biomimetic materials in tissue engineering

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ISSN:1369 7021 © Elsevier Ltd 2010 JAN–FEB 2010 | VOLUME 13 | NUMBER 1-2 14 Biomimetic materials in tissue engineering Cells receive numerous signals from their immediate microenvironment, the extracellular matrix (ECM) 1 . Within a biomechanical context provided by this elastic milieu 2 , cells adhere by receptor-mediated interactions with ECM components such as fibronectin and laminin (among many others, as reviewed elsewhere), mediated by specialized adhesion receptors such as integrins and others 3 . These receptors transmit stress from the ECM, through the membrane (the receptors are transmembrane proteins), to the cytoskeleton within the cell in a dynamic and concerted manner 4 . The adhesion receptors do much more than transmit stress, however; in particular within clusters of adhesion receptors in the membrane, biochemical signal transduction takes place through kinase activation and other mechanisms 3,4 . In addition to adhesion proteins, the ECM also sequesters and presents a number of morphoregulatory molecules including growth factors, which control processes of cell division, differentiation, and multicellular morphogenesis 2,5 . The growth factors bind ECM components such as heparan sulfate proteoglycans 6 and fibronectin 7 , among others, and reside there until released by enzymatic processes or dissociation. These Biomaterial matrices are being developed that mimic the key characteristics of the extracellular matrix, including presenting adhesion sites and displaying growth factors in the context of a viscoelastic hydrogel. This review focuses on two classes of materials: those that are derived from naturally occurring molecules and those that recapitulate key motifs of biomolecules within biologically active synthetic materials. For biologically derived materials, methods are being sought to gain molecular-level control over biological characteristics and biomechanics. For synthetic, biomimetic materials, chemical schemes are being developed to enable in situ cross-linking and protease- dependent degradation and release of incorporated growth factors. These materials will open new doors to biosurgical therapeutics in tissue engineering and regenerative medicine. Jennifer Patterson 1,+ , Mikaël M. Martino 1,+ , and Jeffrey A. Hubbell 1,2 * 1 Institute of Bioengineering, Ecole Polytechnique Fédérale de Lausanne (EPFL), Lausanne, Switzerland 2 Institute of Chemical Sciences and Engineering, Ecole Polytechnique Fédérale de Lausanne (EPFL), Lausanne, Switzerland *E-mail [email protected] + These authors contributed equally.

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ISSN:1369 7021 © Elsevier Ltd 2010JAN–FEB 2010 | VOLUME 13 | NUMBER 1-214

Biomimetic materials in tissue engineering

Cells receive numerous signals from their immediate

microenvironment, the extracellular matrix (ECM)1. Within a

biomechanical context provided by this elastic milieu2, cells

adhere by receptor-mediated interactions with ECM components

such as fibronectin and laminin (among many others, as reviewed

elsewhere), mediated by specialized adhesion receptors such as

integrins and others3. These receptors transmit stress from the

ECM, through the membrane (the receptors are transmembrane

proteins), to the cytoskeleton within the cell in a dynamic and

concerted manner4. The adhesion receptors do much more

than transmit stress, however; in particular within clusters

of adhesion receptors in the membrane, biochemical signal

transduction takes place through kinase activation and other

mechanisms3,4. In addition to adhesion proteins, the ECM

also sequesters and presents a number of morphoregulatory

molecules including growth factors, which control processes of

cell division, differentiation, and multicellular morphogenesis2,5.

The growth factors bind ECM components such as heparan sulfate

proteoglycans6 and fibronectin7, among others, and reside there

until released by enzymatic processes or dissociation. These

Biomaterial matrices are being developed that mimic the key characteristics of the extracellular matrix, including presenting adhesion sites and displaying growth factors in the context of a viscoelastic hydrogel. This review focuses on two classes of materials: those that are derived from naturally occurring molecules and those that recapitulate key motifs of biomolecules within biologically active synthetic materials. For biologically derived materials, methods are being sought to gain molecular-level control over biological characteristics and biomechanics. For synthetic, biomimetic materials, chemical schemes are being developed to enable in situ cross-linking and protease-dependent degradation and release of incorporated growth factors. These materials will open new doors to biosurgical therapeutics in tissue engineering and regenerative medicine.

Jennifer Patterson1,+, Mikaël M. Martino1,+, and Jeffrey A. Hubbell1,2*1Institute of Bioengineering, Ecole Polytechnique Fédérale de Lausanne (EPFL), Lausanne, Switzerland 2Institute of Chemical Sciences and Engineering, Ecole Polytechnique Fédérale de Lausanne (EPFL), Lausanne, Switzerland

*E-mail [email protected]+These authors contributed equally.

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factors, when released and sometimes also when matrix-bound8,

bind to cell-surface receptors and trigger signaling, principally

through kinase activation. Thus, the ECM serves as a reservoir of

signaling molecules, both adhesion molecules and growth factors,

to instruct cell decision processes.

These characteristics of the ECM guide the design of biomaterials

for use in tissue engineering and regenerative medicine, with the

overall goal of mimicking these key features: the presentation of

adhesion molecules and the sequestration, display and release of

growth factors. These features may be accompanied by aspects of

formation and removal of the biomaterial ECM analogs. Namely, the

materials should be formable in situ, in the presence of cells without

damaging them and biomolecules without altering them. Further,

cells should be able to exert their enzymatic processes to remodel

and eventually replace the matrices, principally through matrix

metalloproteinases (MMPs)9, as they do with the natural ECM.

Thus, we organize this review from the perspective of biomimicry.

Having above begun, by very briefly outlining some of the biological

goals of biomimetic material design in tissue engineering, below we

proceed to review some of the biologically-derived materials that form

our design basis for biomimicry, following which we move to chemical

approaches that are being employed to actually mimic these materials.

Thus we proceed from biological, through biosynthetic, to completely

synthetic material systems.

Biologically derived and biologically produced materials CollagenCollagen type I, the most abundant protein in mammals, has a triple-

helical structure made of three polypeptide chains containing repeating

Gly-X-Y triplets in which the X and Y positions are frequently occupied

by proline and 4-hydroxyproline, respectively10, the latter of which

is very important for intermolecular hydrogen bonding. Collagen can

be readily purified from animal tissues, such as skin and tendon, and

from discarded human tissues, such as placenta. This ECM protein

can be reconstituted into a fibrillar matrix (Fig. 1a) by increasing the

Fig. 1 Scanning electron micrographs of biologically-derived and biologically-produced materials. (a) Fibrillar structure of collagen type I hydrogel. Reprinted with adaptation with permission from94. © 2007 Mary Ann Liebert, Inc. (b) Fibrin network. Reprinted with adaptation with permission from94. © 2007 Mary Ann Liebert, Inc. (c) Cross-linked adipic acid dihydrazide-modified hyaluronic acid hydrogel. Reprinted with adaptation with permission from95. © 2005 Elsevier. (d) Self assembling peptide RADA16-II hydrogel. Reprinted with adaptation with permission from96. © 2008 Wiley Periodicals, Inc. Scale bar = 500 nm.

(b)(a)

(c) (d)

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pH and temperature of a precursor solution. Widely used, collagen

type I is frequently treated with proteases to remove small nonhelical

telopeptides that are present at the ends of the triple-helical domain

and that contribute to most of the protein’s cross-species immunogenic

character. Nevertheless, for clinical application of this material,

concerns of immunogenicity and disease transmission remain11. To

avoid these risks, methods for the recombinant expression of collagen

have been developed in several eukaryotic expression systems12, and

recombinant human collagen types I and III are commercially available.

Because in vivo application of collagen gels is limited by a deficiency

in mechanical strength, several procedures have been described in

order to form collagen matrices that possess sufficient mechanical

proprieties to at least partially resist cell-induced contraction. For

example, chemical glycation can adjust, in a rather controllable manner,

the elastic character of collagen gels13. Collagen hydrogels can be

enzymatically cross-linked14 or compressed to relatively high density15.

Heat and chemical treatments have also been developed to produce

cross-linked collagen sponges for bone and cartilage repair16, although

these materials become so dense that they are no longer really

hydrogels.

Interestingly, the dipole moment of fibrillar collagen gives rise to

the possibility to obtain permanent microscopic and even macroscopic

alignment of fibrillar collagen matrices. For instance, collagen fibril

alignment under a strong magnetic field has been demonstrated to

impart special characteristics for inducing directed cell migration, such as

neurites which grow preferentially in the direction of fibril alignment17.

Fig. 2 Improvement of biologically-derived and biologically-produced materials with biomolecules. (a) Fibrin matrices can be functionalized with recombinant ECM fragments or growth factors. (i-v) The factor XIIIa enzymatic substrate sequence NQEQVSPL is fused to bioactive peptides or recombinant proteins, allowing these molecules to be covalently incorporated during the fibrin’s natural polymerization process via factor XIIIa-induced transglutamination31,97. The molecules are engineered in a manner that depends on their biological function and local cell-induced proteolysis. (ii) For example, growth factor release by cell-demand can be accelerated by incorporating a protease sensitive sequence after the factor XIIIa enzymatic substrate sequence69. (iii) Because ECM peptides or fragments are mostly active when they are bound to the matrix, the molecule is designed without a protease sensitive sequence (PHSRN and RGD are the critical sequences from fibronectin for engagement of integrins)33. (iv) Since molecules such as heparin naturally bind certain growth factors, systems combining this interaction and transglutamination have been developed. For example, a heparin-binding peptide can be fused to the factor XIIIa enzymatic substrate and cross-linked into the matrix with affinity-bound growth factor. A complex between the three molecules is formed, and the growth factor release depends mostly on its affinity for heparin. (v) Fibrin matrices can also be functionalized noncovalently with growth factors fused to a fibrin binding sequence such as Gly-Pro-Arg-Pro (GPRP)29,30. (b) Hyaluronic acid derivatives can be functionalized with biomolecules such as protease sensitive sequences or ECM fragments. For example, a recombinant ECM fragment engineered with a single free cysteine is first reacted with divinylsulfone for further cross-linking into thiol modified hyaluronic acid hydrogel by Michael-type addition41. (c) Similarly to fibrin matrix, collagen matrix can be functionalized with a growth factor fused to a collagen binding sequence derived, for example, from fibronectin, von Willebrand factor (vWF) or collagenase. The growth factor release depends on the binding sequence affinity for collagen19-22. (d) Hydrogels based on self assembling peptides can be functionalized with biomolecules such as protease sensitive sequences or ECM fragments fused to the self assembling building block peptide (to both or to one extremity), as long as it does not disturb hydrogel self-assembly51-53,57.

(b) Hyaluronic acid

(a) Fibrin

(c) Collagen

(d) Self-assembling peptides

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While collagens bind several cellular receptors that modulate cells’

behavior18, collagen gel bioactivity can be increased with growth

factors that have been engineered to bind the matrix in a noncovalent

manner. Collagen binding sequences derived from collagenase, von

Willebrand factor or fibronectin have been recombinantly fused

to growth factors in order to delay their release from collagen

scaffolds19-22 (Fig. 2c). For example, injection of collagen binding

vascular endothelial growth factor (VEGF) is able to improve cardiac

function after an acute myocardial infarction23.

FibrinClinically available from autologous sources and from cryoprecipitated

pooled human blood plasma, fibrin is a specialized protein network that

is formed principally in spontaneous tissue repair. The fibrin matrix

forms spontaneously by polymerization of fibrinogen, a circulating

glycoprotein homodimer of a heterotrimer, in the presence of thrombin

protease. Thrombin cleaves the so-called fibrinopeptides on fibrinogen

that prevent physicochemical self-assembly or polymerization of the

molecule. The resulting network is chemically cross-linked by the

blood transglutaminase factor XIIIa24, and its complex fibril structure

(Fig. 1b) and cross-linked character depend upon the details of its

formation25-27. Although fibrin is not an ECM in the usual sense,

since cells in the local environment do not produce it, the material is

nevertheless a crucial member of the body’s repertoire of matrices and

serves the role of a provisional matrix, being remodeled and replaced

with ECM molecules.

In contrast to fibrillar collagen matrices where cell migration

proceeds both through mechanisms that are dependent and

independent of proteolytic degradation, cell migration in fibrin is

almost exclusively dependent upon cell-associated proteolytic activity

(e.g., from plasmin and MMPs24). This distinction in cellular behavior

from fibrillar collagen probably results from the smaller mesh size of

the fibrin matrices and the stronger fibril–fibril interactions, owing to

the nature of network formation and covalent stabilization.

Several proteins are naturally incorporated into fibrin matrix

during coagulation, such as fibronectin and alpha-2-plasmin inhibitor,

which are covalently cross-linked into the matrix by factor XIIIa24.

Other biomolecules, such as fibroblast growth factor-2 (FGF-2) bind

noncovalently to fibrin and are able to provide growth factor-specific

bioactivity, such as potentiating of endothelial cell proliferation,

even when bound28. Nevertheless, fibrin matrices remain naturally

poorly active for most cell types, leading to their functionalization

with ECM peptides, recombinant ECM protein domains or growth

factors. Similarly to collagen matrices as described above, engineered

growth factors or ECM fragments can be incorporated into fibrin

non-covalently by fusing a fibrin-binding domain to the recombinant

protein29. For example, the native knob:pocket interactions during

fibrin assembly can be exploited to delay the release of recombinant

proteins30. However, to have even better incorporation and controlled

release, biomolecules can be covalently incorporated. In a powerful

approach, the factor XIIIa enzymatic substrate sequence of alpha-2-

plasmin inhibitor, Asn-Gln-Glu-Gln-Val-Ser-Pro-Leu (NQEQVSPL), can

be fused to bioactive peptides or recombinant proteins, allowing these

molecules to be covalently incorporated during the fibrin’s natural

polymerization process via factor XIIIa-induced cross-linking31. For

example, VEGF, fibronectin fragments and parathyroid hormone 1-34

(PTH1-34) have been engineered and covalently incorporated to be

released in a manner that depends on local cell-induced proteolysis32-

34 (Fig. 2a). In the case of the PTH1-34 fusion, the hormone is active

only when released: modification of its N-terminus with the factor XIIIa

substrate fusion inhibits the activity of the hormone and this activity

is regained after cell-induced proteolysis between the fusion and the

PTH1-34 domain. This engineered peptide in fibrin is currently in human

clinical evaluation for local bone repair.

GlycosaminoglycansThe structural proteins of the ECM are augmented in their

biomechanical and biochemical functions by long unbranched

polysaccharides, the glycosaminoglycans. In most cases, these are

components of proteoglycans of the ECM, except in the case of

hyaluronic acid (also called hyaluronan), which is not covalently

attached to a protein core and is entangled within the extracellular

space. These strongly anionic polymers absorb water, which provides

compressive strength to the ECM, while the glycosaminoglycans also

directly affect tissue organization via interactions with cell-surface

receptors35.

Hyaluronic acid can be isolated from animal tissue, such as

the rooster comb, and can be biotechnologically produced using

Streptococcus bacterium. Because the material absorbs enormous

amounts of water at equilibrium, it forms a hydrogel that is non-

fibrillar (Fig. 1c), and owing to entanglement associated with its high

molecular weight, up to several million Daltons, the gel dissolves only

very slowly.

Several chemical hyaluronic acid derivatives have been prepared

and by controlling the functional group (e.g., pendant hydrophobic

groups), the type of covalent bond (e.g., stable or hydrolytically-

sensitive), and the cross-linking density, it is possible to create a wide

range of physically diverse materials. A number of modifications of the

carboxyl and hydroxyl groups of hyaluronic acid have been developed,

to crosslink the material into an elastic gel that resists dissolution or

rendering the polymer controllably more hydrophobic and thus less

soluble36.

Hyaluronic acid hydrogels have been used for various applications,

including keratinocyte transfer for dermal wound healing and chondrocyte

transplantation for cartilage repair37,38. While hyaluronic acid interacts

with at least three cell surface receptors (CD44, RHAMM, and ICAM-1)39,

its biological activity can be significantly augmented by the incorporation

of other functional biomolecules. For example, hyaluronic acid gels have

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been functionalized with peptides or protein fragments derived from

fibronectin to improve fibroblast proliferation and wound healing40,41

(Fig. 2b). Similarly, cellular infiltration into the hydrogel can be

improved by using MMP-sensitive cross-linkers, since hyaluronic acid gel

degradation is principally due to hyaluronidase42 whereas cell migration is

driven principally by the activities of various MMPs.

Self-assembling polypeptides Inspired by the understanding of protein self-assembly, important

progress has been made using supramolecular self-assembly of

biomolecules to form nanofibrillar matrices in situ. These approaches

use non-covalent intermolecular interactions to fabricate higher order

structures by self-assembly of oligomeric peptide, nucleotide and non-

biological amphiphilic building blocks43.

Whereas many of these systems require nonphysiologic conditions

for self-assembly, several can gel at cell-tolerable conditions. Zhang

and coworkers developed a class of nanofibrillar gels with very high

water content (> 99%) cross-linked by assembly of self-complementary

amphiphilic peptides in physiological medium44 (Fig. 1d). These ionic

self-complementary peptides are characterized by a periodic repetition

of alternating ionic hydrophilic and hydrophobic amino acids. Upon

exposure to aqueous solutions with neutral pH, they form stable

β-strand and β-sheet structures, partitioning the side-chains to two sides,

one polar and the other nonpolar. They self-assemble to form nanofibers

with the nonpolar residues inside and positively and negatively charged

residues forming complementary ionic interactions, like a checkerboard.

A number of similar scaffolds have been reported and customized to

deliver growth factors and cells43,45. For example, gels of RAD self-

assembling peptide have been used to encapsulate several growth factors

to accelerate dermal wound healing and myocardial repair46-48.

Self-assembled peptide hydrogels are also being developed as

a tool for three-dimensional cell culture. For example, cell types

such as chondrocytes49 and neural stem cells50 have been cultured

within these scaffolds. However, although these gels biomechanically

organize cells in a three-dimensional fashion, they show no specific

cell interaction, because they are not equipped with any specific

biofunctional ligands. To engineer receptor-mediated biospecificity,

active and functional motifs from the ECM have been employed to

significantly enhance their interactions with cells and tissues. Stupp and

coworkers have designed self-assembling oligomeric-amphiphiles that

allow incorporation of specific biomolecular signals without inhibiting

the self-assembling properties and nanofiber formation. Using

scaffolds presenting the laminin-derived peptide Ile-Lys-Val-Ala-Val,

encapsulated neural progenitor cells were observed to differentiate into

neurons51. Following this trend, a variety of self-assembling peptides

with different functional motifs such as cell adhesion sites or protease

sensitive sequences have been produced50,52 (Fig. 2d).

Another class of polypeptides that forms hydrogels is derived

from the Val-Pro-Gly-X-Gly pentapeptide repeat (where X is a guest

amino acid other than proline) found in human tropoelastin. Called

elastin-like-polypeptides (ELPs), the chains are soluble in aqueous

solution, but as the solution temperature is raised, ELPs become

insoluble and aggregate at a critical temperature, termed the inverse

transition temperature53. By changing the composition and chain

length of guest residue(s), the transition temperature can be precisely

tuned between 0-100°C for specific applications such as recombinant

protein purification54 or cell culture55. For example, ELPs have been

demonstrated to promote the synthesis and retention of cartilaginous

matrix from encapsulated chondrocytes and adult stem cells when

cultured in vitro56. Similarly to self-assembling peptides, ELPs have also

been modified with ECM ligands derived from fibronectin, in order to

promote better cell attachment57.

Synthetic materials to mimic biological functionalityThe biological materials described above serve as a point of departure

for biomimicry, the use of synthetic materials to recapitulate salient

materials features of natural ECM molecules, such as in situ cross-

linking, presentation of adhesion ligands, binding of growth factors,

and susceptibility to cell-derived proteases. Synthetic analogs can

offer several advantages. The chemistries used for matrix formation

and functionalization are becoming increasingly straightforward and

easy to control. Many reactions can be performed under gentle, often

physiological, conditions that allow incorporation of cells or biological

molecules with little loss of viability or function, respectively. The use

of entirely synthetic materials eliminates the purification problems

that can occur with naturally derived materials as well as reduces

the potential for an immune response or pathogen transmission.

Finally, polymeric-based hydrogels, which represent a large class of

synthetic materials used for this application, mimic the highly hydrated

viscoelastic properties of the natural ECM, allow for transport by

diffusion and interstitial flow, and can present soluble, affinity-bound,

or covalently bound biological factors.

Synthetic cell-responsive hydrogels for use in tissue engineering

can be formed from a number of hydrophilic synthetic polymers or

polysaccharides, including poly(ethylene glycol) (PEG), poly(vinyl

alcohol) (PVA), poly(hydroxyethyl methacrylate) (poly(HEMA)),

alginate derivatives, and others, and have been reviewed extensively

in the literature58-61. Here, we focus on the mechanisms of hydrogel

formation, the engineering of biological functionality into hydrogel

constructs, and the presentation of multiple signals in a temporally

dynamic and/or spatially patterned manner.

Mechanisms of hydrogel formationPolymers can be covalently cross-linked into hydrogel networks by

several mechanisms, which can be broadly grouped into chain-growth

polymerizations, such as photopolymerization, and step-growth

polymerizations, such as Michael-addition reactions. Examples of

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photopolymerization include photo-initiated reactions between PEG

diacrylate (PEG-DA) molecules62 (Fig 3a) or between thiol-acrylates63

or thiol-enes64. Lutolf and Hubbell developed vinyl sulfone modified

multi-arm PEG macromers (PEG-VS) to allow a Michael addition

reaction between the acrylated PEG and a thiol group (Fig. 3b), which

is most often presented as a free cysteine on a peptide or protein65.

In this system, the PEG-VS is first functionalized with mono-cysteine

species, such as cell adhesion ligands, growth factor binding ligands,

or even growth factors themselves. The functionalized PEG-VS is then

cross-linked into a biodegradable gel network using peptides that

contain a protease-sensitive substrate sequence flanked by cysteine-

containing domains. Michael addition reactions are not limited to

acrylates and also occur between maleimide and thiol groups (Fig. 3c),

as has been shown for the crosslinking of a p-maleimidophenyl-

modified dextran66.

Recently, sequential copper-free click chemistry (Fig. 3d) has

been applied for the formation of hydrogels and their subsequent

patterning with biomolecules67. Specifically, an azide can react

with an alkyne to form a triazole using a di-fluorinated cyclooctyne

(DIFO3) moiety during hydrogel formation. Orthogonal thiol-ene

photocoupling, such as with peptides containing the photoreactive

allyl ester Fmoc-Lys(alloc)-OH, can then be used to pattern

biomolecules. Anseth’s group has used this click chemistry scheme

to prepare PEG hydrogels with MMP-cleavable sequences and

patterned incorporation of Arg-Gly-Asp (RGD) peptides to provide

localized cell attachment67. In addition to Michael addition reactions

and click chemistry, step-wise polymerizations can also be achieved

through enzyme-catalyzed crosslinking of peptide-functionalized

materials. Sperinde and Griffith first used transglutaminase to cross-

link multifunctional glutaminyl-PEG with polypeptides containing

alternating lysine and phenylalanine residues68. Alternatively, a

mixture of multi-arm PEG molecules, each conjugated with one of

two different counter-reactive peptide substrates for factor XIIIa,

can be crosslinked in the presence of this transglutaminase69 (Fig

3e). Similar to the approach used to covalently functionalize fibrin

matrices, biofunctional peptides tagged with one of these factor XIIIa

substrates can also be incorporated into PEG hydrogels in a simple “one

pot” reaction. Increasingly sophisticated means have been developed

to provide triggered activation of transglutaminase to initiate the

crosslinking process in situ, both for peptide-grafted polymers70 and

for alginates or fibrinogen71. In these systems, thermally triggered

release of calcium from phospholipid vesicles was used to activate the

Fig. 3 (a) Cyclic RGD-modified PEG-DA macromer for photopolymerization, after62 (single letter amino acid nomenclature). (b) Michael addition macromers for reaction between a vinyl sulfone group and a thiol, after79. (i) Multi-arm PEG vinyl sulfone (PEG-VS). (ii) Peptide crosslinker containing free thiols from cysteine residues. (c) Michael addition macromers for reaction between maleimide and thiol groups, after66. (i) p-maleimidophenyl isocyanate (PMPI) conjugated dextran. (ii) Peptide crosslinker containing free thiols from cysteine residues. (d) Functionalized macromers for click chemistry reactions, after67. (i) PEG tetra-azide. (ii) Bis(DIFO3) di-functionalized peptide crosslinker. (e) Peptide functionalized macromers for transglutaminase catalyzed crosslinking, after69. Reactive functionality is either FKGG or NQEQVSPL peptide tag (single letter amino acid nomenclature). (f) Silk-inspired multiblock copolymers that self-assemble, after72. “(g) Photolabile cross-linking macromer for the formation of photodegradable hydrogels, after93.

(b)(a) (c)

(g)

(e)(d) (f)

i i

i

ii ii

ii

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factor XIII (which is calcium ion-dependent) and initiate rapid hydrogel

formation.

In addition to the covalent crosslinking mechanisms described

above, hydrogels can also be formed by physical or ionic interactions

between molecules. This behavior is observed in the self-assembly of

peptide amphiphiles into fibrillar β-sheet structures, as described above,

and during the complexation of polymers or polysaccharides with ions.

Taking inspiration from nature, hybrid hydrogels have been developed by

replacing the amorphous peptide domain of N. calvipes silk with PEG72.

Poly(Ala) self-assembling domains, as in the silk, have been combined

with short-chain PEG segments (Fig. 3f) to form self-assembling hydrogels

with good mechanical properties and retention of the natural β-sheet

structure72. Complexation via the divalent cation calcium leads to

hydrogel formation for alginate or modified alginate polysaccharides73-

75. Alginate is composed of D-mannuronic acid and L-guluronic acid in

a multiblock copolymer form, the guluronic acid units of which allow

for crosslinking by the formation of an “egg-box” conformation upon

the complexation with calcium76. Altering the bulk composition or

distribution of mannuronic and guluronic acid units as well as the overall

molecular weight of the alginate affects the final hydrogel properties75,77.

Engineering biological functionality in synthetic hydrogelsPeptide-conjugated polymers provide a facile means for the

presentation of ECM-derived biomolecular signals. Extensive work has

distilled key elements of proteins, such as cell adhesion sites, protein

binding sites, or protease substrate sites, down to short peptide

sequences that display similar functionality. These peptides can be

conjugated to polymer networks using the chemistries described above,

or incorporated using other strategies that do not interfere with the

reactive functional groups involved in network formation62, and have

been shown to impart biological activity to otherwise inert materials.

The RGD peptide sequence mentioned above, found in several ECM

proteins and known to bind the integrin family of cell-surface adhesion

receptors78, has frequently been used as a cell adhesion ligand and has

been shown to influence 3D fibroblast migration in functionalized PEG

gels (Fig 4a) in a concentration-dependent manner79. More recently,

RGD incorporation has been shown to cause polarization of epithelial

cell aggregates cultured in Michael addition-crosslinked PEG gels; this

effect was similar to that seen with gels functionalized with the full-

length protein laminin80. The specific conformation of these signals can

also matter. Incorporation of cyclic RGD into photocross-linked PEG-DA

hydrogels has resulted in improved endothelial cell adhesion compared

to similar hydrogels with a linear version of the peptide62.

Additionally, other ECM-derived peptides can be used to target

specific cell populations. For example, oligo(PEG fumarate) hydrogels

have been modified with an osteopontin-derived peptide for the

attachment of osteoblasts81, and PEG-DA hydrogels have modified

with a collagen mimetic peptide, (Pro-hydroxyPro-Gly)7, to enhance

chondrogenic differentiation of embedded mesenchymal stem cells

Fig. 4 (a) Fibroblast invasion into MMP-sensitive Michael addition PEG gels. Reprinted with adaptation and permission from79, © 2003, National Academy of Sciences, USA. Phase contrast images after 1, 3, 5, and 7 days (left-to-right, top-to-bottom); scale bar = 250 μm. Confocal image (right) of fibroblasts stained for cell membranes and cell nuclei; scale bar = 150 μm. (b) Outgrowths from neural cells cultured in hydrolytically degradable PEG hydrogels after 14 (left) or 16 (right) days in culture. Scale bar = 20 μm. Reprinted with adaptation and permission from84, © 2005, Elsevier Ltd. (c) Outgrowths from aortic rings into alginate hydrogels without growth factors (left) or with sequential delivery of VEGF and PDGF-BB (right). Reprinted with adaptation and permission from89, © 2007, European Society of Cardiology. (d) Increased cell spreading of hMSCs in hyaluronic acid based hydrogels with (left-to-right, top-to-bottom): (i) no RGD, no MMP-sensitivity; (ii) RGD, no MMP-sensitivity; (iii) no RGD, MMP-sensitivity; and (iv) RGD, MMP-sensitivity. Scale bar = 50 μm. Reprinted with adaptation and permission from42, © 2008, Springer Science+Business Media, LLC. (e) Increased collagen type II expression by hMSCs encapsulated in photopolymerized PEG-DA hydrogels functionalized with a collagen mimetic peptide (right) compared to controls without peptide (left). Scale bar = 100 μm. Reprinted with adaptation and permission from82, © 2008, Mary Ann Liebert, Inc. (f) Effect of RGD and spatial patterning of RGD on 3T3 fibroblasts cultured within PEG hydrogels formed by click chemistry (left to right): no RGD, uniform incorporation of RGD, patterned incorporation of RGD within marked square. Reprinted with adaptation and permission from67, © 2009, Macmillan Publishers Limited.

(b)(a) (c)

(e)(d) (f)

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(MSCs) (Fig. 4e) and collagen retention within the gels82. Further,

these cell adhesive materials have been used in vivo and shown to have

improved efficacy compared to non-biologically active materials. As

one example, this has been demonstrated in a rat myocardial infarct

model treated with alginate hydrogels where the gels modified with

RGD had the greatest angiogenic response83.

Another key parameter for the design of cell responsive synthetic

materials is their ability to degrade. In tissue engineering, the ultimate

goal is to have the implanted materials completely remodeled and

replaced by living tissue. Whether cells are incorporated into the

matrix before implantation or the scaffold simply provides signals

and physical support for host cells to invade, it is important that the

cells can have space to secrete their own ECM to eventually replace

the initial support provided by the hydrogel. Nonspecific hydrogel

degradation can be implemented by the incorporation of hydrolytically

degradable segments, such as poly(glycolic acid) (PGA) or poly(lactic

acid) (PLA) domains, within the hydrogel cross-links84 (Fig. 4b) or by

using hydrolytically degradable polymers, such as partially oxidized

alginate75, as the backbone of the matrix.

In vivo, ECM molecules are enzymatically, rather than hydrolytically,

degraded, with matrix degradation coupled to cell migration through

the expression of cell-associated or cell-secreted proteases. Thus, most

biologically derived hydrogels are naturally degradable, such as the

hyaluronic acid and fibrin hydrogels described above. Cell-mediated

control of degradation can be engineered into synthetic hydrogels

by the incorporation of protease substrate sequences42,52,66,79

(Fig. 4a, d). In these examples, peptide sequences derived from the

protease cleavage site in type I collagen85 or from combinatorial library

screening86 have been utilized to render covalently cross-linked42,66,79

or self-assembled52 hydrogels enzymatically degradable. Going one

step further, exogenous control of degradation can be achieved by

the incorporation of enzyme labile units into hydrogels. Degradation

of photocross-linked PEG gels containing caprolactone units within

the linkers occurred only in the presence of lipase, and the timing and

duration of lipase exposure affected the deposition and organization of

type II collagen by encapsulated chondrocytes in vitro87.

In addition to functionalities that allow cells to interact with the

hydrogel scaffolds, either by binding to or degrading the matrix,

truly biofunctionalized scaffolds will also provide signals necessary to

influence cell behavior. Thus, many groups seek to provide controlled

release of morphogens by physical entrapment, affinity-binding,

or covalent binding within hydrogels. The current state of the art

recognizes that delivery of a single growth factor may be insufficient to

recapitulate natural processes, and thus attention has turned to novel

delivery approaches that allow the release of multiple growth factors

at different doses and with different release rates88. In the area of

angiogenesis, where different growth factors stimulate vessel growth

and maturation, sequential delivery from hydrogel scaffolds has been

achieved by several different strategies. Affinity binding by alginate-

sulfate incorporated into alginate hydrogels resulted in different

delivery rates for bound VEGF, platelet-derived growth factor (PDGF)-

BB, and transforming growth factor (TGF)-β1 and led to increased

vessel density and maturity in scaffolds implanted subcutaneously74.

Improvements in cardiac function and neo-vessel maturity were also

seen with alginate hydrogels sequentially delivering VEGF and PDGF-BB

that were injected in the border zone of induced myocardial infarcts in

rats89, and sprouting was observed from aortic rings embedded in these

gels (Fig. 4c).

Spatially patterned or temporally dynamic hydrogels for presentation of multiple signalsIncreasingly complex hydrogel-based systems allow for presentation of

multiple signals, with the possibilities of spatial patterns of molecules

and/or temporal variation in exposure or release. Drawing from

microfluidics and microfabrication technologies, it is becoming possible

to engineer hydrogels down to the microscale90. Modular approaches

have been used to create in vitro engineered tissues, as seen with

the assembly of cell-embedded collagen microgels91, and processing

conditions such as tissue printing92 can be used to control the

architecture and spatial composition of cell-laden polymeric scaffolds.

The sequential click reactions described above67 also allow for localized

photopatterning of biological molecules after hydrogel formation

(Fig. 4f).

Using stimulus-sensitive linkers, protecting groups, or other

exposing mechanisms, it is possible to create temporally dynamic in

addition to spatially patterned matrices. One recent example comes

from Anseth’s group where gel properties can be manipulated in situ

by photodegradable segments (a nitrobenzyl ether-derived moiety)

incorporated within PEG hydrogels (Fig. 3g)93. The hydrogel itself

is cleaved, resulting in reduced local network cross-link density by

release of PEG; changes in stiffness, water content, and diffusivity;

and eventually complete erosion by release of poly(acrylate) chains.

To allow space for cell spreading or ECM production, the hydrogels

can be eroded to produce channels, leading to directed cell migration

or cell-cell connectivity, and three-dimensional features such as

interconnected channels can be produced by two-photon methodology.

Further, local modification of the chemical environment can be

achieved by the incorporation of photolabile tethered biologically

active functionalities, which was demonstrated to lead to enhanced

chondrogenic differentiation of encapsulated human MSCs upon

removal of RGD93.

SummaryIn the sections above, we have illustrated some of the goals and

challenges of creating biofunctionality in synthetic materials:

the biological materials from which the design concepts arise,

approaches to use those biological materials directly without and with

modification, and schemes by which to mimic the key features of the

MT131-2p14_22.indd 21 01/02/2010 10:35:57

REVIEW Biomimetic materials in tissue engineering

JAN–FEB 2010 | VOLUME 13 | NUMBER 1-222

natural ECM in completely synthetic implementations. Some of these

approaches are being evaluated in clinical trials and more are making

their way through preclinical investigation toward clinical studies.

These materials will enable new biosurgical approaches to tissue

engineering and regenerative medicine for repairing and replacing

damaged or diseased tissue, such as broken bones, acute and chronic

dermal wounds, damaged articular cartilage, and infarcted cardiac

muscle.

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