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8/11/2019 Druk Advances in Polymeric Systems for Tissue Engineering and Biomedical Applications 2012 Macromolecular Bioscience
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Advances in Polymeric Systems for TissueEngineering and Biomedical Applications
Rajeswari Ravichandran, Subramanian Sundarrajan,*Jayarama Reddy Venugopal, Shayanti Mukherjee, Seeram Ramakrishna*
1. Introduction
The prerequisites of polymer scaffold materials for tissue
engineering applications are manifold and highly challen-
ging. These include (i) biocompatibility of the polymer
material (ii) the material must not elicit unnecessary
inflammatory response (iii) the material should not
demonstrate any adverse immune response or cytotoxicity
(iv) similar to allmaterials in contact with human body, the
scaffolds must be easily sterilizable to prevent infection.
In addition, the mechanical properties of the polymeric
scaffold must be compatible andshould notcollapse during
surgical implantation or during the patients regular
activities. There is an extensive list of criteria a polymer
has to satisfy in order to be applied safely as polymer
therapeutics or as an agent in tissue regeneration. A
polymeric biomaterial intended for tissue engineering
application can be characterized as a material intended to
Review
R. Ravichandran, S. Sundarrajan, J. R. Venugopal,
S. Mukherjee, S. Ramakrishna
Healthcare and Energy Materials Laboratory, Nanoscience and
Nanotechnology Initiative, Faculty of Engineering, National
University of Singapore, Singapore
E-mail: [email protected]; [email protected]
R. Ravichandran, S. Sundarrajan, S. Ramakrishna
Department of Mechanical Engineering, National University of
Singapore, Singapore 117576
The characteristics of tissue engineered scaffolds are major concerns in the quest to fabricate
ideal scaffolds for tissue engineering applications. The polymer scaffolds employed for tissue
engineering applications should possess multifunctional properties such as biocompatibility,
biodegradability and favorable mechanical properties as it comes in direct contact with the
body fluids in vivo. Additionally, the polymer system should also possess biomimetic archi-tecture and should support stem cell adhesion, proliferation and differentiation. As the
progress in polymer technology continues, polymeric biomaterials have taken characteristics
more closely related to that desired for tissue engineering and clinical needs. Stimuli
responsive polymers also termed as smart biomaterials respond to stimuli such as pH,
temperature, enzyme, antigen, glucose and electrical stimuli that are inherently present in
living systems. This review highlights the exciting advancements in these polymeric systems
that relate to biological and tissue engineering appli-
cations. Additionally, several aspects of technology
namely scaffold fabrication methods and surface
modifications to confer biological functionality to
the polymers have also been discussed. The ultimateobjective is to emphasize on these underutilized adap-
tive behaviors of the polymers so that novel appli-
cations and new generations of smart polymeric
materials can be realized for biomedical and tissue
engineering applications.
286
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interface with biological systems to evaluate, treat,
augment, repair or replace any tissue, organ or function
of the body.[1] In tissue engineering applications the
polymeric scaffold material serves as a biomimetic
template for cell adhesion, proliferation, differentiation,
extracellular matrix (ECM) formation and mineralizationthereby providing a favorable environment for the
regeneration of damaged tissue. Similarly in the case of
drug delivery applications, if the polymer is not a drug
itself, it often provides a passive function as a drug
carrier, reducing immunogenicity, toxicity or degradation,
while improving circulation time and potentially a passive
targetingfunction. In this case thepolymer hasto be water-
soluble, non-toxic, non-immunogenic and it needs to be
safe before and after the drug has been released, including
a safe excretion. If the polymer is non-degradable (e.g.,
polymethacrylates), the size needs to be below the renal
threshold ensuring that it is not accumulated in the body.
If the polymer is degradable (e.g., polyesters), the toxicity
and/or immune response of the byproducts have to be
taken into consideration. As the line of development
continues, polymeric biomaterials will take on character-
istics more closely related to that of pharmaceutical and
clinical needs, where the polymer material is designed
to function in a biospecific manner by interacting with
specific biological and biochemical pathways in vivo. The
objective of this review is to put in evidence the evolution
and potentiality of the emerging polymer approaches and
the advancements made to the existing polymer systems
for tissue engineering applications. This review highlights
the exciting developments in polymer technology thatrelate to biomedical applications. Toward these ends,
several aspects of technology have been highlighted,
namely scaffoldfabrication methods,surface modifications
to confer biological functionality to the polymers and the
advanced polymeric systems that can be employed in
combination for several tissue engineering and biomedical
applications.
2. Scaffold Fabrication Techniques forBiomedical Engineering
Current efforts in polymer scaffold fabrication techniques
have been focused on custom designing and synthesizing
polymers with tailored properties for specific biomedical
applications; for instance (i) developing novel polymeric
materials with unique chemistries to increase the
diversity of polymer structure, (ii) developing biosynthetic
processes for fabricating biomimetic polymer structures
and (iii) adopting combinatorial and computational
approaches for scaffold design. Other highly desirable
features concerning scaffold fabrication include near-net-
shape fabrication and scalability for cost-effective produc-
tion. Polymers are the primary materials for scaffold
fabrication and the requirements for following certain
fabrication techniques has ranged from their potential
ability to create scaffolds with controlled porosities,
Rajeswari Ravichandran received her B.Tech in
Biotechnology from Anna University, India. She
then did her M.Eng in Bioengineering from
National University of Singapore (NUS) and is
currentlypursuingher PhDfrom NUS. Herresearch
interests include cardiac tissue engineering, bio-materials and stem cell biology.
Subramanian Sundarrajan graduated with an
M.Sc. in Inorganic Chemistry from the University
of Madras, India in 1996. Later, he worked in a
project at Indian Institute of Science, India till
1999. He his PhD from the University of Madras
in2003and joined the NUS in2003and currently
working as Senior Research Fellow and he is
working on electrospinning of nanofibers, metal
oxide nanoparticles and nanofibers for tissue
regeneration applications.
Shayanti Mukherjee completed her B.Tech inIndustrial Biotechnology fromDr. MGRUniversity,
Chennai, India and is now pursuing her Ph.D in
National University of Singapore, Singapore in
myocardial tissue engineering. Her researchinter-
est lays in translational regenerative medicine.
Jayarama Reddy Venugopal received his PhD in
neuroendocrinology from University of Madras,
India. At present, he holds an appointment with
National university of Singapore as senior research
fellow. His research experience spans over fabrica-
tion of biocompatible nanofibrous scaffolds for
skin, nerve, bone, vascular and cardiac tissue
engineering. His research interests include celland molecular biology and nanomedicine.
Seeram Ramakrishnareceived hisPhD in Materials
Science & Engineering from the University of
Cambridge and General Management training
from the Harvard University. He is a Professor
at the department of Mechanical Engineering in
NUS. He is a former Dean of NUS Faculty of
Engineering from 2003 to 2008, and founding
Co-Director of NUS Nanoscience & Nanotechnol-
ogy Initiative from 2003 to 2010. He is a Fellow of
major professional societies in Singapore, UK and
USA including the American Society for Materials
International (ASM); American Society for Mech-anical Engineers (ASME); American Institute for
Medical & Biological Engineering (AIMBE); Institu-
tion of Mechanical Engineers (IMechE) UK; Insti-
tute of Materials, Minerals & Mining (IOM3) UK;
and Institution of Engineers Singapore (IES).
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encapsulation of growth factors, for controlled delivery of
pharmaceutical agents, the removal of residual solvents
following other fabrication methods, and to produce
biomimetic scaffolds for tissue regeneration.[2] Materials
fabrication and design will be more closely related to
and actuated by specific clinical needs and biologicalphenomena. Some of the widely employed scaffold
fabrication techniques for tissue engineering applications
are discussed in the following section.
2.1. Thermally Induced Phase Separation (TIPS)
3D resorbable polymer scaffolds with very high porosities
can be produced using TIPS technique to develop scaffolds
with controlled macro- and microstructure architectures
suitable for nerve, muscle, tendon, ligament, intestine,
bone, and teeth applications.[35] This procedure requires
the use of a solvent with a low boiling point that is easy to
evaporate. For instance, dioxane could be used to dissolve
aliphatic polyesters and subsequently phase separation is
induced through the addition of a small quantity of water.
Thisleadstotheformationofapolymer-richandapolymer-
poor phase, which upon cooling leads to the formation of a
porous scaffold. The polymeric scaffolds obtained via TIPS
are highlyporouswithanisotropic tubular morphology and
extensive pore interconnectivity. Using TIPS technique the
pore morphology, mechanical properties, bioactivity and
degradation rates can be controlled by varying the polymer
concentration in solution, volume fraction of the secondary
phase, quenching temperature and the polymer and
solvent used.[3]
Injectable scaffolds for tissue engineeringapplications to fill voids in damaged tissue are fascinating
due to their ability to conform to the implant site, whereas
preformed scaffolds require prior cognition of the defect to
be filled and those with irregular size and shape can prove
problematic. A microsphere network withsufficiently large
interstices allows tissue to infiltrate the network to bear
the mechanical loads as the scaffold degrades. A novel
application of TIPS in tissue engineering is in the rapid
formation of porous microspheres, wherein the pore
morphology and the pore size can be tailored to facilitate
surface modifications like the incorporation of bio-
molecules. Such porous biodegradable microspheres are
desirable for tissue engineering and drug delivery applica-
tionsbecausetheconstituentamountofpolymerisreduced
compared with solid microspheres, yet the scaffold volume
is kept constant and the degradation mechanism is more
predictable.[6] A study suggested that TIPS microspheres
produced from liquidliquid phase separation of PLGA/
dimethylcarbonate(DMC)andbyanemulsionroute,PLGA/
silver-doped phosphate-basedglasses by solidliquidphase
separation, and dispersion of protein particles by phase
separation of PLGA/DMC in the presence of fluorescently
labeled antibody; as a suitable candidate for localized drug
delivery, tissue regeneration/augmentation and tissue
engineering.[6] Chitosan and collagen are other proteins
that have been used to fabricatemicrospheres using TIPS.[7]
The microspheres developed by TIPS enable control over
both the open pore structure (determined by microsphere
size) and the internal structure, which could be matched tothe desired tissue by adjusting the processing parameters.
A network of biologically active microspheres fabricated
using TIPS could be applied as a tissue engineering scaffold,
or could actas a filler material for inaccessiblesoft andhard
tissue repair/augmentation.
2.2. Solvent Casting and Particulate Leaching
Solvent casting of polymeric scaffolds involves the
dissolution of the polymer in an organic solvent, mixing
with porogen granules like sugar, inorganic salt, paraffin
spheres, and casting the solution into a predefined 3D
mould. The size of the porogen particles used will affect the
scaffold porosity, while the polymer to porogen ratio is
directly correlated to the amount of porosity of the final
structure. The solvent is later allowed to evaporate and the
porogen particles are removed by leaching, following
the main processing step. The main advantage of this
fabrication technique is the ease of fabrication without the
need for any specialized equipment; however organic
solvents must be fully removed in order to avoid any
possible damage to the cells seeded on the scaffold. A
study reported an enhanced solvent casting/particulate
leaching (SCPL) method developed for preparing three-
dimensional porous polyurethane (PU) scaffolds forcardiac tissue engineering applications.[8] It involved a
combination of SCPL with centrifugation; with the
aim to enhance the pore uniformity and the pore
interconnectivity. These scaffolds showed uniform distri-
bution of the human aortic endothelial cells, useful for
cardiac tissue engineering applications.
2.3. Solid Freeform Fabrication Techniques (SFFT)
Thistechniqueis employed to fabricatehighly reproducible
scaffolds with completely interconnected porous net-
works.[9,10] Using digital data produced by imaging soft-
ware such as computer tomography or magnetic resonance
imaging enables appropriate design of the polymeric
scaffold structure.[10] Solid freeform (SFF) manufacturing
coupled with conventional foam scaffold fabrication
procedures (phase separation, emulsion-solvent diffusion
or porogen leaching) may be used to fabricate materials
with controlled micro- and macroporous architectures.
Suchbiomimetic internal architectures may provevaluable
for multi-tissue and structural tissue interface engineering.
Unlike conventional computerized machining techniques
which involve the removal of materials from a stock,
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SFF process employs the underlying concept of layered
manufacturing,[11] whereby three-dimensional objects
are fabricated with layer-by-layer building via the proces-
sing of solid sheet, liquid or powder material stocks. For
instance, Xiong et al.[8] fabricated composites of PLLA/TCP
with porosities of up to 90% as shown in Figure 1, withmechanical properties close to human cancellous bone by
using low-temperature deposition based on a layer-by-layer
manufacturing methodof SFFfabrication. Theflexibility and
manufacturing advantages of SFF have been employed for
biological applications ranging from the production of
scale replicas of human bones[12] and body organs[13] to
the advancedcustomized drugdelivery devices[14]and other
areas of medical sciences including medical forensics.[15]
Three-dimensional printing (3D-P) is the most widely
investigated SFF technique for scaffold fabrication.
Kim et al.[16] employed 3D-P with particulate leaching
technique for creating porous scaffolds using polylactide-
co-glycolide (PLGA) powder mixed with salt particles and a
suitable organic solvent. Electron microscopy results
performed 2 days after the in vitro cell culture with
hepatocytes (HCs) revealed the successful attachment of
largenumbersofHCsontothescaffolds.Similarly,Zeltinger
etal.[17] employed 3D-P-fabricatedporous poly(L-lactic acid)
(L-PLA) disc shaped scaffolds measuring 10mm (diameter)
by 2 mm (height) to investigate the influence of pore size
and porosity on cell adhesion, proliferation and matrix
deposition. The scaffolds were constructed with two
different porosities (75% and 90%) and four different pore
size distributions (
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the polymer solution to create a gel. Water is then used to
extract the solvent from the gel; the gel is cooled to a
temperature below the glass transition temperature of the
polymer and freeze dried under vacuum to produce a
nanofibrous scaffold.[2022] The desired architecture can be
obtained through the addition of various porogens topolymer solution during the phase separation process. This
provides engineers with a significant amount of control in
tailoring both the pore sizes and interconnectivity of the
resultant polymeric material by altering the concentration,
size, and geometry of the porogens used.[21] Unlike other
techniques, phase separation is a simple process that does
not require much specialized equipment. Besides, it is also
easy to achieve batch-to-batch consistency; and tailoring of
scaffold mechanical properties and architecture is easily
achieved by varying polymer/porogen concentrations.
However, thisfabricationprocess islimited tobeing effective
with only a selective number of polymers and is strictly a
laboratory scale technique.[20]
2.5. Electrospinning
Inthelastdecade,electrospinningtechniquehasattracteda
great interest as it allows the fabrication of fibrous non-
woven micro/nano fabrics for tissue engineering applica-
tions, mainly due to the structural similarity to the tissue
extracellular matrix (ECM). Electrospinning technique
involves the application of high voltage to a polymeric
solution, in order to create an electrically charged jet
randomly collected onto a grounded target.[26,27] Electro-
spinning is a simple and versatile method to prepare ultrathinfibersfrompolymersolutionsormelts.Itutilizesahigh
voltage source to inject charge of a certain polarity into
a polymer solution or melt, which is then accelerated
toward a collector of opposite polarity. As the electrostatic
attraction between the oppositely charged liquid and
collector and the electrostatic repulsions between like
charges in the liquid become stronger, the leading edge of
the solution changes from a rounded
meniscus to a cone (the Taylor cone). A
fiber jet is eventually ejected from the
Taylor cone as the electric field strength
exceeds the surface tension of the liquid.
The fiber jet travels through the atmo-
sphereallowing the solvent to evaporate,
thus leading to the deposition of solid
polymer fibers on the collector.
By modifying variables such as the
distance to collector, magnitude of
applied voltage, or solution flow rate,
researchers can dramatically change the
overall scaffold architecture depending
on the desired application. The fibres
produced by this technique usually have
a diameter from several nanometers to a few micrometers.
Electrospun polymer nanofibres possess many extraordin-
ary properties including small diameters, favorable biomi-
metic architecture, concomitant largespecificsurface areas,
large surface to volume ratio which favors enhanced
protein adsorption, a high degree of structural perfectionandgoodmechanicalproperties.Inordertomoreaccurately
mimic the natural ECM, research has also examined the
electrospinning of natural materials such as: collagen,[28]
chitosan[29] and gelatin.[30] However, these materials often
lack the desired physical properties or are difficult to
electrospin on their own, which hasled to the development
of hybrid materials, which consist of a blend of synthetic
and natural polymers.[3032] Stitzel etal. studied the use of a
hybrid blend of type I collagen (45%), PLGA (40%), and
elastin (15%) to form a vascular prosthesis using electro-
spinning.[31] The addition of PLGA was shown to improve
mechanical properties such as burst strength and com-
pliance of the prosthesis in comparison to scaffolds
composed solely of type I collagen and elastin alone.
Electrospun fibers can be oriented or arranged randomly,
givingcontroloverboththebulkmechanicalpropertiesand
the biological response to the scaffold. For example, in
designing scaffolds meant to replace highly oriented tissue
such as the medial layer of a native artery it is desirable to
generate aligned nanofibers. In the medial layer both
smooth muscle cells and ECM fibrils are aligned circumfer-
entially, which allows for vasoconstriction and vasodila-
tion in response to corresponding stimuli. Xu et al.
developed an aligned nanofibrous scaffold using electro-
spinning of a poly (L-lactide-co-e-caprolactone) (P(LLA-CL))(75:25) copolymer for vascular tissue engineering.[32]
Smooth muscle cells attached and migrated along the axis
of the aligned fibers. Furthermore, the proteins comprising
the cytoskeleton of the smooth muscle cells were aligned
parallel to the aligned fibers, demonstrating the cells
proclivity to organize along oriented fiber topography.
Figure 2 shows a SEM image of electrospun gelatin
Figure 2. SEM image of electrospun gelatin nanofibers at a) 500x and b) 5000xmagnification.
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nanofibers at 500x and 5000x magnification. A detailed
review concerning the performance of nanofibrous poly-
meric materials in guiding cells to initially adhere
and spread over the material, as well as further triggering
them to differentiate and secrete the appropriate ECM
biomolecules targeted to skin, blood vessel, cartilage,muscle, adipose, nerve and bone tissue engineering
applications is dealt elsewhere.[33] The various advantages
and disadvantages of the scaffold fabrication techniques
aresummarizedin Table1. Whilethe size scale,porosity and
orientation of scaffolds fabricated by these techniques can
be modified to influence cell functions such as adhesion,
proliferation, and migration, even greater enhancement
over the control of cellular function can be achieved by
attachingbioactivemoleculesto the surface of the scaffolds
as discussed in the following section.
3. Surface Functionalization of PolymericMaterials for Tissue Engineering
Over the years, several advancements in polymeric
materials have addressed biological aspects of increased
complexity, starting on the level of ion interactions
and moving to growth factor and stem cell incorporation.
The polymeric materials were extended from purely
Table 1. Various scaffold fabrication techniques, their advantages and disadvantages and their potential applications.
Fabrication technique Advantages Disadvantages Applications
Thermally induced
phase separation
Can control the porosity
and pore morphology
Achievable sizes range
from only 10 to 2000 mm
in diameter
Proteins and drug delivery
and higher drug encapsulation
efficiency
Solvent casting
and particulate
leaching
Simple operation, control
of the pore size and
porosity by selecting the
particle size and the
amount of salt particles
Distribution of salt particles
is often not uniform within
the polymer solution, and
the degree of direct contact
between the salt particles is
not well controlled,
interconnectivity of pores
in a final scaffold cannot be
well controlled, limited
membrane thickness, lack
of mechanical strength,
problems with residual
solvent, residual porogens.
Cardiac and vascular tissue
engineering applications
Solid freeform
fabrication
techniques
Customized design,
computer-controlled
fabrication, anisotrophic
scaffold microstructures,
processing conditions
Lack of mechanical strength,
limited to small pore sizes
Production of scale replicas
of human bones and body
organs to advanced
customized drug delivery
devices
Phase separation Allows incorporation ofbioactive agents,
highly porous structures
Lack of control overmicro-architecture,
problems with residual
solvent, limited range
of pore sizes
Drug release and proteindelivery applications
Electrospinning Easy process, High porosity,
High surface
area to volume ratio
Limit range of polymers,
lack of mechanical strength,
problems with residual
solvent, lack of control over
micro-architecture
Bone, skin, nerve and cardiac
tissue engineering
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synthetic materials to material/biologic hybrids, engi-
neering at the same time bioactivity and biodegradability
by imparting biological cues into the polymer. For
instance, in order to apply biodegradable polyester based
nanocomposite in tissue engineering, their surfaces have
to be chemically and physically modified with theincorporation of bioactive molecules and cell binding
proteins. This provides the desired biomimetic micro-
environments for cell adhesion, proliferation and differ-
entiation. Therefore, many approaches to functionalize
the surface of biodegradable polymer scaffolds have been
undertaken in order to introduce useful surface character-
istics to the polymer for tissue engineering applications.
The scaffold environment should be able to present
and deliver combinations of biomolecules such as cell
adhesion motifs, protein molecules, growth factors,
angiogenic factors, differentiation factors and immuno-
suppressive or anti-inflammatory agents.
3.1. Stem Cells
Stem cell incorporation into polymeric scaffolds is of
immense potential for creating next-generation syn-
thetic/living composite biomaterials that feature
high adaptability to the biologicalenvironment. Scaffolds
seeded with stem cells allow the cells to undergo
differentiation to adapt the desired tissue engineering
approach. This approach enables the polymeric scaffold
surface to mimic complex biological functions leading to
in vitro and in vivo growth of tissues and organs. Acombination of mesenchymal stem cells, growth factors,
and bioresorbable polymers canprovide a solution for the
treatment of difficult tendon injuries. A knitted PLGA
matrix populated with allogeneic bone-marrow-derived
MSCs was used to bridge Achilles tendon defects in adult
rabbits.[34] In this study the specimens treated with MSCs
showed a higher rate of tissue regeneration and remodel-
ing at 2 and 4 weeks after surgery compared with the
group treated with PLGA alone.Similarly in a recentstudy
byourgroup,acombinationofMSCswithcardiomyocytes
cultured on an elastomeric poly(glycerol sebacate)/
gelatin nanofibrous scaffolds was shown to be favorable
for the regeneration of the infarcted myocardium. In the
study the incorporation of stem cells into the polymeric
material induces paracrine signaling effects, reducing
the cell death of cardiomyocytes.[35] Figure 3 shows the
double immunostaining image for both MSC marker
protein and cardiac marker proteins expressed by stem
cells which have undergone cardiogenic differentiation.
Thus the incorporation of stem cells into polymeric
materials lead to the production of next generation
biomimetic scaffolds for various tissue engineering
applications.
3.2. Biomolecules
There is a significant scope in the application of surface
modifications, through the use of protein biomolecules to
provide more cues for cell adhesion, proliferation and
differentiation. Integrins, laminin and Arg-Gly-Asp (RGD)
proteins and also several natural proteins like collagen,gelatin and fibrinogen were shown to be essential for cell
attachment to polymeric material surfaces devoid of any
cell recognition sites.[36,37] The immobilization of these
proteins to polymers not only promotes cell adhesion and
proliferation but also increases hydrophilicity of hydro-
phobic polymers such as aliphatic polyesters. One such
surface functionalization for biopolymer substratesurfaces
is attachment of RGDpeptides that isthe most effectiveand
often employed peptide sequence for stimulating cell
adhesion on synthetic polymer surfaces. This peptide
sequence can interact with integrin receptors at the focal
adhesion points. Once the RGD sequence is recognized bythe integrins, it will initiate an integrin-mediated cell
attachment pathway and activate signal transduction
between the cell and ECM, thus influencing various cell
behaviors on the substrate including proliferation, differ-
entiation, survival and migration.[38] In another study,
Jiang et al.used a coaxial electrospinningset-up tofabricate
biodegradable core/shell fibers with PCL as the shell and
BSAcontainingdextranasthecore. [39]BothBSAloadingand
its release profile could be controlled by varying the feed
rate of the core solution, with higher feed rates giving
higherBSA loading andacceleratedrelease. Addition of PEG
to the shell was used to further control the release profile,
and was shown to increase release of BSA. By varying the
inner solution feed rate as well as the PEG content of
the shell the authors were able to vary the release period
from 1 week to approximately 1 month, for drug delivery
applications.
3.3. Growth Factors
Scaffolds functionalized with immobilized growth factors
is of utmost importance in several tissue engineering
applications as the embedded growth factors (i) could
be released in response to several cell-mediated stimuli,
(ii) could create a highly regulated network of signaling
molecules for mediating several biological pathways,
(iii) able to orchestrate cell attachment, migration, organi-
zation and proliferation finally giving rise to functional
tissue. The focus of studies during the past decade has been
on numerous growth factors that promote soft-tissue
regeneration suchas platelet-derived growth factor (PDGF),
epidermal growth factor, basic fibroblast growth factor,
insulin-like growth factor-I, bone morphogenetic proteins
(BMPs) and transforming growth factors.[4045] Polymeric
systems that provide a gradual and controlled release of
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growthfactors to thesite of injuryis of extreme importance
for tissue repair and regeneration. Growth factors like
BMPs were shown in in vivo studies to be osteoinductive.
Immobilizing these growth factors on the scaffold
surface might significantly shorten the bone healing
process, enhance osseointegration and reduce patient
recovery time. Chew et al. examined the release ofb-nerve
growth factor (NGF) stabilized in BSA from a copolymer
consisting ofe-caprolactone and ethyl ethylene phosphate
(PCLEEP).[46] Due to its relatively hydrophobic backbone,
PCLEEP has a slow degradation rate demonstrating a mass
loss of approximately 7% over a 3-month period. Using this
system, the authors observed a sustained release of NGF
over a period of 3 months. Due to the relatively small
amount of mass loss over this period it was inferred
that NGF release was occurring primarily via diffusion,
demonstrating that a biodegradable system can be used to
obtain a desirable release profile while still eliminating the
need for a second surgery for implant removal. A detailed
review of various growth factors and their significance for
tissue engineering has already been discussed.[47,48]
3.4. Surface Modification Techniques
Surface modification techniques such as plasma treatment,
ion sputtering, oxidation and corona discharge affect the
chemical and physical properties of the polymer surface
without significantly changing the inherent bulk material
properties. For example, using plasma processes, it is
possibleto change thechemicalcompositionand properties
of the polymer system such as wettability, surface energy,
metal adhesion, refractive index, hardness, chemical
inertness and biocompatibility.[49] Plasma techniques can
easily be used to induce the desired groups or chains onto
Figure 3.Dual immunocytochemical analysis for the expression of MSC marker protein CD 105 (a, d, g) and cardiac marker protein actinin(b, e, h) in the coculture samples and the merged image showing the dual expression of both CD 105 and actinin (c, f, i); on the TCP (a, b, c),gelatin nanofibers (d, e, f), and PGS/gelatin core/shell fibers (g, h, i) at 60x magnification. Nucleus stained with DAPI.
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the surface of a polymer.[5052] Appropriate selection of
the plasma source provides the introduction of diverse
functional groups on the polymer surface to improve
biocompatibility or to allow subsequent covalent immo-
bilization of various bioactive cues. For instance, plasma
treatments with oxygen, ammonia, or air can generatecarboxyl groups or amine groups on the polymer sur-
face.[53,54] Plasma treatment affects the chemistry of the
biodegradable polymer surface, but at the same time it
also introduces significant changes in topography.[55,56] A
variety of ECM protein components such as gelatin,
collagen, laminin, and fibronectin could be immobilized
onto the plasma treated surface to enhance cellular
functions.[57,58] In a recent study, it was noticed that
fibroblasts proliferation, morphology, CMFDA dye expres-
sion and secretion of collagen were significantly improved
on plasma-treated PLACL/gelatin scaffolds compared to
PLACL nanofibrous scaffolds, proving that the plasma-
treated PLACL/gelatin nanofibrous scaffold is a potential
biocomposite material for skin tissue regeneration.[59]
4. Advancements in Polymer Systems forCell Culture
Living systems respond to external stimuli adapting
themselves to changing environmental conditions. Poly-
mer scientists recently have been trying to mimic this
behavior for creating smart polymeric systems for tissue
engineeringapplications.Thesesmartpolymersaredefined
as polymers that undergo reversible large, physical orchemical changes in response to small external changes in
the environmental conditions, such as temperature, pH,
light, magnetic or electric field, ionic factors, biological
molecules, etc. Smart polymers have very promising
applications in the biomedical field as delivery systems
of therapeutic agents, tissue engineering scaffolds, cell
culture supports, bioseparation devices, sensors or actua-
tors systems. Hoffman et al.[60] demonstrated, in a very
elegant design, that theaction of an enzymatic receptor can
be modulated when this kind of polymer isconjugated close
to its active place. The authors were able to switch on-off
the receptor using the transition between extended and
coiled form of the molecule.[61] Two different kinds of
bioconjugates including stimuli-responsive polymers can
be prepared for biological applications by: a). Random
polymer conjugation to lysineaminogroups of a protein. b).
Site-specific conjugation of the polymer to genetically
engineered specific amino acid sites. The placement of
stimuli-sensitive polymers near the active place of a
recognition protein can provide a highly environmental-
sensitive system. Some of the widely employed smart
polymeric systems have been dealt with in detail in the
following sections.
4.1. Nanocomposites Embedded Polymer Systems
Nanocomposites are an efficient strategy to upgrade
the structural and functional properties of synthetic
polymers. Aliphatic polyesters such as polylactide (PLA),
poly(glycolides) (PGA), poly(caprolactone) (PCL) have
attracted wide area for their biodegradability and bio-compatibility in the human body. However all these
desired properties cannot be achieved from a single
polymer system. In fact, although several polymeric
materials are available and have been investigated for
tissue engineering, no single biodegradable polymer can
meet all the requirements for biomedical applications.
Therefore, the design of multi-component polymer
systems represents a viable strategy in order to
develop innovative multifunctional polymeric materials.
A consequence has been the introduction of organic
and inorganic nanofillers into biodegradable polymers
to produce nanocomposites based on hydroxyapatite,metal nanoparticles or carbon nanotructures, to prepare
advanced polymeric systems with enhanced properties.
This combination of bioresorbable polymers and
nanostructures opens new perspectives in the
application of nanomaterials for biomedical
applications with desirable mechanical, thermal and elec-
trical properties.
4.1.1. Polymer/HA Nanocomposites
Natural bone matrix is an organic/inorganic composite
material of collagen and apatites. Composite material
systems based on inorganic nanoparticles, showed astrongly enhanced polymer degradation rate when com-
pared to pure polymer systems. Studies have shown that
tricalcium phosphate filled polymers showed deposition of
small, 10 nm sized hydroxyapatite crystals on the surface
of the PLGA polymer composite, while for pure PLGA/
nanohydroxyapatite formation was observed during
degradation, indicative of enhanced osteoconductive
properties of PLGA nanocomposites. The fast degradation
and the superior osteoconductivity make these nanocom-
posites a promising material for application in ortho-
pedics.[62,63] These polymer/HA nanocomposites can also
be surface modified for the release of biomolecules. For
instance, Nie and Wang examined the release of DNA from
electrospun scaffolds consisting of a blend of PLGA and
hydroxyapatite (HAp) with various HAp contents (0%, 5%,
and 10%) for bone tissue engineering applications.[64] DNA
wasincorporated into thescaffolds in three ways: (1)naked
DNA, (2) adsorption of DNA/chitosan nanoparticles onto
scaffoldsafter scaffold fabrication by dripping, or (3)blend-
ing DNA/chitosan nanoparticles with the PLGA/HAp
solution prior to electrospinning. They noticed that higher
HAp contents led to faster DNA release for both free and
encapsulated DNA. This may be due to the hydrophilic
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nature of HAp, which caused the DNA/chitosan bind to
HAp particles in the presence of dichloromethane
during the emulsion procedure. These techniques not only
increase encapsulation efficiency, as was noted by the
authors, but it would also increase the release rate.
As the HAp nanoparticles diffuse out of the PLGA fibersthey leave pores through which the DNA/chitosan
particles can easily diffuse through the scaffolds.
The authors noted that encapsulated DNA/chitosan nano-
particles enhanced transfection efficiency leading to
higher cell attachment and viability in an in vitro study.
Thus, it was demonstrated that the encapsulation of
DNA/chitosan nanoparticles in PLGA/HAp electrospun
scaffolds has the potential to augment bone tissue
regeneration. In a recent study by our group n-HA was
precipitated by calcium-phosphate dipping method on
PLLA/PBLG/Col scaffolds. The incorporation of n-HA drives
the adipose derived stem cells to osteogenic lineage on
these scaffolds, in the absence of any induction medium,
for bone tissue engineering.[65]
4.1.2. Polymer Embedded Nanoparticles
Silver (Ag) has been known to have a disinfecting effect
and has been applied in traditional medicines. Several
salts of Ag and their derivatives are commercially
employed as antimicrobial agents. Hence, Ag/polymer
nanocomposites have been investigated for their anti-
bacterial property.[66,67] Theuseofnanoparticleshasbeen
limited by difficulties associated with handling and
processing of nanoparticles. Embedding of nanoparticlesinto biodegradable polymer matrices represents a valid
solution to these stabilization problems and permits a
controlled effect.[68] For instance, Liang et al. examined
the incorporation of DNA into PLGA scaffolds fabricated
by electrospinning.[69] Plasmid DNA was condensed in a
poor solvent mixture, and was then encapsulated in
micelles composed of a triblock copolymer (PLA-PEG-PLA)
givingencapsulatedDNAnanoparticles.
The micelles were then dissolved in a
solution of DMF with PLGA and electro-
spun, resultingin theformation of PLGA
fibers containing encapsulated DNA
nanoparticles. The DNA was encapsu-
lated in the PLA-PEG-PLA triblock copo-
lymer in order to protect it from
degradation during electrospinning
with the PLGA copolymer. An in vitro
release study demonstrated that
approximately 20% of the encapsulated
DNA was released after a period of 7 d.
Such systems are very useful for
gene delivery anddrug delivery applica-
tions.
4.1.3. Polymers Incorporating Carbon Nanostructures
Polymers incorporating carbon nanostructures have been
investigated by several groups for a variety of biomedical
applications.[7072] Carbon nanotubes (CNT) have the
potential to provide the needed structural reinforcement
for biomedical polymer scaffold. By dispersing a smallfraction of carbon nanotubes into a polymer, significant
improvementsin the polymer/CNT composites mechanical
strength has been observed. The role of the interface
between the CNT and polymer matrix is essential in
transferring the load from the polymer to the nanotubes,
thereby enhancing the mechanical and electricalproperties
of the composite. The electrical conductivity of CNT based
nanocompositesis a usefultoolin order todirect cell growth
and cell differentiation since they can conduct electrical
stimulus into the desired tissue, thereby stimulating the
healing process in nerve, bone and cardiac tissue engineer-
ing applications. For example when an alternating currentis applied to the nanocomposites of poly(lactic acid) and
MWCNTs, it showed increased osteoblast proliferation and
calcium production, favorable for bone regeneration as
shown in Figures 4a and b respectively. [73]
4.2. Polymer Blended Hybrid Systems
When a singlepolymerdoes nothave theproperties desired
for a tissue engineering application, a copolymer or blend
(simple mixture) of polymers may be employed to achieve
the desired properties. Studies have shown that PLA:PCL
blends and copolymers possess great solvent flexibility andalso exhibit a percent increase in elasticity while main-
taining an ultimate tensile strength that is analogous to
that of pure PLA scaffolds. It was shown that by adding a
small amount of PCL (as little as 5 wt%) to PLA, the strain to
failure of a scaffold increased from less than 25% to
more than 200%.[26] Similarly, Kwon et al. successfully
Figure 4.A) Osteoblast proliferation under electrical stimulation: & without electricalstimulation; & with electrical stimulation. Values are mean SEM; n4; p
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electrospun poly(L-lactide-co-e-caprolactone) (PLCL) at
various weight percents from methylene chloride in ratios
of 70:30 (9wt%), 50:50 (7wt%), and 30:70 (11 wt%).
It was found that at 25 8C the 70:30 ratio was a hard solid,
the 50:50 ratio was an elastomer, while the 30:70 ratio
was a gummy solid. The two ratios that resulted in fibrousscaffolds, 70:30 and 50:50, were determined to have a
Youngs modulus of 14.2 and 0.8MPa, respectively,[74]
showing a great deal of promise for use as an arterial graft.
This is because an engineered vascular graft must be
strong enough to accommodate a large pressure increase
while having enough elasticity to passively expand to
allow blood flow downstream. Similarly in another study
Xu et al. electrospun a nanofibrous scaffold of P(LLA-CL)
(75:25). This scaffold exhibited a tensile modulus of
156 MPa, tensile strength of 5 MPa, and strain at break of
127%. This compares favorably to the mechanical proper-
ties of native coronary artery having a tensile strength of
1.4011.14 MPa and a strain at break of 4599%.[75] As is
true with copolymers, blending of natural polymers and/or
synthetic polymers serves to enable further capability of
tuning a material to attain the desired properties for
tissue engineering applications. For instance, Huang et al.
electrospun collagen type I and poly(ethylene oxide) to
tailor fiber morphology and mechanical properties of
scaffolds.[76] Additionally collagenGAG scaffold has been
utilized extensively as dermal regeneration templateswith
unprecedented biological activity to achieve enhanced
healing response.[77,78]
4.3. Stimuli-responsive Polymers
Stimuli-responsive polymers are also termed as smart
polymers. Interest in stimuli-responsive polymers has
persisted over many decades, and a great deal of effort
has been crafted to fabricate new smart materials. Stimuli-
responsive polymers show an acute change in properties
upon a slight change in environmental condition. This
property can be utilized for the preparation of smart
polymeric systems, which mimic biological response
behavior to a certain extent. Stimuli-responsive polymers
mimic biological systems in a primitive way where an
external stimulus results in a change in properties. This
can be a change in conformation, change in solubility,
alteration of the hydrophilic/hydrophobic balance or
release of biomolecules or combination of two or more
responses. Many advances in stimuli response polymers
are advantageous in the biomedical fields due to their
specificity and their ability to respond to stimuli that
are inherently present in vivo. The physical or chemical
stimulus that triggers specific responses is limited to the
formation or destruction of secondary forces (hydrogen
bonding, hydrophobic effects, electrostatic interactions,
etc.), simple reactions (e.g., acidbase reactions) of moieties
on the polymer backbone, and/or osmotic pressure
differentials that result from such phenomena.[79] Besides,
theresponsecanalso beexpandedto include more dramatic
alterations to the polymeric structure. For instance,
degradation of polymeric hydrogels upon specific stimulus
can occur by reversible or irreversible bond breakage of thepolymeric backbone or cross-linking groups. Such systems
may facilitate the application of smart polymers in drug
delivery, diagnostics, separations and other clinical appli-
cations. The major advantage of these polymer systems is
their ability to apply these stimuli in a non-invasive
manner in living body. The stimuli may occur internally
(such as change in pH or temperature) or externally
(external stimuli such as magnetic or electric field, light
and ultrasound). Figure 5 showsa schematicrepresentation
of the various stimuli and their responses. [79]
4.4. Conducting Polymers
Conducting polymers (CPs) have attracted much interest as
suitable matrices for biomolecules owing to their high
electronic and ionic conductivities, which has been used to
enhance the stability, speed and sensitivity of various
biomedical devices. The electrically active tissues like the
brain, heart and skeletal muscle provide opportunities to
couple electronic devices and computers with human or
animal tissues to create therapeutic bodymachine inter-
faces. The conductive and semiconducting properties of CPs
make them an important class of materials related to such
applications. The origin of electrical conduction in CPs has
been ascribed to the formation of nonlinear defects suchas solitons, polarons or bipolarons formed during either
doping or polymerization of a monomer.[80] For example,
Abidian et al. has demonstrated the use of the CPs
polypyrrole (PPy) and poly(3,4-ethylenedioxythiophene)
(PEDOT) for nerve tissue engineering applications, by
culturing neuronal cells using an in vitro dorsal root
ganglion model.[81] Electrical stimulation with neural
electrodes is used clinically to improve conditions such as
hearing loss(cochlearimplants). It wasfound that, whenPPy
CP was coated onto cultured substrates on which cochlear
explants were cultured, the neurite growth improved upon
the incorporation of neurotrophins like NT-3 and brain-
Figure 5.Potential stimuli and responses of synthetic polymers.
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derived neurotrophic factor (BDNF). Such CP-based smart
biomaterials provide a biocompatible substrate to help
protect auditory neurons from degradation after sensor
neural hearing loss and encourage neurite outgrowth
towards the electrodes.[82] Conductive neural interfaces
tailored forcellinteraction by surfacefunctionalization withincorporation of bioactive factors produce superior neuro-
prosthetic devices with improved charge transfer capabil-
ities. A study examined the effect of entrapping NGF within
the CP PEDOT during electrodeposition to create a polymer
capable of stimulating neurite outgrowth from proximal
neural tissue.[83] The incorporation of NGF can modify the
biological interactions of the electrode without compromis-
ing on the conductive properties. Another study demon-
strated the use of PEDOT nanotubes polymerized on top of
electrospun PLGA nanofibres for the potential release of the
drug dexamethasone. Here, dexamethasone was incorpo-
rated within the PLGA nanofibers and then PEDOT was
polymerized around the dexamethasone-loaded PLGA
nanofibers. As the PLGA fibres degraded, dexamethasone
molecules remained inside the PEDOT nanotubes. These
PEDOT nanotubes favored controlled release upon electrical
stimulation.ThiswasbecauseofthechangeinvolumeofthePEDOT nanotube upon electrical stimulation owing to the
expulsionofanions.Figure6demonstratestheincorporation
and release mechanism of dexamethasone from PEDOT
nanotubes due to electrical stimulation.[84] For cardiovas-
cularapplications,Lietal.hasdemonstratedthepotentialfor
using PANIas anelectroactivepolymer. Theystudiedvarious
advancements in CPs by covalently attaching oligopeptides
to PANI and electrospinning PANIgelatin blend nanofiber
scaffold. These scaffolds were analyzed as potential candi-
dates for cardiac tissue engineering applications using H9c2
myoblast cells.[85] A detailed review on the application of
Figure 6. Schematic illustration of the controlled release of dexamethasone: A) dexamethasone-loaded electrospun PLGA, B) hydrolyticdegradation of PLGA fibers leading to release of the drug, and C) electrochemical deposition of PEDOT around the dexamethasone-loadedelectrospun PLGA fiber slows down the release of dexamethasone (D). E) PEDOT nanotubes in a neutral electrical condition. F) Externalelectrical stimulation controls the release of dexamethasone from the PEDOT nanotubes due to contraction or expansion of the PEDOT. Byapplying a positive voltage, electrons are injected into the chains and positive charges in the polymer chains are compensated. To maintainoverall charge neutrality, counterions are expelled towards the solution and the nanotubes contract. This shrinkage causes the drugs tocome out of the ends of tubes. G) Cumulative mass release of dexamethasone from: PLGA nanoscale fibers (a), PEDOT-coated PLGAnanoscale fibers (b) without electrical stimulation, and PEDOT-coated PLGA nanoscale fibers with electrical stimulation of 1 V applied at thefive specific times indicated by the circled data points (c). H) UV absorption of dexamethasone-loaded PEDOT nanotubes after 16 h (a), 87 h(b), 160 h (c), and 730 h (d). The UV spectra of dexamethasone have peaks at a wavelength of 237nm. Data are shown with a standarddeviation (n 15 for each case).
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conductingpolymer for biomedical applications has already
been discussed elsewhere.[86] CPs thus represents a class of
smart polymeric materials that provides an excellent
opportunity for fabricating highly selective, biocompatible,
specific, economic and handy biomedical devices. However,
challenges facing CPs include poor stability and mechanicalproperties as well as poor control of the mobility,
concentration and presentation of bioactive molecules.
4.5. Glucose-responsive Polymers
While a variety of polymer systems have been reported for
diagnostic applications, polymers that respond to glucose
have received considerable recognition because of their
potential application in both glucose sensing and insulin
delivery applications. Diabetes mellitus, commonly
referred to as diabetes, is a chronic disease characterized
by deficient production of insulin. Treatment for diabetic
patients generally involves regular monitoring of blood
sugar concentrations and subcutaneously administering
insulin every day. This need for continuous patient
vigilance often leads to poor compliance with the
prescribed treatment. One potential route proposed taking
advantage of the advancements made in polymer systems,
is the developmentof smart delivery of polymer systems in
whichinsulindeliveryisautomaticallytriggeredbyarisein
blood glucose levels in vivo. While a variety of approaches
can be visualized to achieve this objective, considerable
research by material scientists has been dedicated to
developing self-regulated insulin delivery systems based
on glucose-responsive polymers.[87]
The majority of reportsdetailing glucose-responsive polymers are based on the
GOx-catalyzed reaction of glucose with oxygen. Typically,
glucose-sensitivityisnotduetodirectinteractionofglucose
with the glucose responsive polymer, but rather by the
response of polymer to the byproducts that result from
the enzymatic oxidationof glucose. The enzymatic reaction
of GOx on glucose is highlyspecific and leads to byproducts
suchasgluconicacidandH2O2. Therefore, theincorporation
of a polymer that responds to either of these byproducts
can indirectly trigger a glucose-responsive system. Another
type of glucose-responsive system as reported by Brownlee
and Cerami utilizes competitive binding of glucose with
glycopolymerlectin complexes.[88] Such glucose responsive
polymers are of potential importance for diabetic patients
owing to their specific insulin delivery applications.
4.6. pH responsive Polymers
pH-sensitive polymersare polyelectrolytes that bear in their
structure weak acidic or basic groups that either accept or
releaseprotonsin responseto changesin environmental pH.
Different organs, tissues and cellular compartments may
have large differences in pH, which makes the pH a suitable
stimulus for biological applications. When weak acids
(carboxylic acids, phosphoric acid) and bases (amines) are
linkedto the polymer structure,they exhibit a change in the
ionisation state upon variation of the pH. This leads to a
conformational change in thecaseof solublepolymers and a
change in the swelling behavior is observed in the case ofhydrogels. The pH range that a reversible phase transition
occurs can be generally modulated by two strategies:
1. Selecting the ionizable moiety with a pKa matching the
desired pH range. Therefore, the proper selection
between polyacid or polybase should be considered
for the desired application.
2. Incorporating hydrophobic moieties into the polymer
backbone and controlling their nature, amount and
distribution. When ionizable groups become neutral
non-ionized- and electrostatic repulsion forces disappear
within the polymer network, hydrophobic interactions
dominate. The introduction of a more hydrophobic
moiety can offer a more compact conformation in the
uncharged state and a more accused phase transition.
Ionisable polymers with a pKa value between 3 and 10
are ideal candidates for pH-responsive systems. Some of
the most widely studied pH responsive polymers are
poly(acrylamide), polyacrylic acid, poly(methacrylic acid),
poly(dimethylaminoethyl methacrylate) (PDEAEMA),
poly(dimethylaminoethyl methacrylate) (PDMAEMA). For
example, the change in pH along the gastro-intestinal
tract[89] from acidic in the stomach (pH 2) to basic in the
intestine (pH 58) has to be considered for oral deliveryof any kind of biomolocule, but there are also other, more
subtle changes within different tissues in the body.
Polymers are usually taken up into cells by fluid-phase
pinocytosis or receptor-mediated endocytosis. Within the
early endosome towards the lysosomes the pH drops from
6.2 to 5.0 giving a large change in proton concentration
inside these compartments. This drop in pH has been
utilizedin orderto release biomoleculesfrom thelysosomes
tothecytosol.[90] Intracellular delivery of oligo/poly(nucleic
acids) usually uses cationic polymers, which complex the
negatively charged nucleic acids. These cationic polymers
are then deprotonated within the endosomes, which
triggers endosome membrane disruption and release into
the cytosol before reaching lysosome with its hydrolytic
enzymes as shown in Figure 7.[91] Thus, tailoring the
protonation/deprotonation by altering the polymer struc-
ture can largely allow fine-tuning of the response in a
specific compartment with respect to the change in pH.
Similar to glucose responsive polymers, a pH responsive
polymer can also be applied for insulin delivery applica-
tions. For example a pH responsive polymer loaded or
conjugated with GOx, triggers the GOx-catalyzed reaction.
The gluconic acid byproduct that results from the reaction
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with glucose induces a response in the pH-responsive
macromolecule, thereby triggering the release of insulin
biomolecule. For applications specifically intended for
diabetes therapy, the pH-response generally causes swel-ling or collapse of the polymer matrix, releasing insulin.
Imanishi and co-workers have reported the covalent
modification of a cellulose film with GOx-conjugated
poly(acrylic acid) (PAA) for insulin delivery applications.[92]
Duncan et al. designed poly(amidoamine)s by combining
both positive and negative charges within the polymer
backbone.[93] On one side of which a very unique profile in
size changes upon protonation/deprotonation was found
with neutron scattering and NMR experiments.[94] The
amphoteric backbone yields an expanded shape at low pH,
which slowly collapsed when neutral pH is approached.
This seems to be the reason that these polymers exhibit
endosomolytic properties, which makes them very inter-
esting candidates in cancer therapy, e.g. delivery of non-
permeant toxins like gelonin. The pH-responsive swelling
and collapsing behavior has been used to induce controlled
release of model compounds like caffeine,[95] drugs like
indomethacin,[96] or cationic proteins like lysozome.[95,97]
For instance, Poly(L-histidine)-b-PEG in combination with
PLLA-b-PEG and adriamycin as drug was studied for an
extracellular tumor targeting. The system shows a very
sharp transition from non-protonated and hydrophobic
at pH 7.4, where the mixed micelles are stable, to ionized
and micelle-destabilizing at pH 6.6.
Adriamycin is rapidly released from the
micelles at this pH value.[98] Several other
groups have developed polymeric pro-
drugs (polymers in which the drug is
covalently attached to the macromole-cular chain) susceptible to hydrolytic
cleavage dependent upon the pH and
hence suitable for colon drug delivery.
This is the case of poly(N-metha-
cryloylaminoethyl5-aminosalycilamide)
or poly(methacryloylethoxyethyl 5-amino-
salycilic acid)[99] or the copolymeric
system developed based on 2-acryl-
amido-2-methylpropane sulfonic acid
(AMPS) and a methacrylic derivative of
an antiaggregant drug called Triflu-
sal.[100] The system depends on colonic
microflora for liberation of entrapped
drug, which seems most suitable, i.e.,
glycosidase activity of the colonic micro-
flora is responsible for liberation of drugs
from glycosidic prodrugs and the pre-
sence of azoreductase from the anaerobic
bacteria of the colon plays a main role
in the release of drug from azo bound
prodrugs.[101,102] Researchers have
designed more sophisticated pH-sensitive polymers in
order to take advantage of the pH changes that occur in
nature. These materials are inspired by living organisms
trying to mimic their response mechanisms. For instance,Sauer et al.[103] reported the synthesis of pH-sensitive
hollow nanocontainers inspired in virion particles. The
poly(acrylic acid) vehicles were synthesized by vesicular
polymerization and emulsion polymerization. These nano-
capsules combined the protective ability of the nanocon-
tainers in combination with controlled permeability and
therefore can be used to trigger the release of encapsulated
materials from the inner core. Kataoka[104] recently
communicated the development of polymeric micelles
as nanocarriers for gene and drug delivery based on
doxorubicin-conjugated block copolymer poly(ethylene
glycol)-poly(aspartame hydrazinedoxorubicin) [(PEG-
p(AspHid-dox)]. The polymer retained drugs and genes at
physiological pH and released the drugs as pH decrease
below 6.0.
Another most promising application of pH-sensitive
polymers is as nonviral gene carriers. Naked DNA is very
difficult to incorporate into thecellsbecause it is negatively
charged and it has a very large size at physiological
conditions. Godbey and Mikos reviewed some of the
advances in non-viral gene delivery research[105] describing
the use of poly- (ethylenimine) (PEI) and poly(L-lysine)
(PLL) as two of the most successful candidates for this
Figure 7. Schematic of DNA condensation and encapsulation into polymeric depotsystems. A) DNA complexation with cationic polymers leads to the formation ofnanometer sized polyplexes. B) Condensed or uncondensed DNA can be encapsulatedinto polymeric scaffolds for sustained delivery.
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application. PEI is a highly polycationic
synthetic polymer that condense DNA in
solution, forming complexes that are
directly endocytosed by several cell
types. Chitosan, a biocompatible and
resorbable cationic aminopolysacchar-ide, has also extensively been used as
DNA carrier.[106,107] Lim et al.[108] pre-
pared a self-destroying, biodegradable,
polycationic polyester, poly(trans-4-
hydroxy-L-proline ester) (PHP ester), with
hydroxyproline, a major constituent of
collagen, gelatin, and other proteins,
as a repeating unit. PHP ester formed
soluble polymer/DNA complexes with
average diameters of less than 200 nm.
These complexes could transfect the
mammalian cells, being comparable to
the transfection obtained with PLL, the
most commonpolymer for gene delivery,
at suitable pH conditions.
4.7. Enzyme-responsive Polymers
A relativelynew area of research in polymeric systemsis the
design of materials that undergo macroscopic property
changes when triggered by selective enzymatic reac-
tions.[109,110] These polymeric systems consists of the
following characteristics (i) sensitivity of this system is
unique because enzymes are highly target specific (ii) they
can operate even under mild conditions in vivo, and (iii) arevital components in many biological pathways. Enzyme-
responsive polymericmaterialsare composed of an enzyme-
sensitive substrate and another component that directs or
controls interactions that lead to macroscopic property
variations.[110] Catalytic reaction of an enzyme on a
substrate can lead to changes in geometry, supramolecular
architectures, swelling/collapse of gels, or variations of
surface characteristics.[110] Xu and co-workers reported the
use of enzymatic dephosphorylation to induce a solgel
transition. In the reaction sequence the small molecule
fluorenylmethyloxycarbonyl (FMOC)-tyrosine phosphate
was exposed to a phosphatase, and the resulting removal
of phosphate groups led to a reduction in electrostatic
repulsions, supramolecular assembly and eventual gelation
of the polymer system.[111,112]Figure 8 illustratesthe design
of a visual assay, wherein the precursor, which acts as the
substrate of an enzyme, transforms into a hydrogelator
when the enzyme catalyzes its conversion. Then, the self-
assembly of the hydrogelators in water induces the
formation of hydrogel. When inhibitors competitively bind
with the enzyme active site and block the conversion of the
precursor catalyzed by the enzyme, no hydrogel forms,
asshowninFigure8. [112] Inasimilarexample,Kaplanandco-
workers reported the modification of a genetically engi-
neered variant of spider dragline silk via enzymatic
phosphorylation and dephosphorylation.[113] Enzyme-
responsive polymers upon exposure to a specific enzyme,
undergoes changesin their macroscopic properties owing to
the creation of new covalent linkages. For instance, Ulijin
and co-workers used proteases to cause self-assembly of
polymeric hydrogels via reversedhydrolysis of peptides.[114]
Transglutaminase, a blood clottingenzyme, hasthe abilitytocross-link the side chains of lysine (Lys) residues with
glutamine (Gln) residueswithin oracross peptidechains.[110]
This behavior was exploited by Griffith and Sperindefor the
synthesis of hydrogels of cross-linked functionalized PEG
and lysine-containing polypeptides.[115,116]
4.8. Temperature-responsive Polymers
Temperature-responsive polymers and hydrogels exhibit a
volume phase transition at a certain temperature, which
causes a sudden change in the solvation state. When
hydrogels are prepared by cross-linking temperature-
sensitive polymers, the temperature sensitivity in water
results in changes in the polymer hydration degree. Below
the transition temperature the polymer swells up to
equilibrium hydration degree, being in an expanded state.
By increasing thetemperature above thetransition hydrogel
deswells to a collapsed state. This transition is usually
reversible and can be applied in a pulsatile manner making
the polymer to behave as an on-off system when the
stimulus is applied or removed. Polymers, which become
insoluble upon heating, have a lower critical solution
temperature (LCST). Systems, which become soluble upon
Figure 8. The illustration of the design for identifying inhibitors of an enzyme byhydrogelation.
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heating, have an upper critical solutiontemperature (UCST).
The change in the hydration state, within the aqueous
environment as the human body, causes the volume phase
transition, where intra- and intermolecular hydrogen
bonding of the polymer molecules are favored compared
to a solubilisation by water. Typical LCSTpolymersarebasedonN-isopropylacrylamide (NIPAM),[117,118]N,N-diethylacry-
lamide (DEAM),[119] methylvinylether (MVE),[120,121] andN-
vinylcaprolactam (NVCl)[122,123] as monomers. A typical
UCST systemis based on a combination of acrylamide (AAm)
and acrylic acid (AAc).[124] Most biomedical applications use
the change from room temperature to body temperature in
order to induce a change in the physical properties for e.g.
gelation, especially in applications using injectable biode-
gradable scaffolds. Polymers with LCST have been tested in
controlled drug delivery matrices and in on-off release
profilesinresponsetoastepwisetemperaturechange.Inthis
sense, polyNIPAAm hydrogels form a thick skin on their
surface above the LCST in the collapsed state, which reduces
the transport of bioactive molecules out of the hydrogels.
NIPAAm has also been copolymerized with alkyl methacry-
lates in order to increase the hydrogels mechanical proper-
ties, without compromising on the temperature sensitivity.
Poly(NIPAAm-co-butyl methacrylate) (poly(NIPAAm-co-
BM))was studied forthe deliveryof insulinin a temperature
on-off profile, below and above the LCST that in this system
wasloweredtoabout258C.[125]The temperature-responsive
polymers have the transition temperature in the region
2040 8C, which is interesting for biomedical applications.
However the transition temperature can be strongly
dependent on factors such as solvent quality, salt concen-tration and molecular weight. PNIPAM is a very interesting
temperature-responsive material, having good biocompat-
ibilityand thepositionof the LCST at 3233 8C,whichisideal
for biological applications. The LCST of PNIPAM is indepen-
dent of the molecular weight and the concentration,[126]but
it can be changed upon shifting the hydrophilic/hydro-
phobic balance, by copolymerization with a second mono-
mer. Hydrophobic co-monomers increase the LCST, whereas
hydrophilic co-monomers have the opposite effect.[127,128]
PNIPAM copolymers have been mainly studied for cardiac
tissue engineering and oral delivery of biomolecules like
calcitonin and insulin. In the latter case, the peptide or
hormoneisimmobilizedinPNIPAMbeads,whichstaystable
while passing through the stomach. Then in the alkaline
intestine the beads disintegrate and the drug is released.
Serres et al.[129] and Ramkisson-Ganorkar et al.[130] synthe-
sized P(NIPAM-co-BMA-co-AAc) for the intestinal delivery of
human calcitonin. Meanwhile, Kim et al. investigated the
delivery of insulin.[131] In both cases the combination of
the hydrophobic BMA moiety (butylmethacrylate) and
acrylic acid (AAc), which is non-ionized at low pH, prevents
disintegration of the beads in the acidic environment of the
stomach. At elevated pH the beads disintegrated due to the
solubilisation by the now ionized AAc. Thus in this case
besides temperature, the pH also plays an important role in
stimulating the drug release. Besides PNIPAM, Poly(methyl
vinyl ether) also has a transition temperature exactly at
37 8C, which makes it very interesting for biomedical
application. Poly(N-vinyl caprolactam) (PVCa) has not beenstudied as intensively as PNIPAM, but it also possesses very
interesting properties for medical and biotechnological
applications, e.g. solubility in water and organic solvents,
biocompatibility, high absorption ability and a transition
temperature within the settings of these applications
(338C).[123] Several approaches have been performed in the
tissueengineering fieldas temperature sensitive scaffolds or
surface modifications for the manipulation of cell sheets.
Poly(NIPAAm-co-acrylic acid) (poly(NIPAAm-co-AA)) gels
have been applied as extracellular matrix for pancreatic
islets in biohydrid pancreas.[132]
4.9. Antigen-responsive Polymers
Similar to enzymatic reactions, antigenantibody interac-
tions are also highly specific and are associated with
complex immune responses that help to recognize any
foreign bodies in the blood stream. Binding reaction
between antigens and antibodies can rely on a variety of
non-covalent interactions, such as hydrogen bonding, van-
der-Waals forces, and electrostatic and hydrophobic
interactions. Antibodies are employed in a number of
immunological assays for the detection and measurement
of biological and non-biological substances,[133] and the
highaffinity and targetspecificity of theirinteractions with
antigens have been harnessed to yield a variety of antigen-
responsive synthetic polymeric systems for biological
applications. In most cases, antigenantibody binding
has been used to induce responses in polymeric systems
prepared by either physically entrapping antibodies or
antigensin networks, chemicalconjugation of theantibody
or antigen to the network, or using antigenantibody pairs
as reversible cross-linkers within networks.[134] Miyata
et al. prepared antigen-sensitive hydrogels by coupling
rabbit immunoglobulin G (Rabbit IgG) to N-succinimidy-
lacrylate(NSA). The modified monomer was polymerized in
the presence of goat anti-rabbit IgG as an antibody and
acrylamide, which results in the formation of a hydrogelcross-linked both covalently and by antigenantibody
interactions. Upon the addition of rabbit IgG as a free
antigen, competitive binding of the goat anti-rabbit IgG
antibodies resulted in the breakage of antigenantibody
cross-linkers and a change in the morphology of the
hydrogel characterized by swelling of the hydrogel.[135]
4.10. Redox/thiol-responsive Polymers
Redox/thiol sensitive polymers are another class of
advanced polymers that are of immense importance in
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bioengineering applications especially in the field of
controlled drug delivery.[136,137] The interconversion of
thiols and disulfides is a key step in many biological
processes as it plays a significant role in the stability and
rigidity of native proteins in living cells.[138] Since disulfide
bondscanbereversiblyconvertedtothiolswhenexposedtovarious reducing agents and/orundergo disulfide exchange
in the presence of other thiols, polymers containing
disulfide linkages can be considered for both redox- and
thiol-responsive applications.[139,140] For instance, Glu-
tathione (GSH), is the most abundant reducing agent in
living cells. It has an intracellular concentration of about
10 mM, whereas an extracellular concentration of about
0.002mMin the cell exterior.[141] This significant variation
in concentration has been utilized to design thiol/redox-
responsive polymer drug delivery systems that specifically
release therapeutics into cells. For example, Lee and co-
workers synthesized polymer micelles with shells cross-
linked via thiol-reducible disulfide bonds as shown in
Figure 9. These served as carriers that preferentially release
anticancer drugs under reducing conditions typical of
cancer tissues.[142]
4.11. Shape Memory Polymers
Shape memory polymers (SMPs) can rapidly change their
shapes from a temporary shape to their permanent shapes
under appropriate stimulus such as temperature, light,
electric field, magnetic field, pH, specific ions or enzyme.
Schematic representation of thermally induced shape
memory effect is given in Figure 10, where heating a
sample above the switching transition temperature
(T Trans) induces the recovery of the permanent shape of
polymer. SMPs have the advantages of light weight, low
cost, good processability, high shape deformability, high
shape recoverability and tailorable switching tempera-
ture.[143145] For example, poly(D,L-lactide) is a good shape
memory biomaterial having good biodegradability and
biocompatibility; and have been utilized for several tissue
engineering applications. Zheng et al.[146] reinforced
poly(D,L-lactide) (PDLLA) with hydroxylapatite for hard
tissue engineering applications. The composite showed
good biodegradation, biocompatibility and shape memory
properties. The study showed that PDLLA/hydroxylapatite
composites at a suitable fraction ratio of between 2.0:1 and
2.5:1 had much higher shape recovery ratios and recover
speed than pure PDLLA. By using a crystalline polymer as
the fixing phase and a second crystalline polymer as the
reversible phase, SMP blends were created for bioengineer-
ing field. In that context, Behl et al. [147] reported binary
biodegradable polymer blends from two crystalline poly-
mers poly(p-dioxanone) (PPDO) and poly(caprolactone)
with poly(alkylene adipate) mediator as a compatibilizer.
The crystalline PPDO provides the hard segment to
determinethepermanentshapeandthepoly(caprolactone)
provides the soft segment to determine the switching
temperature. Besides, Zhang et al.[148]
introduced poly-(ethyleneglycol) dimethacrylate (PEGDMA) to PLGA/iso-
phorone diisocyanate (IPDI). This system showed good
shape memory and hydrophilic properties useful for
biomedical applications. Further, Zhu et al.[149] improved
the radiation efficiency of polycaprolactone by blending it
with polymethylvinylsiloxane (PMVS)
before radiation cross-linking, for biome-
dical applications.
4.12. Electro-responsive Polymers
Electro-responsive polymers (ERPs) are
very beneficial for biological applications
because of their potential to being direc-
tional, which can give rise to anisotropic
deformation. ERPs can be used to prepare
materials that swell, shrink, or bend in
response to an electric field.[150,151] Since
ERP can transform electrical energy into
mechanical energy, they have promising
applications in biomechanics, artificial
muscle actuation, sensing, energy trans-
duction, sound dampening, chemicalFigure 9. Illustration of shell cross-linking in drug -loaded polymer micelles and facili-tated drug release in response to cellular GSH.
Figure 10. Schematic representation of the thermally inducedone-way shape-memory effect.
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separations, and controlleddrug deliveryapplications.[150153]
Geldeformation, which involves bending in an electric field
is influenced by a number of factors like variable osmotic
pressure based on the voltage-induced motions of ions in
the solution, pH or salt concentration of the surrounding
medium, position of the gel relative to the electrodes,thicknessorshapeofthegel,andtheappliedvoltage.[153,154]
Transforming the application of an electric field into a
physical response by a polymer generally results in certain
changes in the properties of the polymer matrix like,
collapse of a gel in an electric field, electrochemical
reactions, electrically activated complex formation, ionic
polymermetal interactions, electrorheological effects, or
changes in electrophoretic mobility.[150] Typically, ERPs
have been investigated in the form of polyelectrolyte
hydrogels[150,151] which are capable of deformation under
an electricfield dueto anisotropicswellingor deswelling, as
charged ions are directed toward the anode or cathode
sideofthegel.[150]Forexample,underanelectricfield,inthe
case of hydrolyzed polyacrylamide gels, mobile H ions
migrate toward the cathode while the negatively charged
immobile acrylate groups in the polymer networks are
attracted toward the anode, creating a uniaxial stress
withinthegel.Theregionsurroundingtheanodeundergoes
the greatest stress while the area in the vicinity of the
cathode exhibits the smallest stress. This stress difference
contributes to the anisotropic gel deformation under an
electric field.[155] The natural polymers used to prepare
electro-responsive materials include chitosan,[156] chon-
droitin sulfate,[157] hyaluronic acid,[158] and alginate.[159]