Biomimetic Biomaterials || Biomimetic coatings for biomaterial surfaces

© Woodhead Publishing Limited, 2013

91

4 Biomimetic coatings for biomaterial surfaces

E. MÁZL CHÁNOVÁ and F. RYPÁČEK, Academy of Sciences of the Czech Republic, Czech Republic

DOI: 10.1533/9780857098887.1.91

Abstract: This chapter outlines the role of surface modifi cations for biomaterials used in tissue engineering, and reviews approaches to creating biomimetic coatings on several types of biomaterial. In the fi rst section, the principles used in preparing nonfouling surfaces with minimized non- specifi c interactions with proteins are discussed, followed by the creation of biomimetic structures which promote controlled cell- surface interactions. The following sections go on to provide examples of key methods used in relation to specifi c materials, including polymers, metals and ceramics. Finally, methods for the characterization of coating effi ciency are reviewed.

Key words: tissue engineering, biomimetic coating, cell– biomaterial interactions, biomaterials.

4.1 Introduction

Current research trends in regenerative medicine are focused on methods promoting tissue self- regeneration, or the replacement of damaged tissue with a novel tissue developed ex vivo through tissue engineering. Tissue engineering is an interdisciplinary approach that combines material engineering methods with knowledge of biology and medicine (Langer and Vacanti, 1993). In tissue engineering, an artifi cial, three- dimensional (3D) biomaterial structure is combined with living cells to serve as a temporary support, for example as a scaffold for growing tissue. Materials used in tissue- engineering scaffolds can be based on natural or synthetic polymers, metals, ceramics and/or their composites (Park, 1995).

The requirements for scaffold design generally depend on the intended application and arise from the biological demands of the specifi c cells or tissues to be regenerated. Besides biocompatibility, non- toxicity, suitability of the 3D structure and matching mechanical properties, the quality of surfaces belongs to key scaffold properties. Many scaffolds used in the area of tissue engineering are prepared from synthetic polymers, the structure and physical properties of which can be tailored by rational synthesis. In addition, various processing techniques can be applied to produce scaffolds with unique characteristics, such as desired shape, 3D pore architecture and mechanical properties, to suit the macroscopic requirements of a particular application. Figure 4.1 shows a typical example of a 3D scaffold made of poly-L-lactide (PLLA) with variable internal structure.

92 Biomimetic biomaterials


Various approaches to scaffold manufacturing have been thoroughly reviewed in the literature (Lanza et al. , 2000, Atala and Lanza, 2002).

In addition to its shape- forming role and mechanical strength, the scaffold supports the seeded cells and growing tissue primarily through the interaction of the cells with the scaffold surfaces. In a broader sense, the scaffold represents a synthetic analogue of extracellular matrix (ECM). Therefore, to actively promote formation and/or regeneration of a functional tissue, the scaffold surfaces must also provide the additional cues that would mimic the ECM-cell signalling mechanisms. To this end, the biomaterial used to construct the scaffold should have a capacity to expose the surface, on which such biomimetic cues can be provided through adjustment of the structure, reactivity and presence of specifi c functional groups.

When a biomaterial is brought into a biological environment, such as the human body or a cell culture, a series of events occur between the biomaterial surface and the living cells. These interactions are critical for the success of any biomedical device. In the fi rst instance, the proteins contained in biological fl uids (blood plasma, tissue liquor, tissue culture media) adsorb to the surface of the device. Typically, by the time the cells approach the surface, it is already covered with a layer of proteins. Therefore, the approaching cells do not recognize the material itself, ‘seeing’ instead a dynamic layer of proteins (Hubbell, 1995, Lee et al. 1995). Any discussion of the role of biomaterial surfaces in cell adhesion must therefore take into account the effects the surface properties of a material have on protein adsorption.

4.1 Three- dimensional polymer matrix for cell culturing made by spinning of poly-L-lactide solution and subsequent pressing (a) with diameter of 1 cm and the detail of internal structure (b) observed by scanning electron microscopy.

Biomimetic coatings for biomaterial surfaces 93


4.2 Issues being addressed through biomimetic

coatings

The phenomena occurring at biomaterial– living system interfaces will now be discussed, and approaches to preparing controlled microenvironments on biomaterial surfaces, for cell- surface interactions that mimic the processes in natural tissues, will be reviewed.

4.2.1 Controlling protein adsorption and the creation of nonfouling surfaces

The microenvironment of immobilized functionalities has been proven to be an important factor in effective biomaterial/cell receptor interactions. However, to study the effect of attached biomimetic groups, the surface must fi rst be resistant to non- specifi c protein adsorption. Without this, the biomimetic motifs present would be masked by an adsorbed layer of proteins that occurs at the fi rst instant after the biomaterial is brought into contact with protein- containing biological fl uids. The subsequent events would then be primarily determined by an adsorbed layer of proteins. A modifi cation of the surface with such materials as hydrophilic polymers, for example, which suppress the non- specifi c protein adsorption, is therefore required (Chen et al. , 2008).

Protein adsorption onto surfaces is a complex and dynamic process which involves non- covalent interactions, including hydrophobic or electrostatic interactions, hydrogen bonds and van der Waals interactions. The adsorption process is infl uenced by protein properties (e.g. primary structure, size, structural stability and diffusion) as well as by the character of the biomaterial surface properties (including surface energy, roughness or chemical composition). Additionally, multi- component systems, such as blood plasma, exhibit dynamic profi les of adsorption, where the more concentrated proteins or those with higher mobility adsorb on the surface in the fi rst instance, and may be replaced over time with proteins of higher surface affi nity.

It is also interesting to note that many proteins undergo conformational changes when they adsorb on a solid surface, which can change their biological activity (García, 2006). In this respect, a surface coating with a hydrophilic polymer layer such as poly(ethylene oxide), PEO, preferably in a brush conformation, can create an energetic barrier resulting in good surface shielding against protein adsorption (Elbert and Hubbell, 1996, Chen et al. , 2008). Theoretical studies dealing with the polymer brush model, its performance in contact with protein molecules, and its ability to suppress protein adsorption are given by Szleifer (Szleifer, 1997, Carignano and Szleifer, 2000) and Halperin (Halperin, 1999, Leckband et al. , 1999). Surface- tethered PEO was proven to be effi cient in the creation of a nonfouling surface for a number of biomedical applications, as well as for conjugates of drugs. Its resistance to protein adsorption is explained by steric repulsion and



barrier effects, which result from the solvation of PEG chains with water molecules. Besides PEO, a range of other biocompatible polymers have been studied for their nonfouling properties, including poly(vinylalcohol), poly(hydroxyethyl methacrylate), poly(hydroxypropyl methacrylamide) and poly(acrylamide). Growing attention is being placed on polybetaines, with particular focus on carboxybetaine or sulfobetaine, as well as polymers based on fosforylcholine (Chen et al. , 2008). Structures of selected polymers studied as antifouling agents are depicted in Fig. 4.2 .

In general, the surfaces of 3D polymer scaffolds can be modifi ed with nonfouling polymers either through chemical (covalent) binding or physical methods. Irrespective of the method used, the fi nal nonfouling layer should be uniform, contiguous across the entire scaffold surface (including internal parts) and stable under physiological conditions. The immobilization process and resultant layer should not signifi cantly change the internal architecture and mechanical properties of the original 3D scaffold. In addition, the nonfouling layer should exhibit a capacity to be functionalized with bioactive molecules.

Chemical/covalent modifi cation

PEO is particularly suitable for coating with a layer of end- tethered chains, because of the presence of its end- functional groups. Monofunctional or heterobifunctional PEO derivatives carrying various functional groups (amine, thiol, propargyl) can be prepared, and some such derivatives are already commercially available. On fl at substrates, such as glass or Si wafers (Si/SiO 2 ), the nonfouling layers of PEO can be formed by spin- casting and/or surface

4.2 Structure of selected polymers used for preparation of nonfouling surfaces: (a) poly(ethylene oxide), (b) poly(hydroxyethyl methacrylate) and (c) poly(carboxybetaine).



deposition of Langmuir-Blodgett fi lms. The creation of self- assembled monolayers (SAMs) of thiols on gold substrates, with the subsequent covalent immobilization, is a well- established method. Such model surfaces allow quantitative analysis using surface plasmon resonance (SPR) or ellipsometry, for characterization of PEO layer (thickness), as well as determination of dynamic processes, such as polymer layer stability or protein adsorption in situ (Kim et al. , 2008, Fragneto, 2009). However, although these methods for immobilization are very useful in the preparation of model systems for fundamental and theoretical studies, they cannot usually be applied to 3D polymer matrices with a complex shape or different material properties.

Covalent grafting of polymer chains allows for stable modifi cation, but requires, however, the presence of reactive functional groups on the surface. In the ‘grafting- to’ approach, existing polymers carrying a corresponding functional group can be immobilized. The ‘grafting- to’ approach has some limitation, particularly with respect to the chain density that can be achieved. When the modifi cation is carried out from a solution, the polymer chain typically assumes a coil conformation with a certain hydrodynamic volume of the coil. The dimensions of the molecules in the solution determine the density at which the polymer molecules can occupy the surface. As the polymer layer develops, the incoming polymer chains have to diffuse through the layer of already bound chains, and the effect of steric barriers increases in line with the increasing thickness of the polymer layer. As such, the grafting density achieved through ‘grafting- to’ from a polymer solution is typically insuffi cient to form a real polymer brush. Nevertheless, the immobilization of PEO precursors remains the most frequently used approach to the production of stable poly(ethylene oxide) coatings (Gong et al. , 2000, Chen et al. , 2008). The effect of polymer free volume can be minimized by carrying the ‘grafting- to’ reaction from a polymer melt. A substrate- independent modifi cation based on a dopamine- melanine anchoring layer and subsequent grafting of PEO chains from the melt has been developed by Pop-Georgievski et al. (2011). Grafting of PEO from melt in a temperature range 70–110°C produced densely packed PEO layers showing exceptionally low protein adsorption when exposed to whole blood serum or plasma.

The ‘grafting- from’ process appears to be more effective in producing dense polymer brushes. This process requires that polymerization initiator groups are fi rst immobilized on the surface, followed by in situ polymerization of hydrophilic monomers. The steric hindrances are not as signifi cant, because of the small size of the approaching monomer molecules. Depending on the surface density of the polymerization initiator, the distances between the neighbouring growing polymer chains can be as low as to be comparable with the dimensions of monomer molecules. In addition to the self- assembled monolayers (SAM) on model substrates previously mentioned, the plasma treatment or discharge in the presence of amine, and/or UV irradiation are often used to introduce surface functional groups for the binding of polymerization initiators to polymers.



Controlled radical polymerization techniques, such as atom transfer radical polymerization (ATRP) or TEMPO-mediated polymerization, are frequently used for in situ polymerization. Although the controlled radical polymerization is technologically a more demanding process, requiring precise control of each step, it provides polymer brushes with a well- controlled molecular weight of tethered polymer chains and consequently facilitates well- controlled thickness of the coating layer. (Zhao and Brittain, 2000) Polymer brushes of hydroxy- and methoxy- capped oligoethyleneglycol methacrylate and carboxybetaine acrylamide can be grafted from bromoisobutyrate initiators attached to different surfaces, including polypropylene and Ti-Al-V surface by ATRP (Rodriguez-Emmenegger et al. , 2011a). Well- controlled polymerization kinetics made it possible to control the thickness of the brushes at a nanometre scale, and zero fouling from single protein solutions and a reduction of more than 90% in fouling from blood plasma on coated surfaces was achieved. The developed coating can be built on various classes of substrate, and its antifouling properties are preserved even in undiluted blood plasma. This is therefore useful for fabrication of biotechnological and biomedical devices with tailor- made functions, as has been demonstrated through functionalization with bioactive compounds, via the covalent attachment of streptavidin onto a poly(oligoethyleneglycol methacrylate) brush and the subsequent immobilization of model antibodies and oligonucleotides.

Rodriguez-Emmenegger et al. (2011b) studied poly[N-(2-hydroxypropyl) methacrylamide], poly(HPMA) and brushes grafted from gold surfaces, to which ATRP initiator groups were introduced through functional SAM. The protein adsorption from solutions of single plasma proteins, as well as full blood plasma, was investigated via the SPR method. The effect of the poly(HPMA) brush was compared with other antifouling brushes, such as poly(carboxybetaine acrylamide), poly(CBAA), poly(oligo(ethylene glycol) methyl ether methacrylate), poly(MeOEGMA) and with the widely used antifouling SAMs of hexa(ethylene glycol) undecanethiol on gold (OEG6). Although the poly(HPMA) surface exhibited only moderate hydrophilicity ( θ adv = 40°), it was found that together with highly hydrophilic poly(CBAA) with ( θ adv = 23°), both of these polymers were capable of preventing fouling from blood plasma. These results, together with previous observations from the same laboratory (Rodriguez-Emmenegger et al. , 2009) indicate that neither hydrophilicity nor surface wettability alone can explain the blood plasma fouling phenomena.

Physical binding techniques

A physical entrapment of polymers is the simplest method for immobilization of PEO or any other hydrophilic polymers onto the polymer biomaterial surface. In general, the biomaterial surface layer is fi rst swollen with a mixture of good and poor solvent (or precipitant) for the polymer to be modifi ed. The solvent mixture composition is selected in such a way that PEO is soluble in both components of



the solvent. After immersion of the biomaterial to solvent mixture, the surface layer of the biomaterial is swollen and the polylactide (PLA) chains become more mobile, thus enabling diffusion and penetration of the PEO macromolecules. The collapse of the swollen surface layer is caused by addition of precipitant, and the PEO chains become physically entrapped in the superfi cial layer of the biomaterial (see Fig. 4.3 for the scheme). Surface- modifi ed PLA has been prepared using this method. The PLA surfaces modifi ed via either physical entrapment of PEO or poly(L-lysine) PLL were compared. Although the entrapment of PEO showed good accord between the density of PEO on the surface and the period of contact with solution, producing a resultant surface layer exhibiting uniform surface properties, the PLL formed a heterogeneous surface. A stable modifi cation can therefore be obtained, but the properties of the resulting surface layer depend on the compatibility of the polymers used (Quirk et al. , 2000).

Polymer blends with PEO block copolymers take advantage of the preferential segregation of one phase on the surface of the blend, thus revealing another method for the preparation of a surface- immobilized layer of hydrophilic polymer. The segregating phase exposed to the surface is fi xed in the bulk through chain entanglements, which contribute to its stability. However, the addition of block copolymer into the original polymer usually changes such material properties as mechanical or processing parameters. Modifi cation of poly(methyl methacrylate), PMMM, through blending with PMMA/PEO graft copolymers has been studied by Griffi th’s group (Walton et al. , 1997, Irvine et al. , 2001). The blends of PMMA with a copolymer content above 2 wt.% exhibited surface enrichment with PEO. The adsorption of albumin on such surfaces decreased by about 50% in comparison with pure PMMA, and cell adhesion was completely suppressed (Walton et al. , 1997). Analogously, blends of PLA with PMMA/PEO graft copolymer were prepared and effi cacy in obtaining nonfouling properties was demonstrated. In addition, when a fraction of the copolymer was substituted with a functionalized

4.3 Scheme of physical entrapment of poly(ethylene oxide) (PEO) chains onto polylactide (PLA) surface.



sample of identical molecular parameters, an increased fi broblast adhesion was achieved, in contrast to such a surface prepared by blending PLA with a non- functionalized copolymer (Irvine et al. , 2001).

Deposition of amphiphilic block copolymers represents another method for the creation of a hydrophilic coating on a hydrophobic polymer material. In this concept, the segment of block copolymer, having higher affi nity to the modifi ed surface, adsorbs to the substrate through physical bonds and serves as an anchor for the copolymer, whereas the hydrophilic block of the copolymer remains exposed to the aqueous solution. Amphiphilic block copolymers containing PEO as the hydrophilic component can easily be prepared with a variety of polymers acting as the hydrophobic blocks, and some of these are already commercially available. Although PEO itself has a low affi nity to hydrophobic surfaces, the introduction of PEO in a block copolymer with a hydrophobic polymer allows its immobilization through hydrophobic interactions (Currie et al. , 2003). Amphiphilic block copolymers in aqueous solutions or selective solvents, above a critical concentration, self- assemble into associated supramolecular structures – micelles – in order to minimize contact between the hydrophobic block and aqueous environment or solvent. In a micellar solution, the copolymer aggregates are in equilibrium with the single polymer chains (unimers), and when the solution is brought into contact with a hydrophobic surface or a polymer compatible with the micelle core, both the micelles and unimers can adsorb on the surface (see Fig. 4.4 for the scheme). In addition, adsorbed micelles can relax on the surface, resulting in tight contact between the micelle core polymer and the substrate surface. The process of adsorption and micelle relaxation depends on the structure and molecular parameters of the block copolymer, the conditions of deposition, the solvents and the surface properties. The adsorption of micelles occurs quickly when the micelle core exhibits an affi nity to the substrate, whereas the adsorption of unimers is signifi cantly slower (Munch and Gast, 1990).

4.4 Adsorption of amphiphilic block copolymers from aqueous solution or a selective solvent.



The mechanism of A-B block copolymer adsorption on the surface has been studied, taking into account copolymer concentration, critical micellar concentration and the infl uence of the solution’s copolymer concentration on the total amount of copolymer adsorbed (Zhan and Mattice, 1994a). This has been modelled using the Monte Carlo method (Zhan and Mattice, 1994b). The effects of lyophobic block length and solvent quality on the critical micellar concentration of copolymer in selective solvents (thus the mechanism of adsorption on the surface) have been established using Self-Consistent Field theory (Lent and Scheutjens, 1989). Polymer chains in micelles are less mobile than low- molecular weight surfactants. The length of block forming the micelle core, and the effect of this on the unimer exchange has been established and confi rmed practically. A core of the micelle can be ‘frozen’ in a glassy state, for example, when suitable selective solvent composition and molecular copolymer parameters are used. The exchange of unimers can be suppressed or eliminated, and the micelle stability prolonged from seconds, minutes and hours up to a state of permanence in micelles with cross- linked cores (Förster and Antoniemi, 1998). The effects of copolymer composition (block lengths) and temperature on unimer exchange have been studied in relation to various copolymers via steady- state fl uorescence measurement (Gohy, 2005). In the context of biomaterial modifi cation, the rate of unimer exchange in the system of PLA- b -PEO block copolymer micelles in an aqueous environment was proven to decrease with increasing length of the insoluble PLA block (Popelka et al. , 2005).

Di- block copolymers of poly(ethylene oxide) and poly(propylene oxide), PEO-PPO and commercially available PEO-PPO-PEO tri- blocks (Pluronics) are of interest in experimental adsorption studies. The effect of copolymer composition on the amount and chain density of adsorbed copolymer and, consequently, the resulting resistance of the adsorbed copolymer layer to albumin adsorption has been studied on various surfaces, including hydrophobized SiO 2 (Norde and Gage, 2004), polystyrene PS (Su et al. , 2008) and others (Currie et al. , 2003). Block copolymers composed of PEO and aliphatic polyester, such as polycaprolactone (PCL) or PLA, frequently have been studied. Such amphiphilic block copolymers extend the applicability of their building components in terms of biocompatibility and biodegradability. PLA-PEO or PCL-PEO copolymers can be readily prepared with well- defi ned molecular parameters through controlled ring- opening polymerization of lactones, with PEO as a macromolecular co- initiator (Popelka et al. , 2005). PLA-PEO copolymers in aqueous solutions (or in selective solvents) form nanocolloids with a hydrated poly(ethylene oxide) shell shielding the hydrophobic PLA core. The hydrophobic micelle core can solubilize low- molecular-weight compounds, for example hydrophobic drugs, making such nanocolloids applicable to drug delivery. While the micelle’s core can carry the drug payload, the PEO coating prevents the nanoparticles from non- specifi c protein adsorption, thus preventing their capture by cells of the reticulo- endothelial system (RES) (Hagan et al. , 1996, Novakova et al. , 2002, Otsuka et al. , 2003).



Furthermore, the nanoparticle surface can be modifi ed with specifi c molecules, facilitating their targeting to particular tissues (Bae et al. , 2009). Block copolymers with different functional groups at the end of the PEO block can be prepared and used in the preparation of functionalized surfaces for cell studies (Cannizzaro et al. , 1998, Black et al. , 1999, Tessmar et al. , 2002, Tessmar et al. , 2003). The modifi cation of PLA surfaces using PLA-PEO block copolymers with different block lengths has been evaluated by its effect on PLA surface wettability, using contact angle measurements and determination of albumin adsorption (Otsuka et al. , 2000). Additionally, the copolymer can be also modifi ed at the end of the PLA chain, making it possible to stabilize the core of copolymer micelles through cross- linking and resulting in copolymer micelles with higher stability. When deposited on the surface, these will adhere without deformation (Emoto et al. , 1999).

Because the PLA block can be prepared with various degrees of stereoregularity, depending on the use and ratio of L-lactide or D,L-lactide, the degree of crystallinity of the PLA phase in a block copolymer layer can be modifi ed accordingly. In cases where blocks of copolymers are too short or copolymers are composed of amorphous blocks, their phase separation is driven by the incompatibility of the phases. In addition, the morphology and phase separation of PLA-PEO block copolymers can also be affected by the degree of crystallinity (Sommer and Reiter, 2006). The morphology of PLA-PEO block copolymers has been extensively studied with respect to the conditions of deposition (solvent, concentration, temperature), molecular parameters of copolymers (molecular weight, PLA or PEO block length, stereoregularity of PLA block) and the nature of the substrate (hydrophilic or hydrophobic, PLA surface and mica, for example) (Fujiwara et al. , 2000, Muller et al. , 2000a, Muller et al. , 2000b, Fujiwara et al. , 2001, Huang et al. , 2008, Huang et al. , 2009). These studies have shown that controlled deposition of PLA-PEO copolymer nanocolloids can result in nanoscale organized surfaces, which can be tuned through adjustment of the overall process conditions.

4.2.2 Biomimetic coating to control biomaterial– cell interactions through integrin- targeting and focal adhesions

The ECM, a complex system of polysaccharides and proteins produced by cells, is a fundamental structural element of living tissues. It provides a mechanical support to cells as well as binding ligands, through which cells are attached to the ECM by their surface receptors. Extensively studied groups of receptors involve integrins, also called adhesive receptors, which mediate cell interactions with ECM and, potentially, cell adhesion to artifi cial materials. Integrins are transmembrane proteins, which in its extracellular part contain binding sites for specifi c interactions with ECM molecules, for example fi bronectin, laminin or



collagen (Alberts et al. , 1994). It is well known that the activity of integrin/ECM interactions can be simulated by the interaction of integrins with small peptide fragments derived from ECM proteins, for example fi bronectin; the most studied RGD sequence (Pierschbacher and Ruoslahti, 1984).

Besides RGD, other sequences have been proven to provide specifi c integrin interactions. Some integrin receptors, for example, require additional interaction with another sequence in a molecule of fi bronectin in order to provoke effective ligand interactions. Such a sequence is PHSRN, proline- histidine-serine- arginine-asparagine, that itself alone is negligible. However, together with RGD, this sequence can signifi cantly increase cell adhesion (Redick et al. , 2000). The crucial aspect of the synergic effects of RGD and PHSRN is spatial arrangement, as it has been proved that their increasing distance disrupts specifi c interactions with integrin, preventing the occurrence of the effective signal transmission (Grant et al. , 1997).

Other peptide motifs found in ECM proteins are believed to enhance the affi nity of particular cell types. As examples, REDV sequence (arginine- glutamic acid- aspartic acid- valine) from fi bronectin was reported to specifi cally promote adhesion of endothelial cells, SIKVAV sequence (serine- isoleucine-lysine- valine-alanine- valine) derived from laminin has been shown to increase adhesion of neural cells, their migration and neurite growth (Yamada and Kleinman, 1992), and KRSR sequence (lysine- arginine-serine- arginine) has been demonstrated to elevate adhesion of osteoblasts while inhibiting the adhesion of fi broblasts (Dee et al. , 1998).

Besides adhesion proteins, ECM also contains a number of morphoregulatory molecules, such as growth factors, cytokines and chemokines. The mechanism of action and the combination of growth factors with materials are described in the literature (Ito, 2009, Gibson et al. , 2009, Leslie-Barbick et al. , 2009, Jorissen et al. , 2003, Bab and Chorev, 2002, Mann et al. , 2001, Maheshwari et al. , 2000). The growth factors can be immobilized both as whole molecules and as fragments mimicking a particular binding domain, as has been described for human fi broblast growth factor-2 (Lee et al. , 2007).

4.3 Approaches to the creation of biomimetic

surfaces

In order to control cell adhesion, growth and differentiation on an artifi cial biomaterial scaffold, the biomaterial– cell interface must be designed to provide an analogous environment for material– cell communication, as it is provided using ECM. This requires a biomaterial surface modifi cation, exposing structural surface motifs specifi c to cell integrins and/or other cell membrane receptors. In order to control the effect of such modifi cation, we need to suppress non- specifi c interactions with the biomaterial surface of proteins that could otherwise mismatch the added specifi c cues. The approaches to prevention of non- specifi c protein adsorption have been discussed in previous paragraphs.



Techniques used in the treatment of biomaterial surfaces differ, depending on the material and its application. Some biomimetic modifi cation strategies can be applied to both polymer biomaterials and metal surfaces. A seemingly straightforward approach to the introduction of chemical cues for integrin- mediated cell adhesion is to coat the biomaterial surface with native ECM or blood plasma proteins. Adhesive molecules of ECM, such as fi bronectin, collagen or laminin, have been applied to a variety of surfaces by either physical adsorption or covalent binding. However, the results obtained when comparing covalently and physically bound proteins are not consistently positive, as reviewed by Chen et al. (2008). Some researchers observed that covalently conjugated proteins, such as fi bronectin, are favourable to the behaviour of some cells. In contrast to another study, the physically adsorbed fi bronectin promoted cell adhesion, while the interaction of a covalently bound protein with cells was inhibited. These results indicate that during the immobilization of whole proteins to the surface, various infl uences, which may hardly be controlled, can modify the protein activity. The composition of the adsorbed protein layer, orientation of the protein molecules, conformational changes upon binding and lack of control of the surface distribution or pattern can be identifi ed as the most important factors.

These diffi culties prompted many researchers to use short, well- defi ned synthetic peptide fragments (as described in previous paragraphs) instead of whole ECM proteins. Although the effi ciency of peptide ligands in promoting integrin- mediated cell adhesion has been confi rmed by many studies, there is increasing evidence that it is not only the presence of peptides, but also their spatial distribution on the surface which is important. The research on surface modifi cation using RGD fragments was reviewed by Hershel et al. (2003). In 1991, Massia and Hubbell had already reported their model study, in which an originally non- adhesive surface was conjugated with different concentrations of RGD peptides. When the average distance of peptides was calculated from the average surface concentration, it was found that a certain minimum concentration and, consequently, a certain maximum distance of RGD peptides is required to produce focal adhesions (Massia and Hubbell, 1991). In this work, the binding of RGD on randomly activated surface was conducted from solution, therefore meaning a rather random statistical distribution of peptide on the surface can be assumed, with actual RGD distance having some statistical distribution.

More recently, Arnold et al. (2004) prepared model surfaces on which the distance of RGD peptide motifs was precisely controlled within a very narrow distribution range. They showed that the optimal spacing of individual RGD molecules to ensure activity is in the range of 50–73 nm. Although surfaces with a larger distance between RGD motifs, 110 nm for example, still contained a lot of RGD, they exhibited the same cell response as the control surface with no RGD at all. When the RGD distance increased, the focal adhesions (thus the involvement of cell actin fi laments) dramatically decreased. These results clearly demonstrated that it is not only the presence, but also a spatial distribution in the tens of



nanometres scale, that is crucial for controlling cell- surface interactions. A proposed model of an ideal surface of polymer biomaterial is schematically depicted in Fig. 4.5 .

Research into biomimetic coatings also includes different approaches to surface modifi cations that are driven by the nature of the material and the purpose of application. One frequently studied strategy deals with the surface modifi cation or incorporation of hydroxyapatite structures onto polymer, ceramic or metal surfaces designed for bone and dental tissue engineering (Engel et al. , 2008). Besides molecular structure, the surface topography, roughness and the presence of surface features have been reported to infl uence the cell- surface interactions (Anselme et al. , 2010). One can derive that in living tissues, basal membranes are not fl at structures but contain hills and pores of micron- or submicron- size which interact with adjacent cells (Abrams et al. , 2000). In addition, ECM proteins form 3D networks featuring nanometre- scale structures. Topography- induced cell orientation is known as contact guidance. This phenomenon has been proven using various cell types on different surface features of metals, inorganic compounds and polymers (Kearns et al. , 2010, Bettinger et al. , 2009, Flemming et al. , 1999, Curtis and Wilkinson, 1997).

Other studies, also proceeding from principles occurring in the living body, are focused on the infl uence of applied mechanical stimuli, as this is a natural part of the bone remodelling process. The effect of mechanical loading on osteoblast proliferation, differentiation and other cell responses is under extensive investigation (El Haj et al. , 2005, Yang et al. , 2004, Yang et al. , 2002, Anselme, 2000). These studies also explore the key role played by the biomimetic coating that facilitates integrin- mediated cell adhesion; a prerequisite for the transfer of mechanical stimuli to the cytoskeleton.

4.4 Range of biomaterials

4.4.1 Polymers

Polymers are widely used as biomaterials. Taking advantage of their rational synthesis and easy processing, synthetic polymers allow for preparation of tailor- made materials with a range of desired chemical and mechanical properties.

4.5 Schematic drawing of biomaterial surface with biomimetic motifs.



Current trends focus largely on biodegradable polymers that can decompose into non- toxic products within the living body in a predictable way. The degradation of the polymer chain can proceed either through hydrolytic reactions, typical for aliphatic polyesters such as PLA, PGA or PCL, or by enzymatic degradation, for example in the case of materials cross- linked with peptide linkages or composed from poly(amino acid)s (Ulery et al. , 2011, Griffi th, 2000, Park, 1995, Škarda et al. , 1993, Pytela et al. , 1994, Sedlačík et al. , 2011).

Rational synthesis can afford a broad range of polymer structures, making it useful in achieving the bulk material properties required for particular applications. However, it is desirable to limit the modifi cation of biomimetic structures only to surfaces, or individual superfi cial layers of the material, in order to prevent deterioration of the bulk properties. Surface modifi cation or surface coating thus becomes an integral part of polymer biomaterial research. Typically, the coating serves two purposes: prevention of undesired interactions or surface fouling, and the introduction of specifi c biomimetic structures. Some general aspects of these processes have been discussed above; in the following section we shall focus on some specifi c problems particularly relevant to modifi cation of polymers.

Usually, direct immobilization of a bioactive structure onto solid hydrophobic surface would result in a loss or decrease of its activity, because of the limited accessibility caused by steric hindrance, change of conformation or denaturation in the case of biomacromolecules. As such, spatial separation of the bioactive structure from the solid surface through some fl exible spacer is often required. When designing a nonfouling surface which is to additionally carry biomimetic motifs, success can be achieved if some of the tethered polymer chains of the nonfouling layer carry suitable functional groups. Such functional groups can be the ultimate bioactive components, or can be a reactive group for subsequent functionalization.

The effect of the microenvironment on effi cient RGD enforcement was studied using self- assembly of alkanethiol- conjugates with PEO and a small fraction of an identical conjugate with the RGD peptide sequence. The surface character was altered by varying the number of ethylene oxide units in the polymer chain and by concentrating the conjugates carrying the peptide sequence. The attraction of surfaces was evaluated by assessing the number of adherent cells and their spreading (Houseman and Mrksich, 2001). The effect of binding conditions and surface character on the interaction of various cells type with bound RGD adhesion sequences and other bioactive peptide motifs has been the subject of many studies (Koo et al. , 2002, Hershel et al. , 2003, Shin et al. , 2003, Arnold et al. , 2004, Cavalcanti-Adam et al. , 2006, García, 2006, Kuhlman et al. , 2007, Bacakova et al. , 2007, Ma et al. , 2007, Petrie and García, 2009).

The role of a fl exible polymer chain, for example PEO, as a spacer in improving the specifi city of interactions of coupled biomacromolecules or peptides has been demonstrated in biosensors. Using PEO fl exible spacers, the effi ciency of interaction can be enhanced, as has been reported for various peptides and



immunogens (Ikeda and Nagasaki, 2011) improving the specifi city of biosensor interactions (Goddard and Hotchkiss, 2007). PEO can easily be functionalized, and, in addition, heterobifunctional PEOs of various molecular weights are also commercially available.

Strategies for bioactive molecule conjugation to polymers

For conjugation of bioactive compounds to fl exible polymer chains, organic reactions of thiols, aldehydes, carboxylic acids, alcohols and primary amines are used. The progress in organic chemistry and development of bi- and multi- functional reagents that can serve as spacers of different lengths have extended the strategies for coupling reactions (Hermanson, 2008). One of the most frequently used strategies for the coupling of a peptide is based on the reaction of its N -terminal group with an active ester at the end or side chain of a polymer.

When using a polymer with a carboxyl group, the active ester can be prepared in situ via a reaction with N -hydroxysuccinimide, as it is shown in Fig. 4.6 . A survey of PEO derivatives suitable for reaction with amino- groups is provided in the literature (Lutz and Börner, 2008, Roberts et al. , 2002, Veronese, 2001). The oriented coupling of a peptide through its terminal amino group can be complicated by the presence of other nucleophilic groups in the side chains, such as lysine or arginine, for example. Unless they are selectively protected, this can lead to a mixture of conjugates with different peptide orientations, meaning that they consequently exhibit different affi nities for interactions with cell receptors (Kinstler et al. , 2002).

4.6 Conjugation strategies for immobilization of bioactive compound onto the fl exible polymer chain such as PEO: (a) conjugation via active ester, (b) Michael addition, (c) Huisgen cycloaddition.



The Michael- type addition (Michael, 1887), based on the reaction of a thiol group of the peptide cysteine with the double bond of α , β -unsaturated compound, such as maleimide or a vinylsulfone group, is more specifi c than conjugation through amino groups. The addition of a thiol group to a double bond proceeds quickly under slightly acidic conditions and is an order of magnitude faster than the competitive addition of an amino group (Mather et al. , 2006). Furthermore, the cysteine residue is not found frequently on the protein surface, particularly in comparison to lysine, thus affording a higher specifi city of the coupling reaction. In addition, cysteine residue can simply be incorporated at the end of the chain during the peptide synthesis (Roberts et al. , 2002).

‘Click chemistry’ reactions developed during the last decade have become very powerful tools for the specifi c conjugation of biomimetic structures to various types of supports. The term ‘click chemistry’ refers to several highly selective reactions that effi ciently avoid side products, and are capable of producing a wide range of functional molecules. To be effi cient, the reaction must proceed quickly under mild conditions, and must be insensitive to oxygen and water. Typically, the reaction is stereospecifi c, producing high yield and limited (if any) by- products, removable by non- chromatographic processes (Hoyle and Bowman, 2010, Kolb et al. , 2001).

Among ‘click reactions’, the most popular is Huisgen Cu(I) catalyzed azide- alkyne cycloaddition, which proceeds in aqueous solution ( Fig. 4.6c ). Various click reaction settings have already been used for modifi cation of surfaces with bioactive motifs (Mackova et al. , 2011, Guan et al. , 2011, Lahann, 2009). Thiol- ene reactions, including the Michael thiol addition reaction, exhibit many attributes of click reactions and are routinely referred to as ‘thiol click reactions’.

In addition to covalent binding, a non- covalent immobilization of biomimetic motifs can be carried out by taking advantage of the strong interactions between some ligands and certain proteins. Low- molecular-weight biotin (vitamin H) and high- molecular-weight protein avidin, form one of the strongest protein- ligand complexes, with dissociation constant K d = 10 −15 M. Avidin, a tetramer protein, contains a specifi c binding domain for biotin in each peptide chain. Thanks to the tetravalency of avidin and the spatial arrangement of binding sites for biotin, it can be used as a bridge between biotinylated bioactive molecules (Black et al. , 1999, Cannizzaro et al. , 1998) or between a biotinylated polymer substrate and a biomimetic molecule (see Fig. 4.7 . for scheme). Biotin can be covalently linked through its valeric moiety to a fl exible spacer, for example PEO, with preserved capability to bind to avidin (Ren et al. , 2009, Chao et al. , 2008, Rosano et al. , 1999). In an analogous way, multi- layers of positively charged avidin and a polycationic polymer can be prepared, despite the repulsive forces of cations (Anzai et al. , 1999). Immobilization of avidin and its analogues, such as streptavidin and neutravidin, has been described and the effects of the immobilization of avidin activity towards biotin studied (Wolny et al. , 2010, Reznik et al. , 2001).



Surface patterning

Various techniques have been used in creating specifi c surface patterns of structural or chemical cues on the micron and submicron scale. With fl at model substrates, the surface patterning methods can be based on lithographic techniques. Rugged surfaces can be prepared by photo- lithography, using shielding lattice and UV radiation. Surface patterns can be duplicated from SiO 2 wafers, used in microelectronics, onto polymer biomaterial surfaces using stamps from polydimethylsiloxane (PDMS). This technique, called ‘soft lithography’, can be used for both structural patterning of surfaces and coating with specifi c biomolecules, using the desired solution as ink.

Micro- contact printing ( μ CP) and other lithographic techniques such as scanning probe lithography for biomolecular patterning, etching, plasma and ion- beam treatment have been reported (Schmidt and Healy, 2009, Engel et al. , 2008, Falconnet et al. , 2006, Curtis and Wilkinson, 2001, Flemming et al. , 1999). Scanning probe lithography methods can be used for surface patterning with functional domains, down to a range of 10 1 –10 2 nm (Kramer et al. , 2010, Hyun et al. , 2002). Development of surface morphology mimicking that found in nature has been approached by direct irradiation of materials using ultrafast laser pulses (Stratakis et al. , 2011). Simultaneously, surface topographical and chemical effects on cell adhesion have been also studied (Sardella et al. , 2008). However, many of these techniques are limited by the size and shape of the substrate being modifi ed, or by the nature of the substrate. Advanced lithographic techniques, capable of creating a pattern with a resolution on the scale of tens of nanometres, are very signifi cant tools for experimental model studies. However, such techniques require highly specialized and expensive technical equipment and they cannot yet be successfully applied to 3D scaffolds made of a biodegradable polymer material.

The term ‘nanocolloid lithography’ is used to describe surface patterning techniques based on self- assembling processes and the deposition of nanocolloids,

4.7 Surface immobilization of a bioactive peptide through the biotin–avidin interaction.



for example block copolymer micelles. These approaches use objects with dimensions of individual macromolecules, meaning they can reach a scale range below 100 nm, and include such methods as SAMs, block- copolymer lithography and colloidal lithography (Cox et al. , 1999, Zhang and Yang, 2010, Lohmüller et al. , 2011). Solutions of colloids or block copolymers can be used for the deposition of metal nanoparticles, the preparation of regular, ordered structures and/or as templates for further modifi cation. Combinations of different techniques have also been reported, such as colloidal lithography with plasma polymerization for nanopatterns of covalent grafting PEO (Singh et al. , 2011), or colloidal lithography with chemical vapour deposition (Valsesia et al. , 2006).

Block copolymers consist of covalently linked polymer chains, made of two different monomers. The blocks are usually immiscible and, because of their microphase separation, can form a variety of ordered structures either in bulk, on surfaces or in solutions, dependent on the present molecular parameters and conditions (Gohy, 2005, Abetz and Simon, 2005, Hadjichristidis and Pispas, 2006). Micellization of block copolymers, adsorption of micelles on solid surfaces and adsorption of copolymers applicable in tissue engineering are described in the literature (Emoto et al. , 1999, Riess, 2003, Albrecht et al. , 2006, Popelka et al. , 2007). Lateral phase segregation in ultra- thin fi lms of block copolymers results in ordered structures at interfaces, and is therefore also of great interest (Albrecht et al. , 2006, Darling, 2007).

The deposition of block copolymers, discussed above as a method for the immobilization of antifouling polymer chains, such as PEO, can also be used for surface patterning. The surface morphology of block copolymers of PEO with aliphatic polyesters, such as PLA or PCL, for the potential modifi cation of biomaterials, has been investigated in various studies with respect to molecular parameters and conditions of depositions (Fujiwara et al. , 2000, Huang et al. , 2008, Giacomelli and Borsali, 2006, Huang et al. , 2009, Hsu et al. , 2007). Moreover, the deposition of nanocolloids of amphiphilic block copolymers of PLA and PEO from selective solvents has been used for both structural and chemical surface patterning of PLA-based biomaterials, where a portion of the copolymer chains carried RGDS adhesion peptides.

Surface topography of PLA modifi ed with PLA- b -PEO copolymers, as seen in atomic force microscopy (AFM), differs with changes in PEO block length and can be attributed to the ability of PEO to crystallize. The accessibility of functional groups, incorporated at the end of some PEO chains for further interactions with large objects such as proteins and cells, has been demonstrated. This proves that through deposition of block copolymers, the PLA surface can expose biomimetic cell- adhesion structures, while the overall character of the surface remains hydrophilic and repellent for non- specifi c protein adsorption (Tresohlava et al. , 2010a). In addition, a specifi c surface pattern of biomimetic peptide motifs can be created via controlled deposition of mixtures of functionalized and non- functionalized PLA-PEO nanocolloids.



The specifi c surface pattern was proved by peptide labelling and AFM detection, and has been shown to be favourable for the initial attachment of bone cells (Tresohlava et al. 2010b). Using this technique, multiple bioactive molecules can be step- by-step conjugated through rich PEGylation chemistry, and the cluster distance and concentration can be optimized by the ratio of functionalized and non- functionalized copolymers, their molecular parameters and coating condition. Adsorption of block copolymers and their nanocolloids can be accomplished simply by immersion of the material into nanocolloid solution. This coating can therefore be easily applied to 3D-scaffolds with complex shape and geometry.

4.4.2 Metallic biomaterials

Metals have traditionally been used as biomaterials, mainly because of their excellent mechanical properties. However, they exhibit unmatched electrical and thermal conductivity. The types and uses of metallic biomaterials have been thoroughly reviewed in a variety of studies (Park, 1995). However, the application of metallic biomaterials must take into account not only their mechanical properties, but also metal corrosion processes. Tissue fl uids, in addition to water, contain proteins, various ions, dissolved oxygen and reactive oxygen species, thus forming a corrosive environment for metal implants. The corrosion process can be accelerated by the mechanical stress applied to implants. Implant corrosion can result in local pain and swelling in the region of the implant, along with such additional effects as fl aking or cracking of the implant (Park and Kim, 2003). Some metallic materials, magnesium alloys, for example, may undergo complete degradation in this environment, which could be an advantageous feature in some applications, such as degradable bone implants or cardiovascular stents (Keim et al. , 2010).

For successful performance, the interaction of a metallic biomaterial with the surrounding tissue is also important. In dentistry and orthopaedic bone implants, good fi xation of the implant to bone, osseointegration, can prevent loosening of the implant, and, as such, is a key functional requirement (Reyes et al. , 2007). Surface fi nishing and the modifi cation/coating of metal implants are therefore widely studied. Surface treatments with chemicals such as acids, alkalis and hydrogen peroxide have been widely conducted as simple and effective methods offering nanostructured surfaces for titanium implants, for example, without compromising biocompatibility. Physicochemical properties, roughness and topography of thus obtained surfaces are investigated with respect to their capacity to enhance adhesion and proliferation of osteogenic cells, deposition of hydroxyapatite, and expression of genes and proteins (Liu et al. , 2010a; Kubies et al. , 2011).

Coatings with synthetic hydroxyapatite (HA) are used to provide osteoconductive properties to the surfaces of metal implants, as HA has some similarity to the mineral phase of bone and tooth material. HA and other calcium phosphate



surfaces are used for stabilization of the implant in the surrounding tissue, via direct binding and the improvement of bone apposition (Tsang et al. , 2011, Palmquist et al. , 2010, Habibovic et al. , 2005). Several methods for making HA coatings on metals have been reported, such as plasma spraying (Latka et al. , 2010), electrochemical deposition (Zhao et al. , 2009, Yang et al. , 2010) and investment casting (Escobedo et al. , 2006).

The direct growth of nanoapatite coating on Ti surfaces from aqueous solutions, and the capability of thus prepared bioactive surfaces has also been reported (Li, 2003). Moreover, organic- inorganic nanocomposite coatings based on multi- layers of polymer and Ca-P are also under investigation (Schade et al. , 2010).

Coatings for metal implants can be used also for the local delivery of drugs which could reduce side effects or facilitate integration with the surrounding tissue. In this respect, Ca-P coatings on Ti surfaces have been adapted to deliver bone morphogenetic proteins (BMPs) and other growth factors in a localized fashion, thus enhancing the effi ciency of osteointegration (Majid et al. , 2011, Verron et al. , 2010). Covalently bound polymer coatings composed of aliphatic polyesters were developed for metallic vascular stents, facilitating the controlled release of drugs that attenuate smooth muscle cell proliferation in stented vascular walls and additionally prevent restenosis (Rypáček et al. , 2005, 2007).

4.4.3 Ceramics

Ceramic materials are mainly used in the regeneration of hard tissue, such as bone tissue engineering or dentistry, because of their natural and mechanical properties. Similar to other materials used in medical applications, ceramics were fi rst used as bioinert materials, having no reaction with living tissues. This approach has changed signifi cantly during recent decades and the new generation of bioactive ceramic implants was designed to react with the surrounding environment and actively infl uence new tissue formation. Bioactive ceramics, bioceramics, can be resorbable or non- resorbable. These are prepared from aluminium or zirconium oxide, carbon, calcium phosphates and silica compounds (glasses) among others. State- of-the- art designs are focused on porous ceramics acting as scaffolds for cells and introducing active molecules for improved self- regeneration of tissues (Billotte, 2003, Vallet-Regí and Ruiz-Hernández, 2011, Rahaman et al. , 2011).

Taking advantage of their chemical similarity to the mineral content of mammalian bones and teeth, calcium phosphates are widely used in bone implants. Calcium phosphates are non- toxic and bioactive, resulting in an intimate bond between the implant and bone, that is osteointegration. They have been proven to facilitate osteoblast adhesion and proliferation, and to be osteoconductive. However, because of their brittleness and poor fatigue resistance, calcium phosphates in biomedical applications are primarily used as coatings (Dorozhkin, 2010, Vallet-Regí and Ruiz-Hernández, 2011). Frequently used substrates for coating are titanium alloys and commercial implants coated with hydroxyapatite



(HA), whereas other calcium phosphates are produced by plasma spraying. In addition, other techniques, such as physical or chemical vapour deposition, pulsed laser deposition and sol- gel based dip coating have been reported. These processes allow control over the coating thickness and crystallinity of phases. Moreover, the coatings can be loaded with drugs for local delivery at the site of the implant (Laurencin et al. , 2001, Liu et al. , 2010b, Verron et al. , 2010, Vallet-Regí and Ruiz-Hernández, 2011). Bioactive glasses, as well as calcium phosphates, can also be doped or modifi ed with elements such as Cu, Zn, Sr or Mg, that are known to be benefi cial for bone growth (Rabadjieva et al. , 2011, Rahaman et al. , 2011).

The composites of various polymers with calcium phosphates, bioactive glasses and HA have also been extensively studied (Pattanayak et al. , 2011). The addition of biodegradable polymers can improve the degradability of ceramics and alter such mechanical properties as tensile strength and brittleness, while maintaining good biocompatibility (Li and Kawashita, 2011). As an example, composites with aliphatic polyesters, such as polylactide, polyglycolide (Dinarvand et al. , 2011, Bae et al. , 2011, Zhao et al. , 2011) or polyacrylic acid have been studied (El-Bahy et al. , 2011).

From another research viewpoint, the grain size of ceramics has been proven to be a signifi cant parameter for the bioactivity of bioceramics. After three and four weeks in cultures of nanophase ceramics containing grains of tens of nanometres, in comparison with conventional ceramics with grains of hundreds of nanometres up to micrometres, better osteoblast adhesion and proliferation was observed, in addition to a signifi cant enhancement in the synthesis of alkaline phosphatase and deposition of calcium- containing minerals by osteoblasts (Webster et al. , 2000). The effect of nanostructured surfaces on bone cells has also been reported for aluminium (Webster et al. , 2001) and titanium surfaces (Gutwein and Webster, 2004). In a comparative study of osteoblast adhesion onto surfaces of compacted (pressed particle) nanoamorphous calcium phosphate, nanocrystalline HA and conventional HA, either with or without functionalization using a RGD-model peptide, the effect of ceramic nanosize was clearly demonstrated. The osteoblast adhesion was shown to be signifi cantly higher on the nanoamorphous calcium phosphate compact, even without the RGD-peptide, in comparison with a conventional HA carrying RGD sequence. This indicates the signifi cant effect of nanostructure and crystallinity on osteoblast adhesion (Balasundaram et al. , 2006).

4.5 Evaluation of coating effi ciency

Several methods can be used to monitor surface modifi cation and dynamic surface processes, as well as to evaluate changes in surface properties quantitatively, either in absolute or, at least, in relative values. The most frequently used methods are reviewed below. Usually, no individual method should be used alone; the data obtained via several different methods should be combined in order to provide a deeper understanding.



4.5.1 Surface properties: wettability, contact angle, surface energy

The thermodynamic characterization of biomaterial surfaces and their properties at solid–liquid interfaces has become an essential step in the development of biomaterials for tissue engineering. Measurement of the contact angle (CA) of tested liquid, in most cases water, on a solid surface is a direct method for evaluating surface energy that cannot be obtained by other physical methods. CA measurement provides information about the outermost few Ångstroms of the solid (3–20 Å). The obtained data can be used for characterization of the solid surface itself, or for characterization of processes such as wetting, liquid penetration through porous solids and adsorption. Nevertheless, the obtained information is non- specifi c; it does not indicate the chemical composition of the tested surface, and it ranks among common methods for thermodynamic characterization of biomaterial surfaces (Davies, 1996).

When a liquid comes into contact with a surface, the interface at the solid–liquid– vapour boundary is characterized by a contact angle θ at the three phase junction, as is depicted in Fig. 4.8 . This contact angle results from a balance of lateral forces between the interfacial tensions at the solid– vapour ( γ sv ), liquid–vapour ( γ lv ) and solid– liquid ( γ sl ) interfaces, and is described by Young’s equation:

γ sv − γ sl = γ lv cos θ

This equation is valid for an ideal surface (a smooth, uniform and homogenous, rigid isotropic solid surface, not interacting with used liquid). However, polymer and other material surfaces rarely fulfi l such requirements, instead exhibiting contact angle hysteresis. Although the liquid front advancing across a new surface may exhibit a larger contact angle (the advancing contact angle θ A ), the same liquid receding from an already- wetted surface may have a much smaller contact angle (the receding contact angle θ R ) (Davies, 1996, Myers, 1999, Lam et al. , 2002, Gao and McCarthy, 2006).

A number of techniques are used to measure contact angles. The selection of the method depends mainly on the sample geometry (compare fl at surfaces, fi bres or hydrogels), surface heterogeneity (low/high roughness, for example) and on the information required (Davies, 1996, Myers, 1999). Although a static CA

4.8 Drop of liquid on a solid surface showing the contact angleθ.



measurement based on a drop goniometry measurement can provide fast characterization, it is believed that more accurate data can be obtained by dynamic measurements providing θ A and θ R values, thus avoiding any data misinterpretation produced by local surface properties (Kwok and Neumann, 1999). When analyzing data, it is necessary to consider the surface characteristics which can lead to hysteresis, such different chemical compositions (local variation of γ s ) or surface roughness.

The most frequently used methods are ‘sessile drop’ and the Wilhelmy plate technique. The latter can also be used to determine liquid surface tension, adsorption of protein on a surface or for monitoring of dynamic processes occurring at the solid– liquid interface. The ‘sessile drop’ method is particularly convenient for quick characterization of fl at surfaces, and only one side of a plane substrate is required for the measurement. However, several factors have to be considered for reproducibility, such as surface roughness and different local compositions, as well as interaction with the tested liquid, which can lead to a signifi cant deviation of obtained CA values. The Wilhelmy plate technique is based on weighing performed with high accuracy. This method is most suitable for surfaces with higher roughness and heterogeneity. However, the surface tension of the liquid used and the perimeter of the sample have to be known; the plane sample must have the same composition on both sides and the layer must be stable. If even a minor dissolution occurs, it can affect the CA calculation. This method can also be used for the characterization of porous materials, powders and fi bres by using special holders (Davies, 1996).

For biomaterial applications, water is typically used as the wetting liquid. When surfaces demonstrate a high affi nity to water and thus low contact angles are observed, we consider them to be hydrophilic, high- energy surfaces. Surfaces exhibiting a low affi nity to water, characterized by high contact angles, are considered hydrophobic, low- energy surfaces. Water contact angles of various polymers are available in literature.

4.5.2 Coating quantifi cation

Surface coatings may need to be characterized in terms of coating thickness, homogeneity and stability under both ambient and physiological conditions, as well as in terms of biomolecular interactions, such as extent and dynamics of protein adsorption. For this purpose, SPR, ellipsometry, infrared spectroscopy and the quartz crystal microbalance technique (QCM) among others, have been widely used and combined, depending on the information and resolution required.

SPR is an optical, highly sensitive method that is used for the characterization of processes occurring at or near interfaces. It is based on total internal refl ection from an interface between a material with a higher refractive index (e.g. a metal fi lm on glass) and a medium with a lower refractive index (such as a liquid). The generated evanescent wave extends from the interface into the medium, and its



analysis provides information about the adsorbed layer and its thickness. The measurement is performed in situ in real time and can therefore provide dynamic kinetic information about the studied process. This technique is particularly useful for monitoring the changes in the composition of the surface layer, to follow factors including the adsorption of proteins, antibody– antigen interactions and binding kinetics, (Slavik et al. , 2002, Rodriguez-Emmenegger et al. , 2009, 2011a).

Ellipsometry is a versatile optical technique based on measuring the change of polarization of light upon refl ection. The nature of the polarization change is determined by sample properties, such as thickness, complex refractive index or dielectric function tensor. Besides these properties, ellipsometry can provide information about surface roughness, composition and optical anisotropy of thin fi lms. This method is extremely sensitive, enabling information about layer thickness ranging from 10 –1 to 10 2 nm. Ellipsometry can be performed on various substrates, including silicon, gold, titanium and stainless steel, in addition to other substrates relevant for biomedical application. Thin fi lm polymers, self- assembled layers and/or Langmuir-Blodgett fi lms can also be characterized (Tompkins and Irene, 2005). A detailed description of the physical methods of coating quantifi cation cannot be given in the breadth of this chapter, and interested readers should refer to the following specialist literature dealing with each method, and books and reviews focused on surface characterization techniques (Davies, 1996, Vadgama, 2005, Marx, 2003, Merrett et al. , 2002, Schasfoort and Tudos, 2008).

4.5.3 Coating morphology and topography

In addition to average coating properties, such as coating composition, thickness and wettability, detail morphology, topography and viscoelastic properties of the surface are of great importance. The following section briefl y describes some microscopic techniques that are typically used for imaging coating surfaces.

Scanning electron microscopy, SEM, is commonly used when studying surface morphology, the 3D structure of hydrogels, fracture areas, the inner structure of different materials and the cellular response to biomaterials. SEM is performed in a vacuum environment, where a precisely focused electron beam interacts with a sample while scanning a selected area point by point and row by row. During this process, electrons leaving the surface are collected, providing a micrograph. In high- vacuum SEM, the sample must be conductive or covered with a conductive metal layer, typically of platinum or gold, to avoid charging with incident electrons. Low vacuum SEM (LVSEM) can be used for specimens containing frozen water, and environmental SEM (ESEM) can even be used for samples containing liquid water. These operational modes enable characterization of various biomaterials and biological structures, almost in their native state. Moreover, when the electron beam hits the sample surface various signals are produced and can be detected, allowing the use of special detectors to provide additional micrographs, complementary to SEM surface morphology. Besides



secondary electron detection, the detection of backscattered electrons provides a material contrast for samples composed of two or more materials. Drawing on similar methodology, another technique increasing in popularity is energy dispersive X-ray analysis, EDX, based on the detection of X-rays and enabling element analysis (Ziegler, 2005, Reimer, 1998).

AFM is a powerful technique providing high- resolution surface topography, both under ambient conditions and in liquids, over a wide range of temperatures. AFM belongs to scanning probe microscope (SPM) techniques that apply a probe to scan a sample surface, and map its properties based on the particular probe used. In AFM, a sample characterization is based on the mapping of forces as the sharp tip at the end of a cantilever scans the surface. The principal of mapping forces is based on cantilever defl ection during scanning, depending on attractive or repulsive forces acting on the approaching tip. The cantilever defl ection is detected by sensitive laser sensors, as schematically depicted in Fig. 4.9 , and the signal can be transferred to a topography image via the controller computer. The surface forces depend on the tip– sample distance (the shape of the curve is depicted

4.9 Scheme of principle of atomic force microscopy (AFM) (a) with a typical shape of force curve based on tip– sample distance (b).



in Fig. 4.9 ) and are relevant for any sample surface, regardless of its conductivity. Physical forces such as van der Waals interactions, capillary or short- range (Pauli) forces are employed in cantilever bending. Cantilever with AFM tips are usually prepared from silicon or silicon nitride, by lithography, for example, with the construction dependent on the applied mode. Currently, cantilevers with a wide range of stiffness (spring constant k =10 −2 –10 1 Nm −1 ) and resonant frequencies are available on the market, as well as tips with different geometries and dimensions. However, special tips, such as sharpened tips, coated tips for the additional introduction of various chemicals, or custom- made tips have become more popular for state- of-the- art surface analysis (Schönherr and Vancso, 2010).

Such measurements can be realized in a number of modes, depending on the application. The key possible modes are static modes (also called contact modes) and dynamic modes. The modes differ in the forces applied at the tip, thus also in tip– sample interactions and signals for feedback. Generally, dynamic modes are more suitable for soft materials, for example polymer surfaces under ambient conditions. Contact mode, meanwhile, can often be used in aqueous solutions without hindrance or negative effects on the sample surface, tip or course of measurement (Schönherr and Vancso, 2010, Ellis et al. , 2006).

AFM provides true 3D topography images of a sample surface with lateral resolution 1 nm and vertical resolution less than 0.1 nm. Besides characterization of surface topography, AFM is widely used for visualization of DNA, RNA or other biological molecules or biomaterial surfaces (Ellis et al. , 2006). Furthermore, adsorption of biomacromolecules onto the substrate surface or deposition of block copolymer micelles can be observed in situ (Almeida et al. , 2004, Hamley et al. , 2004). The developments in AFM instrumentation have led to a growing interest in advanced applications, such as single molecule recognition based on tip– sample interaction measurements in real time, or determination of local mechanical or elastic properties (Ebner et al. , 2009).

AFM can be combined with other optical methods such as fl uorescence microscopy, thus allowing different information to be obtained, resulting in growing interest in AFM for studies in the fi eld of biological membranes, cell surfaces and processes. However, one disadvantage of AFM is the size of the image; a maximum scanning area is usually limited to 100 μm, with the height only to the order of micrometres. For this purpose, a combination of AFM with other microscopic techniques, such as SEM or optical microscopy, is usually required.

4.6 Conclusion

The performance of many medical devices, for example tissue- engineering scaffolds, depends on both the bulk properties of the biomaterials from which they are made, and the properties and composition of their surfaces. The coating applied to the surface of biomaterial can signifi cantly improve the properties of



the device as it helps to control the interactions between the material and the environment of living tissues. Suitably designed polymer coatings may affect the surface properties of the device to such extent that the chemical nature of the underlying material becomes irrelevant and the biological response to the device is governed merely by the properties of extremely thin surface layer. This allows for independent engineering of the bulk material and its surface characteristics. For example, the surface characteristics of the material can vary to a large extent while the mechanical properties of the bulk can be almost completely preserved.

In many applications it is desirable that surfaces exhibit certain features that mimic those of native tissues or facilitate certain interactions with tissue components. Such biomimetic features can conveniently be introduced through biomimetic coating. For example, coating with polymers carrying peptide sequences derived from the proteins of extracellular matrix can induce specifi c cell adhesion and even provide signals for cell differentiation. In order to take full advantage of such targeted modifi cation, the non- specifi c interactions, for example a spontaneous adsorption of proteins from the media, which would otherwise mask the biomimetic motifs of the coating, fi rst have to be minimized. Therefore, the problems of making surfaces nonfouling and, at the same time, selectively interactive through their biomimetic motifs, are always intertwined in the development biomaterials for tissue engineering.

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Biomimetic Biomaterials || Biomimetic coatings for biomaterial surfaces