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Acta Biomater. Author manuscript; available in PMC 2015 January 01.
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Published in final edited form as:
Acta Biomater. 2014 January ; 10(1): . doi:10.1016/j.actbio.2013.08.022.
Electrospun Scaffolds for Tissue Engineering of Vascular Grafts
Anwarul Hasan, MDa,b, Adnan Memicc, Nasim Annabia,b, Monowar Hossain, MDd, Arghya
Paula,b, Mehmet R. Dokmecia,b, Fariba Dehghanie, and Ali Khademhosseinia,b,f,g,*
aCenter for Biomedical Engineering, Department of Medicine, Brigham and Women’s Hospital,
Harvard Medical School, Cambridge, Massachusetts 02139, USA
bHarvard-MIT
Division of Health Sciences and Technology, Massachusetts Institute of
Technology, Cambridge, Massachusetts 02139, USA
cCenter
of Nanotechnology, King Abdul-Aziz University, Jeddah, 21589, Saudi Arabia
dDepartment
eSchool
of Medicine, Lyell McEwin Hospital, University of Adelaide, SA 5112, Australia
of Chemical and Biomolecular Engineering, The University of Sydney, Sydney, 2006
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fWyss
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Keywords
Institute for Biologically Inspired Engineering, Harvard University, Boston, Massachusetts
02115, USA
gWorld
Premier International – Advanced Institute for Materials Research (WPI-AIMR), Tohoku
University, Sendai 980-8577, Japan
Abstract
There is a growing demand for off-the-shelf tissue engineered vascular grafts (TEVGs) for
replacement or bypass of damaged arteries in various cardiovascular diseases. Scaffolds from the
decellularized tissue skeletons to biopolymers and biodegradable synthetic polymers have been
used for fabricating TEVGs. However, several issues have not yet been resolved, which include
the inability to mimic the mechanical properties of native tissues, and the ability for long term
patency and growth required for in vivo function. Electrospinning is a popular technique for the
production of scaffolds that has the potential to address these issues. However, its application to
human TEVGs has not yet been achieved. This review provides an overview of tubular scaffolds
that have been prepared by electrospinning with potential for TEVG applications.
electrospinning; tubular scaffolds; vascular grafts; tissue engineering; burst strength; compliance;
mechanical properties
1. Introduction
Each year 1.4 million patients in US need arterial prostheses [1]. The available options for
replacement of vascular grafts have limited clinical success with a cost of more than US$25
billion [1]. Particularly, the pathologies affecting small and medium sized blood vessels are
© 2013 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.
*
Corresponding author: alik@rics.bwh.harvard.edu.
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the primary cause of death [2, 3]. Atherosclerosis is one of fatal diseases that causes buildup of plaques under the intimal layer, reducing the cross-sectional area available for blood
flow and thereby resulting in decreased flow of blood to the tissues downstream to the
plaque [4]. Eventually, cardiac and peripheral bypass surgeries become necessary, requiring
replacement of a segment of blood vessels. In blue baby syndrome, only one of the two
ventricles of an infant functions properly, and a “Fontan Operation” becomes necessary [5].
In “Fontan Operations”, an engineered blood vessel is required to connect the right
pulmonary artery with the inferior vena cava so that the deoxygenated blood can bypass the
heart and travel straight to the lungs. Similarly coronary artery diseases and peripheral
vascular diseases very often require the replacement of diseased and damaged native blood
vessels [6]. The currently available options for these transplants are autologous grafts (e.g.
coronary artery bypass graft with autologous mammary arteries and saphenous veins),
allografts (donor/cadaveric), xenografts (e.g. bovine or porcine pulmonary valve conduit),
artificial prostheses or synthetic vascular grafts made of expanded polytetrafluroethylene
(ePTFE) and polyethylene terephthalate (PET) [7].
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The use of autografts and allografts are limited due to the lack of tissue donors, previous
harvesting or anatomical variability [8, 9]. Besides, there are concerns about their long-term
functionality due to the use of strong detergents and decellularizing agents. Xenografts
suffer from their relatively shorter life span. As an example, a bovine or porcine graft may
last for up to fifteen years, which is a major issue for pediatric patients, which will require a
new implant replacement at every ten to fifteen years interval. Other issues include poor
control over physical and mechanical properties, inflammation and calcification [10, 11].
Synthetic prosthetic grafts are rejected within a few months by the immune system of the
body if the diameter of the vessel is smaller than 6 mm. This rejection arises from the
associated re-occlusion problem caused by thrombosis, aneurysm, and intimal hyperplasia
due to mismatch of compliance (compliance is opposite of the stiffness, measured as the
strain/expansion or contraction of the graft with force) [8, 10, 12, 13].
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Tissue engineering is an alternative approach for creating new vascular grafts. In this
approach cells are seeded or encapsulated in scaffolds fabricated from a biodegradable
polymer. In tissue engineering, it is anticipated that cells produce extracellular matrix
(ECM) while the polymer is degraded gradually creating the intended tissue. Extensive
research has been conducted on TEVG over the past few decades, and as a result significant
progress has been made in terms of achieving the remodeling of the tissue in the TEHV
constructs similar to the native blood vessels, [14–17] as depicted in Figure 1. Several
reviews have already been published that discuss about different methods for tissue
engineering of vascular grafts [3, 18], design of natural and artificial arteries as well as
vascular networks [19], and application of stem cell and other human cells for tissue
engineering of blood vessels [20]. Main approaches for fabrication of scaffolds include
methods that are based on molecular self-assembly, hydrogels, solvent casting-particulate
leaching technique, thermally induced phase separation and electrospinning process. The
latter is a versatile technique for fabrication of nano-micro scale fibers, which has a great
potential for mimicking the microenvironment of natural ECM. The success of an
implantable tissue engineered vascular graft is governed, among other factors, by the
development of a scaffold that mimics the ECM [21]. It is well known that in natural tissues
the ECM is a three dimensional (3D) network of 50–500 nm diameter structural protein and
polysaccharide fibers. Electrospinning has evolved to allow the fabrication of nanofiber
scaffolds in this size range and beyond [13].
Numerous review articles are available on fundamentals and applications of electrospinning,
its historical development, and various modifications [22–30]. These reviews illuminate the
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capability of electrospinning technique for fabrication of fibers from broad range of
polymers including purely natural to synthetic to composites mixtures of ECM analogs and
interstitial constituents from organ specific extracts [31–33]. These fibers with nano-micron
sized pores and large surface area have been used as catalysts, filtration systems, protective
clothing, drug delivery depots, optical wave guides, electronics, and tissue engineering
scaffolds [31, 32, 34–39]. The versatility of this technique makes it also possible to process
various categories of non-polymeric materials including ceramics and their composite with
polymer [34]. Electrospinning is a flexible technique in which the mechanical and biological
properties of fibers can easily be tuned by varying the composition of a mixture which is not
easily possible in other scaffold fabrication methods.
Due to these advantages, in recent years, the interest in electrospinning technique for the
fabrication of tissue engineering scaffolds has increased exponentially. In the current review
we present a summary of the most recent application of electrospinning for fabrication of
tissue engineered vascular grafts. The scope is limited to small and medium sized blood
vessels (diameter ≤ 5 mm), while large blood vessels (diameter > 5 mm), microvasculature
and capillary network are beyond the scope of this review. The aspects that are critical for
the design of TEVG such as structure of vascular graft and function are also discussed.
2. Structure of Vascular Grafts
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The native blood vessels in the human body have complex structures with distinct features.
The arterial wall is composed of three different layers: (i) intima, (ii) media and (iii)
adventitia, (Figure 2). Intima, the innermost of the three concentric layers, consists of a
continuous monolayer of endothelial cells (ECs) directly attached to the basement
membrane which is mainly comprised of connective tissues. Media is the middle layer
comprised of dense populations of concentrically organized smooth muscle cells (SMCs)
with fibers or bands of elastic tissues. Adventitia is consisted of a collagenous ECM that
mainly contains fibroblasts and perivascular nerve cells. In smaller arterioles and capillaries
some of these layers might be less obvious or absent [3]. An internal and an external elastic
lamina separate the intima, media and adventitia from each other. Each layer serves a
specific function: the collagenous adventitia functions to add rigidity while the elastic
lamina provides elasticity to the vessel walls.
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A confluent monolayer of EC in the lumen of the blood vessel is critical for preventing the
clotting of the blood, infection and inflammation of adjacent tissues. The EC monolayer is
also important for regulating the gaseous and molecular (oxygen and nutrients) exchange. It
also controls the signaling to the media, particularly to its muscular component [4]. The
SMCs in the middle layer (media) of the blood vessel walls have a unique contractile
function. When the signals from EC or cytokines stimulate the SMCs, the latter cells dilate
and contract in a coordinated manner. This leads to the dilation and contraction of the
vessels as the pressure in the vessels change. The concentric layer-wise structure of ECM,
the spatial organization and alignment of the EC and SMCs, and the interplay between the
cells and ECM structures are, therefore, important factors that should be taken into account
when designing tissue engineered blood vessels.
3. Functional Requirements for Vascular Grafts
It is crucial that TEVG functions, i.e. gets integrated to the adjacent blood vessels, sustains
the load from blood pressure, and allows blood flow without leakage, immediately after
implantation. Biocompatibility and bioactivity are other primary requirements for
engineering vascular grafts. In addition, the mechanical properties, adhesive ligands, growth
factor presentation, transport, and degradation kinetics of the materials used for the scaffold
should mimic the relevant ECM environment to a reasonable extent. It is favorable to use
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cost-effective and cytocompatible materials with tunable properties for a particular tissue
[37]. An exact replication of the native tissue structure is not always necessary, however, as
a functional construct, it is necessary that an engineered vascular graft fulfills certain basic
criteria [3, 14]. For example the burst strength above 260 kPa is required for a TEVG [6,
14], to prevent rupture due to variation in blood pressure. In addition, adequate elasticity is
crucial to withstand cyclic loading in which no dilation occurs in constructs within a month
of in vitro cyclic loading within physiological ranges [14]. The engineered vascular graft
should be compatible with the adjacent host vessel and provide an anti-thrombotic lumen
(autologous endothelium) [4]. The ability of the scaffold to provide initial mechanical
function for the vascular graft is another important factor, even though the structural and
mechanical characteristics of the native vessels are expected to be gradually acquired
through remodeling, repair, and growth upon implantation. Finally, it is also important that
the implanted vascular grafts minimize intimal hyperplasia and allow for regeneration of
arterial tissues.
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The uniaxial tensile strength of several native human blood vessels is presented in Table 1.
The properties of native grafts, as listed in Table 1, can be used as guidelines to target as
properties for tissue engineered grafts. Some safety factors may also be incorporated in the
design. There are some apparent discrepancies in literature among the reported values of
various mechanical properties for the same tissue by different researchers. For example, the
uniaxial tensile values of 43 MPa [40], 4.2 MPa [41], and 2.25 MPa [54] are reported for
human saphenous vein in the circumferential direction. This discrepancy is due to the
inconsistency in methods for extracting the results from the experimental load-displacement
data. The values for the elastic modulus in Donovan et al. [40], for example, was taken as
the maximum linear slope of the curve before failure while that in Stekelenburg et al. [41]
was taken from the initial linear elastic region of the curve. A reconciliation of literature
data is therefore important when comparing experimental results from different groups.
In developing any tissue engineering product including TEVG it is also crucial that
researchers take into account the requirements behind the FDA’s approval process to ensure
that FDA’s approval can be obtained for releasing the end product in the market and that it
can be used in the clinic. Researchers need to contemplate the safety, efficacy, purity and
identity of biomaterials in the engineering and design of products.
4. Application of Electrospun Scaffolds in Tissue Engineering of Vascular
Grafts
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Electrospinning offers the ability to fine-tune mechanical properties during the fabrication
process, while also controlling the necessary biocompatibility and structure of the tissue
engineered grafts. The ability of electrospinning technique to combine the advantages of
synthetic and natural materials makes it particularly attractive for TEVG where a high
mechanical durability, in terms of high burst strength and compliance (strain per unit load),
is required. In addition, incorporation of natural polymers, with abundance of cell binding
sites, can promote the formation of a continuous monolayer of EC in the lumen and
proliferation of other cell types in the matrix of the graft’s wall. The electrospinning
technique also offers precise control over the composition, dimension, and the alignment of
fibers that have impact on the porosity, pore size distribution and the architecture of
scaffolds. This method allows for engineering of a wide range of tunable structural and
mechanical properties as required for specific applications. Moreover, aligned nanofibers
can be used for orienting cells in a specific direction necessary to provide the anisotropy
encountered in certain organs including blood vessels. Companies have recently commenced
fabrication of electrospun grafts for transplantation of trachea, and other tissue engineered
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conduits. The world’s first tissue engineered tracheal transplant was successfully used in a
clinical trial in Sweden in June 2011 [42].
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4.1. Electrospinning Process and Parameters
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The electrospinning process is based on stretching of a viscoelastic solution into nano-micro
fibers using a high electrostatic force. In-depth reviews about the electrospinning process
can be found elsewhere [15–23]. In brief, the material to be electrospun is loaded into a
syringe and is pumped at a slow flow rate by a syringe pump (Figure 3a). A high DC voltage
is applied to the solution causing a repulsive force within like charges in the liquid. Under
the large applied electric field, the tip of the liquid droplet makes a cone shape, also called
the Taylor cone. When the applied voltage is high enough to overcome the surface forces
acting on the Taylor cone, a narrow jet of liquid generates from the Taylor cone and travels
toward the collector. An electrode of either opposite polarity or neutral (grounded) charge is
located nearby to attract and collect the fibers. As the liquid jet travels through the ambient
toward the collector, the solvent from the fiber jet evaporates and a solid fiber is deposited
on the collector. Schematics of three different modified experimental setups of
electrospinning method that can be used for fabricating multilayer, composite, and hybrid
scaffolds are presented in Figure 3. In one of these setup, a multilayered composite scaffold
is prepared by forming an electrospun tubular scaffold using a rotating mandrel as shown in
Figure 3a, followed by molding of a concentric layer of hydrogel around the electrospun
scaffold. It is also viable to prepare hybrid scaffolds from two types of fibers collected
simultaneously by electrospinning using setup in Figure 3b that contains a single mandrel A
modified approach has also been developed in which a hydrogel prepolymer is electro
sprayed concurrently with the electrospinning of fibers [1, 43] as shown Figure 3c. These
modifications allow fabricating scaffolds with multilayer structures, enhancing mechanical
and biological properties through the use of hybrid and composite structures. The
simultaneous electrospinning – electrospraying approach can be used to combine the
advantages of hydrogels and electrospinning, including uniform cell distribution throughout
the scaffold, enhanced cell attachment, spreading and proliferation, and improved
mechanical properties resulting from electrospun fibers.
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The morphology of electrospun fibers is affected by various parameters including the
density, viscosity, electrical conductivity, molecular weight, surface tension, applied
voltage, flow rate, distance of the collector from the tip, and environmental parameters such
as humidity and temperature [32, 44–46]. In brief, an increase in the concentration of solute
increases the fiber diameter in a power law relationship, which in turn enhances the porosity.
As an example, micron size fibers generate a more porous scaffold compared to nano fibers.
Similar trends are observed for the effect of polymer molecular weight and viscosity, raising
these parameters also increases the fiber diameter, and hence the pore size. However, prior
to fiber formation it is critical to determine the range for each of these variables for the
formation of uniform, continuous and stable fibers.
An increase in the electrical conductivity of solution generally decreases the fiber diameter.
Contradicting results were reported for the effect of applied voltage [32]. While some
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