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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.

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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,


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,


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


11/26

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.

302 Macromol. Biosci.2012, 12 , 286311


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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]

Druk Advances in Polymeric Systems for Tissue Engineering and Biomedical Applications 2012...

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