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observed to form bridges similar to focal contacts with the nearest fibers as
seen in Fig. 15.7b. The expression and accumulation of cytoskeletal proteins
such as a-actin (Fig. 15.7c) and myosin (Fig. 15.7d) inside the SMC in relation
to the nanofiber orientation revealed the oriented fiber matrix effect on SMC
expression. This is further assumed that cytoskeletal protei ns expressed may
have guided SMCs to orient along the fiber direction. In a native blood vessel,
tunica intima consists of ECs aligned parallel to blood flow and in tunica media
SMCs aligned perpendicular to the blood flow. Therefore, this particular study
is very important as it could successfully align SMCs in a preferred direction so
that the tissue engineered vascular grafts could better mimic the structure and
also the cellular functionality of a native blood vessel.
Adding a cellular recognition layer to the internal surface of a vascular
construct could significantly reduce intimal hyperplasia following scaffold
implantation. Chemical and topographical simulation of the ECM may control
ECs spreading, proliferation, and phenotype retention that are of great interest
when designing biodegradable nanostructures for vascular graft applications.
Therefore, an effective approach would be modifying the nanofiber mat surface
with biocompatible layers such as gelatin and collagen to impart biofunction-
ality that ultimately prevent thrombosis and improve graft patency rate.
Several studies thus far addressed mimicking the ECM topographically;
however, Ma et al. grafted gelatin on to PCL nanofibers and created artificial
basal lamina with both ECM-mimicking structure and cell-compatible surfaces
[89]. Gelatin immobilization on random and oriented PCL nanofiber scaffolds
was achieved via air plasma treatment to introduce COOH groups on PCL
nanofiber surface, followed by covalent grafting of gelatin molecules. Gelatin
grafting increased the wettability of PCL nanofiber mats and significantly
improved EC spreading and growth, and also maintained their phenotypic
character. Endothelialization of the vascular graft is one of the best ways to
avoid coagulation and make these scaffolds ideal for blood vessel tissue
engineering. For this reason, human coronary artery endothelial cells
(HCAECs) were seeded onto the collagen-blended [P(LLA-CL), 70:30]
electrospun nanofibers [90]. Cells on collagen-blended fibers showed better
phenotypic spreading as compared to P(LLA-CL) nanofibers. Further,
HCAECs observed to interconnect well with the nanofibers and also the
pseudopods of the cells were oriented along the collagen-blen ded nanofibers.
This study has confirmed that biocompatible layers such as gelatin and
collagen in conjunction with nanofiber topography induce better phenotypic
retention for longer time periods, which is a desirable feature for a vascular
graft scaffold.
Angiogenesis, the process of forming new blood vessels from preexisting
vessels, is essential in the repair and regeneration of tissues. For this reason,
various scaffolds have been designed with the ability to deliver high
concentrations of angiogenic growth factors locally [91]. Another important
strategy is using heparin as a delivery agent for growth factors where vascular
endothelial growth factor (VEGF) and fibroblast growth factor-2 (FGF-2) are
396 BIOMEDICAL NANOSTRUCTURES
bonded through their he parin-binding domains [92, 93]. To better understand
and control molecular-scale matrixheparin microstructure, Rajangam et al.
used heparin nucleated PA self-assembled nanofibrous scaff olds for effective
binding of angiogenic growth factors [94]. In this process, aqueous solution
containing heparin and growth factors is mixed with PA (specially designed to
bind heparin) aqueous so lution to form nanofibrous scaffold. Positively
charged PA molecule aggregate is nucleated by negatively charged heparin and
later bond to the resultant cylindrical nanostructure. The growth factor that is
bonded to the rigid heparin on the nanofiber causes less Brownian motion that
leads to a high ligandreceptor binding. This hypothesis was proven in vitro by
a slow release of FGF-2 (57.1% in 10 days in presence of heparin) from a
network of PAheparin nanofibers compared to a faster release from
PANa
2
HPO
4
gel (98% in 10 days in the absence of heparin). A rat cornea
angiogenesis assay was performed to test the angiogenesis capacity of
heparinPA system in vivo. It was observed that heparin-binding PA
nanostructures combined with growth factors VEGF and FGF-2 promoted
vascularization that is significantly higher than all the controls involved in this
study. Thus, the self-assembled nanostr uctures cause extensive new blood
vessel formation and might have potential as scaffolds for vascular tissue
engineering.
15.5.3 Neural Tissue Engineering
Neural tissue regeneration is necessary in cases where scar tissue is formed after
injury, gaps in the nervous system formed during phagocytosis, and failure of
growth and extension of axons in the mature mammali an central nervous
system [95]. As described in Section 15.2.3, a peptide hydrogel consists of
standard amino acids (1% w/v) and 99% water. Under physiological
conditions, the peptide component self-assembles into a 3D hydrogel that
exhibits a nanometer-scale fibrous structure. Self-assembled peptide nanofiber
scaffolds (SAPNS) are superior over the currently available scaffolding
materials because they form a network of nanofibers that mimic the ECM,
break into natural amino acids that can be effectively used by surrounding
tissue, and are immunologically inert since they are free from chemical and
biological contaminants.
Self-assembled scaffolds support the growth of PC12 cells and the formation
of functional synapses in vitro with rat hippocampal neurons on arganine,
alanine, aspartate, and alanine nanofiber scaffolds [96]. Extensive neurite
outgrowth and active synapse formation encouraged similar studies in an
in vivo model to demonstrate the suitability of these scaffolds for nano neuro
knitting [97]. In this model, a wound was created in the mid brain superior
colliculus (SC) region and SAPNS is inflicted to regenerate an optic tract in
young and adult hamsters. The cavity that is seen with the control group
(where 10 ml of saline solution is injected) at 30-day time point showed the
failure of tissue healing (Fig. 15.8a), whereas the site of the lesion has healed as
NANOSTRUCTURES FOR TISSUE ENGINEERING/REGENERATIVE MEDICINE 397
evident from Fig. 15.8b with the injection of 10 ml of 1% SAPNS. In the case of
SAPNS, axons have grown through the treated area and the tissue appeared to
be knitted with little sign of the original lesion. Figure 15.9c clearly shows the
innervation density, where the white regions correspond to axons or axon
terminals labeled from the retina. This study has successfully demonstrated the
feasibility of using SAPNS for axon repair and regeneration.
After CNS trauma, inhibition of astrocyte proliferation is important in
preventing glial scar tissue formation. This is an important step since the newly
formed scar tissue acts as a barrier to axon elongation [98]. Recently, many
studies have aimed at achieving neural progenitor cell (NPC) differentiation
into neurons while suppressing astrocyte differentiation. Silva et al. studied
selective differentiation of NPCs using nanofibrous neurite promoting laminin
epitope isolucinelysinevalinealaninevaline (IKVAV-PA) self-assembled
scaffolds [99]. Murine NPCs were mixed with 1 wt% of IKVAV-PA in 1:1
volume ratio in media or physiological fluids. The resulting transparent gel-like
solid containing NPC clusters were encapsulated within a 3D network of
nanofibers formed via self-assembly of PA molecules. NPCs survived the self-
assembly process and showed differentiated neurons that were significantly
larger than those of controls at both 1 and 7 days. In addition to the
nanofibrous structure that simulates the ECM-like morphology, high content
of water (99.5%) present in these scaffolds serves as a medium for nutrient
diffusion and for cell migration. Differentiation and migration of NPC
FIGURE 15.8 Dark field images of parasagittal sections from animals at day 30 after
lesion and treatment. (a) Cavity in the retinal projections shows the failure of the tissue
healing. (b) A similar lesion section injected with 10 ml of 1 % SAPNS. Axons that
are indicated by light color have grown through the lesion and helped in tissue healing.
Enlarged view of boxed area in (b) is shown in (c), which is an area of dense termination
of axons (traced with cholera-toxin labeling) that have crossed the lesion. Scale bars,
100 mm (SC, superior colliculus). (Reproduced from Reference [99] with permission,
from The National Academy of Sciences of the USA.)
398
BIOMEDICAL NANOSTRUCTURES
neurospheres encapsulated in IKVAV-PA gel wer e observed to be significantly
higher than the nonphysiological sequence glutamic acidglutamineserine
(EQS) gel and poly(
D-lysine) controls. In addition to the rapid selective
differentiation, these scaffolds were further proven to be in vivo compatible and
might find applications in neural tissue engineering where controlled neural
differentiation and migration are required for treating CNS trauma conditions.
15.5.4 Cardiac Tissue Engineering
Myocardial infarction, fibrotic scar tissue formation (massive cell loss due to
ischemia), and impaired cardiac function are common problems that lead to heart
failure. This occurs because of the fact that cardiomyocytes (CMs) are terminally
differentiated cells and unable to regenerate into myocardial tissue after
infarction. Although heart transplantation is the ultimate cure to heart failure,
lack of organ donors and complications associated with immune suppressive
treatments are forcing scientists and surgeons to look for new strategies to
regenerate the wounded heart [100]. Gene delivery techniques followed by cell
transplantation techniques were considered in the mid-1990s to spark the failing
heart [101, 102]. Recently, cardiac tissue engineering has emerged as a novel
approach to repair and reinforce wounded cardiac tissue using embryonic or
neonatal cardiomyocytes in combination with natural and synthetic biodegrad-
able scaffolds [103]. This approach has been further modified by developing
nanofeatured scaffolds that mimic certain features of the natural ECM and
improve the regenerative capability of the tissue engineered cardiac constructs.
Collagen and gelatin-based tissue engineered patches have been recently
reported [104, 105]. They exhibited good contraction and function, but are not
popular for use due to the lack of stability required for handling. An alternative
approach is using synthetic polymers that require physical properties such as
stability and resistance to contractions [103]. Zong et al. seeded cardiomyocytes
FIGURE 15.9 (a) A 10-mm-thick fibrous mesh suspended across a wire ring that acts
as a passive load to condition cardiomyocytes during in vitro culture. Inset shows the
SEM of an electrospun mesh with an average fiber diameter of 250 nm. (b) Cross-
sectional SEM recorded on a cardiac nanofibrous mesh after 14 days of in vitro culture
showing cardiomyocyte presence throughout the entire mesh. (Reproduced from
Reference [107] with permission, from Elsevier)
NANOSTRUCTURES FOR TISSUE ENGINEERING/REGENERATIVE MEDICINE 399
on electrospun nanofiber scaffolds made of PLLA, PLGA (PLA10/
PGA90) + PLLA, and PLGA(PLA75/PGA25) + PEGPLA [106]. This study
compared growth of cardiomyocytes on randomly oriented fiber scaffolds, and
also on locally oriented fiber scaffolds where the microscale fiber orientation was
achieved by uniaxial stretching. It was observed that the fiber orientation
provided guidance for CM growth and showed sensitivity to the composition
and degradation rate of the electrospun PLGA-based scaffolds. CM cell density
was observed to be higher on relatively hydrophobic surfaces that in general
exhibited slower degradation. Myocardial functional studies of CMs on different
synthetic scaffolds showed PLLA as a better scaffold as compared to other
candidates involved in this study.
Shin et al. electrospun PCL nanofiber meshes and seeded with cardiomyo-
cytes isolated from neonatal Lewis rats [107]. ECM-like mesh with an average
fiber diameter of 250 nm and a mesh thickness around 10 mm (inset of
Fig. 15.9a) was suspended across a wire ring as shown in Fig. 15.9a.
Cardiomyocytes attached well to these scaffolds and started contracting after 3
days of culture. After 14 days of culture, cardiomyocytes retained their
phenotype, penetrated the entire scaffold, and stained positive for cardio
typical proteins such as actin, tropomyosin, cardiac troponin-I, and connexin
43. Also, a cross-sectional view clear ly confirms the presence of cells through
the entire nanofibrous scaffold as presented in Fig. 15.9b. This study has
established a versatile in vitro system to develop functional myocardial patches
that may be used for healing and regeneration of wounded hearts.
As described in the previous sections, nanotechnology for tissue engineering/
regenerative medicine is actively perused recently for multiple applic ations
including the regeneration of bone, cartilage, neural, v ascular and cardiac
tissues. In addition, research is currently underway to regenerate skin, bladder
and liver tissues using nanostructured scaffolds [108, 109, 110].
There have already bee n successful tissue engineering products that are
clinically relevant, have achieved FDA approval, and are currently being
marketed for tissue replacement. The most successful product is tissue
engineered skin that is currently in clinical use (Dermagraft
1
and Apligraf
1
)
for treating burn victims and patients with diabetic ulcers. Some of the
products at advanced clinical stage are contractile pa tches for damaged heart,
autologous chondrocytes in a polyme r gel for treating urinary reflux in
children, and urinary stress incontinence in adult women [111]. In addition,
various other tissue engineered products are at preliminary stages of
investigation (in vitro or animal studies) for regenerating tissues/organs such
as bone, ligament, heart valves, heart muscles, blood vessels, and bladders
[112]. Thus far, efforts have been on designing materials in combination with
cells and factors to meet the targeted clinical application. The success of this
approach can be further enhanced by creating cellular recognition nano/
microtopographical features that aid in the control of cell functions and may
lead to better tissue engineer ed products that meet future clinical needs.
400 BIOMEDICAL NANOSTRUCTURES
15.6 CONCLUSIONS
The need for mimicking the ECM has led to the development of various
scaffolding methodologies. These methods have helped in the creation of 3D
scaffolds consisting of fibers, grooves, pits, and pillars in nanoscale for tissue
engineering and biomedical applications. Due to the in vivo-like environment,
cells seeded on nanofeatured scaffolds showed improved adhesion, prolifera-
tion, migration, differentiation, and phenotypic expression. These scaffolds
with improved cellular compatibility can be directly translated into successful
tissue constructs for restoration, repair, and regener ation of bone, cartilage,
neural, vascular, and cardiac tissues. Even though scaffolds with nanofeatures
have generated lot of interest and excitement, the future of tissue engineering
will incorporate scaffolds with topographical and biochemical cues that would
exactly duplicate the in vivo environment. Nanostructured scaffolds are still at
their infancy stage and further work is required to answer questions such as
what size is right for a desired cell response and how different cells react to the
same topography.
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