Fisher John P. e.a. (ed.) Tissue Engineering
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26-8 Tissue Engineering
26.5 Bioreactors for Mechanical Conditioning
In vivo, the pulsatile nature of blood flow imposes radial pressure upon the vessel wall, which subjects
SMCs within the medial layer to cyclic strain. Thus, a great deal of research has examined SMC behavior
in response to cyclic stretch and found such stimuli important in the fabrication of vascular tissue,
particularly with respect to ECM synthesis and tissue organization. For example, SMCs seeded on purified
elastin membranes and exposed to 2 days of cyclic stretching (10% beyond the resting length) have been
shown to incorporate hydroxyproline into protein three to five times more rapidly than stationary controls,
indicating increased collagen synthesis in response to strain [Leung et al., 1976]. Cyclicstrain also increased
the synthesis of collagen types I and III and chondroitin-6-sulfate without stimulating DNA synthesis.
Another study also detected enhanced matrix production in collagen constructs seeded with rat aortic
SMCs and subjected to 7% cyclic strain [Kim et al., 1999]. Over 20 weeks of culture under cyclic strain,
SMCs upregulated expression of elastin and collagen type I. Elastin content from these SMCs increased
49% over unstretched controls. Furthermore, organization of the tissue was observed, as evidenced by
perpendicular alignment of SMCs to the direction of the applied strain.
Similar results have been obtained in three-dimensional constructs. Tubular collagen constructs seeded
with SMCs were cultured over thin-walled silicone sleeves and subsequently exposed to regulated intralu-
minal pressures to stretch the vessel in a repeatable fashion for up to 8 days. The 10% cyclic distension
in diameter caused the scaffolds to contract, SMCs and bundles of collagen fibers to align circumferen-
tially around the vessel, and improvement of the scaffold’s mechanical properties [Seliktar et al., 2000].
Moreover, this model system was employed to investigate the remodeling capacity of these constructs
via the activity of matrix metalloproteinases (MMPs) known to cleave solubilized type I collagen frag-
ments [Seliktar et al., 2001]. Constructs mechanically conditioned for 4 days contained five times higher
amounts of MMP-2 compared to static controls and increased MMP-2 activity. The increases in MMP-2
levels correlated favorably with improvements in mechanical strength and material modulus as a result
of cyclic strain. When a nonspecific inhibitor of MMP-2 was added to the culture media, MMP-2 levels
decreased and mechanical properties were reduced, negating the benefits of mechanical conditioning.
These studies indicate that strain-mediated remodeling of collagen scaffolds is essential for improved
construct of mechanical properties.
Because of the profound effects of cyclic strain on SMC orientation, ECM production, and tissue organ-
ization, preculture of vascular graft constructs in a pulsatile flow bioreactor system may help recreate the
natural structure of native vessels and allow one to better achieve the mechanical properties required
of the construct. A schematic of a typical pulsatile flow bioreactor system is shown in Figure 26.3. The
mechanical stimuli from pulsatile flow could generate the cyclic strain necessary to alter ECM production,
thereby creating a histologically organized, functional construct with satisfactory mechanical characterist-
ics for implantation. To develop a blood vessel substitute, Niklason et al. [1999], cultured PGA constructs
in a pulsatile blow bioreactor generating 165 beats per minute (bpm) and 5% radial strain. The pulse
frequency of this system was chosen to mimic a fetal heart rate, believed to possibly provide optimal
conditions for new tissue formation. However, most mechanical conditioning investigations mentioned
above conducted strain studies at 60 bpm, more representative of an adult heart rate, with promising
outcomes. Therefore, the optimal bioreactor culture conditions for the development of a TEVG remain to
be elucidated. Nevertheless, such a system shows promise for the production of a blood vessel substitute
with the necessary mechanical and biochemical components.
26.6 Conclusions
In the past couple of decades, a great deal of progress on TEVGs has been made. Still, many challenges
remain and are currently being addressed, particularly with regard to the prevention of thrombosis and
the improvement of graft mechanical properties. In order to develop a patent TEVG that grossly resembles
native tissue, required culture times in most studies exceed 8 weeks. Even with further advances in
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Tissue Engineered Vascular Grafts 26-9
5% CO
2
Compliance
chamber
Graft
chamber
Pulsatile
pump
Reservoir
Check valves
TEVG
FIGURE 26.3 Diagram of a typical pulsatile flow bioreactor for culture of TEVGs.
the field, TEVGs will likely not be used in emergency situations because of the time necessary to allow
for cell expansion, ECM production and organization, and attainment of desired mechanical strength.
Furthermore, TEVGs will probably require the use of autologous tissue to prevent an immunogenic
response, unless advances in immune acceptance render allogenic and xenogenic tissue use feasible.
TEVGs have not yet been subjected to clinical trials, which will determine the efficacy of such grafts in the
long term. Finally, off the shelf availability and cost will become the biggest hurdles in the development of
a feasible TEVG product.
Although many obstacles still exist in the effort to develop a small-diameter TEVG, the potential
benefits of such an achievement are exciting. In the near future, a nonthrombogenic TEVG with sufficient
mechanical strength may be developed for clinical trials. Such a graft will have the minimum characteristics
of biological tissue necessary to remain patent over a time period comparable to current vein graft
therapies. As science and technology advance, TEVGs may evolve into complex blood vessel substitutes.
TEVGs may become living grafts, capable of growing, remodeling, and responding to mechanical and
biochemical stimuli in the surrounding environment. These blood vessel substitutes will closely resemble
native vessels in almost every way, including structure, composition, mechanical properties, and function.
They will possess vasoactive properties, able to dilate and constrict in response to stimuli. Close mimicry
of native blood vessels may ultimately aid in the engineering of other tissues dependent upon vasculature
to sustain function. With further understanding of the factors involved in cardiovascular development
and function combined with the foundation of knowledge already in place, the development of TEVGs
should one day lead to improved quality of life for those with vascular diseases and other life threatening
conditions.
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27
Cardiac Tissue
Engineering: Matching
Native Architecture
and Function to
Develop Safe and
Efficient Therapy
Nenad Bursac
Duke University
27.1 Introduction.............................................. 27-1
Cellular Cardiomyoplasty • Tissue Cardiomyoplasty
27.2 Cardiac Architecture and Function..................... 27-3
27.3 Current State of Cardiac Tissue Engineering .......... 27-4
Cardiogenesis In Vitro • In Vivo Implantation for Cardiac
Repair
27.4 Design Considerations .................................. 27-10
Cell Source and Immunology • Cellular Composition •
Tissue Architecture • Tissue Thickness • Electrical
Function and Safety • Spontaneous Activity
27.5 Future Work .............................................. 27-15
27.6 Conclusions .............................................. 27-16
Acknowledgment................................................ 27-16
References ....................................................... 27-17
27.1 Introduction
After an acute myocardial infarction, lost cardiomyocytes are replaced by a noncontractile fibrous tissue.
Although it is suggested that heart has a small regenerative potential via cell proliferation [1], or stem
cell recruitment [2], the rate of renewal is insufficient to compensate for myocyte loss. As a result, altered
workload of a surviving myocardium may ultimately lead to deterioration in contractile function and
congestive heart failure (CHF). Besides traditional pharmacological therapies (diuretics, β-blockers,
angiotensine, and aldosterone inhibitors) [3] or heart transplant [4], investigators are evaluating innovat-
iveapproachesfor treatment of CHF including mechanical assist devices [5], dynamic cardiomyoplasty [6],
27-1
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27-2 Tissue Engineering
transmyocardial laser revascularization [7], and artificial heart [8]. Nevertheless, in end stage disease, heart
transplant remains the only option with good long-term results [4]. However, inadequate availability of
donor organs (∼10% of current needs [9]) requires new strategies for treatment of increasing number of
heart failure patients.
One promising approach is augmentation of the number of functional myocytes in the diseased heart
using methodologies for cardiomyocyte cell cycle activation [10], adult stem cell mobilization [11], or
cellular transplantation [12]. Cellular transplantation at the site of injury in the heart can be accomplished
either by injecting isolated cells (“cellular cardiomyoplasty”) [13], or by implanting a cardiac tissue patch
engineered in vitro (“tissue cardiomyoplasty”) [14].
27.1.1 Cellular Cardiomyoplasty
More than 10 years ago, pioneering studies in the laboratory of Lauren Field have shown feasibility of
cell transplantation in the heart [15,16]. Since then, different investigators have used cardiac [15] or
skeletal myoblast cell lines [16], fetal [17,18], neonatal [19], and adult [20] cardiac myocytes, autologous
[21,22] and syngeneic [23] skeletal myoblasts, smooth muscle cells [24], endothelial cells [25], native [26]
or genetically altered fibroblasts [27], embryonic [28], bone marrow [29], mesenchymal [30], or heart
derived [31] stem cells, as potential donor cells. As a result, treated hearts have shown improvement in
diastolic function, almost independent of transplanted cell type [26]. The improvement in the systolic
function (generation of active force) on the other hand required use of cells with the myogenic (contractile)
potential [26,32]. The possible therapeutic benefit of cellular cardiomyoplasty stems from structural
remodeling of scar region [33], enhancement of myocardial revascularization [12], and direct structural
and functional integration of donor cells with the host myocardium [18]. Presently, clinical studies in the
United States, Europe, and Asia are under way to investigate feasibility and safety in using autologous bone
marrow derived stem cells and skeletal myoblasts in treatment of postinfarction left ventricular disfunction
[13,34,35]. Initial results are promising, but reveal risk for ventricular arrhythmias [34], limiting in some
studies the pool of patients to only those that already have internal defibrillators. Since no systematic
studies have been performed to assess the electrical performance of the heart postcardiomyoplasty, and
little data exists on electrical interaction between donor and host cells in vivo or in vitro [36], the causes
of the postoperative electrical instability are not known. Plausible explanations include inflammatory
response and subsequent fibrosis at implantation site [35], possible electrical coupling in conjunction with
different electrophysiology between implanted cells and cardiomyocytes [18,34], and possible stimulation
of sympathetic nerve sprouting and overexpression of neurotransmitters after the cell transplantation [37].
27.1.2 Tissue Cardiomyoplasty
Some of the major hurdles in restoring heart function by cellular cardiomyoplasty include limited survival
of injected cells in the region of scar tissue, and no architectural repair of the infarcted area. An alternative
approach is tissue cardiomyoplasty, which involves in vitro cultivation of compact three-dimensional (3D)
cardiac tissue patch, and subsequent implantation over or instead of the infracted scar tissue. Although
surgically more challenging compared to cellular cardiomyoplasty, this methodology has a potential to
improve efficiency and localization of tissue repair in larger size cardiac injury such as infarction of major
coronary vessels, or congenital heart defects [38]. Ideally, based on the location, shape, and size of injury
and architecture of surrounding tissue (assessed by ultrasound, MRI, or other noninvasive technique
[39]), functional cardiac patch with needed geometry and 3D structure is engineered in vitro starting
from selected cell type, natural or synthetic scaffold, and appropriate culturing vessel (bioreactor). Inside
the bioreactor, cells attach to biocompatible (and possibly degradable) scaffold, interconnect, and assemble
in three dimensions to reconstitute an in vivo-like cardiac tissue equivalent (construct). The combination
of biochemical and physical stimuli during culture is designed to best mimic physiological state of tissue,
and to support cell differentiation or transdifferentiation and desired 3D tissue architecture. At a proper
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Cardiac Tissue Engineering 27-3
time point, tissue construct is removed from the bioreactor and surgically implanted in the site of injury
in order to restore or improve electrical and mechanical function of diseased heart.
In reality, however, the successful reconstitution of cardiac-like tissue patch in vitro starting from
single cells is an extremely challenging problem due to limited proliferation potential and high metabolic
demand of cardiac cells, as well as complex anisotropic architecture and electro-mechanical function
of native cardiac tissue. While recent reviews on cardiac tissue engineering [14,40] have focused on
scaffold biomaterials, bioreactors, and cultivation conditions, this chapter will provide emphasis on the
electrophysiological considerations and role of tissue architecture in the development of functional cardiac
patch. It is this author’s view, that these factors will play an important role in the design of efficient and
safe therapies, despite the fact that they are frequently neglected in current in vitro and in vivo studies.
27.2 Cardiac Architecture and Function
The crucial architectural and functional feature of cardiac muscle tissue (Figure 27.1) is anisotropy, that is,
anatomically and biophysically, properties of cardiac muscle vary in different directions [41]. Microscopic
structural anisotropy in cardiac tissue results from the spatial alignment of elongated cardiac myocytes
(Figure 27.1a), and the preferential location of intercellular junctions (e.g., fascia adherence, gap junctions,
desmosomes) in end-to-end vs. side-to-side cell connections [42]. Macroscopic anisotropy is a result of
the presence of aligned cardiac muscle fibers and sheets that transmurally rotate inside the heart wall (180
◦
rotation from endo- to epicardium) [43]. The unique anisotropic architecture of cardiac tissue enables
an orderly sequence of electrical and mechanical activity, and efficient pumping of blood from the heart.
Beside architecture, electrical membrane properties of cardiac myocytes also vary substantially depending
on the location in the heart with distinct differences between atria and ventricles, endo- and epicardial
regions, left and right heart, base and apex, etc. [44–46]. Moreover, the heart contains a large variety
of nonmyocytes (e.g., fibroblasts, endothelial cells, smooth muscle cells, neural cells, leukocytes) with
specific roles in the cardiac function that are still not fully elucidated.
Structural anisotropy and intercellular continuity of the excitable cardiac substrate have profound
effect on electrical and mechanical functioning of the heart. For example, anatomical anisotropy in heart
tissue causes a larger intracellular resistance per unit length in the transverse (across fiber) than longit-
udinal (along fiber) direction, resulting in smaller velocity but larger maximum slope of action potential
upstroke and safer electrical propagation in the transverse direction [47,48]. As a consequence, electrical
stimulation of a small region in the heart tissue results in development of elliptical rather than circular
propagating wavefront [49] (Figure 27.1b). Directly related to anisotropy is evidence that cardiac impulse
conduction at the microscopic level is discontinuous at sites of gap junctions, and even stochastic due to
small local variations in ion channel function, gap junction distribution, and adjoining tissue architec-
ture [50,51]. In contrast to these small physiological variations, larger variations of electrical properties
(e.g., action potential duration, intercellular coupling, electrical load) at the cellular and tissue level may
result in increased susceptibility to propagation slowing and conduction block [46,51]. Slow propagation
velocity and unidirectional block are some of the main prerequisites for the initiation of reentrant cardiac
arrhythmias [51,52].
The degree of anatomical and functional anisotropy depends on location in the heart and age of the
individual, with main determinants being cell size and geometry, type, amount, and distribution of
cell junctions in membrane, and macroscopic tissue architecture [53–55]. Electrical anisotropy changes
in certain cardiac pathologies such as ischemia, infarction, and heart failure [56,57]. This change is a
consequence of the altered gap junction distribution and expression, as well as formation of longitud-
inal collagenous septa between the cardiac fibers which result in discontinuous transverse propagation
(“nonuniform anisotropy”) and increased susceptibility to reentrant arrhythmias [58]. In particular, in
canine hearts, “border zone” between infracted and normal tissue exhibits disarray of cardiac cells and
gross change in anisotropy, resulting in conduction slowing, block, and reentrant “figure of eight” circuits