Improved Biomaterials for Tissue Engineering ...

Current Topics in Medicinal Chemistry, 2008, 8, 341-353

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Improved Biomaterials for Tissue Engineering Applications: Surface Modification of Polymers Rajesh Vasitaa, Kirubanandan Shanmugama and Dhirendra S. Kattia,* a

Department of Biological Sciences and Bioengineering, Indian Institute of Technology - Kanpur, Kanpur - 208016, INDIA Abstract: Tissue engineering approaches that combine biomaterial-based scaffolds with protein delivery systems have provided a potential strategy for improved regeneration of damaged tissue. The success of polymeric scaffolds is determined by the response it elicits from the surrounding biological environment. This response is governed, to a large extent, by the surface properties of the scaffold. Surfaces of polymeric scaffolds have a significant effect on protein and cell attachment. Multiple approaches have been developed to provide micrometer to nanometer scale alterations in surface architecture of scaffolds to enable improved protein and cell interactions. Chemical modification of polymeric scaffold surfaces is one of the upcoming approaches that enables enhanced biocompatibility while providing a delivery vehicle for proteins. Similarly, physical adsorption, radiation mediated modifications, grafting, and protein modifications are other methods that have been employed successfully for alterations of surface properties of polymeric scaffolds. The goal of this review is to provide an overview of the role of surface properties /chemistry in tissue engineering and briefly discuss some of the methods of surface modification that can provide improved cell and protein interactions. It is hoped that these improved polymeric scaffolds will lead to accelerated and functional tissue regeneration.

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Keywords: Tissue engineering, biomaterial, growth factor, protein immobilization, biocompatibility, surface modification, protein delivery systems.

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1. INTRODUCTION

Technological advancements have been instrumental in helping us overcome challenges in healthcare thereby ensuring increased life expectancy. However, this has in turn lead to newer problems due to aging processes in various body parts. With several vital tissues such as bone and cartilage undergoing degeneration, there arises a need to support/augment the aging body through external means such as regenerative therapies. Over the past few decades loss of tissue/organ due to trauma or disease has increased dramatically, thereby augmenting the challenges associated with regeneration [1-3]. Reconstruction surgery and organ transplantation are the currently available strategies to address the problem of tissue / organ loss [1, 2]. However, these methods are associated with clinical limitations. Surgically reconstructed tissue and organs may not be able to completely substitute the biological functions of tissue/organ [2]. Similarly, shortage of donor tissues or organs, donor site morbidity, immunological rejection and side effects in response to use of immunosuppressive agents are major limitations associated with organ transplantation [3]. The limitations associated with the current approaches to alleviate the problems of tissue loss have prompted the scientific community to search for newer therapeutic alternatives with better patient compliance. An improved understanding of tissue function, organization and regeneration has encouraged medical and engineering professionals to develop alternative regenerative approaches. One such approach is tissue engineering, that provides therapeutic solutions for the regeneration of diseased or aging tissue by artificially

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*Address correspondence to this author at the Department of Biological Sciences and Bioengineering, Indian Institute of Technology - Kanpur, Kanpur-208016, Uttar Pradesh, INDIA; Tel: +91-512-2594028; Fax: +91512-2594010; E-mail: [email protected]

1568-0266/08 $55.00+.00

generating native tissue via accelerated proliferation /differentiation of cells either in vitro or in vivo [4].

Tissue engineering, applies the principles of biology and engineering towards the development of functional artificial substitutes for damaged tissue [4,5]. Every tissue is a balanced composition of cells and the extra cellular matrix (ECM) that together maintains tissue homeostasis. Cells synthesize their own ECM which in turn provides physical support and other signaling molecules to cells. Loss of ECM and cells during trauma or disease induces regeneration procedures; however, loss of tissue beyond a certain limit does not lead to complete and functional regeneration. Therefore, in the event of tissue loss beyond this critical limit, there arises a need to induce / support the physiological regenerative process artificially. Three key factors need to be considered for the process of tissue reconstruction via tissue engineering methodologies, namely, - cells, ECM and signaling molecules for cell function. Hence, tissue engineering comprises of three technologies (i) development of scaffolds that are appropriate substitutes for native ECM, (ii) isolation, proliferation and delivery of cells (tissue specific / progenitors) to enable functional tissue regeneration, and (iii) isolation, expression and delivery of suitable growth factors which act as signaling molecules to induce/ modulate cell function [6]. i. Tissue Engineering: Challenges

Essential Components

and

One of the foremost challenges of tissue engineering is the development of scaffolds that are an appropriate mimic of the natural ECM. The ECM not only provides a physical support to the cells but also an environment for them to proliferate and differentiate, which in turn allows for appropriate tissue regeneration [7-10]. In defects of large sizes (critical size defects), providing cells alone does not © 2008 Bentham Science Publishers Ltd.

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induce the regeneration process. This is because, cells in the absence of native ECM cannot enable three dimensional (3D) growth of tissue [11]. Therefore, to induce tissue regeneration at large defect sites, design of an artificial ECM (commonly referred to as scaffold, matrix or construct) that allows cell adhesion followed by 3-D proliferation and differentiation is essential [12]. It is expected that cells seeded on the scaffold and cells from the native surroundings infiltrate the porous scaffold, adhere on it and start proliferating thereby leading to the synthesis of natural ECM [12,13]. As the defective tissue starts taking shape the scaffold gradually degrades away and is removed via metabolic activities of the body. Most scaffolding systems used in tissues engineering are therefore biodegradable and in addition, it is desirable to have scaffolds that produce a minimal immunogenic reaction. Hence, biocompatibility of the implant and proper immobilization/adhesion of cells on the scaffold are key issues [14]. Many polymers (both synthetic and natural), due to their biocompatibility and biodegradability have come to be the materials of choice for scaffold synthesis [15-17]. Successful polymeric scaffolds are those that in addition to being biocompatible, biodegradable, and mechanically appropriate, possess certain specific surface properties and three-dimensional architecture [18,19]. Design of such polymeric scaffolds or constructs whose surface properties can be tailored to meet specific requirements is an important aspect of scaffold design in current tissue engineering strategies [13, 20-23].

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2. SURFACE MATERIALS

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From the above discussion it can be inferred that a combination of several key features and technologies is needed to successfully construct a functional tissue. One of these that is central to tissue engineering is the development of techniques for engineering and design of novel polymeric scaffolds [35, 36].

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BIOMEDICAL

Biomaterial surfaces play a vital role in the success of an implant. Interactions between tissue and biomaterial based implants (scaffolds) occur at the surface of the implant. Surfaces offer advantages such as high accessibility for reactants and hence high surface areas enhance reaction turn over [37]. Unique microenvironments of the biomaterial surfaces enhance reactions and affinities of specific molecules on the surface [37]. The surrounding environment and phase boundaries play a key role in determining the surface behavior of biomaterials. For instance, at the biomaterial-air interface i.e. prior to implantation (outside the body), the non-polar groups on the surface have a tendency to move towards the phase boundary with air due to minimal surface energy and chain mobility. This in turn leads to different surface and bulk properties with low molecular weight components either moving towards or away from the surface [37]. On the other hand, at the biomaterial–aqueous interface after in vivo implantation of the scaffold, polar groups on the surface will interact with the surroundings thereby leading to an altered surface behavior [38,39]. Hydrophilicity of the surface (surface energy) and functional groups present at the surface that can initiate a reaction and thereby elicit a biological response are some of the chemical properties of the surface [40]. Whereas, crystallinity/amorphousness, surface roughness and presence of physical forms/patterns are some attributes that define the topography of a surface [41,42]. These properties of the surface and the biological environment together determine the success of a tissue engineered construct. Better biocompatibility, improved cell behavior, and controlled immobilization of protein/growth factor are the key factors for successful tissue regeneration that can be modulated via the surface properties of the scaffold [35, 43]. This review focuses on two major aspects of scaffolds that are controlled by surface chemistry - cell adhesion and immobilization of proteins.

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Secondly, in cell seeded scaffolds it is many a times observed that providing the cells with the scaffold alone need not necessarily cause regeneration of functional tissue [24]. It has now been understood that this is primarily because of the lack of factors that would promote proliferation and differentiation [25, 26]. These factors, mainly growth factors are protein molecules that are responsible for providing cell signals that in turn prompt specific cell behavior/function such as cell migration, proliferation, differentiation and death [26, 27]. Direct delivery of growth factors at the site of the defect is one approach for tissue regeneration. However, this approach is not very successful because of the short half life of growth factors, loss in their bioactivity due to the diseased condition / diffusion in surrounding tissue, and carcinogenic effects at higher concentrations. Immobilization or incorporation of growth factors in polymeric scaffolds prior to cell seeding presents a potential approach for direct delivery of growth factors [2833]. Immobilization of growth factors on scaffolds provides controlled and sustained release of these bioactive molecules at the site of defect [27-30]. However, this technology also needs to overcome several limitations such as loss of bioactivity during immobilization and reduced control on release kinetics via degradation of carrier devices [34].

PROPERTIES

As described earlier scaffolds provide a substrate for cell adhesion, proliferation and differentiation. The surface properties of the scaffold play a critical role in influencing these cell functions. When the scaffold is exposed to the physiological environment, ECM proteins such as fibronectin and vitronectin are non-specifically adsorbed on the surface of scaffold. The cells indirectly interact with the scaffold surface via the adsorbed ECM proteins. The interaction of cells with the adsorbed proteins is governed by cell membrane receptors known as integrins [12,44]. Integrins bind to specific domains of adsorbed proteins such as arginine-glycine-aspartate (RGD) and Proline-HistidineSerine-Arginine-Asparagine (PHSRN), which hook the cells to the matrix and trigger integrin-based intracellular signaling pathways that in turn alter/promote cell function such as proliferation and differentiation. The availability and accessibility of RGD domains for binding with integrins depends upon the adsorption behavior and quantity of adsorbed ECM proteins which in turn depend on the surface properties of the biomaterial [45]. Therefore, biomaterial surface properties influence ECM protein adsorption and conformation, which in turn influences cell behavior and function [45].

Improved Biomaterials for Tissue Engineering Applications

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Fig. (1). Cartoon depicting the role of surface properties of polymers during cell interactions with the surface.

The size of protein, charges on the protein surface, stability of the protein structure and the unfolding rate of proteins determine protein behavior on the biomaterial surface [46]. Similarly the surface topography, composition of biomaterials, and hydrophobicity are surface properties of the polymer that influence protein mediated cell interaction [46]. Biomaterials with varying compositions inherently have different functional groups on the surface by virtue of which they possess different surface properties (Fig. 1) [4749]. Hence, one way to alter surface properties would be to alter the composition of the biomaterial [50]. Another approach is to modify just the surface of the processed biomaterial i.e. the form in which the biomaterial would be eventually used, such as scaffolds for tissue engineering applications. Modification of the surface to achieve desired surface chemistry is relatively easier for in-vitro studies than for in-vivo tissue engineering. The biological environment is a multi-component system with various proteins and lipids. The scaffold surface tends to provide preferential affinity to certain proteins over others in this multi component system. Further, the composition of the layer of adsorbed proteins on the scaffold surface can change with time until a pseudosteady state is reached. Hence, achieving desired surface chemistries to enable the adsorption of specific proteins from a multi component system is a major challenge for the success of an implant / scaffold.

commonly used for such immobilization studies onto biomaterials such as poly (lactic acid) (PLA), poly (caprolactone) (PCL), and poly(lactide-co-glycolide) (PLGA) being used as the scaffold material. Although the latter are biodegradable, biocompatible polymers with good mechanical strength, the hydrophobic nature of their surface does not allow sufficient adsorption of proteins on it. Therefore, such surfaces lead to poor cell adhesion and are poor delivery vehicles for protein/growth factor. Hence, biomaterial surfaces play an important role that can influence cell behavior as well as protein/growth factor immobilization for specific tissue engineering applications [51]. The ability to tailor surface physical properties can help control cell behavior (adhesion, proliferation, and differentiation), whereas, the ability to control surface chemistry can help in the immobilization and release of growth factors from the polymeric scaffold [43, 47, 48]. The ability to manipulate surface chemistry also has a direct influence on the biocompatibility of the polymeric scaffold. Several methods have been developed for physical or chemical immobilization of these proteins including polymer blending, grafting, and chemical modification of polymers [51, 52]. Therefore, this review shall focus on the role of surface chemistry in tissue engineering, modification of polymers to provide greater control on the surface properties of scaffolds and the current research status in these areas of study.

Several approaches have been explored for the immobilization of desired proteins onto scaffolds to allow for improved cell interaction and/or for controlled delivery of proteins at the site of implantation. Growth factors, therapeutic proteins (insulin) and ECM proteins have been

3. MODIFIED POLYMERS SURFACE PROPERTIES

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Developing new biomaterials for purpose of tissue regeneration of tissues is a tedious procedure that would take more than a decade to reach the clinical trials stage for every


new biomaterial developed. Therefore, most approaches have been developed for improving the biocompatibility and surface properties of existing biomaterials to enable their use in tissue engineering applications. Natural and synthetic biomaterials that have been explored for surface modification include chitosan, collagen, alginate, gelatin, poly (vinyl alcohol) (PVA), poly (glycolic acid) (PGA), PLGA, PLA and PCL [17, 53, 54]. There are several methods that have been reported for fabricating scaffolds such as freezedrying, porogen leaching, gas forming, phase separation and electrospinning [55, 56]. In most of these methods the developed scaffolds possess surface and bulk properties that are not similar and the surface chemistry and topographies obtained are also different. Therefore, there is a need for separate surface modification protocols for each material based on the tissue that needs to be regenerated and the site of implantation (i.e. biological environment). Natural materials such as hyaluronic acid and collagen differ in properties from batch to batch and from one source to another. In addition, the properties of natural materials can be modified only to a limited extent. On the other hand, synthetic materials have drawn more attention because of the relative ease in controlling their physical and chemical properties by tailoring their molecular structure. This in turn enables the manipulation of their biodegradation rate/time, porosity, mechanical properties as well as their surface properties.

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popular use as a biomaterial [60]. In spite of all favorable properties, it is a hydrophobic polymer and PLA based scaffold surfaces have demonstrated less protein and cell interaction capability as compared to hydrophilic biomaterials such as PGA and PVA [60, 61]. Similarly, PLGA is a copolymer of the monomers lactic acid and glycolic acid. Although glycolic acid provides hydrophilicity to the surface, increasing percentage of glycolic acid leads to decrease in mechanical strength and faster degradation of the copolymer [62]. Therefore, numerous attempts have been made to modify the surfaces of these two polymers (PLA and PLGA) while maintaining their desirable bulk properties. In this review we will illustrate the different types of surface modifications and their importance by restricting the discussion to mainly two polymers PLA and PLGA (Fig. 3). CH3 O O

CH C

(a) Poly(lactide) (PLA)

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CH3 O

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(b) Poly (lactide-co-glycolide) (PLGA)

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Fig. (3). Chemical structure of a. Poly (lactide) (PLA) b. Poly (lactide-co-glycolide) (PLGA).

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Amongst the biodegradable synthetic materials PLA and PLGA have been explored extensively for biomedical engineering applications [57-59]. In the past two decades they have also been explored for synthesis of scaffolds for tissue engineering, as wound dressing materials, and as materials for the development of controlled delivery systems [60]. Biocompatibility, significant mechanical strength, and slow degradation rate are the key features of PLA for its

O

Most biomaterials are biodegradable and can be cleaved by hydrolysis, biocatalytic activity (enzymatic), chemical treatment (solvents) and radiation processes such as UV/gamma rays exposure. Therefore, the process for surface modification should be preferably one that provides low temperature conditions, a solvent free environment and low energy irradiation conditions (Fig. 4). To achieve a surface chemistry that would remain stable in biological environments the covalent grafting of functional groups is preferred over physical coating (adsorption only). The surface stability achieved via chemical modification is much higher than that

Fig. (2). Biomaterial surface modification for improved Cell/ Protein/Growth Factor immobilization.



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obtained by the physical absorption of the same biomolecules. In covalent modification, the presence of functional groups facilitates better attachment of bioactive molecules or species onto the surface that leads to a higher surface stability. Secondly, chemically modified surfaces, due to their enhanced wettability, offer greater biocompatibility towards cell growth and flow of body fluids. This is a major advantage over the hydrophobic surfaces of most unmodified biomaterials. i. Modification of Polymeric Biomaterial Surface for Influencing Cell Behavior (a). Physical Modification

Historically, coating of biomaterial surfaces with cell adhesive proteins like fibronectin, vitronectin, collagen, or laminin, or ECM resembling molecules such as chitosan and gelatin has been one of the most common and popular approaches for the development of improved biomaterial surfaces [61, 63]. A variety of biomimetic materials have been developed using this procedure. These biomimetic materials are capable of eliciting specific cellular response and directing new tissue formation mediated by biomolecular recognition [19]. Several studies have been conducted using the protein coating technique and have demonstrated enhanced cell adhesion and proliferation on protein coated surfaces [64, 65]. However, since these proteins are isolated from other organisms they could potentially cause an immune response. In addition, proteins are susceptible to proteolytic degradation in the in vivo environment. Thus,

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Fig. (4). Methods for polymer surface modification.

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long-term applications of such modified materials could be challenging. Furthermore, as described earlier the orientation of the immobilized protein can be influenced by surface topography, surface charges, and wettability [46]. On hydrophobic surfaces proteins are adsorbed with maximum interaction and hence tend to loose their structural conformation. Therefore, there is also a need for orientation of the chemically immobilized ECM proteins. Most of the abovediscussed problems can be overcome either by immobilization of peptides (i.e. the cell binding domain of the protein such as the RGD peptide and PHSRN peptide) rather than the whole protein, or by oriented immobilization of proteins on the biomaterial surface [40,66]. Detailed mechanism of chemical immobilization of these peptides has been described in the section on chemical modification. b. Chemical Modification Chemical modification of polymeric biomaterials is another major approach for immobilization of various biomolecules on their surfaces [67]. A wide range of chemical reactions and reagents have been explored for this purpose [68]. The process of chemical surface modification begins with surface activation, which involves the creation of functionalities on the surface of the polymer. The presence of functionalities on polymeric surfaces enables the surface immobilization of ligands under mild condition. In the past, methods have been proposed for creation of functional groups on polymeric films, however, extending them to 3dimensional (3-D) scaffolds of varying size, shape, and architecture has been demonstrated to be challenging [68]. A


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commonly used chemical modification method is the alkali hydrolysis of aliphatic polyester surfaces. Alkali hydrolysis is driven by small, highly mobile protons that can diffuse relatively easily between the uncharged polymer chains and hence, are capable of penetrating porous 3-D scaffolds. Surface hydrolysis causes the cleavage of the ester bonds and thereby leads to the formation of free hydrophilic carboxyl and hydroxyl functionalities [69]. The creation of these active functionalities allows the possibility of covalent attachment of other moieties on the surface of the polymer.

and heparin. Yuan Lu demonstrated the use of EDAC and NHS for aminolysis of NaOH treated PLLA surfaces. PLLA films that were modified and coated with 1% chitosan demonstrated high cell (chondrocyte) proliferation rates along with increased synthesis of glycosaminoglycan (GAG) and collagen protein which are markers of chondrocytic cells [75]. Similarly, in another study EDAC/NHS modified PLGA surfaces were immobilized with chitosan and heparin. The ability of these surface modified polymers to demonstrate increased wettability, improved blood compatibility, enhanced cell proliferation (hepatocyte) and biocompatibility makes such scaffolds better natural ECM mimics [76].

The resulting carboxyl and hydroxyl groups that are used for immobilization of the ligand (ECM proteins, RGD peptides or bioactive molecules) can be chemically modified for further reaction such as aminolysis. Since most of the ligands used for improving cell interactions have a biological origin, water is often the only choice of solvent as the coupling medium. Primary amine conjugation after hydrolysis is the most preferred route for many biodegradable polymers such as PLA, and PLGA. The conjugation is accomplished by aminolysis using diamines such as ethylnediamine [70], N-aminoethyl-1-3-propanediamine, 1,6-hexanediamine, and 1-ethyl-3-(3- dimethylaminopropyl)-carbodiimide (EDAC) [71]. Carbodiimides find applicability in a wide variety of conjugation reactions in both organic as well as aqueous solvents [72, 73]. This is owing to their ability to form bonds without the addition of other atoms or spacers. Hence, carbodiimides are often referred to as zero length cross-linking agents. For example, carboxylate groups can be activated by carbodiimides by the formation of a highly active species/intermediate O-acylisourea that after further reaction with amine nucleophiles leads to stable amide bonds [72] (Fig. 5). In aqueous conjugation reactions, target molecules/polymers soluble in water can be conjugated with water-soluble carbodiimides such as ethylenediamine [73]. To form more stable Nhydroxysuccinimide (NHS) ester derivatives as reactive acylating agents NHS or N-hydroxysulfosuccinimide (SulfoNHS) can be added to the reaction. The corresponding NHS or Sulfo-NHS esters react readily with nucleophiles to form the acylated product, however, only primary or secondary amines form stable amide or amide linkages, respectively. The aminolysed matrices/scaffolds can then be immobilized with ECM proteins or other signaling proteins with the help of crosslinking agents or simply by adsorption.

Covalent immobilization of RGD peptides / small protein fragments on polymeric surfaces is another approach for surface modification of polymers [65,77]. Some of the studies that have applied this approach include covalent immobilization of RGD peptides on polystyrene surfaces with the help of polyethylene oxide [40], RGD peptidemodified poly(lactic acid-co-lysine) (PLAL) wherein, the RGD peptide was chemically cross-liked with lysine residue of PLAL with the help of 1,1-carbonyl diimidazole [78], and covalent immobilization of RGD peptides, collagen and fibronectin protein on polyvinyl alcohol surfaces [79-81]. These studies demonstrated that covalent immobilization of short peptide sequences derived from ECM proteins provide improved biocompatibility. These short peptides also exhibit higher stability during their immobilization, and sterilization conditions [82]. Previous studies have demonstrated slow enzymatic degradation of linear peptides in vivo [82] and hence these peptide immobilized surfaces exhibited excellent long term stability. In addition, covalent immobilization of peptides required lower surface area and allowed for the quantification of the immobilized protein. Thus, immobilization of RGD peptides covalently is a relatively easier and effective method for modifying the surface properties of polymers.

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Although the wet chemical method for surface modification of polymers has distinct advantages it can be associated with some limitations. The complex morphology of scaffolds can lead to low yields of ligand immobilization due to inaccessible reactive sites on the ligand molecules. Because of this physical constraint in 3-D scaffolds and the buried ligand sites, the immobilized ligand may not have optimal cell-receptor interaction. Moreover, the wet chemical method is non-specific and depends on the molecular weight of the polymer and its crystallinity. Lastly, the removal of excess/unreacted reagents and cross-linkers from the scaffolds can be challenging due to the complex 3-D architecture of scaffolds. In spite of the aforementioned limitations, the wet chemical method of peptide/protein immobilization still remains one of the preferred approaches for protein immobilization.

Several studies on the immobilization of proteins and other ECM molecules via hydrolysis and aminolysis of the polymer have been reported, however, most of them have been performed on polymeric films. The physical, chemical and theoretical aspects of aminolysis and hydrolysis of polyesters has been described in detail by Tristan I. Croll [74]. In addition to surface modification of polymers using proteins, carbodiimide chemistry has also been employed for immobilization of non-protein molecules such as chitosan

R1 R1

O C OH

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Fig. (5). Chemical reaction of carbodiimide mediated amines with carbonic acids.

O + H2N-R4 - Isourea

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Another major approach for covalent immobilization of signaling or therapeutic protein is their conjugation with polymers. For controlled protein delivery systems, it is desirable to have cleavable conjugating systems. There can be two ways in which the immobilized proteins can be released, firstly via the degradation of the carrier and secondly the degradation of the protein-polymer linkage. A wide range of biodegradable polymers that degrade via hydrolytic or enzymatic cleavage of bonds have been explored for protein delivery applications [29]. Similarly, a variety of biodegradable linkers between proteins and polymers have been explored for controlled delivery applications [68]. The side chains present on amino acids are generally used in linker formation. Some of the popular linkages used are amide, thiourea, alkylamine and urethane linkage of side chain of amino group of lysine and amino acids of proteins, thioester linkages from thiol group of free cysteine group, and amide and alkylamine linkage of carboxylic groups of aspartic and glutamic acid [68]. These linkages are normally degradable under physiological conditions (pH 7.4, 37 oC) and thereby enable the successful release of immobilized proteins. c. Modification Via Plasma Treatment Plasma treatment is another efficient and potential technique for modifying the surface properties of biomaterials without affecting their bulk properties. In this process a vessel is first evacuated and then filled with a low-pressure gas such as argon [83], ammonia [84], or oxygen [85] to enable the creation of glow discharge plasma. Excitation of the gas is then achieved using an energy source such as electric discharge, heat, radio-frequency energy, microwaves or alternating/direct current. This leads to ionization of the gas into free radicals, ions, protons, electrons, gas atoms and molecules of different energies. Subsequent bombardment of biomaterial surfaces with these high-energy species leads to the transfer of energy from plasma to the substrate. This transfer of energy in turn causes a series of chemical and physical changes on the surface of the substrate. These energetic species can react with the substrate up to a depth that ranges from several hundred angstroms to 10 microns without causing any changes in the bulk properties.


significant decrease in contact angle from 68º to 52º that potentially made the substrate more suitable for tissue engineering applications [83]. Modification of polymeric surfaces using plasma treatment has several advantages such as enhanced cell adhesion and surface wettability, improved biocompatibility, surface functionalization and molecular immobilization. In addition, adjustment in the treatment parameters and the type of gas used can lead to improved surface hydrophilicity of the polymer [60]. This technique is particularly useful for enhancing the affinity of cells onto scaffolds since it can cause induction of specific groups or species on the polymer surface. d. Photochemical Modification Biomaterial surfaces can be tailored using photochemical techniques that can lead to polymer cross-linking, grafting of biomolecules on polymeric substrates and photopolymerization of monomers via free radical polymerization. The photochemical techniques use high-energy photons (UV rays, x-rays, -rays) to initiate chemical reactions, such as breaking of chemical bonds and generation of free radicals by photo-initiators. These free radicals combine with another monomer molecule that eventually leads to the propagation of the reaction and formation of a polymer chain. Some common photo-initiators are halogens and their organic compounds (I, Br, ICl, IBr), hydrogen peroxide and alkyl hydroperoxides (R-OOH). Photochemical modification is advantageous when compared to other chemical modification techniques since it has the potential to selectively immobilize the target species at specific regions of the biomaterial, with the graft layers (>1μm in thickness) well bonded to the substrate.

Two types of plasma are generally used. cold or low temperature plasma and hot or elevated temperature plasma. Atmospheric pressure arcs can generate hot plasma, whereas, cold plasma is generated using low temperature glow discharge. Due to the thermal motion of surface molecules in high temperature plasma, the plasma treatment effect is often very difficult to maintain. Therefore, low-temperature plasma is preferred since the plasma-generated effects are efficiently maintained [60] Several studies have been reported on the use of plasma treatment for surface modification of polymers. Jian Yang demonstrated the use of anhydrous ammonia gaseous plasma treatment for the surface modification of poly (D, L, lactide). Their cell culture studies with mouse 3T3 fibroblasts demonstrated a two fold increase in number of cells on plasma modified surface as compared to untreated PLA surfaces [84]. In another study, -(13) (16)- glucan was grafted on argon plasma treated surfaces of PLGA for wound healing applications. The surface treatment led to a

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Ma et al. have studied the use of photo-polymerization for creating functionalities on PLLA surfaces for cartilage tissue engineering applications. In this study the authors introduced (-OH), (-COOH), and (-CONH2), functionalities on PLLA by photo-grafting hydroxyethyl methacrylate, methacrylic acid or acrylamide respectively. Their results demonstrated improved cytocompatibility on –OH and – CONH2 modified surfaces as compared to the –COOH modified surfaces [86]. In another approach, the same authors grafted PLLA surfaces with polymethylic acid via photo-oxidation and subsequent UV induced polymerization. Immobilization of proteins on the polymer grafted PLLA films was perfor-med using two approaches. The first approach involved surface activation using methyl sulfonyl chloride followed by adsorption of proteins such as collagen and gelatin. The second approach involved activation of the polymer film via EDAC treatment followed by immersion in gelatin and collagen solution [87]. The results of both the approaches demonstrated improved surface wettability. In another similar study Ma et al have used the previously described collagen-PLLA (EDAC treated) matrix for covalent immo-bilization of b-FGF. Their results demonstrated improved surface wettability, chondrocytic cell viability and spreading on the b-FGF-collagen surfaces as compared to the unmo-dified PLLA matrices [88]. Most photopolymerization studies in the past have been conducted on polymeric films and have demonstrated


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promising results for use in tissue engineering applications. However, the extrapolation of the technology from film surfaces to three-dimensional scaffolds will be essential before it can be practically useful in tissue engineering applications. e. Modification Via Composites and Graft Formation Chemical and high-energy treatments are methods wherein polymers can potentially loose their inherent properties such as mechanical strength due to the possibility of accelerated degradation. Therefore, it is desirable to explore techniques that would enable alterations of surface properties while maintaining the native properties of the polymer. Modification of polymer surfaces via grafting is one such technique. Grafting of polymers can be achieved in multiple ways to produce polymers with specific forms and chemistry, such as block polymer (PLGA), brush grafts and composites of polymers. PLGA is one of the most popular block co-polymers wherein the hydrophobicity of the polymer can be controlled by the extent of glycolic acid incorporated. However, the extent of lactic acid determines important physical properties such as degradation, and mechanical strength of the polymer. Hence, the ratio of lactic acid to glycolic acid can be altered to provide a range of physical properties. However, most tissue engineering applications demand good mechanical properties and degradation times that are not short. Therefore, the extent of glycolic acid in the PLGA polymers is often limited. Hence, PLGA polymer surfaces do not provide appropriate hydrophilicity for cell and protein adsorption. To circumvent this problem several types of other combinations of PLGA with synthetic and natural polymer have been explored including PLGA-g-poly (L-lysine) [89] and PLGA-PEG-NH 2 composites [90] have been developed for the delivery of chondrocytic cells and angiogenic growth factors respectively. Similarly, PLA has also been explored with several type of combinations including PLA-PCL, di- and tri- block PLA-PEG (poly (ethylene glycol)), PLA-PEG multi-block copolymers (Fig. 6), polycaprolactone/polylactide/poly (ethylene oxide) copolymer, star- and dendrimer-like copolylactides, poly(l-lactide-co-RS--malic acid), polylactide-gdextran, and polylactide-g-chitosan [60]. Protein and cell binding affinity of these PLA composites have been studied for tissue engineering applications and the results demonstrated improved hydrophilicity and cell adhesion properties.

hydroxy and carboxyl group on polymer surfaces [86]. Most of these surface modified polymers demonstrated enhanced protein adsorption and cell affinity as compared to the unmodified control. Although blending of synthetic polymers with natural polymers or hydrophilic polymers can influence surface properties, they adversely affect degradation behavior and biomechanical properties which is not desirable for tissue engineering applications. Therefore, the prospect of blending is not very desirable. Apart from the methods described above, there are other approaches that have been explored for the functionalization of biomaterial surfaces. One such method is vapor phase grafting (VPG). Reports, wherein acrylamide, maleic anhydride and vinyl pyrrolidone have been grafted on PLLA surface have demonstrated increase in its surface wettability [91, 92]. In another recent study N-vinylpyrrolidone grafted surfaces demonstrated improved surface wettability as well as cell attachment [91]. The VPG technique, unlike the plasma treatment method, offers the advantage of control on temperature during the reaction which prevents the polymer from degradation. Hence VPG enables polymer grafting at low temperature and could be a potential approach for improving the surface properties of PLLA.

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ii Modification of Polymeric Biomaterial Surface for Protein Immobilization

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Another type of polymer modification used for improvement of hydrophilicity is brush grafting, wherein, hydrophilic polymer brushes bearing alcoholic and hydroxyl groups are introduced onto polyester surfaces by photopolymerization. In addition, hydroxyethyl methacrylate, methacrylic acid and acrylamide have been used to introduce

In the previous sections protein immobilization using physical adsorption, chemical treatment, plasma treatment and photochemical treatment have been discussed. This section discusses the immobilization of proteins with a focus on the development of protein delivery systems. The reason for this focus is the importance of protein delivery systems in tissue engineering. Some of the approaches used under this section could be overlapping with the approaches discussed in the previous sections, however, to enable a proper emphasis on protein delivery systems, they have be reiterated here as and when appropriate.

One of the focuses of tissue engineering is appropriate delivery of therapeutic proteins or cell signaling molecules to augment the process of regeneration [26, 93]. Towards this end, several biodegradable material based carrier systems have been explored in the past two decades [27, 33] such as micro-and nano-particles [94, 95] protein encapsulated hydrogels [96], coating and/or blending of growth factor binding protein and other ECM components such as GAG, heparin, collagen [97, 98] and fibrin [99] during or after scaffold synthesis (Fig. 7). Physical adsorption or entrapment/encapsulation are the two major methods for protein immobilization in all the aforementioned approaches. Most of these studies demonstrated the controlled and

CH3 O

O

Tetraphenyltin

+ O

O

O

H

15 hrs

O O

O CH3 Lactide

O

O O O

PEG

Fig. (6). Synthesis of co-polymer of Poly (lactide)-co-poly (ethylene glycol) (PLA-PEG).

PLA-PEG-PLA



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Fig. (7). Methods of Protein/ Growth factor delivery for tissue engineering application.

sustained release of proteins, however, they are associated with limitations such as random orientation, over-crowding and rapid desorption of adsorbed proteins. Moreover, loss of bioactivity due to the short half-lives of these signaling proteins combined with reduced bioavailability further limit the adsorption or entrapment based delivery systems [34]. Therefore, covalent immobilization of proteins is gaining increased attention for tissue engineering applications [67, 100]. The concept of covalent immobilization of enzymes (functional proteins) and cells has been extensively explored previously for downstream processing and enzyme technology applications [101]. The approach of covalent immobilization of signaling proteins provides oriented immobilization with minimal loss of structure, conformation and protein spreading. The reduced spreading allows for immobilization of higher concentration of proteins as well as quantification of the immobilized protein becomes more realistic. In addition, covalent immobilization also reduces the amount of signaling/therapeutic protein required for triggering an appropriate function as compared to adsorption. Proteins can be covalently immobilized on polymer surfaces via two approaches (i) by modifications in protein and/or polymer structure or (ii) chemically crosslinking the polymer and the protein. Covalent attachment can be classified in two broad categories: random and oriented covalent attachment

[68]. Protein structures contain a variety of functional groups including amino, carboxyl, hydroxyl and thiol. The functional group can readily be used for covalent binding to polymeric surfaces with complementary chemical functionalities on the polymer. In the mid-1990’s Ito et al. reported the chemical immobilization of signaling protein insulin on several polymeric surfaces including hydrolyzed poly (methylmethacrylate) [102], polyurethane, NaOCl-treated polyacrylamide beads and poly(hydroxyethylmethacrylateco-ethylmethacrylate) [100]. The insulin was immobilized onto the polymers with the help of water soluble carbodiimides or dimethyl suberimidate crosslinking agents that form a covalent bond between the carboxy group of the polymer and the amine terminals on proteins. In vivo studies conducted on all the aforementioned polymeric systems demonstrated a stronger signaling effect when compared to freely available insulin thereby indicating an improved performance by the covalently immobilized proteins [52]. Similarly, there are other approaches for cross-linking the available functional groups of protein with those on the polymer surface. Most of these strategies make use of specialized cross-linkers designed for both attachment and physical separation of the protein from the polymer surface, thereby enabling a greater amount of the protein functional domain to be exposed to the solvent [68].


In a recent study Kitajima et al. demonstrated the coexpression of growth factor and growth factor binding protein (GFBP) with the help of recombinant DNA technology. Since growth factors are bound to and released from GFBP in the physiological environment, this approach could potentially eliminate the need for chemical immobilization and significantly reduce the problem of loss in bioactivity and bioavailability [103]. Nishi et al. demonstrated the expression of fusion proteins carrying basic fibroblast growth factor (bFGF) at the N terminal of collagen binding domain (CBD). The collagen bound fibroblast growth factor (CBFGF) stimulated the growth of BALBc 3T3 fibroblasts as much as their infused counterparts. Moreover, the presence of CBFGF was observed up to 10 days post injecting subcutaneously in nude mice, whereas loosely injected FGF was not detectable 24 h after injection [104]. Recently, Kitajima et al. demonstrated the expression of a fusion protein consisting of hepatocyte growth factor (HGF; an angiogenic factor) and a CBD polypeptide of fibronectin (FN) for tissue engineering applications [103]. Their study demonstrated that CBD-HGF produced stronger collagen binding activity than native HGF and promoted the growth of endothelial cells (ECs) to a greater degree than native HGF added to the culture medium. This study demonstrated that the fused CBD moiety helped in the immobilization of HGF on collagen as well as in the stabilization of the fusion molecule. The collagen sponges containing bound CBDHGF were implanted subcutaneously in rats to study the angiogenic activity of CBD-HGF. Their results demonstrated a 46-fold increase in blood vessel formation after 7 days in the CBD-HGF containing sponges as compared to the control sponges. Therefore, fusion protein containing scaffolds can potentially circumvent the need for chemical immobilization of proteins [103]. In another recent study the C terminal of EGF protein was modified using hexa histidine residues. This histidine terminal protein was used for grafting onto Ni (II)-chelated surface of a glass-based substrate for efficient expansion of neural stem cells(NSCs) [105]. The results of this study demonstrated the highly selective subculture and further expansion of neural stem cells (while retained their multipotency) on the developed surface. In another similar study vascular endothelial growth factor was expressed with an N-terminal Cys-tag wherein the free sulfhydryl group of cystine was used for site-specific conjugation with fibronectin [106]. This study demons-trated that site-specific immobilization provides a reliable method for permanent deposition of growth factors via Cys-tag on synthetic scaffolds. Both the aforementioned studies demonstrated co-expression of metal binding residues with growth factor for their covalent immobilization on respective metal bounded polymeric surfaces. Although these studies need in-vivo validation, the preliminary results show potential for development as next generation scaffolds for protein delivery application.

Katti et al.

functional domains of proteins is the most popular approach for chemical modification. There are a number of reagents and cross-linkers that have been reported for site specific modification of protein and attachment [68]. These crosslinkers have been used for degradable and non-degradable crosslinking of proteins. However, very few of them have been explored for tissue engineering application such as EDAC, and ethylenediamine. The aforementioned approaches provide improved delivery vehicles for therapeutic and signaling proteins, however, most of these methods lead to covalent immobilization that is random and non-specific and hence require higher concentrations of protein when compared to oriented immobilization. In addition, the cleavage of covalent bonds for release of protein is a task that can be challenging. Many studies that report the use of these approaches are still under in-vitro investigations and need to undergo in vivo investigations before reaching the clinical trail stage. Therefore, these approaches still have a long way to go before they can be used practically in a clinical setting.

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Oriented covalent immobilization is another popular approach that is used in industries where protein activity is critical and the ability to reuse the immobilized protein multiple times is desirable. However, this approach has not been explored for tissue engineering application as extensively as it has been explored for enzyme technology applications [67]. Modification in side chains of non-

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The fundamentals of protein attachment on biomaterial surfaces have been widely reviewed, however, the prediction regarding attachment still remains a challenge because of the extreme diversity, sensitivity and complexity of signaling/ therapeutic proteins. Moreover, the use of chemicals and crosslinkers for immobilization can potentially induce protein denaturation and hence there is a need for improved approaches that can overcome these limitations as well. 4. CONCLUSION AND FUTURE PERSPECTIVE

Interdisciplinary efforts being made for the development of scaffolds to be used in tissue engineering have made current day scaffolding systems versatile. Better understanding in the aspects of cell biology and biomaterial science such as cell-cell, cell-ECM, cell-biomaterial, and biomaterial-protein interactions, hierarchical arrangement of cell in tissue, and effects of various signaling molecules and their combinations on cell behavior, have contributed significantly to the design of scaffolds for improved tissue regeneration. Identification of functional domains of various ECM proteins have lead to biomimetic approaches that involve precision immobilization of peptides (functional domains) and proteins on the surface of polymeric scaffolds. The ability to control non-specific adsorption of proteins and enable oriented immobilization without denaturing protein conformation is desirable for two reasons (i) improved cell attachment, and (ii) the development of controlled delivery system for therapeutic proteins.

This review briefly describes the different approaches used to modify polymer surface properties via protein/ peptide immobilization. In times to come designing scaffolds with nanometer scale surface features would probably gain more importance. Future studies, both in vitro and in vivo, on such scaffolding systems that possess nanoscale surface features will bring more clarity in the area of biomaterial surface engineering. It is hoped that this clarity will instigate the development of next generation scaffolding systems with nanoengineered surfaces that elicit improved cell behavior and can act as a delivery system for therapeutic proteins. Similarly, fusion protein-based protein delivery systems that


can potentially circumvent the limitation associated with the conventional approaches have demonstrated potential to be developed as the protein delivery systems of the future. Therefore, a combination of physical modification (nanoengineering) and protein immobilization would enable the successful design of scaffolding systems for tissue engineering applications. 5. ACKNOWLEDGMENTS

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One of the authors RV, would like to acknowledge the fellowship received from Council of Scientific and Industrial Research, India. DSK would like to acknowledge the grant received from Department of Biotechnology, India.

[16] [17] [18]

ABBREVIATIONS

[19]

ECM

=

Extracellular Matrix

RGD

=

Arginine-Glycine-Aspartate

[20]

PHSRN

=

Proline-Histidine-Serine-ArginineAsparagine

[21]

PLA

=

Poly (lactic acid)

PCL

=

Poly (caprolactone)

PLGA

=

Poly (lactide-co-glycolide)

PGA

=

Poly (glycolide)

PVA

=

Poly (vinyl alcohol)

EDAC

=

1-ethyl-3-(3-dimethylaminopropyl)carbodiimide

[25]

[26]

[22]

=

N-hydroxysuccinimide

PLAL

=

Poly(lactic acid-co-lysine)

bFGF

=

Basic fibroblast growth factor

VPG

=

Vapor phase grafting

PEG

=

Poly (ethylene glycol)

GAG

=

Glycoaminoglycan

HGF

=

Hepatocyte growth factor

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