Scaffolds for tissue engineering of cardiac valves

Abstract

Tissue engineered heart valves offer a promising alternative for the replacement of diseased heart valves avoiding the limitations faced with currently available bioprosthetic and mechanical heart valves. In the paradigm of tissue engineering, a three-dimensional platform - the so-called scaffold - is essential for cell proliferation, growth, and differentiation as well as the ultimate generation of a functional tissue. A foundation for success in heart valve tissue engineering is recapitulation of the complex design and diverse mechanical properties of a native valve. This article reviews technological details of the scaffolds that have been applied to date in heart valve tissue engineering research.

Full text

Review
Scaffolds for tissue engineering of cardiac valves
S. Jana, B.J. Tefft, D.B. Spoon, R.D. Simari
⇑
Division of Cardiovascular Diseases, Mayo Clinic, 200 First Street SW, Rochester, MN 55905, USA
article info
Article history:
Received 27 November 2013
Received in revised form 25 February 2014
Accepted 12 March 2014
Available online 24 March 2014
Keywords:
Heart valve
Tissue engineering
Scaffold
Fiber
Hydrogel
abstract
Tissue engineered heart valves offer a promising alternative for the replacement of diseased heart valves
avoiding the limitations faced with currently available bioprosthetic and mechanical heart valves. In the
paradigm of tissue engineering, a three-dimensional platform – the so-called scaffold – is essential for
cell proliferation, growth and differentiation, as well as the ultimate generation of a functional tissue.
A foundation for success in heart valve tissue engineering is a recapitulation of the complex design
and diverse mechanical properties of a native valve. This article reviews technological details of the
scaffolds that have been applied to date in heart valve tissue engineering research.
Ó2014 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.
1. Introduction
Valvular heart disease (VHD) is a major health problem that
results in substantial morbidity and death worldwide [1]. In the
western world, 2.5% of the population have a dysfunctional or
diseased valve [2,3]. Secondary to the aging of the population, it
is predicted that there will continue to be an increase in VHD in
industrial nations, owing primarily to an increase in degenerative
pathology [2]. In the UK alone, more than 4 million people from
75 to 84 years of age could develop VHD by 2018, and this figure
could double by 2028 [4]. In developing countries, VHD is primarily
caused by the persistent burden of rheumatic fever rather than degen-
erative pathology, and tends to affect younger individuals [5,6].
The pathophysiology of valvular heart disease is broad and the
specific etiology varies by the particular valve affected. The semilu-
nar valves, consisting of the aortic and pulmonic valves, are com-
monly affected and have distinct primary pathologic mechanisms
of failure. Pulmonic valve disease is most commonly related to con-
genital abnormalities and tends to present early in life. Aortic valve
disease most commonly presents as calcific aortic valve stenosis
secondary to calcific degeneration [7,8], while the presence of a
congenitally bicuspid aortic valve predisposes to subsequent val-
vular stenosis and regurgitation usually at the earlier age [9].
Calcific aortic valve stenosis is the most common valvular
pathology requiring valve replacement and is present to some
degree in 2.8% of adults over the age of 75 years, with a far larger
population showing some evidence of aortic valve thickening,
known as valvular sclerosis [10,11]. Despite the frequency of cal-
cific aortic valve stenosis, our understanding of its pathogenesis
remains incomplete. While there are similarities between the risk
factors and mediators between calcific aortic valve disease and
atherosclerosis, as many as 50% of patients with calcific aortic
valve disease do not show any evidence of significant atherosclero-
sis [12,13]. Recent data demonstrate that valvular calcification is
not a passive process, as originally thought, but rather an active
process that relies on the activation of pro-osteogenic signaling
cascades, such as bone morphogenetic protein and Wnt/b-catenin,
for the induction and progression of disease [14,15]. Additionally,
our understanding of the cellular mediators of valvular calcifica-
tion continues to expand. Conventionally, the differentiation of
valvular interstitial cells into an osteoblast-like phenotype with
the capacity to produce calcification has been thought to be the
primary cellular driver of valvular calcification [16]. Recently,
valvular endothelial cells have been implicated through a process
of endothelial–mesenchymal transformation, as have circulating
progenitor cells through differentiation or paracrine signaling
[17–20]. The calcification process results in the mechanical
disruption of valve function, which can lead to stenosis or
regurgitation, or a combination of the two.
Unfortunately, the treatment of dysfunctional heart valves
requires surgical or interventional repair or replacement.
Replacement options currently include mechanical or bioprosthet-
ic valves. Mechanical valves have excellent durability; however,
the risk of thromboembolism necessitates the use of anticoagulation
therapy and its attendant morbidity. Bioprosthetic valves are less
thrombogenic; however, they are less durable and more prone to
degeneration, particularly when implanted in younger individuals
http://dx.doi.org/10.1016/j.actbio.2014.03.014
1742-7061/Ó2014 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.
⇑
Corresponding author. Fax: +1 507 538 6418.
E-mail address: simari.robert@mayo.edu (R.D. Simari).
Acta Biomaterialia 10 (2014) 2877–2893
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[21]. Bioprosthetic valves are generally treated with glutaraldehyde
(GA) to stabilize the tissue by preventing rejection of the xenogenic
scaffold. However, such treatments stiffen the fiber network and
diminish the cushioning function of the spongiosa layer [22].In
addition, GA is toxic and inhibits the repopulation of cells after
implantation [23]. Both mechanical and bioprosthetic valves share
another disadvantage: they cannot grow and remodel, which may
necessitate sequential surgeries in pediatric patients [24].
Nevertheless, the current generation of bioprosthetic pericar-
dial valves are adequate substitutes for the majority of elderly
patients as they typically do not require anticoagulation and their
durability is usually sufficient for the lifespan of this population. In
the pediatric and young adult populations requiring aortic valve
replacement, the Ross procedure has been used, and has been
shown to have low perioperative mortality and rates of reoperation
[25,26]. In this procedure, a patient’s diseased aortic valve is
replaced with his/her own modified pulmonic valve (autograft)
and then a cadaveric pulmonic valve allograft is used to replace
the pulmonic valve. This procedure has several advantages, includ-
ing minimal thromboembolism, favorable hemodynamics and the
potential for valve growth. A disadvantage of the procedure is
the harvesting of the healthy pulmonic valve, which can lead to
the development of pulmonary valve disease in addition to aortic
valve disease. As an alternative, tissue engineering is a promising
approach for the treatment of defective or diseased heart valves
[27]. In this method, living cells are grown (in vitro or in vivo) onto
a supporting three-dimensional (3-D) biocompatible structure to
proliferate, differentiate and ultimately grow into a functional tis-
sue construct (Fig. 1)[28–30]. Importantly, a tissue engineered
valve may be capable of growth and remodeling, and may mitigate
the need for anticoagulation.
The scaffold is one of the most important entities to be consid-
ered for efficient tissue engineering because its external geometry,
surface properties, pore density and size, interface adherence, bio-
compatibility, degradation and mechanical properties affect not
only the generation of the tissue construct in vitro, but also its
post-implantation viability and functionality [31,32]. All scaffolds
designed for tissue engineering applications must meet basic
requirements, such as biocompatibility, sterilizability and mechan-
ical integrity. Scaffolds intended for heart valve tissue engineering
face additional distinct challenges due to their direct contact with
blood. Specifically, the construct should be resistant to calcification
and thrombosis [33]. In addition, the construct must withstand the
unique hemodynamic pressures and flows of the cardiac environ-
ment from the moment of implantation. These unique challenges
underscore the importance of carefully considering the materials
and design when fabricating a scaffold for tissue engineered heart
valves.
Semilunar valves in human (pulmonic and aortic) consist of
three semicircular leaflets (also called cusps) attached to a
fibrous annulus called the root [23]. The leaflets are less than
1 mm thick and have a flexible structure consisting of three
distinct layers: the fibrosa, spongiosa and ventricularis (Fig. 2).
These layers are composed of valvular interstitial cells (VICs)
within a matrix of collagen, elastin and glycosaminoglycans
(GAGs). Normal leaflets are virtually avascular and obtain nutri-
ents and oxygen from the bloodstream via hydrodynamic con-
vection and diffusion. In contrast, the aortic or pulmonary root
is a bulb-shaped fibrous structure, with intimal, medial and
adventitial layers. They are primarily populated with endothelial
cells in the intima, smooth muscle cells in the media and fibro-
blasts in the adventitia.
Fig. 1. Schematic diagrams of aortic heart valve tissue engineering. Living cells are grown onto a supporting three-dimensional (3-D) biocompatible structure to proliferate,
differentiate and ultimately grow into a functional tissue construct.
2878 S. Jana et al. /Acta Biomaterialia 10 (2014) 2877–2893

In brief, the structure of semilunar heart valves is complex.
To mimic this organization in heart valve engineering, the first
requirement is an appropriate scaffold structure comprising suit-
able morphologies, surface properties, mechanical properties and
pore sizes. In this respect, a broad knowledge of potential heart
valve tissue engineering scaffolds is important for their applica-
tion. However, to date, only a limited number of studies have
examined the outcomes of the different scaffolds used for heart
valve tissue engineering. This review thus focuses on the scope
of usefulness of different scaffolds to regenerate heart valve tis-
sue. It first provides a brief summary of the physiological prop-
erties of native valve that can be mimicked by scaffolds
intended for heart valve engineering. It then covers the most
common types of scaffold and scaffolds with a combination of
various structures used to regenerate heart valve tissue [34].
Finally, it focuses on their drug delivery capabilities, which are
important for regeneration of functional heart valves and for
mitigation of thrombogenicity. The article also outlines the
strategies of in vitro and in vivo populations of cells that have
been tested.
2. Physicochemical properties of native valve
The physicochemical properties of a scaffold, including its
architecture/morphology, mechanical properties and drug delivery
capabilities, are important parameters in tissue engineering as
together they create a microenvironment on which cell adhesion,
proliferation, migration, differentiation and subsequent tissue
development depend [35]. In order to optimize the results of heart
valve tissue engineering, scaffolds should mimic the physicochem-
ical properties of native heart valve. In this respect, a brief discus-
sion on the formation and subsequent physicochemical properties
of native valve could define the set of goals for the design of scaf-
folds intended for heart valve tissue engineering. Heart valve cusps
are thin; however, they reliably perform under high blood flows
and pressures. This is possible due to the organization and orienta-
tion of collagen, elastin and other extracellular proteins within the
heart valve. The valve cusp structure is not just a single organized
extracellular matrix (ECM) layer inhabited by cells. It is more com-
plex and organized into three layers: a layer of inflow surface (ven-
tricularis), with a radially aligned fibrous morphology; a central
layer (spongiosa), containing loose collagen and abundant proteo-
glycans, randomly distributed; and a layer of outflow surface (fib-
rosa), consisting of circumferentially aligned collagen fibers (Fig. 2)
[22,23,36].
During development, heart valves go through several essential
stages: (i) the endocardial cushion is formed and endothelial–mes-
enchymal transformation (EMT) takes place; (ii) endocardial
cushion growth and subsequent formation of the valve primordial;
and (iii) thinning and elongation of the valve primordia to form
valve cusps. At each stage, different signaling and transcriptional
molecules regulate valvulogenesis [36–38]. The transforming
growth factor beta (TGF-b) cytokine superfamily, including bone
morphogenetic proteins (BMPs), acts to support cushion formation
initiation by increasing the ECM synthesis and inhibiting the
expression of chamber-specific genes [39]. BMP-2 and BMP-4 are
responsible for increased deposition of hyaluronan and versican
in cushion morphogenesis regions of the outflow tract [39,40]. Vas-
cular endothelial growth factor (VEGF) from endothelial cells
encourages the proliferation of endocardial cushion endothelial
cells [38]. Endocardial cushion EMT is initiated by the TGF-bfamily
[41]. TGF-breceptors and ligands are also expressed during EMT
[38]. During endocardial cushion EMT, signaling of Wnt/b-catenin
is associated with TGF-band increases endocardial cushion EMT
[42]. The expression of VEGF terminates EMT and promotes endo-
thelial cell proliferation, leading to cushion growth [38]. At the end
of EMT, through cell proliferation and ECM deposition, endocardial
cushions continue to grow, and endocardial cushions ultimately
fuse to form valve primordia [37,43]. BMP signaling and fibroblast
growth factors (FGFs), including FGF4 and FGF receptors 1, 2 and 3,
are responsible for both these developments [44,45]. Wnt/b-
catenin and VEGF/NFATc1 are also involved in these processes
[46,47]. In thinning and elongating of valve primordia into valve
cusps, Notch signaling, BMP2 signaling and EGF4 expression have
Fig. 2. Schematic diagrams of an aortic heart valve. (a) Side section view of aortic valve in the heart. (b) Top view of aortic valve (seen from aortic side). (c) Side view of
splayed-open valve showing semilunar shape of cusps. (d) Section view through the cusp and aortic wall showing the three-layered structure of the cusp. (e) Three cusp
layers – fibrosa, spongiosa and ventricularis – showing their collagen fibril orientations.
S. Jana et al. /Acta Biomaterialia 10 (2014) 2877–2893 2879

been regarded as important [48]. In addition, large numbers of
intermediate signals, such as Tb 20, Twist1, periostin, cadherin-
11, Sox9 and Msx1/2, are responsible for valvulogenesis [49–52].
Thus, the ultimate structure of the adult valve is the result of a
complex interplay between cells, matrix, cytokines and growth fac-
tors over relatively long periods. Engineering approaches must
consider this interplay throughout the process.
Mechanical properties of human aortic valves in both the cir-
cumferential and radial directions have been shown to differ
[53]. The modulus of elasticity, ultimate stress and ultimate strain
of 15.34 ± 3.84 MPa, 1.74 ± 0.29 MPa and 18.35 ± 7.61%, respec-
tively, were found in the circumferential direction. In the radial
direction, these values were 1.98 ± 0.15 MPa, 0.32 ± 0.04 MPa and
23.92 ± 4.87%, respectively. The mechanism responsible for these
dissimilar properties is the orientation of the collagen and elastin
fibers in the valve. The circumferential alignment of collagen in
the fibrosa increases the stiffness and strength in that direction,
whereas the radial alignment of elastin in the ventricularis
increases the elasticity in that direction [54]. Proper functioning
of the valve relies heavily upon these anisotropic mechanical prop-
erties, as too much stiffness can hinder coaptation of the cusps
whereas too little stiffness can allow for distortion and regurgita-
tion. It is also necessary to achieve proper mechanical properties
because VICs are highly sensitive to local tissue strains and
respond by modulating behavior, such as collagen synthesis [55].
Thus, it is imperative to produce scaffolds made of either biological
or synthetic materials with similar anisotropic mechanical proper-
ties when engineering a valve.
3. Types of scaffolds for heart valve engineering
The goal of heart valve tissue engineering requires the design of
a scaffold that provides physiological support for cell attachment,
proliferation and development. The complex structure of a heart
valve includes a spongy middle layer sandwiched between two
outer laminar anisotropic fibrous layers [56]. To mimic the native
heart valve structure, multiple scaffold designs have been pro-
posed and tested. Two main types of scaffold have been developed:
(i) acellular native heart valve scaffolds, from allogeneic/xenoge-
neic sources; and (ii) fully artificial scaffolds, fabricated from syn-
thetic and natural (biological) polymers [57]. The fabricated
scaffolds can be further categorized as porous, fibrous and hydro-
gel scaffolds. The following sections assess these distinct types of
scaffold used in heart valve tissue engineering.
3.1. Decellularized valve scaffolds
Xenogeneic heart valves from pigs, cows and sheep are the main
sources of acellular heart valves because human valves – the
allogeneic valves that are potentially less antigenic compared to
their xenogeneic counterparts – are in short supply, especially for
pediatric patients (Fig. 3a). Decellularized scaffolds retain the origi-
nal valve structure and many of the ECM molecules, providing
potential advantages over fabricated synthetic scaffolds (Fig. 3b).
In this approach, applying one of many tested decellularization
methods, the cells are removed, minimizing damage to the original
structure, to avoid problems during subsequent recellularization or
implantation [58,59]. In some cases, the decellularized tissue scaf-
folds are modified by crosslinkers such as GA and pentagalloyl glu-
cose (PGG) to sterilize the valve, minimize disease transmission,
reduce immunogenicity and avoid negative outcomes, such as cal-
cification and thromboembolism [60].
Different single agents or combinations of agents have been
used to decellularize valve tissues, and each treatment has shown
advantages and disadvantages [61]. The most commonly used sin-
gle-agent treatments are nonionic detergents, ionic detergents and
chelating agents [62,63]. Nonionic detergents such as Triton X-100,
sodium deoxycholate acid (SDC), sodium dodecyl sulfate (SDS), and
deoxycholic acid disrupt lipid–lipid and lipid–protein connections
but not protein–protein connections [64,65]. They lyse nuclear
materials but cannot completely dissolve the fragments, which
leads to residue in the scaffolds. In contrast, ionic detergents
remove cytoplasmatic and nuclear cellular materials completely
[66,67]. The loss of glycosaminoglycans, chondroitin sulfate and
other proteins, including collagen, laminin and fibronectin, is
reduced by treatment with ionic detergents as compared to non-
ionic detergents [68,69]. Thus, the scaffolds are more intact com-
pared to those treated with nonionic detergent. Chelating agents,
such as ethylenedaminetetracetic acid (EDTA), remove the cellular
material from the tissue by isolating bivalent cations such as Mg
2+
and Ca
2+
, which are necessary for cells to attach to collagen within
the ECM [70,71].
In the enzymatic method of single-agent decellularization, pro-
teolytic enzymes such as trypsin cleave peptide bonds to decellu-
larize the valves. Endonucleases and exonucleases degrade both
RNA and DNA [72,73]. Endonucleases catalyze the disruption of
the interior bonds of ribonucleotide or deoxyribonucleotide chains
whereas exonucleases catalyze the disruption of the terminal
bonds of these chains. Trypsin-treated scaffolds showed fragmen-
tation and distortion of elastin fibers [74]. Collagen distribution
was also distorted, and glycosaminoglycans, laminin, fibronectin
and chondroitin sulfate were almost completely washed out after
treatment. However, resynthesis of chondroitin sulfate, laminin
and fibronectin has been achieved by seeding endothelial cells in
the scaffolds. Most of the enzymes used in the decellularization
process come from non-human sources and non-recombinant
sources, raising the possibility of disease transmission through
the decellularization process [75].
Decellularization with a combination of reagents has shown
advantages over single-agent treatments. Cadaveric valves treated
with trypsin and Triton X-100 revealed maintained fibrous scaffold
structure and mechanical strength with low antigenicity [66].
Detergent (e.g. EDTA/sodium deoxycholate and Triton X-100) and
enzymatic (e.g. RNase and DNase) reagents were also used
together to remove cells from cadaveric heart valves. The mechan-
ical properties of the decellularized scaffolds showed no significant
change, with retention of sufficient collagen and elastin along and
glycosaminoglycans [76]. Zhao et al. prepared decellularized por-
cine heart valves in the same manner and cultured them with
canine endothelial cells (ECs) [22]. The results showed that leaflet
scaffolds were covered with cells [77]. A combination of EDTA/
sodium deoxycholate, Triton X-100, RNase and DNase treatments
was used to decellularize porcine heart valves, which were then
cultured with ECs [78]. The ECs formed a layer on the surface of
the scaffolds. Korossis et al. treated fresh porcine aortic valve leaf-
lets with hypotonic buffer, agarose gel, trypsin and SDS to obtain
Fig. 3. Photographs of porcine aortic heart valves. (a) Native. (b) Decellularized.
Table 1 Mechanical properties of PHAs. Table modified from Williams et al. [128].
Reproduced with permission.
2880 S. Jana et al. /Acta Biomaterialia 10 (2014) 2877–2893

acellular scaffolds [79]. The scaffolds showed significant loss of
elastin fibers in the ventricularis layer and increased transition
strain compared to fresh porcine samples. Therefore, in acellular
scaffolds, the mechanical strength and other leaflet competencies
under different systematic treatments were more or less
compromised.
The effect of different decellularization reagents on cells is dif-
ferent, thus these reagents need variable time periods for complete
cell removal. In addition to removing cells, these reagents damage
various proteins in the valve in a time-dependent manner. There-
fore, researchers have investigated different time periods of valve
decellularization to achieve cell removal with minimal damage to
ECM components. For example, Cebotari et al. applied 1% SDC solu-
tion, 1% SDS solution, and a mixture of 0.5% SDC solution and 0.5%
SDS solution for 24 h under continuous shaking at 22 °C[80].
Decellularized valves were then washed with phosphate-buffered
solution in 10 steps, with each step for 12 h at 22 °C under contin-
uous shaking. Sacks and colleagues altered the time periods of
three previously investigated decellularization methods (anionic
detergent-based: SDS; enzymatic agent-based: trypsin; and non-
ionic detergent-based: Triton X-100) to 48 h to compare their
effects on the mechanical and structural properties of decellular-
ized valves [81]. They soaked cadaverous valves in a hypotonic Tris
buffer (10 mM Tris, pH 8.0) with 0.1% EDTA and 10 kIU ml
1
apro-
tinin for 1 h and then treated with 0.1% SDS in the same hypotonic
Tris buffer, protease inhibitors, RNase A (20 mg ml
–1
) and DNase
(0.2 mg ml
–1
) for 48 h (original was 72 h) during continuous shak-
ing at room temperature [82]. Another combination, trypsin and
EDTA (0.5% trypsin and 0.2% EDTA) in a hypotonic Tris buffer with
RNase A (20 mg ml
–1
) and DNase (0.2 mg ml
–1
) was used for
decellularization for 48 h during continuous shaking at 37 °C
[83]. On the other hand, 1% Triton X-100 was mixed with 0.2%
EDTA (Sigma) in a hypotonic Tris buffer with RNase A (20 mg ml
–
1
) and DNase (0.2 mg ml
–1
) and then applied for decellularization
for 48 h (original was 24 h) during continuous shaking in a cell cul-
ture incubator (i.e. 5% CO
2
/95% air atmosphere at 37 °C) [84]. In all
of the above methods, decellularized valves were washed with
copious PBS during shaking, usually at room temperature. The
SDS-based method seemed to best preserve the required mechan-
ical and structural properties of valves. Some investigators have
modified the above methods further to obtain better results in
terms of decellularization with less damage to the existing proteins
[66].
After decellularization, acellular scaffolds are sometimes trea-
ted with crosslinkers such as GA, PGG and nordihydroguaiaretic
acid prior to implantation. This treatment is performed to stabilize
the collagen matrix and decrease antigenicity [85,86]. GA has been
commonly used for crosslinking of ECM molecules in acellular scaf-
folds. However, these scaffolds have reduced cell compatibility as
evidenced by cell toxicity and reduced proliferative capacity of
cells following repopulation. The use of PGG, a collagen-binding
polyphenol crosslinker, results in greater retention of biaxial
mechanical and other biological properties in leaflets when com-
pared to GA-treated leaflets [86]. In vivo, these PGG-treated leaflets
did not calcify, and they supported infiltration by host fibroblasts
and subsequent matrix remodeling. To improve cell attachment
and seeding, cells can be encapsulated in polyethylene glycol
(PEG) hydrogel and then seeded into acellular scaffolds [87]. PEG
hydrogel was shown to assist the scaffold in retaining seeded cells
when the scaffold was tested in a bioreactor with pressurized med-
ium flow. Baraki et al. decellularized ovine aortic valve conduits
with detergents and implanted them as an aortic root in lambs
for up to 9 months [88]. The controls were fresh native ovine aortic
valve conduits. Explants showed trivial regurgitation and no sign of
graft dilation, degeneration or rejection. Minute calcium deposits
were observed at the anastomosis, together with the formation
of thrombi on the leaflets. There was an endothelial monolayer
at the luminal side and neovascularization at the adventitial side.
The control samples also demonstrated some adverse signs, includ-
ing calcification and degeneration. These findings suggest that
methods of decellularization and associated treatment affect the
mechanical and biological functions of implanted scaffolds.
If crosslinking agents are used, scaffolds exhibit less porosity, a
smaller pore size and a stiffer ECM. A harsh treatment, such as ace-
tic acid, can be applied to acellular scaffolds to increase their pore
size and porosity, and the resulting scaffolds can be conjugated
with Arg-Gly-Asp (RGD) for higher cell adhesion [89]. When cul-
tured with human mesenchymal stromal cells in vitro, scaffolds
treated with acid showed increased proliferation and migration
of cells deep into the scaffold.
In addition to chemical and biological reagents, physical meth-
ods, such as temperature, force, flow/pressure and non-thermal
irreversible electroporation, have also been applied to remove cells
from heart valves. A freeze–thaw process has been shown to lyse
cells, while subsequent processing was required to remove the
intracellular contents [90]. This process also disrupted the ECM
ultrastructure, and this disruption must be considered [91,92].
Cells from the surface of a tissue have been removed by physical
or chemical means [92]. Hydrostatic pressure may be more effec-
tive than detergents and enzymes for removing cells from the sur-
face [93]. Electrical pulses have also been applied to kill cells by
inducing the formation of micropores in the cell membrane [94].
Following exposure, the cellular remnants need to be removed
from the tissue.
As decellularized scaffolds are made of ECM, original scaffolds
still degrade and are replaced by new ECM deposition [95,96].
The degradation rate depends on a variety of factors, including
the decellularization treatment, the crosslinking technique and
the use of structural modification agents, such as epigallocate-
chin-3-gallate [97–99]. It is understood that each cell removal
agent and method causes some degree of structural disruption
and alteration of the ECM composition. Thus, the objective of
decellularization should be to minimize the undesirable effects
and to improve the recellularization efficiency so that recellular-
ized bioscaffold implants can be used for tissue engineering.
Despite promising results in ovine implantation models, the
clinical outcomes of decellularized valves have been disastrous
[100,101]. The SynerGraft (Cryolife, Inc., USA) trial involved the
implantation of decellularized porcine valves into the right ventric-
ular outflow tract of four children and all four grafts failed due to
an inflammatory response (at day 2), leading to structural failure
(at day 7) and degeneration of the leaflets (after 6 weeks) and wall
(after 1 year) [100]. The decellularization process was predicted to
remove the majority of antigenic material from the tissue, thus
rendering it relatively inert. However, subsequent analysis of the
pre-implantation SynerGraft scaffolds revealed that they were
inadequately decellularized and still contained immunogenic cell
fragments and DNA [100]. Thus, a strong inflammatory response
occurred and no host cells were found to repopulate the tissue.
Interestingly, favorable results have since been reported with
decellularized allogeneic valves in terms of immunological
response, durability and overall clinical performance [102–105].
There has even been a report of host cell repopulation [106]. Nev-
ertheless, pre-seeding these scaffolds with host cells is likely to
yield more favorable results.
As decellularized constructs relying on host recellularization,
allogeneic valves have proven to be far superior to xenogeneic
valves. Nevertheless, xenogeneic valves have the distinct advan-
tage of being in plentiful supply, and strategies are being explored
to mitigate the immunological response to decellularized xenoge-
neic valves. For example, alpha-galactosidase and gene knock-out
technology have been used to remove the critical alpha-Gal
S. Jana et al. /Acta Biomaterialia 10 (2014) 2877–2893 2881

xenoantigen [107,108]. Antigen masking by gluteraldehyde
fixation is undesirable for a tissue engineered valve and has been
found to be ineffective [109]. Ideally, a decellularized scaffold,
whether xenogeneic or allogeneic, will be seeded with autologous
cells that will degrade the foreign matrix and synthesize an
entirely autologous tissue to eliminate immunological responses
following implantation.
3.2. Fabricated scaffolds
Direct fabrication of a scaffold mimicking the structure and
function of a heart valve from raw synthetic or biological materials
is an approach distinct from decellularization of biological scaf-
folds [34,110]. Potential advantages of fabricated polymeric scaf-
folds may include less immunogenicity and thrombogenicity
[111]. In addition, the biodegradability, durability and mechanical
properties of these scaffolds can be controlled. The disadvantages
are that their structure may not fully mimic the complex structure
and function of native tissue with varied mechanical properties. In
general, three types of scaffold have been fabricated and applied to
heart valve tissue engineering: porous scaffolds, fibrous scaffolds
and hydrogels.
3.2.1. Solid 3-D porous scaffolds
Highly porous 3-D scaffolds with an interconnected homoge-
neous pore network and a large pore size provide a continuous
flow of nutrients and metabolic waste to enable growth and vascu-
larization of engineered tissues [34]. A number of fabrication
methods have been applied to natural and synthetic polymeric
materials to produce porous scaffolds that are suitable for heart
valve tissue engineering. The conventional techniques for scaffold
fabrication are particulate leaching, solvent casting, gas foaming,
vacuum drying, thermally induced phase separation, melt molding,
high internal phase emulsion and microfabrication [112–115].
Although various groups have applied these relatively straightfor-
ward methods, they all have some drawbacks. Most importantly,
the pores are of irregular size and are not interconnected. Solid
free-form (SFF) fabrication is a technique to prepare 3-D porous
scaffolds with interpore connections. The scaffolds are created
using computer-controlled tools for layer-by-layer deposition of
materials [116,117]. The geometry can be obtained from a solid
model file or digital data produced by such imaging sources as
microcomputer tomography and magnetic resonance imaging.
Three-dimensional printing is an advanced form of SFF, with one
or several cartridges containing different biomaterials, biomole-
cules or other required materials [118,119]. With any of these
methods, a broad range of materials, including synthetic or natural
and organic or inorganic, could be used to produce 3-D porous
scaffolds for the engineering of a range of tissues, including heart
valves.
Polyhydroxyalkanoates (PHAs) are perhaps the materials most
often used to produce solid 3-D scaffolds applied to heart valve tis-
sue engineering [120–123]. PHAs are a class of natural polyester
polymers produced by different microorganisms [124]. Initially,
polyglycolic acid (PGA) and polylactic acid (PLA) were used to pre-
pare valve leaflets as promising alternatives to decellularized
biomatrix scaffolds. However, their relatively high stiffness and
rigidity made it impossible to obtain adequate cell proliferation
of applied cells [125–127]. Stock et al. then introduced poly-
hydroxyoctanoate (PHO), a soft elastomeric polymer and member
of the PHA family, to PGA materials to overcome these mechanical
issues [127]. They prepared porous (pore size 80–180
l
m) PHO
leaflets (120
l
m thickness) using the salt leaching method and
coated the construct with PGA felt. The scaffold construct was then
cultured with autologous smooth muscle cells and endothelial
cells, and their progressive cellular and ECM deposition indicated
desirable remodeling of the tissue engineered structure. After
24 weeks of implantation in a sheep model, there was no thrombus
formation and only mild, nonprogressive valvular regurgitation
with the seeded implant, whereas the unseeded implant demon-
strated thrombus formation after 4 weeks.
Besides PHO, other members of the PHA family have demon-
strated promising results in heart valve tissue engineering [23].
Various hydroxy acid monomers have been incorporated to make
PHA polymers, and these monomers can be substituted with other
groups, such as alkyl and aryl groups. PHAs have been of interest in
biomedical applications and the emerging field of tissue engineer-
ing due to their high biocompatibility and tunable mechanical
properties [122]. Several companies, e.g., Metabolix, Inc. and
Tapha, Inc., have been developing PHA-based polymers, including
poly-3-hydroxybutyrate (poly(3HB)), poly-3-hydroxybutyrate-co-
3-hydroxyvalerate (poly(3HB-co-3HV), poly-R-3-hydrooxyoctano-
ate-co-R-3-hydroxyhexanoate (poly(3-HO-co-3HH)), poly-4-
hydroxybutyrate (poly(4HB)) and poly-R-3-hydroxybutyrate-co-
4-hydroxybutyrate (poly(3-HO-co-4HB)) [128]. Their wide range
of thermal properties allows for the use of different processing
methods for a variety of product development needs, including
scaffolds for tissue engineering. Mechanical properties of the PHA
polymers vary widely and thus make them suitable for engineering
both soft and hard tissues (Table 1)[128]. Their mechanical prop-
erties can be varied further by extending the pendant group. For
example, when the hydroxyvalerate group of poly(3HB) is
extended, the polymer becomes more elastomeric (compare col-
umns 2 and 4 in Table 1).
Trileaflet heart valve scaffolds were prepared from porous PHA
foam (Tepha, Inc.) with pore sizes of 180–240
l
m and compared
with acellular biomatrix scaffolds (Fig. 4)[129]. Vascular cells from
ovine carotid arteries were seeded onto these scaffolds, which
were then implanted into a lamb model. Interestingly, all samples
except the acellular scaffolds were covered with tissue and exhib-
ited collagen and ECM deposition. No thrombi formed on any PHA
scaffolds. In another study, vascular cells were seeded into PHA
leaflet scaffolds (Meatabolix Inc, 0.3 mm thick) with pore sizes of
100–240
l
m and similar results were observed [130]. The cells
were viable and ECM deposition was observed following pulsatile
flow exposure within a culture system. Expression of crosslinked
Table 1
Mechanical properties of PHAs.
a
Poly
(3HB)
Poly
(4HB)
Poly(3HB-co-20%
3HV)
Poly(3HB-co-16%
4HB)
Poly(3HO-co-12%
3HH)
Native Cusp in circumferential
direction
Native Cusp in radial
direction
Tensile strength
(MPa)
40 104 32 26 9 NA
b
NA
Tensile modulus
(GPa)
3.5 0.149 1.2 NA 0.008 15.34 (MPa) 1.98 (MPa)
Elongation at
break (%)
6 1000 50 444 380 18.35 23.92
a
Table modified from Williams et al. [128]. Reproduced with permission.
b
NA, not available.
2882 S. Jana et al. /Acta Biomaterialia 10 (2014) 2877–2893

elastin by vascular smooth muscle cells was also observed in PHA
scaffolds [131]. Some researchers tried to improve the biocompat-
ibility of P-3HB-co-3HH porous scaffolds by coating with silk
fibroin and found improved growth of smooth muscle cells on
the silk fibroin-modified hybrid scaffolds [132].
Beside PHAs, other natural polymers, such as chitosan and col-
lagen, have been used to prepare porous scaffolds for heart valve
engineering. A collagen sponge with irregular connecting pores of
58 ± 24
l
m diameter facilitated proliferation of seeded porcine
aortic myofibroblasts [133]. The cells also synthesized a consider-
able amount of ECM, such as collagen and elastin. Chitosan-modi-
fied PCL porous scaffolds were fabricated to improve the
attachment of fibroblast cells for heart valve tissue engineering
[134]. This experiment also demonstrated that chitosan enhanced
the biocompatibility of the PCL scaffolds, which has since been
used extensively in valve and blood vessel tissue engineering.
A composite scaffold prepared in combination with poly(
D
,
L
-lac-
tide-co-caprolactone), poly(
D
,
L
-lactide-co-glycolide) (PLGA) and
type I collagen have been tested to determine their efficacy in heart
valve engineering [135]. The scaffold had a void volume of 80%.
Seeded neonatal rat heart cells showed higher cardiac marker
expression and contractile properties in composite scaffolds com-
pared to controls made of each individual material. Appropriate
porosity, hydrophilicity, structural stability, elasticity and degrada-
bility were the main reasons for the improved results.
Scaffolds with proper morphology can guide cells and conse-
quently the deposition of proteins that replace the gradually
degrading scaffold to achieve the required organization [136].To
obtain native collagen fiber orientation and mechanical properties
in engineered heart valves, a number of approaches have been
used. Engelmayr and colleagues demonstrated the use of microfab-
ricated scaffolds with diamond-shaped pores and observed a corre-
lation between the aspect ratio of the pores and the alignment of
collagen fibers [137]. The microarchitecture of the scaffold influ-
enced the structural orientation of the deposited collagen fibers
through contact guidance of the cellular orientation [138].
Accordion-like honeycomb-microstructured scaffolds made of
polyglycerol sebacate (PGS) were found to generate native-like
ventricular myocardium tissue constructs with proper stiffness
and anisotropy due to the scaffold’s precise biophysical environ-
ment directing the seeded neonatal rat heart cells [139]. The
architecture guided the cell alignment, whereas the controllable
stiffness of PGS in conjunction with electrical stimulation offered
an in vivo biophysical environment to the cells that enabled their
growth and maturation into a desirable tissue construct.
Some current research indicates that PGS could be a worthwhile
material for heart valve tissue engineering [140]. It is a biocompat-
ible, biodegradable and, more importantly, a tough elastomer, with
mechanical and degradation properties that can be controlled by
manipulating the polymerization parameters. Diamond-shaped
microporous PGS scaffolds were used for heart valve tissue engi-
neering by culturing VICs for up to 28 days. The cells were able
to secrete collagen and the deposited collagen was aligned inside
the diamond-shaped micropores [141]. Despite interesting biome-
chanical properties, PGS has poor water uptake capacity (2%),
which limits its utility for tissue engineering, especially of soft tis-
sues. Patel et al. modified PGS with PEG to prepare PGS-co-PEG
block copolymers which had controllable mechanical, degradation
and water uptake properties [142]. Due to the presence of the PEG,
the mechanical properties improved from 13 kPa to 2.2 MPa and
elongation increased sixfold. Coating PGS scaffolds with laminin,
fibrin, collagen-I, fibronectin and/or elastin increased their cellu-
larity, ECM production and regulation of cell phenotype [143].
Although positive results were found during both in vitro and
in vivo testing in terms of cell adhesion, proliferation, differentiation
and ECM deposition on fabricated 3-D solid porous scaffolds, these
tissue engineered valve scaffolds lacked two important features of
native valves: shape and elastomeric flexibility. It is very challenging
to fabricate 3-D porous scaffolds exhibiting the shape and elasticity
of native valves. Bioprinting is a leading technology that may be used
to fabricate scaffolds with the above requirements.
3.2.2. Fibrous scaffolds
Fibrous scaffolds have achieved great importance in tissue
engineering since they resemble the structure of ECM [144]. Due
to the high aspect ratio of fibers, fibrous scaffolds are superior to
Fig. 4. Tissue engineered aortic valve from PHO polymer-based porous scaffold. (a) Photographic image of aortic valve after 5 weeks in vivo implantation. SEM images of
leaflet surfaces after (b) 1 week, (c) 5 weeks and (d) 17 weeks of implantation. (e) Hematoxylin and eosin-stained conduit walls showing ingrowth of vascularized tissue
islands and destruction of polymer (filled arrow). (f) Movat pentachrome-stained leaflet showing significant collagen deposition (open arrow) [129]. Reproduced with
permission.
S. Jana et al. /Acta Biomaterialia 10 (2014) 2877–2893 2883

non-fibrous scaffolds in terms of cell adhesion, migration,
proliferation and differentiation, which are important in tissue
engineering applications [145–147]. Moreover, fibers have a high
growth factor loading efficiency and sustained release capacity to
specific sites of application. For heart valve tissue engineering, fibrous
scaffolds would provide an ideal environment for cells if they could
form 3-D structures with porosity, pore size and mechanical charac-
teristics comparable to native heart valves [148].
Several techniques, including electrospinning, phase separation
and self-assembly, have been applied to fabricate nanofibrous
materials. Electrospinning is the most commonly used technique
to prepare tissue engineering scaffolds due to its versatility, appli-
cability to most polymers, easy handling and cost-effectiveness
[149]. In electrospinning, a high electrical voltage is applied to a
polymer solution and, when the voltage is strong enough, built-
up charges on the surface of the polymer droplets overcome the
surface tension and a liquid jet forms. Then, via electrostatic force,
the jet accelerates toward the grounded collector and is stretched
to produce continuous ultrathin fibers [150]. Applying collectors of
appropriate shapes and arrangements allows the fabrication of
nanofibrous scaffolds with desirable forms and fiber orientations
(e.g. random or aligned).
A majority of the nanofibrous scaffolds intended for heart valve
tissue engineering have been fabricated from PGA polymer. PGA
has many favorable properties, such as biocompatibility and bioab-
sorbility; however, the mechanical stiffness is high (7 GPa) [151].
More importantly, due to its hydrophilicity, PGA monofilaments
can be degraded in less than 4 weeks in vivo, thus providing a con-
structive environment and stimulation for higher collagen and
ECM production [152]. To prepare nanofibers, PGA has been dis-
solved in organic solvents such as 1,1,1,3,3,3-hexafluoro-2-propa-
nol or chloroform to make a solution that can be electrospun by
applying a voltage of 15–20 kV. Hoerstrup et al. cultured human
fibroblast cells on electrospun PGA fibrous meshes with fiber diam-
eters of 12–15
l
m to construct autologous cardiac valves [153].On
a similar kind of PGA mesh, human fibroblast cells followed by
human endothelial cells were cultured to obtain similar constructs
(Fig. 5)[154]. Endothelial cells formed a monolayer, with no forma-
tion of capillaries being detected [125]. In native valves, valvular
endothelial cells (VECs) form the outer layer and work in a similar
fashion; thus the layering of endothelial cells onto a tissue con-
struct seems a rational approach towards functional valve tissue
engineering.
The compatibility of PGA non-woven fibrous meshes for heart
valve tissue engineering was further improved by dip-coating the
meshes with a 1% solution of poly(4HB), a homopolymer that has
a lower degradation rate compared to PGA [155–157], thus the
poly(4HB) restricted the degradation of PGA. After culturing autol-
ogous ovine fibroblast and endothelial cells for 20 weeks, the sur-
face of the tissue construct was smooth, like that of a valve. The
ECM and DNA contents were higher on the modified PGA meshes
compared to on the PGA meshes alone [155]. Cultured bone-mar-
row-derived mesenchymal stromal cells showed differentiation
to the smooth muscle cell lineage, characterized by their expres-
sion of markers such as alpha-smooth muscle actin, desmin and
calponin [157]. Hoerstrup et al. tested PGA/poly(4HB) hybrid scaf-
folds seeded with human bone marrow stromal cells and obtained
results in terms of myofibroblast cell marker expression, tissue and
ECM formation, and biomechanical properties. In addition, the sur-
face of the engineered leaflet was smooth [158]. Umbilical cord
cells exhibited proliferation, growth and differentiation into the
smooth muscle cell phenotype. A layered tissue formed, which
contained collagen I, collagen III and glycosaminoglycans. Collagen
fibril formation was also observed [156]. Due to the higher com-
patibility of the hybrid scaffold compared to the PGA scaffold
alone, the former resulted in engineered tissue constructs with
superior mechanical properties. Further studies on leaflet tissue
engineering with this PGA/poly(4HB) hybrid scaffold revealed that
the addition of the growth factor VEGF and the differentiation fac-
tor TGF-b1 increased proliferation, differentiation and tissue devel-
opment [159]. TGF-b1 was suggested to be responsible for
transdifferentiation from endothelial to mesenchymal cells.
Several investigators have molded PGA mesh into valve-shaped
scaffolds by applying P4HA. Molded scaffolds were then cultured
with appropriate cell lines, including autologous ovine myofibro-
blasts and endothelial cells and autologous bone marrow-derived
mononuclear cells, in vitro and implanted in vivo to determine
the biocompatibility of the scaffold systems [155,160,161]. The
transplanted valve constructs maintained the valvular structure
and showed adequate functionality (leaflet mobility and coapta-
tion) for up to 4 weeks. There were no signs of thrombus formation
or structural damage. From histology and immunohistochemistry,
valve constructs were found to obtain layered endothelialized tis-
sue due to substantial cellular remodeling and in-growth into the
scaffold material.
Fig. 5. Nanofibrous scaffold for heart valve tissue engineering. (a) SEM image of continuous fibroblast/connective tissue on nanofibrous scaffold. (b) CD34 stain
demonstrating the formation of an endothelial monolayer (A) on the surface of a core of fibroblasts (B) and hydrolysis of polymer fibers after 4 weeks (C). Reproduced with
permission from [154].
2884 S. Jana et al. /Acta Biomaterialia 10 (2014) 2877–2893

Besides PGA, other synthetic polymeric materials have been
applied to heart valve engineering. For example, poly(ester
urethane) urea-based fibrous scaffolds with specific anisotropic
mechanical properties could be appropriate for valve tissue engi-
neering [148]. The anisotropic behavior of such scaffolds was
obtained by high-speed electrospinning. PGA/PLLA composite
fibrous scaffolds were studied to evaluate post-implant character-
istics in heart valves [162]. A natural polymer chitosan-based com-
posite scaffold was prepared by incorporating chitosan fibers into a
3-D porous chitosan scaffold. The fibers augmented the modulus
and strength of the scaffold [163]. Heart valves engineered with
this composite scaffolds showed tensile strength as high as
220 ± 17 kPa, which is comparable to the radial values of human
valve leaflets.
The rate of biodegradation of scaffold materials is an important
parameter in tissue engineering because biodegradation reduces
the mechanical properties of scaffold materials, but this may be
balanced by ECM production [34]. Proper material selection or syn-
thesis in a specific tissue engineering application is necessary to
achieve this balance. Sant et al. applied a blend of fast-degrading
polyglycerol sebacate (PGS) and comparatively slow-degrading
polycaprolactone (PCL) to produce scaffolds with controlled bio-
degradation for heart valve tissue engineering [164]. The PGS–
PCL hybrid scaffolds showed a gradual decrease in mechanical
properties due to biodegradation, which was compensated by the
secretion of new ECM by cells within the scaffold.
One of the unmet challenges of nanofibrous scaffolds is their
very small pore size (less than 10
l
m), which prevents cells from
penetrating into the scaffolds. Polymers with appropriate
biodegradable properties could be useful in nanofibrous heart
valve tissue engineering if the rate of biodegradation could be
synchronized with cell proliferation and ECM production. Thinly
coating decellularized scaffolds with nanofibers could possibly
bring their efficiencies together in heart valve tissue engineering.
3.2.3. Hydrogel scaffolds
Hydrogels are swollen hydrophilic polymer chain networks
with a high water content. Due to their structural similarity to
ECM, they have been used as polymeric materials in both tissue
engineering and regenerative medicine [165,166]. Hydrogel mate-
rials generally show high permeability to oxygen, nutrients and
water-soluble metabolites. They can also be delivered percutane-
ously, promoting their use for cell encapsulation in tissue engi-
neering [167]. These materials have many other applications as
well, including drug delivery, biosensors and linings for artificial
implants [168–171].
Hydrogels are synthesized by crosslinking hydrophilic homo-
polymers, copolymers or macromers [172]. Through free radical
polymerization, the polymer chains propagate by consuming vinyl
monomers or other functional groups such as acrylate-based deriv-
atives [173]. The reaction proceeds quickly and uncontrollably,
leading to a wide distribution of molecular weights and inhomoge-
neous properties throughout the hydrogel. An alternative synthesis
method is the conjugate addition reaction based on acrylated
monomers, esters or amides combined with thiols [174–176]. This
method brings the risk of side reactions by competing nucleophiles
from biological compounds, including living cells. Some hydrogels
are prepared by click chemistry, which offers mild reaction condi-
tions with high chemical selectivity [177–179]. Thus, the quality of
hydrogels prepared by this method is far superior to that of hydro-
gels produced by conventional methods. Poly(vinyl alcohol) (PVA),
PEG and polyacrylates such as poly(2-hydroxyethyl methacrylate)
are some of the synthetic monomers used to prepare hydrogels.
Biological hydrogels can be formed from collagen, fibrin, hyalu-
ronic acid, alginate or chitosan.
Encapsulation of cells within a hydrogel is a standard technique
for cell culturing and tissue engineering [167]. The polymer chains
in a hydrogel mimic the structure of proteins and other
biomolecule chains within the ECM; thus, cell-seeded hydrogels
offer high cellular efficiency in tissue engineering through greater
cell adhesion, proliferation and differentiation. PEG-based
hydrogels were used as seeding media for bone marrow mesenchy-
mal stromal cells (BMSCs) in decellularized porcine aortic valves
for better cell attachment [87]. The differentiation of BMSCs into
endothelial and myofibroblast cells was higher in scaffolds with
PEG hydrogel compared to only decellularized scaffolds. Decellu-
larized aortic valves seeded with PEG-encapsulated cells demon-
strated greater tensile strength compared to the aortic valves
seeded without PEG hydrogel. Culture of porcine VICs and VECs
in collagen-GAG hydrogels showed enhanced surface coverage of
VECs on collagen–GAG constructs compared to collagen-only
constructs [180]. Expression of elastin and laminin by VICs and
expression of vasoactive molecules and endothelial nitric oxide
synthase by VECs were higher on collagen–GAG constructs
compared to collagen-only constructs. Bovine aortic valve
endothelial cells, bovine valvular interstitial cells and bovine aortic
endothelial cells were viable in porous amino acid hydrogels; thus,
these hydrogels could be used to engineer heart valves [181].
Hydrogels have weak mechanical properties and their stiffness
decreases further with the addition of cells [59,182,183]. Therefore,
to make cell-encapsulated hydrogels work as a stand-alone scaf-
fold for tissue construct development, the mechanical properties
of hydrogels need to be augmented. A modified PEG hydrogel –
polyethylene glycol diacrylate (PEGDA) – is an attractive material
for heart valve tissue engineering due to its tunable mechanical
and biological properties [184,185]. Three layers of different PEG-
DA hydrogels, each with mechanical properties that correspond
to the three layers of a native heart valve, were sandwiched to pre-
pare cusp scaffolds. Under simulated flow and pressure conditions
in a pulsatile bioreactor, the interfaces and the sandwich remained
intact [186]. Another hydrogel system was developed consisting of
gelatin macromers synthetically modified with methacrylate func-
tionalities. The stiffness could be regulated by adjusting the poros-
ity, which was in turn dependent upon the polymerization rate
[169]. VICs encapsulated within the composite hydrogel were able
to achieve native morphology within 2 weeks of culture. PVA
hydrogels have mechanical properties similar to soft tissue [169].
By regulating freeze–thaw cycles, the PVA hydrogels obtained
mechanical properties similar to those of native porcine aortic
roots within the physiological pressure range [187]. To improve
the mechanical properties of PVA hydrogels directly without
repeated freeze–thaw cycles, PLA has been included for crosslink-
ing [188]. The complex geometry of an aortic valve could be
molded to produce scaffolds for heart valve engineering.
Despite supporting cell adhesion and proliferation, deposited
ECM and collagen have no specific orientation since hydrogels do
not possess any definite structure. Hockaday et al. used a 3-D
printing/photo-crosslinking system with a PEGDA hydrogel to pre-
pare complex and heterogeneous aortic valve scaffolds (Fig. 6)
[189]. The scaffolds were supplemented with VIC-encapsulated
alginate gel. After 21 days in culture, it was found that, due to
the printed orientation and the presence of alginate gel, the engi-
neered scaffold constructs from the blended hydrogel had a stiff-
ness that was 10 times greater than the scaffold constructs
engineered from the original hydrogel. This indicates that the for-
mer scaffold system can withstand dynamic physiological pres-
sures. Duan et al. used a 3-D bioprinter to prepare scaffolds from
alginate/gelatin-blended hydrogels [190]. Aortic root sinus smooth
muscle cells (SMCs) and aortic VICs were encapsulated separately
into alginate and gelatin gels and then mixed to prepare the blend
before printing. After 7 days of culture, the mechanical properties
S. Jana et al. /Acta Biomaterialia 10 (2014) 2877–2893 2885

of the cell-laden scaffolds were much higher than those of the
acellular 3-D printed hydrogels. Bioprinting of cellular microsphe-
roids which can work as bio-ink particles may be useful for fabri-
cating three-dimensional structures like heart valves through the
self-assembly of cells [191]. Challenges remain for the use of extru-
sion-based techniques to produce heart valves due to issues like
the flexibility and mechanical properties of current valves. Modern
printers can use only low-viscosity solutions, which may not be
able to duplicate valve structure and function. Cross-linking of
component materials in a low-viscosity solution could improve
the mechanical properties, but that works against cell compatibil-
ity. A more functional bioartificial valve was made by Tranquillo
and colleagues using a cell-encapsulated fibrin gel generated in a
mold that applied appropriate mechanical constraints to the gel
in order to obtain both proper alignment of deposited fibers and
proper cusp geometry (Fig. 7)[192,193].
Hydrogels are of particular interest in biomedical applications
due to their efficacy in growth factor delivery [168]. The inclusion
of TGF-b1 in metalloproteinase-degradable PEG hydrogels
increased the expression of alpha smooth muscle actin (
a
-SMA)
and collagen-1 after 2 days of culture, indicating the differentiation
of myofibroblasts [194]. Similar observations were made with gel-
atin–methacrylate hydrogels in the presence of TGF-b1[169]. The
incorporation of fibronectin-derived RGD peptides, which promote
integrin binding, increased VIC process extension and integrin
alpha(v)beta(3) binding [194]. A metalloproteinase-degradable
PEG hydrogel system could be useful for characterizing the tran-
sient state of VICs – either fibroblasts or myofibroblasts – a useful
cell type for heart valve tissue engineering [194]. PEG hydrogels
have been shown to reduce the calcification on engineered leaflet
scaffold tissues [195]. The addition of fibrin or fibronectin to PEG
hydrogels further reduced calcification. Furthermore, their addi-
tion did not change the expression of
a
-SMA, a marker for myofib-
roblastic activity.
Promotion of the endothelialization of implanted devices is
another goal for heart valve tissue engineering. Antibody-modified
polysaccharide-based hydrogels are capable of capturing circulat-
ing endothelial progenitor cells (EPCs). Applying this technique,
Camci-Unal et al. prepared CD34 antibodies immobilized on hyalu-
ronic acid (HA) hydrogels and used them for valve tissue engineer-
ing with seeded EPCs [196]. After 2 days, the cells spread better on
modified HA hydrogels compared to unmodified hydrogels. This
technique could be useful for improving the biocompatibility of
implants, including artificial heart valves, by endothelialization.
Hydrogels, including protein gels, were found to be effective
materials for bioprinting complex valve shapes. The elastomeric
flexibility of the printed valves could be adjusted through appro-
priate crosslinking concentration and time, by either radiation or
chemicals. If the cells are encapsulated, then there can be a com-
promise between crosslinking concentration and time for flexibil-
ity and cell survivability. Crosslinked scaffolds can have issues
with sustainability in a dynamic environment. Several weeks of
static incubation of printed hydrogel scaffolds may increase their
mechanical characteristics through cell proliferation and ECM
deposition. Ultimately, dynamic conditioning of the cultured
valves could align the ECM protein fibers and bring the overall
structural morphology closer to that of a native valve. None of
these hurdles have yet been overcome, thus more studies need to
be done to bring printed hydrogel scaffolds to clinical trials, or at
least to in vivo experiments.
3.3. Scaffolds of combined morphologies
Scaffolds may be generated through the combination of materi-
als with distinct morphologies and/or mechanical properties. Spe-
cially, the presence of a nanofibrous structure in combined
scaffolds might provide both nanoscale and microscale structures.
Moreover, the inclusion of a nanoscale fiber coating on scaffolds
caused an increase in mechanical properties compared to original
scaffolds (without a nanofiber coating) when both were cultured
with mesenchymal stromal cells, although cell proliferation was
the same in both scaffolds [110]. Nanofibers present a similar
structure as ECM, and this plays a key role in tissue architecture
by providing structural support and tensile strength [34]. Chitosan
fibers were produced through solution extrusion and a neutraliza-
tion method, and incorporated into a chitosan scaffold [163]. This
fiber-reinforced engineered heart valve scaffold achieved leaflet
tensile strength values of 220 ± 17 kPa, which are comparable to
the radial values of native valve leaflets. A nanofiber coating on
the scaffold surface increased the strength of the scaffold, which
caused greater deposition of collagen fibers, leading to increased
mechanical properties of the tissue engineered valves. In some
Fig. 6. Bioprinting of scaffolds for heart valve tissue engineering. (a) 3-D printer setup showing the printing mechanism. (b) Porcine heart valve isolated for micro-CT scan. (c)
Computerized 3-D valve reconstruction from CT scan data. (d) Computerized slice of the valve reconstruction showing the calculated print head paths. (e) Printed valve where the
root was formed with 700 MW PEGDA hydrogel and the leaflets were formed with 700/8000 MW PEGDA hydrogel. Scale bar = 1 cm. Reproduced with permission from [189].
2886 S. Jana et al. /Acta Biomaterialia 10 (2014) 2877–2893

cases, decellularized scaffolds were coated with nanofibrous
matrix to improve the mechanical properties and degradation
kinetics of the original scaffolds. Hong et al. coated acellular heart
valve scaffolds with poly(3-hydroxybutyrate-co-4-hydroxybuty-
rate) fibers by electrospinning and seeded them with mesenchy-
mal stromal cells (MSC). The hybrid scaffold exhibited better
mechanical properties, higher cell proliferation and increased
ECM production compared to the acellular biomatrix scaffold
[110].
Most of the heart valve tissue engineering substrates produced
so far, including porous 3-D scaffolds, have been based on single-
layer designs because generating appropriate trilayered scaffolds
is a daunting task [197–200]. The results of animal studies have
revealed that, in most cases, only the outer layer was remodeled,
indicating that a more sophisticated scaffold design may be neces-
sary. One of the approaches used to make a scaffold with a trilay-
ered structure consisted of creating a porous spongiosa layer
sandwiched between two collagen layers [30]. The porous spongi-
osa layer was obtained from decellularized and elastase-treated
porcine pulmonary arteries. One of the outer layers, the ventricu-
laris, was also collected from decellularized porcine pericardium,
while the other outer layer, the fibrosa, was created using a fibrous
collagen scaffold. Both outer collagen layers were treated with PGG
to stabilize them against rapid biodegradation without crosslink-
ing. The middle spongiosa layer was seeded with stromal cells
derived from human bone marrow and positioned between the
two outer layers. The final construct was cultured in a heart valve
bioreactor and subjected to physiological pressure. The seeded
stromal cells showed the phenotype of valvular interstitial cells,
indicating that mechanical loads in the bioreactor specific to an
aortic valve induced the required phenotype without the presence
of specific growth factors.
In another approach, trilayered scaffolds were prepared from
synthetic materials with appropriate mechanical properties in
each layer. In the native heart valve, the two outer layers are stif-
fer than the soft middle layer. Tseng et al. prepared such a mul-
tilaminate scaffold consisting of three layers from a hydrogel,
PEGDA, which has tunable mechanical and biological properties
(Fig. 8)[186]. The two outer layers were stiffer compared to the
middle layer. The mechanical properties of the layers were varied
by adjusting the molecular weight and concentration of the poly-
mer. Since all the layers were made of hydrogel, the scaffold sys-
tem did not fail at the interfaces when subjected to force. This
scaffold, with a robust interface, integrated layers of different
mechanical properties, and biofunctionalization, was more
appropriate than single-layered scaffolds for heart valve leaflet
engineering.
Despite these good results, the trilayered constructs were sim-
ply three-layered membrane structures that may not develop into
functional leaflets attached to a heart valve conduit. Indeed, func-
tional leaflets from three-layered membrane structures have not
been produced yet, possibly due to a lack of integrity among the
layers or due to the difficulty of molding it into the shape of a
valve. Since the native heart valve has anisotropic mechanical
properties, fibrous materials could be helpful in preparing scaffolds
with anisotropic properties by aligning the fibers in the required
directions [148].
4. Drug delivery capabilities
In the development process of various tissues, including heart
valve, growth factors are involved at different stages of maturity.
In this process, growth factors are crucial for regulating cell migra-
tion into and differentiation within scaffolds. As cells within heart
valves have limited proliferative capacity, the application of
growth factors might be used to aid functional tissue engineering
in vitro, ex vivo and in vivo. Examples of this approach include
coating of decellularized aortic heart valves with heparin–VEGF,
which not only acted as an antithrombotic agent but also induced
adhesion, proliferation and migration of endothelial progenitor
cells on the decellularized valve [201]. In addition, heparin was
shown to be responsible for a synergistic increase in
a
-SMA
expression in valvular interstitial cells [202]. VEGF acted to
increase the proliferation of endothelial cells and reduce the calci-
fication generated by valvular interstitial cells [201]. In contrast, it
was observed that TGF-b1 alone had a tendency to induce the for-
mation of calcific nodules, which did not form in the presence of
both fibronectin and TGF-b1[203].
The use of nanoscale morphology promotes growth factor
release due to its high surface area with respect to volume,
i.e., its high aspect ratio. Taking advantage of this feature, sev-
eral researchers have modified scaffolds with nanoscale drug
carrier coatings for use in heart valve tissue engineering. For
example, decellularized matrix scaffolds coated with nanofibers
were made by electrospinning a mixture of basic fibroblast
growth factor, chitosan and poly(4HB). The scaffold was cultured
with mesenchymal stromal cells for 14 days. The results
revealed a substantial increase in cell mass and strength due
to the presence of 4-hydroxyproline and collagen in the hybrid
heart valve leaflets [110]. In another example, decellularized
valves were modified with PEG nanoparticles loaded with TGF-
b1. It was observed that, in the presence of both PEG and
TGF-b1, hybrid scaffolds possessed superior biocompatibility,
biomechanical properties and ECM microenvironment compared
to unmodified decellularized scaffolds [28]. Growth factors also
influence the phenotype of delivered and retained cells
[159,204]. Both human and sheep fibroblasts were differentiated
to metabolically active and functional myofibroblasts in the
presence of TGF-b1. Similarly, TGF-b1 differentiated ovine aortic
valvular endothelial cells and circulating endothelial progenitor
cells to a mesenchymal phenotype [159]. VEGF was also applied
within a scaffold system to induce seeded stromal cells (e.g.
human mesenchymal stromal cells) to differentiate into endo-
thelial cells [205].
Fig. 7. Photograph of mold and bileaflet valve after 6 weeks of compaction. (a) Side
view of the bileaflet valve mold. (b) Top view of a valve in static conditions. (c)
Ventricular view of a valve under pressurized conditions (greater than 5 mmHg)
showing coaptation of the leaflets [192]. Reproduced with permission.
S. Jana et al. /Acta Biomaterialia 10 (2014) 2877–2893 2887

When unseeded scaffolds are implanted, a favorable environ-
ment needs to be created to induce host cells to migrate from
the surrounding tissues, to be retained, and to proliferate and dif-
ferentiate into specific cell types [206–208]. In an in vitro study,
Somers et al. employed various growth factor cocktails within
unseeded scaffolds to define their ability to induce proliferation,
migration and invasion of ovine mesenchymal stromal cells
[209]. Interestingly, different combinations of growth factors had
different impacts on proliferation, migration and invasion. Growth
factors may have a negative impact on tissue engineering. Granu-
locyte colony-stimulating factor administration accelerated heart
valve deterioration to a degree similar to that observed in fresh
xenogeneic bioprosthetic valves [210]. Thus, prior to scaffold fabri-
cation, careful planning is needed for the addition of relevant
growth factors at different stages of development and maturity
to ensure desirable results.
5. Thrombogenicity
Thrombogenicity is one of the major issues facing artificial heart
valve scaffolds [211]. Thrombosis occurs due to the adsorption of
blood proteins such as fibronectin, vitronectin, fibrinogen and
von Willebrand factor onto the scaffold surface, causing blood con-
tact activation, platelet activation, and thrombin and fibrin forma-
tion in blood plasma [212]. These phenomena are very
complicated, consisting of a cascade of reactions which can be
divided into two main pathways: intrinsic and extrinsic [213].In
this section, we do not discuss these pathways; rather, we briefly
discuss the properties of the scaffold biomaterial that cause the
thrombogenicity and ways to prevent it.
The surface chemistry of scaffold biomaterials, which correlates
to material properties such as surface energy, intermolecular force,
relative charge distribution on the surface, ionic interaction and
texture, is mainly responsible for the protein adsorption that leads
to thrombogenicity [214]. Blood flow characteristics are also
responsible for blood coagulation to some extent. Proteins are
composed of amino acid subunits that gain or lose charge through
their side chains interacting with the surrounding medium,
depending on the pH of the medium and which of the charged
regions are exposed [214]. Different forces and interactions
between the molecules on a biomaterial surface and the charged
protein molecules in blood are involved in protein adsorption. As
a result, there is a stronger tendency for proteins to adsorb onto
a hydrophobic surface than onto a hydrophilic surface.
Several strategies, such as surface modification, bioactive coat-
ing and endothelialization of the surface, have been applied to
improve the hemocompatibility of scaffold surfaces [212]. Anti-
or noncoagulant materials such as PEG, pyrolytic carbon, elastin,
heparin, thrombomodulin and phosphorycholine can be used for
coating heart valve scaffolds to improve their hemocompatibility
[215]. PEG (and also polyethylene oxide) has a hydrophilic ether
oxygen in its molecular repeat unit which discourages protein
adsorption. To achieve a PEG coating, different PEGylation meth-
ods, including making the surface covalent, chemisorptive or phys-
isorptive, can be used [216]. Absorptive surface PEGylation
methods are superior to covalent surface PEGylation methods
because the latter induce inflammation and lead to thrombosis.
Although all of these anticoagulant materials have achieved suc-
cessful results in laboratory tests, most of them have failed in clin-
ical studies. This is especially true for heart valve scaffolds, likely
due to the presence of a rigorous dynamic environment [212].
Fig. 8. Schematic diagram depicting the fabrication of trilayer quasilaminates. Gel A represents 12.5% 3.4 kDa PEGDA and gel B represents 10% 6 kDa PEGDA. Reproduced with
permission from [186].
2888 S. Jana et al. /Acta Biomaterialia 10 (2014) 2877–2893

Endothelialization is another potentially useful method to pre-
vent thrombogenicity of implanted valves because a layer of endo-
thelial cells offers a non-thrombogenic surface similar to a native
heart valve [217]. Two main methods, in vitro endothelialization
and self-endothelialization, have been applied to form a layer of
endothelial cells on scaffold surfaces [218]. For in vitro
endothelialization, a patient’s endothelial cells are collected and
cultured on the scaffold surface to form a layer of endothelial cells.
Some other cells, such as fibroblasts and/or SMCs, can be co-cul-
tured with endothelial cells to form a better endothelial cell layer
due to higher ECM production from fibroblasts and SMCs [219].
To improve endothelialization, the surface chemistry of scaffolds
can be changed through such chemical phenomena as plasma
vapor deposition [220].
Self-endothelialization of implanted heart valve scaffolds is
possible if the scaffold materials can induce cell adhesion [218].
Several cell types, including endothelial cells and endothelial pro-
genitor cells (EPCs), that circulate in the blood are capable of cell
adhesion [221,222]. To improve cell attachment, the coating of
scaffold surfaces with ECM proteins such as collagen, laminin,
fibronectin and vitronectin may be useful [213]. Scaffold surfaces
modified with RGD peptides have demonstrated improved cell
adhesion. The orientation of immobilized RGD, its spatial arrange-
ment and ligand density were all found to influence endothelial
cell adhesion [223–226]. For example, cyclic RGD is more efficient
compared to linear RGD for cell adhesion. Besides RGD, proteins
specific to EPCs are useful for EPC adhesion. However, due to the
presence of different proteins on scaffold surfaces, unwanted cells
may adhere to scaffold surfaces and differentiate into inflamma-
tory cells [212]. Therefore, scaffolds should be coated with appro-
priate proteins for adhesion of endothelial cells and/or EPCs.
Bioprosthetic valves that are currently used to replace diseased
valves are mildly thrombogenic, so anticoagulants and antiplatelet
agents (such as warfarin and aspirin) are prescribed for some per-
iod of time after implantation [227]. Although decellularized heart
valves lose some proteins during the cell removal process, they
retain most major proteins, including collagen. However, thrombo-
sis is a significant concern when decellularized heart valves are
implanted for in vivo studies [228,229]. Fabricated heart valve
scaffolds that are mainly made of polymeric materials are also
thrombogenic, the degree depending upon the surface chemistry
and texture. Similar to bioprosthetic valves, these fabricated scaf-
folds require anticoagulants when implanted for in vivo studies
[230]. Among all of the strategies for antithrombogenicity, endo-
thelialization is currently seen as the best option for all types of
heart valve scaffolds, including decellularized valves, because this
most closely resembles the native strategy.
6. Conclusions and future directions
Tissue engineered heart valves are a promising alternative
option to current mechanical and bioprosthetic valves. Construct-
ing a heart valve requires a three-dimensional scaffold in which
cells can grow, proliferate and differentiate into a functional tissue
construct. Researchers have attempted to find effective approaches
for scaffold fabrication. Scaffolds should have appropriate mor-
phology with a trilayered structure and mechanical properties
with bipolar stiffness (15.34 and 1.98 MPa in circumferential and
radial directions, respectively), and should include expression (or
delivery) of growth factors and generation of ECM for effective
heart valve tissue engineering. Both decellularized and fabricated
scaffolds have advantages and disadvantages. Decellularized scaf-
folds mostly retain the original valve structure and ECM molecules,
which is a unique advantage over fabricated synthetic scaffolds.
Their mechanical stiffness is close to that of native valve. However,
their low pore size, porosity and cell survivability limit their poten-
tial for valve engineering. Similarly, most fabricated 3-D solid por-
ous scaffolds, fibrous scaffolds and hydrogel scaffolds cannot
properly mimic the trilayered structure of native heart valves,
although there have been some early studies of trilayered scaffolds.
The mechanical properties and morphologies of 3-D solid porous
scaffolds can be tailored according to requirements. Both decellu-
larized and fabricated scaffolds are able to deliver drugs. Building
upon past achievements, we need to optimize and test scaffold sys-
tems that can be applied to heart valve tissue generation.
While the focus of this review has been on scaffolds, any suc-
cessful strategy will need to carefully consider whether cells will
be added prior to implantation or whether ingrowth will simply
be allowed following implantation. If direct application is utilized,
the source and the methods to isolate, culture and deliver these
cells will need to be optimized. Finally, whether the construct is
conditioned or tested in vitro in a bioreactor must be considered.
We anticipate that the coordinated efforts of engineers, biologists,
surgeons and imagers will be required to achieve the goal of a clin-
ically available tissue engineered heart valve.
Disclosures
None.
Acknowledgements
This work is supported by the HH Sheikh Hamed bin Zayed Al
Nahyan Program in Biological Valve Engineering, the Grainger
Foundation, and the Mayo Clinic Center for Regenerative Medicine.
Appendix A. Figures with essential colour discrimination
Certain figures in this article, particularly Figs. 1 and 3–8 are dif-
ficult to interpret in black and white. The full colour images can be
found in the on-line version, at http://dx.doi.org/10.1016/j.actbio.
2014.03.014.
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