear transfer techniques to

a range of mammalian species has brought the possibility of

human therapeutic cloning significantly closer. The objective of

therapeutic cloning is to produce pluripotent stem cells that

carry the nuclear genome of the patient and then induce them to

differentiate into replacement cells, such as cardiomyocytes to

replace damaged heart tissue. In the process, a somatic cell

nucleus from a patient is transferred into an enucleated oocyte

Table 3

Differentiation of human ESCs towards cardiomyocytes

Method of hES differentiation Major results Reference

In vitro: via EBs in suspension ESCs differentiated into cardiomyocytes, even after long-term culture. Upon

differentiation, beating cells were observed after one week, increased in

numbers with time, and retained contractility for >70 days. The beating

cells expressed markers of cardiomyocytes

Carpenter and co-

workers [188]

In vitro: via EBs in suspension ESCs showed consistence in phenotype with early-stage cardiomyocytes,

and expression of several cardiac-specific genes and transcription factors

Gepstein and co-

workers [138]

In vitro: via EBs in suspension ESCs showed a progressive ultrastructural development from an irregular

myofibrillar distribution to an organized sarcomeric pattern at late stages

Snir et al. [193]

In vitro: via EBs in suspension ESC-derive 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. [195]

In vitro: co-culture of differentiated

rat cardiomyocyte and hESC-

derived cardiomyocytes; In vivo:

transplantation of hESC-derived

cardiomyocytes into swine

Tight electrophysiological coupling between the engrafted hESC-derived

cardiomyocytes and rat cardiomyocytes was observed. The transplanted hES

cell-derived cardiomyocytes paced the hearts of swine

Kehat et al. [196]

In vitro: co-culture of

undifferentiated hESC

with mouse endoderm-like cells

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 foetal ventricular cells

Mummery et al. [189]

In vitro: via EBs in suspension 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. [206]

In vitro: co-culture of hESC-derived

cardiomyocytes with rat myocytes;

In vivo: differentiated hESC-

derived cardiomyocytes

transplanted into guinea pig

Electrically active, hESC-derived cardiomyocytes are capable of actively pacing quiescent,

recipient, ventricular cardiomyocytes in vitro and ventricular myocardium in vivo

Xue et al. [197]

In vivo: differentiated cardiac-

enriched hESC progeny

hESCs can form human myocardium in the rat heart, permitting studies of human

myocardial development and physiology and supporting the feasibility of their

use in myocardial repair

Laflamme et al. [192]


(i.e. its nucleus has been removed), as described in the cloning

of Dolly the sheep [207]. The oocyte containing the new

nucleus carries the genetic information of the patient. Using a

tiny pulse of electricity to cause the new nucleus to fuse with the

enucleated oocyte’s cytoplasm, this manipulated oocyte can

develop in vitro into a blastocyst. From this blastocyst,

embryonic stem cells with the genetics of the patient can be

isolated and expanded in vitro, and then differentiate in vitro

into genetically matched cardiomyocytes for transplantation.

Obviously, cloning would eliminate the critical problem of

immune incompatibility.

2.5. Strategies of cell delivery

In cell-based therapy, isolated cell suspensions are directly

injected into injured heart via the pericardium, coronary

arteries, or endocardium. Direct injection of isolated cells

avoids an open-heart surgery. However, it is difficult to control

the location of the grafted cells in the transplantation. To

address these problems, 2D and 3D cell delivery vehicles have

been under development using biomaterials. These strategies

are the core of the present review and will be discussed in great

detail in Section 3.

2.6. Summary of cell-based therapy and their limitations

Cell-based cardiac therapy represents an exciting strategy

in, for example, heart tissue regeneration and heart tissue

engineering. Huge efforts have been invested in the

development of cell sources for myocardial regeneration,

including foetal cardiomyocytes, skeletal myoblasts, bone

marrow stem cells, adipose stem cells, endothelial progeni-

tors, native cardiac progenitor cells, and embryonic stem cells

(ESC) (Table 2). A large amount of studies has also been

carried out to assess the roles of these potential cell types in

the regeneration of myocardial tissue, both in vitro and in

vivo. The translation of these basic scientific studies from

bench to bedside has been progressing. Among these cell

sources, embryonic stem cells possess the greatest potential

because of their intrinsic pluripotency. These pluripotent cells

can self-replicate tirelessly in the undifferentiated state in

vitro and be induced to differentiate into cell derivatives of all

three germ layers, including cardiomyocytes. The ability to

generate human cardiac tissue in vitro using embryonic stem

cells provides therefore the most exciting approach in the

field of cardiac tissue regenerative medicine and tissue

engineering.

Despite the enormous potential of cell-based therapy, a

number of technical issues need to be addressed prior to clinical

application. Key technical issues include (1) scaling up of cells,

(2) cell delivery, (3) efficiency of grafting, (4) suppression of

alternative unwanted cell phenotypes when ES cells are

applied, and (5) immune rejection. These issues, in particular

those associated with the efficiency of cell delivery and

grafting, could be potentially addressed by the strategies of

tissue engineering involving biomaterials.

Table 5

Selected polymeric biomaterials for tissue engineering [79,211–214]

Biomaterial Abbreviation

1. Synthetic polymers

Bulk biodegradable polymers

Aliphatic polyesters

Poly(lactic acid) PLA

Poly(D-lactic acid) PDLA

Poly(L-lactic acid) PLLA

Poly(D,L-lactic acid) PDLLA

Poly(glycolic acid) PGA

Poly(lactic-co-glycolic acid) PLGA

Poly(e-caprolactone) PCL

Poly(hydroxyalkanoate) PHA

Poly(3 or 4-hydroxybutyrate) PHB

Poly(3-hydroxyoctanoate) PHO

Poly(3-hydroxyvalerate) PHV

Poly ( p-dioxanone) PPD or PDS

Poly(propylene fumarate) PPF

Poly (1,3-trimethylene carbonate) PTMC

Poly(glycerol-sebacate) PGS

Poly (ester urethane) PEU

Surface bioerodible polymers

Poly(ortho ester) POE

Poly(anhydride) PA

Poly(phosphazene) PPHOS

Polyurethane PU

Nondegradable polymers

Poly(tetrafluoroethylene) PTFE

Poly(ethylene terephthalate) PET

Poly(propylene) PP

Poly(methyl methacrylate) PMMA

poly(N-isoproplylacrylamide PNIPAAm

2. Natural degradable polymers

Polysaccharides

Hyaluronan HyA

Alginate

Chitosan

Starch


3. Biomaterials for myocardial tissue engineering

3.1. Overview of substrate development for tissue

engineering

The aim of this section is to provide an overview on the

development of substrates (scaffolds, matrices) for tissue

engineering, the principles of which also prevail in the

development of myocardial tissue engineering matrices.

3.1.1. Criteria on tissue engineering substrates

Being a very much fledgling discipline, tissue engineering

encounters a variety of challenges, which can be grouped into

three categories associated with the science and technology of

cells, materials, and interaction between them. The challenges

that the material scientists encounter are caused by the strict

requirements on the tissue engineering substrates, as listed in

Table 4. The first three requirements are directly associated

with the chemical and physical properties of biomaterials, and

the last three criteria are set for the fabrication technologies of

substrates. The criterion list is not exhaustive. Specific

requirements on a specific tissue engineering construct, such

as osteoconductivity of bone tissue engineering scaffolds and

electric conductivity for nerve tissue engineering, are not

included. Some listed criteria might not be essential in certain

approaches. Porosity, for example, is not essentially required in

the heart patch approach (Fig. 4), although a porous heart patch

can be used initially to grow vascular cells before seeding

cardiac muscle cells. Biodegradability is not demanded in the

treatment of congenital heart diseases either.

3.1.2. Polymers used in soft tissue engineering

Polymers are the major type of materials used in soft tissue

engineering. Selected biopolymers are listed in Table 5. They

can be naturally occurring or synthetic. Detailed reviews are

available in literature [79,211–214].

Table 4

Criteria for tissue engineering substrates [208–210]

1. Ability to deliver and foster cells

The material should not only be biocompatible (i.e. nontoxic), but also

foster cell attachment, differentiation, and proliferation

2. Biodegradability

The composition of the material should lead biodegradation in vivo at

rates appropriate to tissue regeneration

3. Mechanical properties

The substrate should provide mechanical support to cells until sufficient

new extracellular matrix is synthesised by cells

4. Porous structure

The scaffold should have an interconnected porous structure for cell

penetration, tissue ingrowth and vascularisation, and nutrient delivery

5. Fabrication

The material should possess desired fabrication capability, e.g., being readily

produced into irregular shapes of scaffolds that match the defects in bone of

individual patients

6. Commercialisation

The synthesis of the material and fabrication of the scaffold should be

suitable for commercialisation

Proteins

Collagen

Gelatin

Fibrin

3.1.2.1. Naturally occurring polymers. Natural extracellular

matrices of soft tissues are composed of various collagens. It is

not surprising that much research effort has been focused on

naturally occurring polymers such as collagen [215,216] and

chitosan [217] for tissue engineering applications. Theoretically,

naturally occurring polymers should not cause foreign materials

response when implanted in humans. They provide a natural

substrate for cellular attachment, proliferation, and differentia-

tion in its native state. For the above-mentioned reasons, natural

occurring polymers could be a favourite substrate for tissue

engineering [211,218]. Table 6 presents major naturally

occurring polymers, their sources and applications. Among

them, collagen, fibrin, gelatin and alginate have been extensively

investigated for myocardial tissue engineering [24,64]. Each of

these polymers will be reviewed in detail in Section 3.2, in terms

of their applications in myocardial tissue engineering.

Table 6

List of naturally occurring polymers, their sources and main application fields [219]

Polymer Source Main application fields

Collagen Tendons and ligament Multi-applications, including cardiac tissue engineering

Collagen-GAG

(alginate) copolymers

Artificial skin grafts for skin replacement

Albumin In blood Transporting protein, used as coating to form a thromboresistant surface

Hyaluronic acid In the ECM of all higher animals An important starting material for preparation of new biocompatible and

biodegradable polymers that have applications in drug delivery and tissue

engineering, including cardiac tissue engineering

Fibrinogen-fibrin Purified from plasma in blood Multi-applications, including cardiac tissue engineering

Gelatin Extracted from the collagen inside

animals’ connective tissue

Multi-applications, including cardiac tissue engineering

Chitosan Shells of shrimp and crabs Multi-applications, including cardiac tissue engineering

MatrigelTM (gelatinous

protein mixture)

Mouse tumor cells Myocardial tissue regeneration

Alginate Abundant in the cell walls of

brown algae

Multi-applications, including cardiac tissue engineering

Polyhydroxyalkanoates By fermentation Cardiovascular and bone tissue engineering


3.1.2.2. Synthetic polymers. Although naturally occurring

polymers possess the above-mentioned advantages, their poor

mechanical properties and variable physical properties with

different sources of the protein matrices have hampered the

progress with these approaches. Concerns have also arisen

regarding immunogenic problems associated with the intro-

duction of foreign collagen [220].

Following the developmental efforts using naturally

occurring polymers as scaffold materials, much attention has

been paid to synthetic polymers. Synthetic polymers are

essential materials for tissue engineering not only due to their

excellent processing characteristics, which can ensure the off-

the-shelf availability; but also because of their advantage of

being biocompatible and biodegradable [220,221]. Synthetic

polymers have predictable and reproducible mechanical and

physical properties (e.g., tensile strength, elastic modulus, and

degradation rate), and they can be manufactured with great

precision. Although they are unfamiliar to cells and many suffer

some shortcomings, such as eliciting persistent inflammatory

reactions, being eroded, not be compliant or able to integrate

with host tissues, they may be replaced in vivo in a timely

Table 7

Selected properties of synthetic, biodegradable polymers investigated as scaffold m

Polymers Melting point,

Tm (8C)

Glass transition

point, Tg (8C)

Degra

1. Bulk degradable polymers

PDLLA Amorphous 55-60 Sutur

PLLA 173–178 60–65 Sutur

PGA 225–230 35–40 Sutur

PLGA Amorphous 45–55 Sutur

PPF <140 �20 Not a

PCL 58 70–72 Not a

PHB 177 4 Much

2. Surface erodible polymers

Poly(anhydrides) 150–200 Surfa

Poly(ortho-esters) 30–100 Surfa

Polyphosphazene �66 to 50 242 Surfa

fashion by native extracellular matrices built by the cells seeded

into them. It has become widely realised that an ideal tissue

engineered substitute should be made from a synthetic scaffold.

Table 5 has listed most synthetic polymers used in tissue

engineering. Table 7 provides selected properties of synthetic,

biocompatible and biodegradable polymers that have been

intensively investigated as substrate materials for tissue

engineering, type I collagen fibres being included for

comparison. Among them, aliphatic polyesters (PLA, PGA,

and PCL) have widely been applied as scaffolding materials for

3D tissue engineering constructs, an approach listed in Table 1.

Degradability is generally a desired characteristic in tissue

engineering substrates because the second surgery to remove

them (such as a heart patch) would be averted if the substrate

could be removed by the physiological system of the host body.

The biodegradable suture is a successful example of the

applications of biodegradable polymers. Moreover, in many

applications 3D tissue engineering constructs cannot be

removed, and the non-degradable biomaterials would act as

barriers to new tissue ingrowth and blood flow. Nonetheless, a

few nondegradable polymers, including PTFE and PET, have

aterials

dation kinetics Reference

es are absorbed completely in vivo in 12–16 months [221,224,225]

es can be absorbed completely in vivo after 24 months [221,224]

es are absorbed completely in vivo in 6-12 months [221,226,227]

es are absorbed completely in vivo in 2-12 months [211]

vailable [211,228]

vailable [229–231]

slower than PLA,PGA and PLGA [232,233]

ce erosion [211,228,234]

ce erosion [211,235]

ce erosion [236,237]

Table 8

Advantages and disadvantages of polymeric biomaterials for tissue engineering

Biomaterial Positive Negative

Naturally occurring polymers Excellent biocompatibility

(nor foreign body reactions)

(1) Poor processability

Biodegradable (with a wide range

of degradation rates)

(2) Poor mechanical properties

Bioresorbable (3) Immunogenic problem (i.e. poor immune compatibility)

(4) Disease transfection

Bulk biodegradable synthetic polymers (1) Good biocompatibility (1) Inflammatory caused by acid degradation products

Poly(lactic acid) (2) Biodegradable (with a wide

range of degradation rates).

(2) Accelerated degradation rates cause collapse of scaffolds

Poly(glycolic acid) (3) Bioresorbable.

Poly(lactic-co-glycolic acid) (4) Off-the-shelf availability

Poly(propylene fumarate) (5) Good processability

(6) Good ductility

Surface bioerodible synthetic polymers (1) Good biocompatibility (1) They cannot be completely replaced by new tissue

Poly(ortho esters) (2) Retention of mechanical integrity

over the degradative lifetime of the device

(2) Concern associated with the long-term effect

Poly(anhydrides) (3) Significantly enhanced tissue

ingrowth into the porous scaffolds,

owing to the increment in pore size

Poly(phosphazene) (4) Good processability

(5) Off-the-shelf availability

Nondegradable synthetic polymers (1) No foreign body reactions (1) Second surgery is required, or

(2) Tailorable mechanical properties (2) Concern associated with the long-term effect if they

have to stay in the host organ for a lifetime

(3) Good processability

(4) Off-the-shelf availability


been investigated for certain tissue engineering approaches,

such as passive diastolic restraint (Fig. 2) [24], congenital heart

disease [222], and especially vascular tissue engineering

[24,52,79].

Table 9

3D fabrication technologies of polymer scaffolds [238]

Fabrication technology Required properties

of materials

Available po

size (mm)

Melting-based

Melt moulding Thermoplastic 50-500

Extrusion/particle leaching Thermoplastic <100

Solvent-based

Solvent casting/particle leaching Soluble 30–300

Paraffin template Soluble 100–700

Emulsion freeze drying Soluble <200

Thermally induced phase

separation/freeze drying

Soluble <200

Electrospinning Soluble 20–100

Membrane lamination Soluble 30–300

Gas foaming

Chemical reactant gas foaming Amorphous 100–700

Physical induced gas foaming/

particle leaching

Amorphous Micropores

Macropores

Rapid prototyping

3D Printing Soluble 45–150

Fused deposition modelling

Direct rapid prototyping

To design a tissue engineering substrate, it is necessary to

weight up the ‘‘pros and cons’’ of the potential precursor

materials, which are summarised in Table 8. None of polymers

is universally suitable for all tissues. Actually no single material

re Porosity

(%)

Porous structure

<80 Spherical pores/low interconnectivity

<84 Spherical pores/low interconnectivity

20–50 Irregular pores

>90% Spherical pores

<97 High volume of interconnected micropores

<97 High volume of interconnected micropores

<95 Fibrous

<85 Irregular pores

>90% High volume of non-interconnected micropores

<50 <97 Low volume of non-interconnected micropores

combined with high volume of interconnected

macropores

<400

<70 100% interconnected macrospores


can provide all necessary properties required by a particular

application. In these instances, a hybrid of natural and synthetic

polymers designed to combine the advantages of both natural

and synthetic materials, such as a porous synthetic polymeric

scaffold filled with naturally occurring biomolecules [223],

may be most appropriate.

3.1.3. Fabrication of tissue engineering substrates

In general, the polymers used for tissue engineering are

either in gel or in solid states. Solid polymers are usually

applied as sheets (or films), knitted meshes, or 3D porous

scaffolds. The processing of polymer sheets/films and knitted

meshes is relative simple, and standard procedures are available

in many polymeric processing handbooks [239]. In the field of

tissue engineering, huge effort has been invested on the

development of fabrication technologies of 3D porous networks

for engineering of tissue constructs, and numerous techniques

have been developed to process porous polymeric, ceramic, and

composite scaffolds. A number of reviews are available in the

literature [238,240–246].

Polymeric scaffold processing methods could be classified

into four groups, as listed in Table 9 which is adapted from Ref.

Fig. 7. Typical porous structures produced by different scaffolding techniques. (a)

Elsevier; (b) extrusion with gas blowing (starch-based), reproduced from ref. [249],

(PLA) http://www.cbmm.lodz.pl/res/sppb_big2.jpg (accessed in May 2007); (d) par

permission from Springer Science and Business Media; (e) chemical reactant gas

drying (PLA), reprinted from ref. [251], with permission from Elsevier; (g) electrosp

May 2007); (h) 3D Printing (PLGA) http://www-bioc.rice.edu/bios576/scaffolds/Sca

reproduced with permission of J. Wiley and Sons Ltd.

[238]: (i) melted-polymer based, (ii) solvent based, (iii) gas-

foaming, and (iv) rapid prototyping (RP). The first three

processing methods are manual techniques, and the last one is

computer-aided technique. Fig. 7 illustrates typical porous

structures of 3D scaffolds fabricated by the techniques listed in

Table 9.

Melted-derived polymer-based methods seldom produce

satisfying porous structure, as shown in Fig. 7a and b. In the

particle leaching processing, it is a tricky task to control the

quality of scaffolds when using salt or sugar, due to the irregular

shape of the particles (Fig. 7c). However, this problem could be

addressed by using nearly perfect microspheres, such as

paraffin microspheres, as illustrated in Fig. 7d. Chemical

reactant gas foaming always produces controlled porous

structure (Fig. 7e), but it is limited to few polymers that could

produce gases during synthesis. It seems that among the

currently available techniques the microsphere-template

technique is the best one to fabricate spherical porous

networks, because it is controllable, low-cost, easy to operate,

and applicable to all soluble polymers, as listed in Table 10.

The freeze-drying and electrospinning methods produce

interesting aligned textures and porous structures (Fig. 7f and

Inject melt moulding (starch-based), reprint from ref. [248] with permission of

copyright by Mary Ann Liebert, Inc., publichers; (c) salt/sugar particle leaching

affin template (paraffin microsphere leaching (PLA) [250], published with kind

foaming (polyurethane); (f) thermally induced phase separation (TIPS)/freeze

inning (PCL), http://www.chem.pku.edu.cn/liangdh/eresearch.html (accessed in

ffolds.htm (accessed in May 2007); (i) fused deposition modelling (PCL) [252],


http://www.chem.pku.edu.cn/liangdh/eresearch.html

http://www-bioc.rice.edu/bios576/scaffolds/Scaffolds.htm

Table 10

Advantages and disadvantages of polymer scaffolding techniques

Techniques Advantages Disadvantages

Melted-polymer based Mass product ability Very poor quality

Solvent-based

Particle-leaching Simple, low cost, and suitable for many soluble polymers Poor in quality control

Microsphere-template Simple, low cost, suitable for many soluble polymers,

controlled quality, and controlled pore size and porosity

Using organic chemicals

Freeze-drying and

electrospinning

Simple, low cost, suitable for many soluble polymers,

and aligned structure

The aligned texture might not

be always desirable

Gas-foaming Simple, low cost, controlled quality, and controlled pore size and porosity Limited to few polymers

Rapid prototyping Controlled quality, controlled pore size and porosity,

and suitable for many soluble polymers

Expensive


g). In order to guide cells in soft tissue engineering, such as

myocardial tissue which has an aligned texture, scaffolds with

an aligned texture are desirable. Specialised cells in these

tissues may maintain their appropriate phenotypes and lay

down extracellular matrix better on the textured scaffolds [247].

Rapid prototyping (RP), also known as solid free-form (SFF)

techniques, are a computer-aided fabrication processes. They

can rapidly produce highly complex 3D objects using date

generated by computer aided design (CAD) systems. An image

of a defect in a patient can be taken, which is used to develop a

3D CAD computer model. The computer can then reduce the

model to slices or layers. The 3D objects are constructed layer-

by-layer using rapid prototyping techniques of fused deposition

Table 11

Overview of biomaterials used in cardiac tissue engineering

Biomaterials Physical state

Naturally occurring (also degradable)

Collagen based Gel

Fibrin glue Injectable gel

Peptide nanofibre Injectable gel

Collagen mesh 3D porous me

Collagen-glycosaminoglycan 3D porous me

Gelatin mesh 3D porous me

Alginate mesh 3D porous me

Synthetic

Biodegradable

PGA and copolymer with PLA 3D porous me

PLLA 3D porous me

PCL and copolymer with PLA 3D porous me

PGS 3D porous foa

Polyurethane (PU) 3D porous foa

Poly(ester urethane) (PEU)

[or poly(ester urethane)urea (PEUU)]

Solid sheet

Nondegradable

Poly(ethylene terepthalate) (PET).

The fibers are manufactured under

trade names Dacron, Terylene & Trevira

Knitted mesh

Solid sheet

Polypropylene (PP) Solid

Poly(tetrafluoroethylene) (PTFE) with or without PGA/PLA Solid sheet

Poly(N-isopropyl acrylamide) (PNIPAAm or PIPAAm) Solid sheet

modelling (FDM) (Fig. 7i), selective laser sintering (SLS), 3D

printing (3D-P) (Fig. 7h), stereo lithography, or extrusion free-

forming [253].

Rapid prototyping was developed as an inexpensive

technology for the manufacture industry. The idea behind this

is that a new machine could be entirely manufactured and

assembled quite inexpensively from the same polymer filament

feed stock that the rapid prototyping machine uses to make

prototypes by the owner of an existing one. Such a self-

replication technique will considerably reduce the cost of

prototyping machines [254]. However, this idea does not fit into

tissue engineering, as no patients have organs or tissues of

exactly the same geometric size and shape. Hence, rapid

s Tissue engineering approaches Reference

Epicardial heart patch [64,256]

Endoventricular heart patch [257]

Endoventricular heart patch [258]

sh 3D tissue engineering construction [259–261]

sh 3D tissue engineering construction [262]



[74,265]

sh 3D tissue engineering construction



m (sponge) 3D tissue engineering construction

and Epicardial heart patch

[268,269]

m (sponge) 2D tissue engineering films that could

be used as heart patches

[270,271]

Cardiovascular grafting [272,273]

Left ventricular constrain [18,274–293]

Cardiovascular grafting [294,295]

Left ventricular constrain [25,26,296,297]

Treatment of congenital heart disease [222]

Cardiovascular grafting [294,298–303]

Scaffold-free cell sheet [304–315]

Fig. 8. Parameters in Laplace’s law. R is the radius of a cylinder-shaped

ventricle, t the thickness of the wall, T the stress in the wall, and DP =

Pinside � Poutside.


prototyping processing can be an expensive fabrication

technique for the application of tissue engineering.

In summary, a variety of fabrication techniques have been

developed to produce polymeric scaffolds for tissue engineer-

ing. The microsphere-template method could be the most

promising one for producing spherical porous networks. The

freeze-drying and electrospinning methods could produce

aligned 3D pore textures, which might be desirable in most soft

tissue engineering applications, such as muscles. Although the

computer-aided technology is expensive, it is probably the most

advanced in terms of precise structural design and quality

control.

3.2. Biomaterials for myocardial tissue engineering

3.2.1. Selection of biomaterials and design of substrates

In general, the selection of biomaterials and the design of

substrates for heart muscle engineering could be guided by the

first three criteria listed in Table 4, i.e. (i) ability to deliver and

foster cells, (ii) biodegradability, and (iii) appropriate

mechanical properties. Porosity is another important factor

to be considered, in particular, for 3D tissue engineering

constructs.

3.2.1.1. Biocompatibility. The first and foremost function of a

substrate for tissue engineering is its role as substratum for cell

attachment. Indeed, the ability to support and foster cells is one

of essential properties of a biomaterial used in tissue

engineering [208,210,255]. Hence, the first step in achieving

a successful substrate is to design and produce an extracellular

matrix-like biomaterial that cells would like to attach to. Hence,

it is not surprising that almost all polymers applied for cardiac

muscle engineering are included in Table 5, which are polymers

initially developed for engineering tissues other than heart

tissue and have been well documented for their good

biocompatibilities.

3.2.1.2. Biodegradability. Biodegradability is not always

essential, depending on the applied approaches for cardiac

tissue engineering (Table 11). In general, nondegradable

polymers are selected in the approach of passive diastolic

constraint and the treatment of congenital heart disease, as well

as cardiovascular grafting. Degradable polymers, including

natural and synthetic, are applied in the heart patch approach

and, in particular, in 3D tissue engineering construction. One of

the important challenges encountered by materials scientists is

to design a material for 3D tissue engineering that could

biodegrade in a rate matching the growth kinetics of a specific

tissue at a specific anatomic position.

3.2.1.3. Mechanical properties. In general, different

approaches of myocardial tissue engineering have different

requirements on the mechanical properties of the biomaterials

applied.

Passive diastolic constraint. The idea behind this approach

is Laplace’s law [18], which gives the relation between the

stress T in a wall and the pressure difference across the wall

(DP = Pinside � Poutside), as illustrated in Eq. (1) and Fig. 8.

T ¼ R � DP

t: (1)

In the case of passive diastolic constraint (diastole is the

process of blood filling into a heart), T is the stress in the heart

wall, being named diastolic wall stress. Assuming that the

blood pressure in the left ventricle at the end of diastole (LVED)

is PLVED, and the supporting pressure provided by a cardiac

support device (CSD) is PCSD, then DP, which is also called

transmural pressure, equals PLVED when no CSD is applied, or

PLVED � PCSD when a CSD is implanted. Therefore, the law of

Laplace’s can be expressed by

diastolic wall stress ¼ R � PLVED

tðwithout a CSDÞ (2)

or

diastolic wall stress ¼ RðPLVED � PCSDÞt

ðwith a CSDÞ: (3)

Myocardial infarction typically results in heart wall thinning

(a reduction in t) and ventricular dilation (an increase in R),

which cause a significant increase in the heart wall stress

according to Eq. (2). Overstressed heart wall leads to

progressive structural and functional changes in ventricles

(ventricular remodelling), predisposing towards the end stage

of heart failure [5]. Hence, to prevent negative ventricular

remodelling, it is important to reduce the heart wall stress. This

could be achieved through a heart-supporting device according

to Eq. (3).

Eq. (3) indicates that theoretically there are no strict

requirements on the mechanical properties of the biomaterials

of a CSD, as long as it could provide a mechanical support at the

end of diastole. Practically, however, a CSD should have

optimal compliance so as to smoothly fit on the surface of heart.

Fig. 9. (a) Representative stress–strain curve of a collagen matrix from skin tested at a strain rate of 38.5% min�1. The stress-strain curve can be separated into three

distinct regions designated ‘‘toe’’, ‘‘linear’’, and ‘‘failure’’, reproduced from ref. [321] with permission of ASME International. (b) A typical stress-strain curve of

collagen fibres from tendon, the strain rate being 0.42% min�1, published from ref. [322] with permission of American Society for Clinical Investigation. Note: stress

level is kPa and MPa, in (a) and (b), respectively.

Table 12

Types I/III collagen ratio in heart muscle

Types I/III collagen ratio Measuring technique Reference

Healthy

heart

Heart with

DCM

0.60 1.27 Cyanogen bromide

analysis

[316]

0.58 1.12 Spectrophotometry [316]

No data 6.4–8.4 mRNA [318]

3.76 6.14 Sirius red positive

staining

[319]

0.3 1.25 mRNA [320]

DCM stands for dilated cardiomyopathy.


The compliance can be achieved, for example, through the

knitting of polymer fibres (Fig. 2). It would also be ideal that the

supporting device exhibits a nonlinear elasticity of heart muscle

such that it could reshape with the heart and thus provide

mechanical support to the heart throughout the beating

processes, rather than only at the end of diastole [18].

Heart patch. One specific benefit of the heart patch approach

could be simple changes in ventricular geometry and mechanics

leading to a reduction of elevated local wall stresses. According

to theoretical simulations, the addition of a heart patch to the

bonder-zone could decrease the heart wall stress [50]. In these

simulations, added materials are noncontractile and have 1–

200% of the average stiffness of passive myocardium. The

reduction in wall stress was calculated to be proportional to the

fractional volume added, with stiffer materials improving this

attenuation better [50].

The simulation work apparently indicates that a heart patch

could be designed to be stiffer than the heart muscle. On the

other hand, it is reasonable to assume that a polymer that is too

stiff may induce diastolic dysfunction. The maximum value of

the stiffness of a heart patch material (i.e. what percentage of

the average stiffness of passive myocardium) is an open

question. Therefore, it would be worthy to carry out in vivo

(e.g. on a suitable animal model) investigations on the effects

of heart patch stiffness on the cardiac functions of an infarct

heart.

Classic 3D tissue engineering approach (scaffold + cells). In

this classic approach of tissue engineering, the regenerative

ability of the host body is enhanced 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. 3D tissue engineering constructs could also be

used to produce in vitro healthy cells for cell-based therapy.

They may also be applied for many biomedical studies, such as

cell biology, organ development, functional cell differentiation

from stem cells, environment–cell interaction, cancer biology

and new drug treatment.

In essence, 3D tissue engineering is a technique of imitating

extracellular matrix. Hence, a construct should display some of

the mechanical properties of native myocardium. This is

another major challenge encountered by materials scientists, in

addition to the controlled degradation kinetics. Since the

extracellular matrix of myocytes is collagen, we hypothesize

that cardiomyocytes would be able to beat adequately in a

scaffolding material with mechanical properties similar

(although impossible to be exactly equal) to those of

myocardial collagen.

The myocardium collagen matrix mainly consists of types I

and III collagens, which form a structural continuum.

Synthesised by cardiac fibroblasts, types I and III collagens

have different physical properties. type I collagen mainly

provides rigidity, whereas type III collagen contributes to

elasticity [316]. The two types of collagens jointly support and

tether myocytes to maintain their alignment, whereas their tensile

strength and resilience resist the deformation, maintain the shape

and thickness, prevent the rupture and contribute to the passive

and active stiffness of the myocardium [317]. The ratio of

collagen types within the heart is of significance. Table 12 shows

values for the types I/III collage ratios reported in the literature.

Fig. 9 shows two typical stress–strain curves of type I

collagen, from which the most important property, elastic

Table 13

Selected mechanical properties of type I collagens

Source of collagen Stiffness Failure stress Failure strain Reference

Calf skin 2–22 kPa 1–9 kPa 0.5–0.7 [321]

Tendon, cartilage,

ligament or bone

25–250 MPa 1–7 MPa 0.24–68 [322,323]


modulus (stiffness or Young’s modulus), as well as two other

mechanical properties, failure stress (also called mechanical

strength) and failure strain, can be determined, as listed in

Table 13.

As regards the mechanical properties of type III collagen,

there has been no reported data, as far as the present authors are

aware of. It is well known that tissues with predominance of

type I collagen (such as tendon) are characterised by relatively

higher strength and stiffness, whereas tissues containing large

amount of type III collagen (such as skin) are characterised by

greater flexibility (compliance) [318]. Hence, it can be

predicted that the increase of collagen type I would result in

an increase in tissue stiffness. This has been demonstrated by

diseased heart muscle with different levels of dilation, as shown

in Fig. 10 [320]. In general, the stiffness of collagen increases

with stress linearly, i.e. ds/de = ks, where the coefficient k is

termed stiffness constant. A larger stiffness constant k gives rise

to a more rapidly rising curve with increasing stresses. Hence

the increase of stiffness constant with increasing ratio types I/III

collagen (Fig. 10) indicates stiffening of heart muscle.

According to Fig. 10, the stiffness constant of healthy heart

muscle with a types I–III collagen ratio of 0.5 is about 2.

A more direct reference for the mechanical design is the

mechanical properties of heart muscle. Typically, the passive

strains of the heart muscle could reach up to 15–22% at the end

of diastole [324–328]. Hence, the heart muscle mainly works

within the nonlinear elastic region, as indicated by Figs. 9 and

11. The stiffness of the heart muscle is around 10–20 kPa at the

early stage of diastole, and rises up to 50 kPa (healthy hearts) or

200–300 kPa (CHF hearts) at the end of diastole [324–328].

Fig. 10. Relation between Collagen I/III ratio and stiffness in myocardium

given by the stiffness constant k [320]. Reproduced with permission from

Elsevier.

These data could be used as the lower and upper bounds of

stiffness in the selection of materials for 3D tissue engineering

scaffolds.

3.2.1.4. Design of porous structures for 3D tissue engineer-

ing. The classic approach of tissue engineering is based on the

supposition that the employment of extracellular matrix-like

scaffold, natural biology (e.g., cells and biomolecules) of the

living body will induce the regeneration tissue or organ ex vivo

[51–53]. However, in the absence of true vascularisation, in

vitro 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 mm without the use of bioreactor, and<200 mm

with the support of bioreactor) [74]. Diffusion barriers that are

present in vitro are most likely to become more deleterious in

vivo due 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 initially thought that a pore size in the range of 50–

100 mm was sufficient to allow the vascularisation of a scaffold

following transplantation [71]. Later Radisic and Vunjak-

Novakovic [74] suggested a larger (�100 mm) pore size for

Fig. 11. Typical passive stress–strain curves of heart muscles of different

murine and human. The stiffness of the heart muscle is around 10–20 kPa at

the early stage of diastole (strain<10%), and rises up to 50 kPa (healthy hearts)

or 200–300 kPa (CHF hearts) at the end of diastole (strain being 15–22%).

Table 14

Investigated or potential biomaterials used in cardiac muscle engineering

Polymer Elastomer (E)

or thermoplastic (T)

Young’s modulus

(or stiffness)

Tensile

strength

Degradation

(month)

Reference

Synthetic, nondegradable

PET T or E 4 GPa Nondegradable [323]

PET/DLA E 2–3 MPa Nondegradable

Synthetic, degradable

PGA T 7–10 GPa 70 MPa 2–12 [214,228,244,334]

PLLA or PDLLA T 1–4 GPa 30–80 MPa 2–12 [214,228,244,334]

PHB E 2–3 GPa 36 MPa Degradable [232,334]

PPD (= PDS) E 0.6 GPa = 600 MPa 12 MPa 6 [334,335]

PU E 5–60 MPa 20–45 MPa Surface errodible [336]

PEUU E Not available Not available Degradable

TMC E 6 MPa 12 MPa Degradable [337]

TMC-PDLLA(50:50) E 16 MPa 10 MPa Degradable [337]

POC T 1–16 MPa 6.7 MPa Degradable [334,338]

PGS E 0.04–0.282 MPa 0.5 MPa Degradable [48,334,339]

Naturally occurring

Collagen fibre

(tendon/cartilage/

ligament/bone)

E 2–46 MPa 1–7 MPa Degradable [322,323,334]

Collagen gel (calf skin) E 0.002–0.022 MPa 1–9 kPa Degradable [321]

Heart muscle

Myocardium of rat E 0.14 MPa (at the end of diastole) 30-70 kPa Not Applicable [340–342]

Myocardium of human E 0.2–0.5 MPa (at the end of diastole) 3–15 kPa Not Applicable [324,343,344]


cardiac tissue constructs. It has been also suggested that pore

size larger than 300 mm is necessary for long-term survival, i.e.

vascularisation [238,329–331]. But excessively large pores

could impair vascularisation, because endothelial cells are

unable to bridge pores greater than a cell diameter [258]. A

potential approach to address this problem is to fill a highly

porous scaffold with a cell-seeded and/or gene-containing

collagen gel, as it has been reported that collagen gels do not

obstruct vascularisation while providing substrata for cells to

attach [77].

3.2.2. Biomaterials used in myocardial tissue engineering

The majority of polymeric biomaterials used in the fields of

cardiac muscle tissue engineering have been listed in Table 11.

Among the natural polymers, collagen, alginate and gelatin

have been under intensive investigation for myocardial tissue

engineering by research groups in Germany [64,66,77], Israel

[63,70,71], and Canada [263,332], respectively. Among the

synthetic biodegradable polymers, PGA and copolymers with

PLLA, and PCL have been studied systematically using

bioreactors for myocardial tissue engineering at MIT

[74,76,269] and Harvard [267,333]. The thermo-responsive

polymer PNIPAAM was applied to cardiac tissue engineering

by researchers in Japan [304–314]. The cardiac devices made

from nondegradable polymers have been under intensive

animal and human clinical trials by surgeons mainly in USA

[18,274–285], Germany [286–291] and Sweden [292,293].

Selected properties of most materials mentioned above are

given in Table 14. Sections 3.2.2.1–3.2.2.5 provide an accurate

and thorough assessment of the state-of-the-art of each of these

materials applied in the field of myocardial tissue engineering.

At the end of each section, critical discussions are given on the

important research issues and topics that remain challenges to

the advancement of the field, aiming to stimulate new insights

and perspectives that may be most useful for material scientists

involved in this field.

3.2.2.1. Materials used in the heart patch approach. In this

application, the major function of a heart patch is to deliver

healthy, functional cardiomyocytes to the infarct area of the

heart (Fig. 4), and the patch is then expected to degrade at the

end of the cell-delivery. These requirements could be achieved

with biological gels [47]. It would be ideal that a heart patch

could also provide mechanical support to the left ventricle. For

this purpose, stiffer synthetic heart patches have also been

applied [48].

Collagen gel. In 1997, Eschenhagen et al. [47] reported, for

the first time, an artificial heart tissue, which was termed

engineered heart tissue (EHT). Several years later, a similar

artificial tissue, ‘Engineered Cardiac Tissue’ (ECT) was

reported by a Chinese research group [345,346]. In Eschenha-

gen et al.’s work, embryonic chick cardiomyocytes were mixed

with collagen solution and allowed to gel. By culturing the

cardiomyocytes in the collagen matrix, they produced a

spontaneously and coherently contracting 3D heart tissue in

vitro. Immunohistochemistry and electron microscopy revealed

a highly organized myocardium-like structure exhibiting

typical cross-striation, sarcomeric myofilaments, intercalated

discs, desmosomes, and tight junctions.

Later the same group reported the long-term survival of

neonatal rat cardiomyocytes in the matrix and its functional and

physiological characteristics in vitro [347]. They detected an

Fig. 12. Different techniques to fix EHTs on the recipient’s heart. (a) Ring-shaped EHTs were placed around the circumference of the heart or sutured onto the heart in

apical-valvular direction, reprinted from ref. [350] with permission from Elsevier. (b) Stacking EHTs were sutured on the recipient’s heart, reprinted from ref. [351]

by permission from Macmillan Publishers Ltd.


increase in beating force up to 18 days after casting the gel

matrix and stable forces thereafter. Up to now this is the only

3D engineered artificial heart tissue that could beat in vitro,

with the maximal contraction force being 2–4 mN (giving a

stress of�13 mN/mm2) [77]. However, this contractility is still

insufficient (isolated adult muscle cells develop a twitch tension

up to 56 mN/mm2). To augment contractile performance, gene

transfer in EHT has been explored [348].

Using a modified protocol (mixing neonatal myocytes with

liquid collagen I, Matrigel1, and serum-containing culture

medium), Zimmermann et al. [349,350] designed EHTs in

circular-shape (inner/outer diameter: 8/10 mm; thickness:

1 mm), and fitted them around the circumference of uninjured

hearts from syngenic rats after 12-day in vitro culture. Most

recently, large (thickness 1–4 mm and diameter 15 mm) EHTs

were produced with neonatal rat heart cells, and were implanted

on heart with myocardial infarction in rats [351]. The surgical

approaches of applying the ring-shaped EHTs to hearts are

shown in Fig. 12.

EHTs were heavily vascularised and retained a well-

organized heart muscle structure after implantation for 14 days

[350]. When evaluated 28 days later, EHTs showed beneficial

effects in terms of electrical coupling to the native myocardium,

prevention of further dilation, and systolic wall thickening of

infarcted myocardial segments. Despite utilisation of a

syngenic model, immunosuppression was necessary for EHT

survival in vivo. To avert immunogenicity of EHT due to the use

of collagen, Matrigel1 and serum, mixed native and

cardiomyocyte enriched heart cell populations (not neonatal

cardiomyocytes) were cultured in serum-free and Matrigel1-

free conditions in order to simulate a native heart cell

environment [352].

The investigations described above are important proof-of-

concept work. The artificial heart tissue developed has

structural, functional, and physiological characteristics of

cardiac tissue. Another significant finding in these studies is

that strong vascularisation takes place in collagen gels in vivo.

However, there is currently 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 cardiomyo-

cytes remains to be established. Immunogenicity and, in

particular, poor mechanical supportive ability of collagen gels

are another two drawbacks associated with this approach. As

discussed in Section 3.2.1.3, an optimal heart patch is expected,

in addition to the cell delivery function, to provide passive

diastolic constraint to the left ventricle so as to improve

regional and global pump function.

Injectable gels. The concept of injectable tissue engineering

scaffolds has been constantly pursued, aiming at minimally

invasive surgery. Endothelial cells cultured within self-

assembling peptide gels form typically capillary-like networks

in vitro, and co-culture of cardiac myocytes with endothelial

cells leads to rapid assembly of capillary-like channels, with

accumulation of cardiac myocytes on the outside of the

endothelial-line channel. The endothelial cells promote

electrical synchrony of the cardiac myocytes [353,354]. These

results suggest that injection of cells in the appropriate self-

assembling peptide could allow cells to assemble in situ [355].

Fibrin glue is another injectable gel used for myocardium

repair.

Anatomically, injectable gels have been applied as an

endoventricular heart patch (Fig. 4b) [50]. It has been shown

that injection of the fibrin glue preserves left ventricular

geometry and prevents a deterioration of cardiac function

following myocardial infarction [50,257,356]. However,

injectable gels lack sufficient stiffness for the application in

human tissues. The stiffness of a variety of possible injectable

materials is in the range of 10 Pa to 20 kPa, such as fibrin

(�50 Pa) [357], Matrigel1 (30–120 Pa) [358], type I collagen

gels (20–80 Pa for 1–3 mg/ml) [359], bioactive hydrogels based

on N-isopropy acrylamide (100–400 Pa) [360,361], alginate

(100 Pa to 6 kPa) [362] and polyethylene glycol (1–3 kPa)

[363]. These materials are significantly softer than the human

heart muscles at the end of diastole, the stiffness of which is

�50 kPa in normal hearts or 200–300 kPa in CHF hearts [324–

328]. Hence, it is unlikely that they could provide sufficient

mechanical support to the diseased heart.

Synthetic elastomer poly(glycerol-sebacate). Wall et al. has

reported a theoretical work [50], which indicates that the


addition of a sheet material to a damaged left ventricular wall

could have important effects on cardiac mechanics, with

potentially beneficial reduction of elevated myofibril stresses,

as well as confounding changes to clinical left ventricular

metrics. The reduction in wall stress was calculated to be

proportional to the fractional volume added, with stiffer

materials improving this attenuation better [50]. Stimulated by

the theoretical simulation of Wall et al. [50], the present authors

have designed a heart patch (Fig. 4a) from a biocompatible and

biodegradable poly(glycerol-sebacate) (PGS) [364].

PGS was recently developed for the field of soft tissue

engineering [339,365]. This polymer is a biodegradable,

biocompatible, and inexpensive elastomer [366], and has

already shown potential in nerve [367] and vascular tissue

engineering [368–371]. PGS also has superior mechanical

properties, being capable of sustaining and recovering from

deformation due to its intrinsic elasticity, and thus it is well

suited to work in a mechanically dynamic environment, such as

the heart [268,269].

The PGS developed by the present authors shows a wide

range of stiffness values (several 10 kPa to 1.2 MPa) that could

be tailored to be either softer or stiffer than the heart muscles of

various animals [364]. Hence, this material could offer the most

convenient tool to carry out in vivo studies on the mechanical

effects of heart patches on the cardiac function of an infarct

heart and to address the following fundamental questions: (1)

Should a heart patch be softer or stiffer than the heart muscle?

(2) What is the maximum stiffness of a heart patch that would

not cause heart dysfunction?

Polyurethane (PU) and poly(ester-urethane) (PEU) rubbers.

Polyurethane rubbers are segmented copolymers having the

characters of thermoplastic elastomers. These copolymers

contain two types of segments: one type of blocks, i.e. A blocks,

is above its glass transition temperature Tg and forms soft

segments, and the other type, i.e. B blocks, is below its Tg and

forms hard segments. In poly(ester-urethane) elastomers, the

hard polyurethane segments tend to aggregate to give ordered

crystalline domains aided by the possibility of forming

hydrogen bonds between adjacent hard segment units. In

contrast, the soft polyester segments remain amorphous at

ambient temperature.

PU can be tailored to have a broad range of mechanical

properties and good biocompatibility, which makes them ideal

for many implantable devices [372,373]. A major problem of

PU composed with di-isocyanate may be the toxicity of

degradation products. However, in vitro and in vivo studies have

reported good cell adhesion and proliferation and no adverse

tissue reaction [374]. Because these materials are subject to in

vivo degradation, they are no longer applicable in long-term

implanted devices. On the other hand, it is because of their

biodegradability and elasticity that polyurethane elastomers

have gained new attention for the field of tissue engineering.

Since the early 1990s polyurethane elastomers have been

studied for soft, in particular, vascular tissue engineering [375–

377]. Recently, the elasticity of polyurethane elastomers has

made them also attractive for cardiac muscle engineering

[49,270,271].

The Young’s modulus of PU is typically in the range of 5–

60 MPa [378], which is much higher than that of the heart

muscle at the end of diastole. However, PU can be processed

into very thin films (10–15 mm in thickness) [270,271] and

highly porous network [49] which makes PU flexible and

compliant. Researchers at Toronto University [270,271] carried

out in vitro studies on the integrity of the cardiomyocyte

patterns on a thin polyurethane film. In these studies, adult or

ESC-derived cardiomyocytes were seeded onto the patterned

films and cultured for up to 4 weeks or 30 days. At the end of

culture, multilayered (approximately 2–3 cell layers) and

patterned cardiomyocytes were formed on the films. The dense

and highly aligned cardiomyocytes were able to contract the

thin polyurethane films [271]. These results indicate that the PU

and PEU films could be promising heart-patch materials,

although this was not explicitly pointed out by the researchers

[270,271].

Most recently, Fujimoto et al. [49] reported a successful

animal trial using a biodegradable, porous PU heart patch. The

heart patch was sutured to the infarct region of the rat heart. The

implanted PU patch promoted contractile phenotype smooth

muscle tissue formation and improved cardiac remodelling and

contractile function at the chronic stage. This work suggests

that the heart patch approach could be a new therapeutic option

against post-infarct cardiac failure.

Major challenges. The most important challenge that

material scientists currently encounter in the heart patch

approach, as well as in 3D tissue engineering scaffold design, is

to achieve, with synthetic polymers, a nonlinear elasticity that

is similar to that of heart muscle. Such a nonlinear elasticity

would ensure that the heart patch can ‘beat’ together with the

recipient heart, thus providing appropriate restraints to the left

ventricle throughout the beating process. To successfully

design such a nonlinear elasticity, material scientists need to

gain fundamental understanding of the nonlinear elasticity of

collagen-based muscles [379–388] and, in particular, the

mechanical behaviours of collagen or collagen gels at a single

molecule level [389,390].

3.2.2.2. Polymers used for 3D tissue engineering scaffold-

s. As discussed in Section 4.1.3, the mechanical requirement

on the materials used for 3D tissue engineering construction is

that they should have mechanical properties similar to those of

native myocardium. The anisotropic, nonlinear elasticity of

heart muscles makes this task the most challenging. In the field

of myocardial tissue engineering, the development of 3D tissue

engineering constructs did not follow what we think today

should be required. Instead, most 3D tissue engineering

constructs investigated for heart muscle engineering were

adopted from other areas of tissue engineering, even from bone

tissue engineering (e.g. bioceramic TiO2 [391]). This is one of

the reasons why the utilisation of artificial matrices to engineer

myocardium has been rather disappointing, with no examples

of human application to date. Nevertheless, it is the following

pioneering work that deepened our understanding of the fate of

implanted materials and cells in the myocardium of both

healthy and diseased hearts.


Collagen fibrous mesh (or collagen sponge). The application

of collagen gel matrix is limited by its insufficient mechanical

strength. Hence, alternatives are being sought. 3D solid

collagen mesh, for example, is commercially available for

scaffold applications, such as Tissue Fleece by Baxter

Deutschland, Heidelberg, Germany [392] and Collagen type

I matrix from Biomaterials Research B.V. Vaals, The Nether-

lands [393,394].

Kofodis et al. [259,260] seeded neonatal rat cardiomyocytes

in vitro into this commercially available collagen sponge, and

called the product artificial myocardial tissue (AMT). The

myocardial tissue generated in the 3D scaffolds in vitro

possesses structural, mechanical, physiological and biologic

characteristics similar to native cardiac tissue [259,260].

Undifferentiated embryonic stem cells were seeded into

collagen mesh scaffolds [395], the construct being called

‘Bioartificial Myocardial Tissue’ (BMT). The bioartificial

mixtures were implanted in the infarct area of rat hearts

(Fig. 13). It was 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. Fractional short-

ening was better in embryonic stem cell-treated animals than in

those which had received matrix implants without cells.

The studies of van Luyn et al. [393] confirmed that

contracting cells (not contracting constructs) were present after

seeding neonatal rat heart ventricular cell fractions and were

able to survive in collagen type I sponges for up to 135 days. In

the subsequent investigation of the same group [394], Collagen-

I scaffolds were directly sutured to healthy or injured left

ventricles of mice without in vitro cell culture. Encouraging

results were reported, in terms of vascularisation, degradation

of scaffolds and foreign body reactions.

To summarize this section, although positive results were

obtained from the histological analysis on the retrieved hearts,

no significant improvement was reported regarding ventricular

function in hearts receiving collagen sponge grafts. Moreover,

no beating constructs have been achieved in vitro with collagen

sponges.

Gelatin mesh. Gelatin mesh is the second type of solid

scaffold applied in cardiac muscle engineering [263]. Gelatin

Fig. 13. Intramural transplantation of a bioartificial, embryonic stem cells-

containing matrix [395]. Reprinted with permission from Elsevier.

mesh is also commercially available as foam (e.g. Gelfoam

prepared from purified pork skin and supplied by Upjohn,

Ontario, Canada). Li et al. [263] seeded foetal rat ventricular

muscle (not isolated cells) into Gelfoam to form grafts. The grafts

were cultured in vitro for 7 days, forming a beating cardiac tissue

in vitro in the grafts. Then the grafts were implanted either into

the subcutaneous tissue of adult rat legs or onto myocardial scar

tissue in a cryoinjured rat heart. In both the subcutaneous tissue

and the myocardial scar, blood vessels grew into the grafts from

the surrounding tissue. The grafts implanted into the sub-

cutaneous 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. Like

in the case of collagen scaffolds, the researchers did not note any

improvement of ventricular function in hearts receiving either

cell-seeded or unseeded gelatin grafts.

In a subsequent investigation, the same research group [396]

seeded various isolated cells into gelatin meshes separately,

including foetal rat ventricular cardiomyocytes, stomach

smooth muscle cells, skin fibroblasts, and adult human atrial

cells, and cultured them in vitro. The important facts revealed

by this investigation include: (1) grafts seeded with foetal rat

ventricular cells exhibited spontaneous rhythmic contractility

in vitro; (2) the gelatin mesh was slowly degraded; (3) all cell

types preferentially migrated to the uppermost surface of each

graft and formed a 300–500 mm thick layer. The last point

indicates a critical issue in 3D tissue engineering, i.e. critical

size. Readers will see from the rest of this section that all 3D

tissue engineering constructs 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 [74]. The usage of a bioreactor can ameliorate the

situation [261,397] but the problem is not totally solved. In fact,

vascularisation is one of the greatest challenges in 3D tissue

engineering construction [398]. This topic will be further

discussed at the end of Section 3.2.2.2 (Challenges in 3D

engineered myocardial constructs).

In another series of interesting studies [399–403], gelatin

and other synthetic (PCL, PGA and PCL-co-PLLA) patches

were combined with foetal, adult rat heart or rat aortic smooth

muscle cells and cultured in vitro to expand the number of cells.

Then the patches were used to replace the right ventricular

outflow tract (RVOT) in syngeneic rats and stayed in vivo for up

to 22 weeks. The gelatin patches, which started to dissolve at 4

weeks, had completely dissolved at 12 weeks [399], while the

patches made from PCL, PGL or PCL-co-PLLA did not thin or

dilate in vivo due to their relatively slow degradation kinetics

[401–403]. The seeded cells survived in the right ventricular

outflow tract. The patches encouraged the ingrowth of fibrous

tissue and remained endothelialized [399–403]. However, a

significant inflammatory reaction was noted in the gelatin patch

after 4 weeks as the scaffold started to dissolve [399]. But the

inflammatory response did not happen with PCL, PGA and

PCL-co-PLLA patches, among which PCL-co-PLLA patches

permitted better cellular penetration in vitro than PCL and PGA

patches [401–403].


In summary, unlike collagen meshes, gelatin meshes seeded

with foetal rat ventricular cells were reported to exhibit

spontaneous rhythmic contractility in vitro. Like collagen

meshes, no improvement of ventricular function was noted in

hearts receiving either cell-seeded or unseeded gelatin grafts.

Critical size is a major issue in the 3D tissue engineering

approach.

Alginate mesh. In addition to protein-based materials, there

is intensive research activity in the area of natural poly-

saccharides. Alginate, a negatively charged polysaccharide

from seaweed that forms hydrogels in the presence of calcium

ions, was initially developed for drug delivery [404,405] and it

is now under development for tissue engineering [406,407].

Alginate sponges with porosity being 90% and pore size in

the range of 50–150 mm were produced by the group of Cohen

and co-workers [408,409], using freeze-drying technique. The

collaborators Leor et al. [264] seeded foetal 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 myocardial infarct, which 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 progres-

sive deterioration in left ventricular contractility, the graft-

treated rats showed attenuation of left ventricular dilatation and

unchanged contractility in the left ventricle [264]. To optimise

the alginate scaffolds, efforts have been invested to achieve 3D

high-density cardiac constructs with a uniform cell distribution

[410] and to enhance the vascularisation of the alginate

scaffolds by incorporating controlled release basic fibroblast

growth factor (bFGF) microspheres [411]. This was the first 3D

scaffold that was reported to improve the cardiac function of the

left ventricle.

Poly(glycolic acid) (PGA) and its copolymer with poly(-

lactic acid) (PLA). Scaffolds made by the electrospinning

technique from PGA, PLA or their copolymers (PLGA) have

long been investigated for the engineering of a variety of tissue

types by the MIT and Harvard Medical School collaboration

since 1991 [65,73,412,413]. The first formally published work on

a synthetic polymer-based scaffold designed for cardiac muscle

engineering was produced by Freed and Vunjak-Novakovic in

1997 [265]. 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 (not constructs),

which consisted of cardiomyocytes with cardiac-specific

structural and electrophysiological properties, contracting

spontaneously and synchronously [265,414]. To the knowledge

of the present authors, this is the first achievement of engineered

beating tissue on synthetic polymers, though the visible

spontaneous contractions in constructs were not comparable

to those of native ventricles and eventually ceased after 4 days of

cultivation [414].

The drawback associated with the critical thickness of the

engineered tissue constructs was investigated systematically by

this group. After one week, a peripheral tissue-like region (50–

70 mm thick) was observed [414]. The small thickness of the

cardiac muscle-like zone in constructs, including both synthetic

and natural polymer scaffolds, was attributed to the poor mass

transportation in the construct centre [74,415]. In the

subsequent studies, efforts were invested to overcome the

limitation on the thickness of engineered tissue through

improving cell seeding and tissue culturing conditions as

follows:

(1) c
ell seeding density [416–418];
(2) c
ell seeding vessel, using rotating vessels (laminar flow)
rather than mixed flasks (turbulent flow) [416,419];

(3) t
issue culture vessel, using dynamic (i.e. rotating) rather
than static conditions [416,419,420];

(4) i
mprovement of gas exchange [421,422];
(5) p
erfusion of medium through the construct (perfusion
bioreactor) [417,422,423];

(6) l
aminin-coated PGA sponge [414,416,419];
(7) u
se of collagen scaffolds instead of PGA scaffolds
[417,418].

In summary, with three investigated scaffolds (including

collagen sponges, gels or PGA), cultivations performed in

dishes, static or mixed (turbulent flow) flasks or rotating

(laminar flow) bioreactors yielded constructs with a thin (0.1–

0.2 mm) peripheral layer [419] of tissue expressing markers of

cardiac differentiation and able to propagate electrical signals.

The non-uniform cell distribution is a result of oxygen

diffusional limitations within the constructs. It was also found

that cultivation with perfusion of culture medium through the

construct could enhance the convective-diffusive oxygen

supply and yield a 1–2 mm thick construct [423] (this equals

to a 0.5–1 mm thick peripheral layer) with physiologically high

and spatially uniform distribution of viable cells expressing

cardiac markers [74,269].

Poly(e-caprolactone) (PCL) and its copolymer with poly(-

lactic acid) (PLA). PCL is another important member of the

aliphatic polyester family. It has been used to effectively entrap

antibiotic drugs, and constructs made with PCL have been

considered as drug delivery systems. 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 oligometric species from

the bulk.

Vacanti’s group [267] demonstrated the formation of

contractile cardiac grafts in vitro using a nanofibrous PCL

mesh. The nanofibrous mesh, which was produced by the

electrospinning 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 alpha-myosin

heavy chain, connexin-43 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 [267].

The subsequent work of the same group [333] demonstrated

the formation of thick cardiac grafts in vitro and the versatility

of biodegradable electrospun meshes for cardiac tissue

engineering. To construct 3D cardiac grafts, the cell-seeded

cardiac nanofibrous PCL meshes were overlaid between days 5

and 7 of the in vitro culture period. In addition to well-attached

and strongly beating cells throughout the experimental period,

constructs with up to five layers could be cultured without any

incidence of core necrosis. The layers adhered intimately, with

morphologic and electrical communication being established

between the layers, as verified by means of histology and

immunohistochemistry as well as synchronized beating. It is

envisioned that cardiac grafts with clinically relevant dimen-

sions can be created by using this approach and combining it

with new technologies to induce vascularisation.

Along with the gelatin meshes, PCL, PGL or PCL-co-PLLA

have also been investigated in vivo to replace the right

ventricular outflow tract due to their relatively slow degradation

kinetics [401–403].

Poly(glycerol-sebacate) (PGS). As discussed in Section

3.2.1.3, the mechanical and structural properties of a 3D tissue

engineered construct should closely resemble those of living

tissues that they intend to replace. To engineer the myocardial

tissue, which beats cyclically and constantly throughout life,

the biomaterial should be as soft and elastic as heart muscle.

These mechanical characteristics are impossible to reproduce

with polyester-based thermoplastic polymers, such as PGA,

PLA, PCL and their copolymers, because they undergo

plastic deformation and failure when exposed to long-term

cyclic strain. The limitation of their use in engineering soft

tissues has made scientists turn to elastomers for cardiac tissue

engineering.

It has been mentioned that PGS with a wide range of

stiffness has been developed for the heart patch approach by the

present authors [364]. Actually this material could also be

applied to build 3D tissue engineering scaffolds, and perhaps is

the most promising material to date. The PGS synthesised at

110 8C has stiffness of several tens kPa, which is in the range of

passive stiffness of the healthy human heart muscle. PGS also

has a tailorable biodegradability, from being degradable in a

couple of weeks to being nearly non-degradable. Initial in vitro

cell culture showed that PGS was biocompatible with

cardiomyocytes, with the human embryonic stem cell-derived

cardiomyocytes beating for at least 2 months when cultured on

the synthesised PGS. Hence, PGS is a very promising soft

elastomer for heart tissue engineering.

Poly(ester urethane) urea (PEUU) and poly(1,3-trymethy-

lene carbonate) (PTMC). 3D PEUU scaffolds have been

fabricated by the electrospinning technique [378,424,425].

Electrospun PEUU is elastomeric and allows for the control of

fibre diameter, porosity, and degradation rate through a variety

of fabrication parameters. The PEUU constructs integrated with

smooth muscle cells were flexible and have anisotropic

mechanical properties, with stiffness ranging form 1.7 to

2.5 MPa on the material axis, which is close to that of the

human heart muscle at the end of diastole [378]. To the

knowledge of the present authors, the use of 3D electrospun

PEUU scaffolds in the myocardial tissue engineering has not

been explored, and would be an interesting topic.

Poly(TMC) is surface erodible elastomers designed for soft

tissue engineering. It has been copolymerised with epsilon-

caprolactone (CL) [337] and DLLA [426,427] to modify its

biodegradability and mechanical properties. The application of

poly(TMC) for heart muscle engineering has been investigated

to a very limited extent [427].

Challenges in 3D engineered myocardial constructs. No

matter what types of biomaterials they are built from, the

utilisation of artificial matrices (scaffolds) to engineer

myocardium has been rather disappointing, with no examples

of human application to date. 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 quality 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) loads. This

mechanical requirement on the engineered construct is proved

to be a tough task for biomaterial scientists, as discussed

below. 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 jeopardized if the scaffold material is

too stiff. Then the question is how soft the material should be?

The obvious answer is that the compliance should be that of

the collagen matrix of the heart muscle. One can see from

Table 14 that most of these biomaterials, including collagen

fibres, are much stronger than myocardium. This explains

why most solid engineered constructs lack contractile

function. An exception is the beating collagen rings produced

by Zimmermann et al. [64]. On the other hand, collagen gels

are too weak. The data in the table indicates that PGS could be

the most promising candidate from the mechanical compat-

ibility point of view. However, there is another hurdle,

nonlinear elasticity of heart muscle. Given the complex

anisotropic mechanical behaviour of heart muscle, it is by no

means an easy task to synthesize a polymer that responses to

stresses the way heart muscle does. Without an adequate

mechanical environment, the engineered tissue would develop

into poor tissue organisation. This is because the potential of

cardiac myocytes to organise into a 3D force-generating unit

is limited when they are forced to seed a structurally defined

environment when solid matrices are employed. The poor

organisation of cardiomyocytes in turn leads to a weak

electrical and functional syncytium.

Fig. 14. Schematic illustration showing surgical procedure of heterotypic heart

transplantation using PTFE. After cold cardioplegia and explantation, (A) a 6-

mm long ventriculotomy was created lateral to the left anterior descending

artery, and (B) the 4 mm � 4 mm � 2 mm patches were implanted into the left

ventricle (C) to enlarge the ventricular cavity [222]. Reprinted with permission

from Elsevier.


Another equally critical issue is size limitation, which is

still a barrier common to most tissue engineering approaches.

Perfusion in vitro may improve the nutrient supply to augment

the size of the engineered constructs. However, once they are

implanted in vivo, the thick tissue will need to be vascularised

so that the cells within will receive the necessary nutrients and

waste products will be removed. Hence, vascularisation will

be a crucial step in the development of any engineered tissue

construct. Previous work [238,329–331] has shown that

angiogenesis, which can occur in many biological tissues

without large pores, does not always succeed with synthetic

scaffolds. It has been reported that necrosis occurs in the core

regions of synthetic scaffolds of low porosity due to the

obstruction of the material [74]. To ensure the smooth

regeneration of the vascular tissue without blocking by the

biomaterial, a 3D scaffold should have reasonably large pore

size (perhaps >300 mm) [238,329–331], as discussed in

Section 3.2.1.4. However, the highly open porosity in the

scaffolds ironically would provide insufficient surface areas

for cells to attach and proliferate. This issue could probably be

addressed by filling the scaffolds with cell-seeded collagen

gels, as previous work has shown that collagen gels do not

obstruct mass transportation and vascularisation [77]. To

promote vascularisation, vascular growth factors (VGF)

incorporated by the gene delivery techniques and an optimal

stem cell type (i.e. mesenchymal stem cells) could be applied

to engineer the constructs. This combinatorial paradigm

might ultimately bring tissue engineering into clinical

application.

3.2.2.3. Nondegradable polymers used for the treatment of

congenital heart diseases. It is well documented that

polytetrafluoroethylene persist at the site of reconstruction

for the life of the repair [222]. For the treatment of certain

congenital heart diseases, Krupnick et al. [222] have developed

an in vivo model of ventricular tissue engineering using

heterotypic heart transplantation and polytetrafluoroethylene

(PTFE)–polylactide meshes combined with types I and IV

collagen gels. In this work, solid scaffolds were made of PGA–

PTFE or PLLA–PTFE. Liquid types I and IV collagen hydrogel

mixture was poured into 4 mm � 4 mm � 2 mm PGA- or

PLLA-based meshes to produce initial scaffolds. Then isolated

mesenchymal progenitor cells were seeded into the above

constructs to form patches. This was followed by overnight

incubation before implantation into syngeneic rats.

The surgical procedure is shown in Fig. 14. In brief, after

cold cardioplegia and explantation, a 6-mm long ventriculot-

omy was created in lateral position to the left anterior

descending artery, and the 4 mm � 4 mm � 2mm cell-seeded

patches were sutured into the left ventricle to enlarge the

ventricular cavity. The assessment duration was up to 1 month.

Successful transplantation and differentiation of mesenchymal

progenitor cells were accomplished using this approach [222].

The outcomes were also encouraging:

(1) T
he constructs resulted in minimal inflammation without
aneurysmal dilatation.

(2) N
o ventricular arrhythmias resulted from this surgical
manipulation and echocardiography revealed both end

systolic and diastolic volume augmentation with ventricular

expansion (note: ventricular expansion is desirable in the

case of systolic dysfunction).

3.2.2.4. Nondegradable polymers used for passive diastolic

constraint. To prevent the negative left ventricular remodel-

ling and left ventricular dilation associated with myocardial

infarction, many studies have examined the use of biomaterial

supports to restrain the left ventricle.

Poly( propylene) (PP) used as heart patches. Kelley et al.

[25] first demonstrated that restraining expansion of acute

myocardial infarctions preserves left ventricle geometry and

function. In this work, a poly(propylene) (Marlex) mesh was

sutured onto the myocardium of sheep at the location of a

subsequently induced myocardial infarction (Fig. 4a without

cells). Bowen et al. [296] demonstrated that restraining acute

infarct expansion with poly(propylene) decreases collagenase

activity in borderzone myocardium. However, the matrix

components of the heart muscle were unchanged within the

infarct region. The work by Enomoto et al. [26] indicated that

stiffening only the infarct region may not be sufficient.

Poly(ethylene terephthalate) (PET) used as heart wrapping

jackets. Other type of left ventricle restraint consisting of a

knitted PET mesh has been developed by Acorn Cardiovascular

Inc. The cardiac support device (CSD) is fitted around both

ventricles (Fig. 2). Investigations with canine and ovine models

showed that CSD decreased left ventricular end-diastolic

volume, myocyte hypertrophy and interstitial fibrosis

[275,276], and increased ejection fraction after myocardial

infarction [27,277,280,282,283,428]. Saavedra et al. [274]

further demonstrated that reverse remodelling with reduced

systolic wall stress and improved adrenergic signalling could be


achieved by passive external support that does not generate

diastolic constriction.

Key issues. Acorn’s clinical trial encompassing 300 patients

initially reported that the CSD reduced left ventricle diastolic

volume, improved the life quality of patients [18,286].

However, the significance of the study has been questioned

because of the partial recruitment of patients in an unblended

way, the influence of missing data, and the conflicting results on

beneficial effects [24]. Therefore, this strategy should be taken

with caution, and there exists a need for more long-term follow-

up data and a more thorough investigation on the exact

mechanisms behind left ventral restraint [24].

3.2.2.5. Scaffold-free myocardial tissue engineering. This

strategy is based on the self-assembling properties of

cardiomyocytes, aiming to generate scaffold-free 3D tissue

constructs. The materials involved in this approach are used as

cell culture supports in vitro.

Polystyrene (culture plates and microcarrier beads). Akins

et al. [429,430] made a first attempt to generate scaffold-free

myocardial tissue constructs, using polystyrene tissue culture

plates and polystyrene microcarrier beads. Kelm et al. [431]

further refined this approach and obtained highly organized 3D

micro-tissue with cardiac phenotypic characteristics. In a

similar approach, Baar et al. [432] obtained spontaneously

beating cylindrical construct.

Thermo-responsive poly(N-isoproplylacrylamide) (PNI-

PAAm). PNIPAAm is a special biocompatible polymer shown

to be temperature-responsive. This material was developed by a

research group in Japan [312,433–439] as a cell culture

substrate ex vivo, rather than used as an implant in the body.

This culture surface is slightly hydrophobic and cell-adhesive

under culture condition at 37 8C and change reversely to

become hydrophilic and non-cell-adhesive below 32 8C due to

rapid hydration and swelling of grafted PNIPAAm. Unlike the

cells undergoing enzymatic digestion (e.g. trypsinisation), cell-

to-cell junction and adhesive proteins within confluent cultured

cells are completely preserved during detachment from

PNIPAAm surface. Eventually an intact cell sheet is produced.

The monolayer sheets can be overlaid to form multi-layered

sheets, which are eventually grafted to the heart without any

artificial scaffolds. So far, this group has produced cell sheets

using various types of cells [310,311,315,440–443], including

cardiomyocytes [310,311,315].

The myocardial cell sheets were first built by Shimizu et al,

and assessed both in vitro and in vivo [310,311,444]. Neonatal

rat cardiomyocyte sheets detached from PNIPAAm-grafted

surfaces were overlaid to construct multilayer cardiac grafts.

Layered cell sheets began to pulse simultaneously and

morphological communication was established between the

sheets. In vivo, layered cardiomyocyte sheets were transplanted

into subcutaneous tissues of nude rats. The microvasculature of

the grafts was rapidly organized within a few days with graft

beatings observed by naked eyes 3 days after transplantation.

The macroscopic beating of the grafts was preserved up to 1

year [444], confirming their ability for long-term survival.

Histological studies showed characteristic structures of heart

tissue, including elongated cardiomyocytes, well-differentiated

sarcomeres, and gap junctions within the grafts. These results

demonstrate that electrically communicative pulsatile 3D

cardiac constructs were achieved both in vitro and in vivo by

layering cardiomyocyte sheets. [310,311,444].

Recently, autologous myocardial cell sheets were grafted to

rats [445] and hamsters [446] with myocardial infarction

(Fig. 3) [46]. The cardiomyocyte sheets became attached to and

integrated with the infarcted myocardium, showed angiogen-

esis, expressed connexin-43, and appeared as homogeneous

tissue in the myocardium. Overall, cardiomyocyte sheets

improved cardiac performance in ischemic myocardium when

compared with fibroblast sheets in infarcted myocardium [445],

myoblast injection, and sham operation [446]. The prolonged

life expectancy was considered due to a reduction in

myocardial fibrosis and re-organization of the cytoskeletal

proteins in hamsters.

Without the use of any additional materials such as carrier

substrates or scaffolds, cell-sheet engineering avoids the

complications associated with traditional tissue engineering

approaches such as host inflammatory responses to implanted

polymer materials. Cell sheet engineering thus presents

significant advantages and can overcome some of the problems

that have restricted tissue engineering with biodegradable

scaffolds [46,447]. The above-described approach holds an

enormous potential in cardiac tissue engineering if combined

with the computer-aided design (CAD) technology [52].

4. Biomimetic tissue engineering—bioreactors

4.1. Brief history of bioreactors

A bioreactor is an apparatus that attempts to mimic

biological conditions in order to maintain and encourage

tissue regeneration [448]. In a bioreactor, a dynamic culture

environment, which mimics in vivo physiochemical (e.g.

temperature, pO2, pH, and pCO2), electrical and mechanical

conditions, is constantly maintained.

The development of bioreactors used for research in tissue

engineering was fuelled by the rather disappointing results of in

vitro tissue regeneration in dishes. Under static culture

conditions, cell growth rates are diffusionally limited due to

increasing cell mass and decreasing effective implant porosity

resulting from extracellular matrix regeneration. It has been

widely recognised that bioreactors are essential for the research

in tissue engineering. In the body, cells are always stimulated

by mechanical, electrical and chemical signals that influence

their biological behaviour. If these signals are inadequate or

non-existent, cells dedifferentiate, become disorganised, and

eventually die. In fact, biological tissues adapt their structure

and composition to surrounding specifically functional

demands [448]. Placing cells alone or only in contact with

materials in culture medium is not enough to obtain a functional

tissue. Therefore, bioreactors are particularly crucial for the

regeneration of complex 3D tissues, such as heart.

The world saw the first tissue engineering bioreactor in

1993, which was invented by Langer’s group for cartilage tissue

Fig. 15. Schematics of four types of tissue engineering bioreactors: (a) static; (b) mixed; (c) rotating; (d) perfusion bioreactors [412].


engineering [412]. Since then, studies on the development and

application of tissue engineering bioreactors have been rapidly

spreading throughout the world, involving actually all types of

tissues in the body. With the use of bioreactors, the products of

in vitro engineered skin and cartilage are commercially

available. Bioreactors have also improved the quality of

engineered corneas, bones, urethras, and pancreatic islets such

that they are currently under in vivo trial in humans [448]. A

number of comprehensive reviews on bioreactors are available

[74,269,448–450].

4.2. Comparison of bioreactors

Bioreactors can be grouped into four types: i.e. static, mixed

(turbulent flow), rotating (laminar flow), and perfusion

bioreactors [269]. Fig. 15, which is adapted from Ref. [451],

shows the four types of bioreactors. The major characters of

these four types of bioreactors are summarised as follows (see

Table 15 as well) [269,449]:

(1) I

Tabl

Cha

Bior

Stati

Mix

Rota

Perf

n a static flask, the cultured tissue is fixed in place and

exposed to motionless medium. Hence, mass transportation

occurs through molecular diffusion, and there is no shear

stress on the cells.

(2) I
n a mixed flask, tissues are fixed in place and exposed to a
well-mixed medium. The mixing is achieved via flask

spinning. Hence, the mass transportation in the medium

occurs via turbulence. Turbulent flow can produce severe

shear stresses on cells at the construct surface such that it

might cause cell damage or death. Within the tissue

construct, the mass transportation is governed by molecular

diffusion.

(3) I
n a rotation chamber, the tissue constructs are suspended in
culture medium. The lateral chamber rotates slowly such

that the tissue constructs can fall by gravity. The slow

e 15

racters of four types of bioreactors

eactor Mass transportation

c No fluid flow in the culture media, so mass transportation occ

ed Mass transportation occurs via turbulent fluid flow in medium

diffusion within the construct

ting Fluid flow is laminar, so mass transportation is by convection

governed by diffusion within the constructs

usion Mass transfer is by pressing a medium to flow through a tissu

transportation is by convection both in the medium and withi

rotation encourages the uniform growth of the tissues. In the

meantime, it decreases mechanical flow stresses in the

medium, and thus alleviates the situation of surface cell

damage and death. The slow rotation also protects fragile

tissues from cracking. However, the mass transportation

with the constructs is still controlled by molecular diffusion.

(4) I
n a perfusion bioreactor, constructs are cultivated with
direct perfusion of culture medium through the cultured

tissue construct, in order to achieve convective-diffusive

oxygen transportation throughout the construct volume

[269].

Most recently, Bilodeau and Mantovani [448] made a specific

review, focusing on the mechanical environment of various types

of bioreactors and its effects on cultured tissue. They concluded

that rotating bioreactors are more suitable for the culture of

fragile tissues, such as cartilage, while static perfusion

bioreactors are better for the culture of tissues with complex

geometry that normally subjects to higher mechanical constrains

in the body. Advantages and disadvantages of these two types of

bioreactors for different tissue cultures are given in Table 16.

4.3. Application of bioreactors in myocardial tissue

engineering

In myocardial tissue engineering bioreactors are very

important to grow the constructs in which the regeneration

of lost cardiac functions is achieved via providing the right

mechanical and physiological properties. However, studies on

the in vitro culture of cardiac muscle constructs are more

complex, less advanced and rarer than those on cartilage. The

effects of culture parameters and their mechanisms are still

largely unknown. The advantage with using bioreactors is the

possibility of obtaining a better understanding of various

phenomena.

Mechanical condition

urs via molecular diffusion No shear stress on the cells

, and is governed by Severe shear stress on the cells

at the surface of the construct

in the medium, and is Mild shear stress

e construct. So mass

n the constructs

Mild shear stress

Table 16

Comparison of bioreactors used for different tissues [448]

Tissues Static/perfusion bioreactor Rotating bioreactor

Cartilage, bone, and ligament Possible application of mechanical stress Adequate exchange of nutrients and glass

Complex shape feasible Insufficient mechanical stresses

Blood vessels Possible application of mechanical stresses and strains Uniform culture

Complex shape feasible Insufficient mechanical stresses and strains

Myocardial tissues Mechanical stress too substantial A free-falling environment protests fragile issues

Heart valves Complex shapes feasible Insufficient mechanical stresses

Possible application of a flow


The bioreactor-based investigations on 3D cardiac muscle

constructs have experienced the processes of applying rotating

and then perfusion bioreactors. The first report on culturing

cardiac muscle constructs in bioreactors was by Freed and

Vunjak-Novakovic in 1997 [265]. In this work, mixed

bioreactors were used. Using this type of bioreactors, the

largest engineered tissue that could be achieved was a 50–

70 mm thick peripheral tissue-like layer [414]. The small

thickness of the cardiac muscle-like zone in constructs,

including both synthetic and natural polymer scaffolds, was

attributed to the poor mass transportation in the construct centre

[74,415]. Later Carrier et al. [416] discovered that a rotating

bioreactor improved cell quality, distribution, and metabolism

compared to static and mixed culture vessel. However, the

thickness of culture muscle tissue was still very limited. With

three types of scaffolds (including collagen sponges, gels or

PGA) cultivated in rotating (laminar flow) bioreactors, the best

result yielded constructs with a thin (100–200 mm = 0.1–

0.2 mm) peripheral layer [419]. The non-uniform cell

distribution was a result of oxygen diffusional limitations

within the constructs, although the mass transportation in the

medium was improved by the rotating condition.

In the subsequent studies, the group invested great efforts to

overcome the limitation on the thickness of engineered tissue

through improving cell seeding and bioreactor design [416–

423], as mentioned in Section 3.2.2.2 (PLA–PGA part).

Eventually, cultivations with perfusion of culture medium

through the construct enhance the convective-diffusive oxygen

supply and yield a 1–2 mm thick construct [423] (note: this

equals to a 0.5–1 mm thick peripheral layer) with physiolo-

gically high and spatially uniform distribution of viable cells

expressing cardiac markers [74,269].

It has been discovered that shear stresses as low as 0.1 Pa are

sufficient to harm cardiac myocytes. Adequate stresses to

encourage growth and differentiation of these cells are around

0.001 Pa, which can be obtained with a laminar flow in a

rotating bioreactor [448].

Fink et al. [452] studied the effects of strain (elongation)

instead of shear stresses. Instead being deteriorated, the cardiac

tissues responded to a 20% strain by hypertrophy. These tissues

had an increase of DNA and protein/cell ratios and an improved

organisation of the cells into parallel arrays. Also, the stretched

cells were longer and wider and the thickness distribution of

cells was more uniform than in the control tissues. Hence, it

would be of importance to investigate 3D heart muscle

regeneration by applying an elastomeric material and dynamic

mechanical conditions, which are similar to the mechanical

properties of the cardiac muscle and the heart beating process,

respectively.

4.4. Key issues

Although bioreactors have improved the in vitro tissue

culture, there are several limitations on their applications. They

do not regenerate completely native-like tissues. During

culture, there is always a period when the properties stop

improving, remain constant for a short periods, and then begin

to decline. Dynamic culture lessens this phenomenon, but it

does not eradicate it. The maximum volume of tissue that can

be regenerated without developing an appropriate vessel tree

into the tissue is approximately 3 mm [448]. From the clinical

point of view, engineered tissue constructs could be used in situ

to enhance its function. The body would be its own bioreactor

and allow for infiltration of cells within the scaffold matrix to

regenerate myocardial muscle and blood vessels. However,

necrosis has been reported at the core regions of implanted 3D

scaffolds. Hence, vascularisation is also one of key issues in the

development of 3D tissue culture system.

In addition, specific tissue culture requires specific design

parameters on bioreactors. The physiological, biochemical,

electrical and mechanical controls, which are essential in cell

and tissue cultures, have not been achieved for most tissue

cultures. For myocardial tissue engineering, the bioreactor that

mimics contractile and electrical activities of the heart is a

particularly important issue.

5. Summary

Heart muscle engineering aims to regenerate functional

myocardium to repair diseased and injured heart. Several

strategies and their approaches have been established over the

last decade (Table 1). Among all types of cells applied so far,

embryonic stem cells and cloned cells seem to be most

promising. Different approaches have different requirements on

biomaterials. Hence, material scientists should select and

design biomaterials according to the specific approach they

have adopted. Myocardial tissue engineering encounters a

variety of challenges, which are grouped into three categories

associated with cells, materials and vascularisation, as

summarised in Table 17. A combinatory paradigm, which

Table 17

Major challenges in myocardial tissue engineering

1. Challenges associated with cells

Cell sourcing, cell scale-up, cell delivery and immune rejection

2. Challenges associated with biomaterials and scaffolds

One of the challenges in myocardial tissue engineering is to design and fabricate mechanically compatible materials and scaffolds that have nonlinear

elasticity similar to the heart muscle and thus can develop synchronically beating with the recipient heart

3. Challenges associated with vascularisation

After transplantation, rapid vascularisation is the critical step for the subsequent events: adequate mass transportation, cell survival, integration and function of

the engineered cardiac constructs. The maximum volume of tissue that can be regenerated without developing an appropriate vessel tree into the tissue is

approximately 3 mm, which is far too smaller than clinical required size. Vascularisation remains the key factor in the transition of in vitro achievements into

effective therapeutic tool


uses stem cells, signalling molecules (such as genes) and novel

biomaterials to enhance cell growth and proliferate, to

encourage vascularisation and to support the damaged heart,

will be needed to bring heart tissue engineering into clinical

application.

Acknowledgements

The authors would like to acknowledge financial support

from the UK Biotechnology and Biological Sciences Research

Council (BBSRC) (grant number BB/D011027/1).

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Myocardial tissue engineering: a review

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