Progress in Polymer Sciencehome.iitk.ac.in/~skishore/tissue/Design of artificial...

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Progress in Polymer Science 36 (2011) 238–268

Contents lists available at ScienceDirect

Progress in Polymer Science

journa l homepage: www.e lsev ier .com/ locate /ppolysc i

esign of artificial extracellular matrices for tissue engineering

yung-Soo Kima,1, In-Kyu Parkb,1, Takashi Hoshibac, Hu-Lin Jiangd,un-Jaie Choie, Toshihiro Akaike f,∗, Chong-Su Choe,∗∗

School of Chemical and Biological Engineering, Seoul National University, Seoul 151-744, South KoreaDepartment of Biomedical Sciences, Chonnam National University Medical School, Gwangju 501-746, South KoreaBiomaterials Center, National Institute for Materials Science, Tsukuba 305-0044, JapanCollege of Veterinary Medicine, Seoul National University, Seoul 151-742, South KoreaDepartment of Agricultural Biotechnology and Research Institute for Agriculture and Life Sciences, Seoul National University, Seoul 151-921, South KoreaDepartment of Biomolecular Engineering, Tokyo Institute of Technology, Yokohama 226-8501, Japan

r t i c l e i n f o

rticle history:eceived 25 May 2010eceived in revised form2 September 2010ccepted 7 October 2010vailable online 20 October 2010

a b s t r a c t

The design of artificial extracellular matrix (ECM) is important in tissue engineering becauseartificial ECM regulates cellular behaviors, including proliferation, survival, migration,and differentiation. Artificial ECMs have several functions in tissue engineering, includ-ing provision of cell-adhesive substrate, control of three-dimensional tissue structure, andpresentation of growth factors, cell-adhesion signals, and mechanical signals. Design cri-

eywords:rtificial extracellular matrixaturally-derived polymertem cellsynthetic polymer

teria for artificial ECMs vary considerably depending on the type of the engineered tissue.This article reviews the materials and methods that have been used in fabrication of artifi-cial ECMs for engineering of specific tissues, including liver, cartilage, bone, and skin. Thisarticle also reviews artificial ECMs used for modulation of stem cell behaviors for tissueengineering applications.

© 2010 Elsevier Ltd. All rights reserved.
issue engineering
ontents

1. Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2391.1. Significance of artificial extracellular matrix in tissue engineering applications . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2391.2. General requirements of artificial ECMs for tissue engineering. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2391.3. Classification of artificial ECMs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 239

2. Liver . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2402.1. Mechanism of specific interaction between hepatocytes and galactose moieties . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 240
2.2. Classification of artificial ECMs for liver tissue engineering
2.2.1. Synthetic polymers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .2.2.2. Naturally-derived polymers. . . . . . . . . . . . . . . . . . . . . . .

2.3. Parameters affecting hepatocellular behavior . . . . . . . . . . . . . .

∗ Corresponding author. Tel.: +81 45 924 5790; fax: +81 45 924 5815.∗∗ Corresponding author. Tel.: +82 2 880 4868; fax: +82 2 875 2494.

E-mail addresses: [email protected] (T. Akaike), [email protected] (C.-S. C1 These authors contributed equally to this work.

079-6700/$ – see front matter © 2010 Elsevier Ltd. All rights reserved.oi:10.1016/j.progpolymsci.2010.10.001

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dx.doi.org/10.1016/j.progpolymsci.2010.10.001

B.-S. Kim et al. / Progress in Polymer Science 36 (2011) 238–268 239

2.3.1. Galactose density and microdistribution of galactose . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2462.3.2. Topology of artificial ECMs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 247

3. Cartilage . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2483.1. Natural material-based artificial ECMs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2483.2. Synthetic material-based artificial ECMs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 251

3.2.1. Poly(ethylene glycol) (PEG) hydrogels . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2523.2.2. Injectable hydrogels . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 252

3.3. Composite artificial ECMs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2534. Bone . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 253

4.1. Artificial ECMs for bone tissue engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2534.1.1. Design criteria of artificial ECMs for bone tissue engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2534.1.2. Non-polymeric materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2534.1.3. Synthetic polymeric materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2544.1.4. Naturally-derived polymers. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2544.1.5. Composite materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 254

4.2. Regulation of cellular functions by artificial ECMs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2544.2.1. Importance of regulation of cellular function by artificial ECMs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2544.2.2. Regulation of osteoblast functions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 254

4.3. Artificial ECMs for osteochondral defects . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2555. Skin . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 255

5.1. Naturally-derived materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2565.2. Synthetic polymeric materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2565.3. Future challenges in skin tissue engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 256

6. Stem cells . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2566.1. Delivery of soluble factors using artificial ECMs. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2566.2. Cell adhesion signal presentation by artificial ECMs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2586.3. Chemistry of artificial ECMs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2586.4. Architecture of artificial ECMs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2596.5. Mechanical properties of artificial ECMs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2596.6. Control of cell shape and size with artificial ECMs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2606.7. Mechanical signal loading on artificial ECMs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 261

7. Concluding remarks . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 262. . . . . . .

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Acknowledgements . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .
1. Introduction

1.1. Significance of artificial extracellular matrix in tissueengineering applications

Tissue engineering, which aims to reconstruct living tis-sues for replacement of damaged or lost tissues/organsof living organisms, has recently emerged as an excitinginterdisciplinary area in the life sciences [1]. To achievethis aim, it is necessary to use cells together with biosig-nalling molecules and extracellular matrix (ECM) into oronto which cells will develop, organize, and behave as ifthey are in their native tissue [2]. Tissue engineering typ-ically uses artificial ECMs to engineer new tissues fromcells [3,4]. Artificial ECMs should be designed to bring thedesired cell types in contact with an appropriate environ-ment and to provide mechanical support until the newlyformed tissues are structurally stable [5]. Therefore, designof artificial ECMs is very important in tissue engineering,because artificial ECMs regulate many cellular behaviors,including proliferation and growth, survival, change of
cell shape, migration, and differentiation [6]. Also, arti-ficial ECMs have multiple functions, such as serving asan adhesive substrate, provision of structure, presen-tation and storage of growth factors, and detection ofsignals [7].
. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 262. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 262

1.2. General requirements of artificial ECMs for tissueengineering

Artificial ECMs should be designed to guide cellattachment, proliferation, and differentiation for tissueregeneration [8]. Material and biological requirementsshould be considered for each tissue engineering applica-tion, even though the exact requirements differ accordingto organ [9–11]. For the material side, there are severalrequirements for artificial ECMs: artificial ECMs should bebiocompatible and non-toxic, and have a desired degrada-tion rate, surface properties, processability, porosity, andmechanical properties. For the biological side, artificialECMs should be non-immunogenic and should not causeforeign body reactions. They should have desired bioactiv-ity, nutrient availability, and access to growth factors.

1.3. Classification of artificial ECMs

Artificial ECMs can be classified into two types: naturalpolymer-based artificial ECMs and synthetic polymer-
based ones. Natural polymer-based artificial ECMs consistof structural and functional proteins, proteoglycans, gly-coproteins, and glycosaminoglycans found in a naturaltissue [12], and they have been implanted for construc-tive remodeling of many tissues in preclinical and human

240 B.-S. Kim et al. / Progress in Polymer Science 36 (2011) 238–268

Table 1General characteristics of commonly used naturally-derived polymers for tissue engineering [modified from Ref. [9]].

Polymers Biocompatibility Disadvantages Biodegradability Application

Collagen Minimal cytotoxicity,Mild foreign bodyreaction, Minimalinflammation

Proteoytic removal ofsmall nonhelicaltelopepties

Bulk, Controllable Skin; cartilage; bone; ligaments;tendons; vessels; nerves; bladder;liver

Hyaluronic acid Minimal foreign bodyreaction, Noinflammation

Highly viscoussolution, Manypurification steps afterchemical modification

Bulk, 1 h to 1 month Skin; cartilage; bone; ligaments;nerves; vessels; liver

Alginic acid Minimal foreign bodyreaction, Noinflammation

Uncontrollabledissolution of hydrogel

Bulk, 1 day to 3 months Skin; cartilage; bone; nerves;muscle; pancreas

Chitosan Minimal foreign bodyreaction, Noinflammation

Uncontrollabledeacetylation andmolecular weight

Bulk, 3 days to 6 months Skin; cartilage; bone; nerves;vessels; liver; pancreas

Gelatin Minimal cytotoxicity,Mild foreign bodyreaction, Minimalinflammation

Weak mechanicalproperty

Bulk, Controllable Skin; bone; cartilage; ligaments;breast

Fibrin Minimal cytotoxicity,Mild foreign bodyreaction, Minimalinflammation

Weak mechanicalproperty

Bulk, Controllable Skin, bone, cartilage; liver,tendons;ligaments; vessels

Poly(hydroxyalkanoate) Minimal cytotoxicity,Mild foreign bodyreaction, Minimalinflammation

Pyrogen removed Bulk, Controllable Skin; bone; tendons; cartilage;nerves; ligaments; heart; vessels;muscle

of serici

coptamstE

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Silk Minimal cytotoxicity,Mild foreign bodyreaction, Minimalinflammation

Inflammation

linical trials. Naturally-derived polymers have advantagesver synthetic ones because they have excellent biologicalroperties, including cell adhesion, mechanical proper-ies similar to those of natural tissues, biodegradability,nd biocompatibility [12]. General characteristics of theost commonly used naturally-derived polymers for tis-

ue engineering are shown in Table 1. A recent review onissue engineering using natural polymer-based artificialCMs is available [10].

Synthetic polymer-based artificial ECMs are generallyabricated from biocompatible, biodegradable polymers tovoid chronic foreign body reactions. Synthetic polymersay be obtained by reproduction on a large scale and can

e processed into artificial ECMs in which the mechan-cal properties and degradation time of some syntheticolymers can be controlled [12]. However, the greatestisadvantage of synthetic polymers is the lack of cell-ecognition sites and the possibility of a pro-inflammatoryesponse after implantation. General characteristics ofommonly used synthetic polymers for tissue engineeringre shown in Table 2. Recent reviews on synthetic polymersor tissue engineering scaffold are available [9,11,12].

. Liver

The liver plays an important role in a broad spec-rum of physiological functions, including metabolism,
torage, and synthesis and release of vitamins, carbohy-rates, proteins, lipids, and cyclic tetrapyrroles [13]. Also,he liver detoxifies and inactivates endogenous and exoge-ous substances, and activates precursor molecules [13].herefore, the design of artificial ECMs for liver tissue engi-
n Bulk, Controllable Skin; ligaments; bone; cartilage;tympanic membrane; vessels;tendons

neering is critically important, because hepatocytes areattachment-dependent cells and lose their liver-specificfunctions without an optimal ECM. This section discussesmechanisms of the specific interaction between hepatocyteand galactose moieties, classification of artificial ECMs, fac-tors affecting cellular behaviors of hepatocytes, and surfacemodification.

2.1. Mechanism of specific interaction betweenhepatocytes and galactose moieties

Asialoglycoprotein receptors (ASGPR) of hepatocyteswere first identified by Pricer et al. [14]; circulatingasialoglycoproteins (ASGPS) bind to and are degraded byhepatocytes. Hepatocytes have cell surface receptors thatrecognize and bind molecules with exposed galactose,N-acetyl galactosamine, or glucose residues [15]. Afterbinding to receptors, ASGPS are internalized via coatedpits and coated vesicles and subsequently appear in acomplex arrangement of larger smooth-surface vesicles[16]. Hepatic plasma membrane receptors localized on thesinusoidal face of hepatocytes mediate specific binding anduptake of partially deglycosylated glycoproteins throughreceptor-mediated endocytosis [17]. Therefore, mimickingthe biological environment, the process of incorporationof old proteins by the liver has been imitated; however,
hepatocyte transplantation using poly(l-lactic acid) (PLA)[18], poly(vinyl alcohol) (PVA) [19], poly(glycolic acid)(PGA) [20], poly(lactic-co-glycolic acid) (PLGA) [21], col-lagen [22], alginate [23], and polyurethane [24] has beenachieved in vivo.


Table 2General characteristics of commonly used synthetic polymers for tissue engineering [modified from Ref. [9]].

Polymers Biocompatibility Disadvantages Biodegradability Application

Poly (lactic acid) Minimal cytotoxicity, Mildforeign body reaction,Minimal inflammation

Local inflammation,Random chain hydrolysis

Bulk, 24 months Skin; cartilage; boneligaments; tendons;vessels; nerves; bladder;liver

Poly (glycolic acid) Minimal cytotoxicity, Mildforeign body reaction,Minimal inflammation


Bulk, 6–12 months Skin; cartilage; boneligaments; tendons;vessels; nerves; bladder;liver

Poly (lactic acid-co-glycolic acid) Minimal cytotoxicity, Mildforeign body reaction,Minimal inflammation


Bulk, 1–6 months Skin; cartilage; boneligaments; tendons;vessels; nerves; bladder;liver

Poly (caprolactone) Minimal cytotoxicity, Mildforeign body reaction,Minimal inflammation

Hydrophobic Bulk, 3 years Skin; cartilage; bone;ligaments; tendons;vessels; nerves

Poly (ethylene oxide)terephthalate-co-butyleneterephthalate

Mild foreign body reaction,No inflammation

Complex biodegradability Bulk, 1 month-5 years Skin; cartilage; bone;muscles

Polyanhydrides Minimal foreign bodyreaction, Minimalinflammation, Minimalcytotoxicity

Limited mechanicalproperty

Surface erosion,Controllable

Bone

Poly (propylene fumarate) Mild foreign body reaction,Minimal inflammation

Weak mechanical property Surface erosion, 1week–16 months

Bone

Poly (orthoester)s Mild inflammation, Mild Weak mechanical property Bulk ∼ several months Ear; bone; cartilage

oad molstributio

foreign body reactionPolyphosphazene Minimal foreign body

reaction, Minimalinflammation

Brdi

2.2. Classification of artificial ECMs for liver tissueengineering

Two types of artificial ECMs for liver tissue engineeringwill be discussed: one is derived from synthetic polymers,and the other is derived from natural sources.

2.2.1. Synthetic polymersSynthetic polymer applications in liver tissue engineer-

ing have attracted attention in recent years due to theirexcellent mechanical properties, processability, low cost,and controllable degradation time. However, the great-est disadvantage of synthetic polymers is the absence ofhepatocyte-recognition sites. Therefore, many efforts havefocused on incorporation of galactose ligands, the cell-adhesion sites for synthetic polymers, into the design ofbiomimetic polymers for hepatocytes.

Galactose ligand was first coupled to polyacrylamide(PAAm) by Weigel et al. [25]. PAAm gels were syn-thesized by copolymerization of N-succinimidyl acrylate,acrylamide, and bisacrylamide. Next, 6-aminohexyl �-d-galactose was reacted with the activated copolymer gelthrough amide linkage [26]. Hepatocyte adhesion on thegalactosylated PAAm gels was verified [27]. The resultsindicated that rat hepatocytes bound to galactosylatedpolyacrylamide in a Ca2+- and temperature-dependentmanner via a patch of ASGPR in rat hepatocytes, but not
chicken hepatocytes [26,27]. Also, cell binding to the poly-mer was inhibited by asialo-orosomucoid, indicating areceptor-mediated mechanism [28]; however, the effectof galactose-coupled PAAm gel hydration on hepatocyteadhesion was not investigated.
ecular weightn

Surface erosion, 1week–3 years

Skin; cartilage; bonenerves; ligaments

Akaike and colleagues synthesized galactose-carryingpolystyrene (PS) in order to guide hepatocyte adhe-sion. Galactose-carrying PS [poly(N-p-vinylbenzyl-4-O-�-d-galactopyranosyl-d-gluconamide), PVLA] was synthe-sized according to three steps, as shown in Fig. 1 [29].Synthesis of PVLA is relatively simple, protection of thehydroxyl groups of lactose is unnecessary, and the yieldof each step is high, even though it is expensive to preparecompared to other synthetic polymers. The authors inves-tigated the use of PVLA as an artificial ECM for hepatocytes.The results indicated that hepatocytes adhered to the PVLAthrough a unique ASGPR-galactose interaction between theASGPR of hepatocytes and highly concentrated galactosemoieties along the polymer chains [30]. The round mor-phologies of hepatocytes on PVLA that adhered at highconcentrations of PVLA triggered spheroid formation ofhepatocytes in the presence of epidermal growth factor(EGF), leading to enhanced differentiation functions [31].In addition, due to its amphiphilic character, this polymeris stably adsorbed onto the tissue culture dish [32].

Lopina et al. synthesized galactose-modified starpoly(ethylene oxide) (PEO) hydrogels by reacting tre-sylated PEO hydrogels with 1-amino-1-deoxy galactose[33]. They investigated hepatocyte attachment on theseconstructs. Hepatocytes were found to exhibit galactose-specific adhesion to the galactose-modified PEO gels,adhering to gels with galactose via the spreading
morphologies of hepatocytes at low concentrations ofimmobilized ligands, but not glucose [34].
Donati et al. synthesized galactosylated poly(styrene-co-maleic acid) (PSMA) by combining PSMA with 1-amino-1-deoxy-�-d-galactose (or 1-amino-1-deoxy-�-d-lactose)

242 B.-S. Kim et al. / Progress in Polymer

Fr

acwi

abeEmesPgac

ig. 1. Synthetic scheme of galactose-carrying polystyrene [29]. Copy-ight 1985, Nature Publishing Group, London, UK.

nd testing HepG2 adhesion [35]. The results showed thatell adhesion on galactosylated PSMA or lactosylated PSMAas approximately five times higher than that of PSMA

tself.Hepatocyte–ECM interaction is a surface phenomenon,

nd is affected by surface properties of ECM; however,ulk properties of the ECM dictate the mechanical prop-rties of the ECM. Therefore, surface modification of theCM with galactose ligand is often necessary for opti-ization of ECM for hepatic tissue engineering. Yoon

t al. synthesized galactosylated PLGA film or galacto-
ylated PLGA sponge by reacting the carboxylic acid ofLGA at the terminal group and N-(amino butyl)-O-�-d-alactopyranosyl-(1 → 4)-d-glucoamide or by reacting themino groups of PLGA-PEG and lactobionic acid for hepato-yte culture [36], even though the albumin secretion rates
Science 36 (2011) 238–268

of the galactosylated PLGA surface in vitro were relativelylow compared to those of other culture systems. However,using the perfusion culture system based on galactose-modified macroporous PLGA scaffolds under optimal flowconditions, the rate of albumin secretion was significantlyincreased due to spheroid formation of hepatocytes in thescaffolds [36].

Mao and colleagues immobilized galactose ligandson arylic acid-graft-copolymerized poly(ethylene tereph-thalate) (PET) film by plasma treatment and evaluatedhepatocyte function [37]. The results indicated that hepato-cytes cultured on the galactosylated surface exhibited goodalbumin and urea synthesis, which were comparable tothose of hepatocytes cultured on the collagen-modified PETsurface. Also, they prepared the galactosylated PET surfaceby surface-grafting poly(acrylic acid) on plasma-pretreatedPET film to increase the density of the immobilized galac-tose ligands [37]. The results indicated that due to highligand density, better maintenance of albumin secretionand urea synthesis function was obtained compared toculture on a collagen-coated surface. Further, they pre-pared galactosylated poly(vinylidene difluoride) (PVDF)by coating with galactose-tethered Pluronic and eval-uated hepatocyte function [38]. Results demonstratedthat hepatocytes attached to galactosylated PVDF formedmulti-cellular spheroids after one day of culture and exhib-ited higher albumin and P4501A1 detoxification than theunmodified PVDF membrane and collagen-coated sur-face. However, due to only hydrophobic interaction inthe adsorption of galactose ligands, the stability of thecoating after long culture periods is a substantial con-cern. Kang et al. prepared galactose-grafted PS surfacesby grafting of N-p-vinylbenzyl-4-O-�-d-galactopyranosyl-d-gluconamide to a PS dish using oxygen plasma glowdischarge treatment and evaluated hepatocyte adhesion[39]. The results indicated that compared to PVLA, hepa-tocytes adhered more slowly to the galactose-grafted PSsurface than to PVLA during the first 2 h of incubation dueto the low galactose density of the galactose-grafted PS.

Higashiyama et al. prepared fructose- and galactose-modified polyamidoamine dendrimers by imine reactionbetween polyamidoamine dendrimer and galactose (orfructose) and evaluated hepatocyte function [40]. Theresults indicated that simultaneous modification of thedendrimer with fructose and galactose had a synergisticeffect on spheroid formation of hepatocytes, higher ureasynthesis, and increased albumin gene expression thanthose cultured on single-ligand modified dendrimers; theligands introduced in the dendrimer were not verified.Synthetic polymers derivatized with galactose ligands aresummarized in Table 3.

2.2.2. Naturally-derived polymersNaturally-derived polymers composed of proteins,

polysaccharides, and nucleic acids have several inher-ent merits, including bioactivity, the ability to present
receptor-binding ligands to cells, biodegradability, and sus-ceptibility to natural remodeling, even though antigenicity,instability/deterioration, complexity of purification, andrisk of disease transmission should be considered [41].This section discusses the liver tissue engineering applica-


Table 3Synthetic polymers derivatized with galactose ligands.

Polymers Type of galactose Linkage Characters References

Poly (acryl amide) 6-aminohexyl �-d-galactose Amide Firstly coupled temperaturedependence of endocytosis, No resultsof effect of hydration on cell adhesion

[25–28]

Polystyrene �-d-lactose Amide Amphiphilic and micellar, Spheroidformation of hepatocytes,Non-degradable

[29–31]

Poly (ethylene oxide) 1-amino-1-deoxy-�-d-galactose SN1 reaction Radiation-crosslinked PEO starhydrogel spreading of hepatocytes,Spatial distribution of galactose ligand

[33–34]

Poly (styrene-co-maleic acid) 1-amino-1-deoxy-�-d-galactose1-amino-1-deoxy-�-d-lactose

Amide Appearance of circular dichroismbands, No mechanism of cell adhesion

[35]

Poly (lactic acid-co-glycolic acid) N-(aminobutyl)-O- �-d-galactopyranosyl-(1 → 4)-d-glucoamide Lactobionicacid

Amide Introduction of PEG spacer, Spheroidformation of hepatocytes, Perfusionculture system

[36]

Acrylic acid-graft-poly(ethyleneterephthalate)

1-O-(6′-aminohexyl)-d-galactopyranoside

Amide UV-induced graft copolymerization,Spheroid formation of hepatocytes,High albumin secretion and ureasynthesis

[37]

Poly(vinylidene difluoride) �-d-lactose Imine reaction Unstable coating due to physicaladsorption, Buffering of proteins byPEO

[38]

EtIm

Polysyrene �-d-lactosePolyamidoamine dendrimer Galactose+fructose

tions of collagen and fibronectin, which are commonly usedin integrin-mediated hepatocyte adhesion, and anothergroup of naturally-derived polymers, including alginate,chitosan, gelatin, silk, and xyloglucan, which are regularlyused in receptor-mediated hepatocyte adhesion.

Collagen is the most abundant protein in the body, andis responsible for maintenance of the structural integrity ofvertebrates and several other organisms [42]. More than 20different types of collagen have been identified thus far inthe human body; the most common collagen is Type I–IV.Among these, Type I collagen is the single-most abundantprotein in the body after cartilage [12]. Characteristics ofcollagen, such as good biodegradability, low antigenicity,and cell-binding properties have made it a prospect for usein ECMs for liver tissue engineering, even though collagenuse is limited due to its low mechanical strength, rapidbiodegradation, and the high cost of pure collagen [12].Collagen gels have been used for culture of hepatocytesfor application to the bioartificial liver [43]. The resultsindicated that the viability of rat hepatocytes cultured incylindrical collagen hydrogels in vitro remained stable withstable albumin synthesis throughout the seven day incuba-tion period, and by 10 days culture, aggregated hepatocytesmaintained an albumin production that was two-fold thatof single hepatocytes [43]. Recently, Zhao et al. useda collagen gel to generate engineered hepatic units forreconstitution of a three-dimensional, vascularized hep-atic tissue in vivo [43]. Staking of hepatic units under thesubcutaneous space was found to result in significant cellengraftment with formation of a large fused hepatic sys-
tem containing blood vessels, suggesting that this is anefficient method for engineering of hepatic tissue in vivo.To overcome the weak mechanical property of the colla-gen gels, collagen/Poloxamine [44] and collagen/chitosan[45] hybrids were prepared as artificial ECMs for liver
her Redox initiation, Graft polymerization [39]ine reaction Uncertain chemistry, Synergistic effect

by fructose[40]

tissue engineering. Production of collagen/Poloxamine-methacrylate hybrid networks exhibited the high stiffnessof the hybrid gels due to the crosslinked Poloxamine hydro-gels in addition to enhanced cell adhesion and high survivallevels [44]. Shape and volume of ammonia-treated colla-gen/chitosan sponges were constant over a period of 50days in enzyme solutions until 25 days postseeding [45],suggesting the potential of these artificial ECMs for use inliver tissue engineering.

Fibronectin is a family of multifunctional ECM knownto play key roles in cell attachment and migratory cellularbehaviors, such as embryogenesis, malignancy, homeosta-sis, wounding healing, and maintenance of tissue integrity[46]. The ability of fibronectin to serve as a substratefor cell adhesion is based on the RGD tripeptide [47].However, due to the high cost, research on the use offibronectin in tissue engineering, except for the coatingof fibronectin on PS [48] and PLGA surfaces [49], has beenlimited. However, significant research has been conductedon the role of fibronectin in regulation of switching fromdifferentiation to growth [50], tubulin monomer growth[51], cell cycle status [52], cell spreading [53], and motil-ity [54]. The effect of hepatocyte growth factor (HGF) [55]in hepatocytes was studied for elucidation of its basicmolecular biology. Bhadriraju et al. [48] compared hep-atocyte adhesion, growth, and differentiated function on2.3 kD and 73 kD RGD peptide-coated PS dishes to that onfibronectin-coated ones. Of particular interest, a 2.3 kD RGDpeptide induced a round cell shape, enhanced differenti-ated function, and inhibited DNA synthesis; however, the
73 kD RGD peptide induced cell spreading, dedifferentia-tion, and enhanced DNA synthesis (similar to the impactof fibronectin), suggesting that RGD peptide conformationdetermines the specificity of cellular response. Fiegel et al.[49] investigated the effect of three-dimensional culture


inate [5

oulsi

sl(badginatngofah

gsti

Fig. 2. Reaction scheme of galactosylated alg

f rat hepatocytes on fibronectin-coated PLGA scaffoldsnder flow cell conditions. The results indicated that cellu-

ar functions of hepatocytes, such as growth and albuminecretion, were increased significantly by fibronectin coat-ng and flow conditions.

Alginate, the primary structural component of browneaweed, is the monovalent form of alginic acid, and is ainear polymer of � (1 → 4)-linked d-mannuronic acid and1 → 4)-linked l-guluronic aid [56]. Alginate is widely usedecause it has low toxicity, is easy to modify chemically,nd undergoes mild ionotropic gelation in the presence ofivalent cations; however, degradation of alginate hydro-el occurs via a slow and unpredictable dissolution processn vivo [57]. Yang et al. [58] prepared galactosylated algi-ate via the reaction of carboxylic acids of alginate andmine-modified lactobionic acid (Fig. 2) and evaluatedhe liver function of hepatocytes in galactosylated algi-ate microcapsules. The results indicated that binding ofalactosylated alginate microcapsules with hepatocytesccurred via the ASGPR patch on hepatocytes and that liverunction of hepatocytes was enhanced compared to that inlginate treated hepatocytes due to spheroid formation ofepatocytes in galactosylated alginate microcapsules.

Chitosan is a linear polymer of �(1 → 4)-linked d-lucosamine, and is obtained by deacetylation of chitin; it isuitable as a substrate for biomimetic glycopolymers due tohe similarity of its structure to glycosaminoglycans foundn native tissue [57]; however, it is only soluble in diluted

8]. Copyright 2002, Elsevier Inc., Oxford, UK.

acids. Free amino groups in chitosan are easily modifiedby covalent attachment of molecules using carbodiimidechemistry. Park et al. prepared galactosylated chitosan (GC)by reacting lactobionic acid with amino groups of chitosanusing EDC/NHS as coupling agents (Fig. 3) and evaluated thesubsequent hepatocyte adhesion [59]. The results indicatedthat hepatocytes adhered via galactose-specific recogni-tion between GC and ASGPR of hepatocytes, and hepato-cytes that adhered to the surface at high concentrationsof GC exhibited substantial spheroid formation after 24 hexposure to EGF. Yang et al. also prepared a hybrid artificialECM composed of alginate and GC because the mechanicalproperties of GC were insufficient [60,61]. Hybrid GC andalginate were found to improve cell adhesion and stabilityof the artificial ECM, which retained differentiated cellu-lar functions better than alginate itself. Also, pore size ofthe artificial ECM could be controlled by the content andmolecular weight of GC and the freezing temperature ofalginate/GC hydrogel. Introduction of heparin in the algi-nate/GC artificial ECM enhanced the level of albumin secre-tion from hepatocytes in the presence of HGF due to rapidformation of stable spheroids in the alginate/GC/heparinartificial ECM, which is closely related to connexin 32 and
E-cadherin genes in cell-to-cell interactions [62].
Gelatin is a naturally-derived polymer obtained fromcollagen and has been used in medical and pharmaceuticalapplications because of its biodegradability and biocom-patibility [63]. Hong et al. prepared galactosylated gelatin


. Copyri
Fig. 3. Synthesis scheme of GC [59]
by reacting lactobionic acid with amine-incorporatedgelatin (Fig. 4) and evaluated it for use in hepatocyte culture[64]. The results indicated that the survival time of hep-atocytes cultured on galactosylated gelatin artificial ECMwas longer than that of hepatocytes cultured on collagen-coated monolayers, and liver function (e.g. secretion ofalbumin, synthesis of urea) was maintained due to thespecific interaction of galactose moieties in galactosylatedgelatin with ASGPR of hepatocytes.

Silk is naturally produced by spiders and Lepdopterai.Silk has been used in ligament tissue engineering due toits unique mechanical properties and good biocompati-bility [65]. Gotoh et al. [66] prepared galactosylated silkfibroin by conjugation of lactose with silk using cyanuricchloride as a coupling spacer, and evaluated its ability tofacilitate hepatocyte attachment. The results indicated thathepatocyte attachment on galactosylated silk-coated PSdishes increased by approximately eight-fold compared tothat of uncoated dishes, and hepatocytes exhibited a round
morphology; however, the investigators did not assessalterations to liver function caused by three-dimensionalgalactosylated silk fibroin.
Xyloglucan, a polysaccharide derived from tamarindseeds, is composed of glucose units in the main chain and

ght 2003, Elsevier Inc., Oxford, UK.

xylose and galactose units in the side chains. Seo et al. [67]studied the specific interaction between galactose moietiesin the side chain of the xyloglucan and ASGPR of hep-atocytes. The results indicated that hepatocyte adhesionto xyloglucan-coated PS surfaces was dependent on thepresence of Ca2+, and spheroidal hepatocytes formed at ahigh concentration of xyloglucan in the presence of EGFdue to galactose-specific recognition. In addition, alginatemicrocapsules prepared with xyloglucan as an artificialECM enhanced multicellular spheroidal hepatocyte forma-tion due to the specific interaction between the galactosemoieties of xyloglucan and ASGPR of hepatocytes [68].Naturally-derived polymers used for liver tissue engineer-ing are summarized in Table 4.

2.3. Parameters affecting hepatocellular behavior

Long-term and stable hepatocyte culture systems arevery important for liver tissue engineering and for develop-
ment of bioartificial liver devices. Morphology, attachment,growth, differentiation, and survival of hepatocytes arehighly dependent on galactose density and microdistribu-tion of galactose in galactose-carrying polymers, structureof carbohydrate, topology of ECM, coculture, and cell


F atin (MGW

st

2g

cmlorr

ddhDoh

TN

ig. 4. Chemical structure of MGLA [a galactose-containing, modified geliley & Sons, Inc, Hoboken, NJ, USA.

ource. This section addresses the effects of these parame-ers on hepatocellular behavior.

.3.1. Galactose density and microdistribution ofalactose

Ligand density is a primary determinant of hepato-yte attachment, morphology, and function; triantennaryolecules with three terminal galactose residues bind to

ectins with higher affinity than biantennary ones with oner two terminal galactose residues, and a multi-subuniteceptor of hepatocytes is responsible for binding galactoseesidues on desialylated glycoproteins [17].

Kobayashi et al. [30] studied the effects of galactoseensity in galactose-carrying PS on the morphology and
ifferentiation of hepatocytes. The results indicated thatepatocytes exhibited higher 3H-thymidine uptake (theNA synthesis indicator) with wider spread morphol-gy at the lower galactose density (0.5 �g/ml); however,epatocytes showed lower 3H-thymidine uptake with
able 4aturally-derived polymers for liver tissue engineering.

Polymers Type of galactose Linkage Characters

Collagen RGD sequence Amide Weak mechanical proFibronectin RGD sequence Amide High cost, Coating of fiAlginate Lactobionic acid Amide Gelation, Eassy chemi

hydrogel in vivoChitosan Lactobionic acid Amide Easy chemical modific

molecular weightGelatin Lactobionic acid Amide Minimal inflammationSilk Lactose Ether Good mechanical prop

residual sericinXyloglucan Galactose – Biocompatible, Therm

) that is modified with lactobionic acid (LA)] [64]. Copyright 2003, John

a round morphology at higher galactose concentrations(100 �g/ml). The group did not determine an optimal con-centration of round hepatocytes. Ise et al. [69] reported thathepatocytes adhered to galactose-carrying PS with coatingdensity lower than 20 ng/ml exhibited higher proliferativeability than those attached to galactose-carrying PS withcoating density higher than 50 ng/ml. Yin et al. [37] alsoreported formation of hepatocyte spheroids one day aftercell seeding with better albumin secretion and urea synthe-sis at a high galactose density of 513 nmol/cm2 on the PETsurface. Kim et al. [70] proposed a mechanism for hepato-cyte behavior according to the coating density of the PVLA(Fig. 5). They explained that ASGPR on hepatocyte mem-branes are clustered within sites of focal adhesion in a large
patch, and that this clustering prevents participation ofintegrin receptors during the adhesion process at high den-sity. Conversely, hepatocytes allow integrin receptors toparticipate in the adhesion process within the space wherenatural ECMs are secreted during cell culture. Thus, inte-
References

perty, Fast enzymatic biodegradation, High cost [42–45]bronectin for application [46–47]

cal modification, Unpredictable dissolution of [58]

ation, Difficult of control in deacetylation and [59–62]

, Clinically approved, Weak mechanical properties [64]erties, Good biocompatibility, Immune reaction by [66]

ally-reversible gelation [67–68]


Fig. 5. Illustration of hepatocyte behavior on PVLA surfaces. ASGPR-dependent adhesion occurs initially within the interface between hepatocytes andthe PVLA surface at both (A) high coat density and (B) low coat density of PVLA. However, ASGPR-independent adhesion takes place more rapidly at lowcoat density than at high coat density during culture. It is expected that ASGPR-independent adhesion would be induced by integrin receptors that have

ed fromrough inrd, UK.

the opportunity to participate in cell adhesion mediated by ECMs secretadhesion that occurs at low coat density causes hepatocytes to spread thdensity inhibit integrin signaling. [70] Copyright 2003, Elsevier Inc., Oxfo

grin receptors play a role in turning round morphologiesinto spread ones and in focal adhesion kinase phosphory-lation due to involvement of the integrin-mediated signalpathway.

The interaction between galactose ligand and ASGPRis not only affected by galactose density on the ECM butalso by spatial microdistribution of galactose. Cho et al.[71] studied the effect of orientation on hepatocyte attach-ment to a PVLA surface prepared by the Langmuir-Blodgett(LB) method. It was found that hepatocytes that adheredto the LB surface of the polymer even at a very low galac-tose density were round due to the spatial orientation ofgalactose; these cells were spread on the standard coatedPVLA surface. Griffith et al. [34] also reported that spa-tial microdistribution of galactose in the ECM affectedmorphology and function of hepatocytes with the acces-sibility of galactose clustered in spatial microdomains asthe hepatocyte receptors organized into aggregated struc-tures within the cell membrane, suggesting that the spatialmicrodistribution of the ligand in the ECM and control ofthe microenvironmental niche are important for regulationof cellular behavior and successful tissue engineering.

2.3.2. Topology of artificial ECMsTopology of artificial ECMs is another important

parameter that influences cell morphology, function, andphysiological responsiveness due to modulation of cellpolarity by the topology of the ECM [72]. It is com-monly accepted that three-dimensional artificial ECMsmore effectively induce differentiated hepatocyte func-

hepatocytes following initial adhesion. Finally, the ASGPR-independenttegrin signaling. The concentrated ASGPR-ligand complexes at high coat

tion than two-dimensional ones due to the provision ofbetter model systems for physiologic situations. Berthi-aume et al. [72] reported that hepatocytes cultured ina sandwich configuration of collagen exhibited a dra-matic reorganization of the cytoskeleton, adoption ofin vivo-like morphology and polarity, and expression ofa wide array of liver-specific functions due to the invivo-like configuration; however, this configuration isdifficult to scale-up and has nutrient transport limita-tions. ECM topology in the micro- and nanometer rangeshas been shown to affect cellular behavior, includingadhesion, migration, proliferation, and gene expression[73,74]. Ranucci et al. [75] reported that PLGA foamswith subcellular size voids (approximately three microme-ters) induced two-dimensional hepatocyte reorganization,whereas foams with subcellular size voids (approximately67 micrometers) promoted three-dimensional aggrega-tion due to enhanced cell-cell contacts. At intermediatevoid sizes (approximately 17 micrometers), both two-dimensional and three-dimensional reorganization waspromoted, and hepatocytes that adhered to collagenfoams with a void size of 82 micrometers exhibiteda high degree of spreading with high albumin secre-tion due to three-dimensional intercellular contacts[22].

Park et al. [76,77] reported that hepatocytes cultured onmicrogrooved glass substrates with a 100 micrometer-highchannel between each substrate had significantly betterliver-specific function than those in substrates withoutmicrogrooves because hepatocytes were protected from

2 Polymer

de

wd9tlw

Etattmcccecodmhbftg

doapf

3

tatrlftohtttiwmrcEtm,a

48 B.-S. Kim et al. / Progress in

etrimental shear stress and maintenance of oxygen deliv-ry.

Glicklis et al. [78] reported that hepatocytes seededithin three-dimensional porous alginate sponges withiameters of 100–150 micrometers exhibited more than0% hepatocyte aggregation within 24 h post-seeding dueo the non-adherent nature of alginate, and they secretedarge amounts of albumin (60 �g/106 cells/day) within a

eek, even though they did not proliferate.Recently, a three-dimensional nanofibrous artificial

CM with high porosity and high spatial interconnec-ivity prepared by electrospinning has been increasinglypplied for tissue engineering [79,80] because its struc-ure is similar to that of natural collagen fibers [81]; andhe nanofibrous artificial ECM can promote cell attach-

ent, proliferation, migration, and differentiation, and itan attract more integrin-binding proteins than film artifi-ial ECMs [82]. Chua et al. [83] reported that hepatocytesultured on galactosylated poly(�-caprolactone-co-ethylthylene phosphate) nanofiber artificial ECMs exhibitedellular function with the enhanced mechanical stabilityf hepatocyte spheroids similar to that on the two-imensional galactosylated substrate; however, the cellorphologies were different. Feng et al. [84] reported that

epatocytes cultured on galactosylated chitosan nanofi-rous artificial ECMs with an average diameter of 160 nmormed flat, spheroidal hepatocytes and exhibited bet-er liver-specific function than spheroidal hepatocytes onalactosylated chitosan films.

In summary, ECM for liver tissue engineering should beesigned to have high galactose density for achievementf higher affinity in hepatocytes, to have three-dimensionalrtificial ECMs in order to provide better model systems forhysiologic situations, and to form hepatocyte spheroidsor enhancement of liver-specific function.

. Cartilage

The connective tissue comprising articular cartilage ofhe knee is highly specialized for reducing joint frictiont the interface of two long bones, and it merely con-ains chondrocytes in an avascular structure [85]. Theegeneration of damaged articular cartilage remains chal-enging due to its poor intrinsic capacity for repair. Thusar, no surgical procedure has been able to reproducehe biological composition and biomechanical propertiesf native cartilage [86]. The specific treatment strategiesave been dictated by the nature or size of lesions andhe preference of the operating surgeon [86]. However,he advent of tissue engineering has provided revolu-ionary potential for treating cartilage-related diseases. Its believed that tissue engineering of articular cartilage

ill overcome the current limitations of surgical treat-ent by offering functional regeneration in the defect

egion. This technology involves ex vivo culturing ofhondrocytes from autologous or allogenic sources in
CM-based constructs and subsequent implantation intohe cartilage defect. Although recent progress has been
ade to engineer artificial cartilage via tissue engineeringthe challenges remain significant. Selection of the properrtificial ECMs and incorporation of growth factors or

Science 36 (2011) 238–268

mechanical stimuli is of primary importance to success-fully produce artificial cartilage for tissue repair. In thissection, artificial ECMs and suitable cell sources used inarticular cartilage tissue engineering will be introduced,and growth factors and mechanical stimuli used as cues forinducing differentiation to chondrocytes (Fig. 6) will also bediscussed.

Artificial ECMs ideally should furnish chondrocytes withoptimal physical conditions mimicking the natural ECMmicroenvironment of cartilage. Construction of artificialECMs involves fabrication of a three-dimensional networkof artificial ECM similar to that of the original structure andthe provision for structural support and internal space forthe residing cells to adhere, proliferate, and differentiate.Therefore, general requirements of three-dimensional arti-ficial ECMs include high porosity, controlled degradation,mechanical stiffness and strength, and biocompatibility,and they also should exhibit adequate surface propertiesfor proper tissue formation of chondrocytes [87]. Higherporosity allows for the migration and proliferation ofadhering cells and the facilitated exchange of nutrientsand waste products. Controlled biodegradability also isimportant so as not to hinder the formation of newly regen-erating tissue within the artificial ECMs. Biocompatibilitymay facilitate cellular attachment and differentiation.

In most cases, biodegradable polymers including syn-thetic or natural polymers have been used in cartilagetissue engineering. These polymers usually are formedin sponges, hydrogels, or nanofibers. Recently, compositeartificial ECMs composed of a mixture of different poly-mers were designed to provide chondrocytes residing inthe matrix with the combination of merits that each com-ponent contributes.

3.1. Natural material-based artificial ECMs

Natural materials have been preferred over syntheticmaterials because of their intrinsic advantages of bio-compatibility, biodegradability, and an improved capacityfor cell attachment. They can be divided in two groups:protein-based artificial ECMs (e.g. collagen, fibrin, silk)and carbohydrate-based artificial ECMs (e.g. hyaluronan,alginate, agarose, chitosan). Applications of these naturalmaterials-based artificial ECMs for in vitro and in vivo car-tilage tissue engineering are summarized in Table 5.

Collagen is a natural component of cartilage and isknown to play an important role in cellular adhesion anddifferentiation through specific interactions between lig-ands on collagen chains and adhering cells [86]. Therefore,it is widely used for engineering artificial tissue in a broadspectrum of organs. Type II collagen also plays an essentialrole in the maintenance of chondrocyte function. BovineBMSCs seeded on type II collagen represent the mostprominent phenotype of chondrocyte differentiation withthe addition of transforming growth factor (TGF)-�1 in atime-dependent manner [88]. In addition, chondrogenic
differentiation only was detected in three-dimensionalhydrogels, not in monolayer cultures. In another in vitrotest, it was found that a collagen sponge could providea superior microenvironment for the formation of ECM,such as proteoglycans, when human intervertebral disc


Fig. 6. Components required for chondrocyte tissue engineering [modified from Ref. [86]].

Table 5Application of natural materials-based scaffolds for cartilage tissue engineering.

Scaffold Component Forms Seeding cells Biofactors Model/target site Results References

Collagen A bilayermembrane

Autologouschondrocyte

Human/Deep cartilagedefects

Hyaline-like cartilageformation at the defect

[93]

Collagen A bilayermembrane

Autologouschondrocyte

Sheep/chondral defects Reparative tissueformation

[91]

Fibrin Hydrogel Autologous ADSC Rabbit/chondral defects Hyaline-like cartilageformation at 8-weekfollow-up completehealing to subchondralbone

[135]

Alginate/agarose Hydrogel Autologous ADSC In vitro Enhanced chondrogenicdifferentiation

[90]

Alginate/agarose Hydrogel Autologouschondrocyte

Human/osteochondraldefects

Predominant hyalinecartilage formation at2-year follow-up

[106]

Alginate Hydrogel Allogeneic BMSC Rabbit/Osteochondraldefects

Increase in aggrecan andtype II collagen

[105]

Chitosan Hydrogel Autologous Wholeblood

Sheep/Chondral defects Significantly more hyalinerepair tissue formation

[108]

Chitosan Hydrogel Autologous Wholeblood

Rabbit/Bilateral trochleardefects

Formation of a moreintegrated and hyalinerepair tissue completerestoration of GAG levels

[109]

HA Sponge Autologous BMSC FGF Rabbit/kneeosteochondral defect

Formation of cartilagetissue with expressed typeII collagen

[101]

HA esterifiedderivative

Sponge Autologouschondrocyte

Human/chronic cartilagelesions

Continued increase ofclinical performance Nomajor adverse eventsduring 3-year follow-up

[103]

HPMC Hydrogel human nasalchondrocytes

Nude mice/subcutaneousimplantation

Formation of acartilage-like tissue withGAG and type II collagen

[110]

Peptide amphiphile Hydrogel hMSC TGF�-1 Rabbit/full thicknesschondral defect

Sustained release ofTGF�-1 from peptidehydrogel Promotedregeneration of articularcartilage

[109]

HA, Hyaluronic acid; HPMC, hydroxypropyl methylcellulose; PLGA, poly(lactic acid-co-glycolic acid); PEG, poly(ethylene glycol); BMSC, bone-marrowstromal cells; ADSC, adipose-derived stem cells; ESC, embryonic stem cell; FGF, fibroblast growth factor; IGF-1, insulin-like growth factor-I; GAG, gly-cosaminoglycan; hMSC, human mesenchymal stem cells.

2 Polymer

cIcdocgocaawpctcscccsITilc

tondaeagmdflrtthoatlnt

ewwdWsesIso


ells were cultured [89]. Artificial ECMs formed from typeI collagen were prepared for neo-cartilage synthesis byrosslinking with genipin, because ECM in cartilage is pre-ominantly composed of type II collagen and networksf proteoglycans (PG), such as hyaluronic acid (HA) andhondroitin sulfate (CS). Their study assessed the effect oflycosaminoglycans (GAGs) added to the culture mediumn proliferation and matrix synthesis of human chondro-ytes [90]. The results indicated that the addition of CSnd HA further up-regulated gene expression of aggrecannd collagen II in chondrocytes. In an animal experimentith sheep, cultured autologous chondrocytes seeded in aorcine collagen I-III bilayer membrane were implanted inhondral defects [91]. Microfracture treatment was usedo enhance the healing response, and it was reported thatombination treatment of microfracture and chondrocyte-eeded collagen membranes further improved the overallartilage repair. A clinical trial also was performed toheck the efficacy of matrix-induced autologous chondro-yte implantation, in which autologous chondrocytes wereeeded on three-dimensional membranes of a bilayer type-III collagen for implantation in deep cartilage defects [92].he presence of hyaline-like cartilage was observed at themplantation site via magnetic resonance, and chondrob-asts and type II collagen also were observed inside theollagen membrane.

The fibrin artificial ECM is a network of fibrous pro-eins naturally polymerized from fibrinogen in the processf blood clotting in response to injury, and it forms aatural, artificial ECM mesh. It was reported that this three-imensional structure of fibrin can be used as a gel-typertificial ECM to encapsulate cells for delivery [86]. Sev-ral animal experiments were performed using fibrin gluertificial ECMs to evaluate the therapeutic effects of autolo-ous chondrocytes, mesenchymal stem cells, and bioactiveolecules on cartilage repair. For autologous adipose-

erived stem cells (ADSCs) loaded in a fibrin artificial ECMor implantation in the treatment of full-thickness carti-age defects of rabbits, immunostaining, Western blotting,everse transcriptase polymerase chain reaction, and quan-itative assessment revealed that articular surface defects,reated with an ADSC-loaded artificial ECM, healed withyaline-like cartilage; however, very little healing wasbserved in the control group. Cartilage markers includingggrecan and collagen type II mRNA also were identified inhe treated group. However, concerns regarding immuno-ogical reactions have been raised; if autologous fibrin isot used, the method requires additional procedures andime necessary for blood collection [93].

Silk is an artificial ECM that has been studied by sev-ral groups because of its high mechanical strength in theet state and its low inflammatory properties [94]. It alsoas reported that fibroin, a primary component of silk,emonstrates good cell attachment and biocompatibility.ang et al. reported that human chondrocytes seeded in

ilk fibroin artificial ECMs could be re-differentiated after
xpansion on culturing dishes and deposited in cartilage-pecific ECM for in vitro cartilage tissue engineering [95].n another in vitro experiment, better-defined cartilage tis-ue with increased GAG amounts and cell density werebserved in silk fibroin sponge hydrogels compared to
Science 36 (2011) 238–268

that of collagen gel [96]. These characteristics also wereobserved with bone marrow-derived mesenchymal stemcells (BMSCs) seeded on silk fibroin artificial ECMs [97].In particular, BMSCs proliferated more rapidly on the silkfibroin artificial ECM, and homogeneous distribution ofcartilage-like tissue was achieved only with silk artificialECM samples among silk, collagen, and cross-linked colla-gen artificial ECMs. Use of exogenous silk protein for tissueengineering might be problematic due to its stimulationof immune responses, such as induction of significant TNFrelease from macrophages [98].

Hyaluronan is a primary physiological component ofECMs in articular cartilage, which can be chondrogenicto mesenchymal stem cells. In a transplantation studywith autologous BMSCs embedded in a hyaluronic acidsponge, BMSC-loaded artificial ECMs were implanted infull-thickness osteochondral defects of the rabbit knee [99].Histological findings revealed that newly formed carti-lage tissue at the implantation site was very similar tothe surrounding normal tissue, which is much better thanthat of the untreated group. It is usually chemically mod-ified for easy fabrication into solid artificial ECMs becauseof its highly hydroscopic nature [100]. The most com-mon form of the modification is the esterification of acarboxylic acid present at the C6 position of hyaluronicacid. The benzyl ester of hyaluronic acid has been com-mercialized as Hayff-11 and tested clinically. Autologouschondrocytes isolated from patients were seeded on Hyaff-11 and implanted in the chronic cartilage lesions of theirknees [101]. There were no major adverse events reportedduring three-year follow-up, and clinical outcomes wereprominent with the treated group. However, the chem-ical modification of hyaluronic acid might compromiseits biocompatible property. In addition, the degrada-tion of cross-linked hyaluronan caused chondrolysis[102].

Alginate is an anionic polysaccharide, which forms ahydrogel instantly in the presence of divalent cations, suchas calcium ions. Alginate beads encapsulating cells are com-monly produced when cells dispersed in alginate solutionare dropped into a calcium chloride solution. In an in vitrostudy using rabbit BMSCs, it was reported that alginate andagarose gels containing BMSCs induced a greater increasein the expression of cartilage-specific markers, such asaggrecan and type II collagen, than cells in type I collagengels [103]. Furthermore, when rabbit BMSCs encapsulatedin alginate beads were deployed in the cartilage defects ofrabbit, the beads remained satisfactorily within the defectregions, which were progressively replaced by the regen-erating tissue. Histological findings showed that viablechondrogenic cells filled in the defects. In a clinical studyfollowing 17 patients (inclusion criteria: an isolated lesionof the femoral condyle (grades III and IV)), significantimprovement was observed in patients with lesions largerthan 3 cm2 when autologous chondrocytes were isolatedand suspended in an alginate-agarose mixture solution and
subsequently implanted in the lesions of patients [104].However, alginate gel suffers from instability in physio-logical solution due to a loss in mechanical strength andintegrity induced by the replacement of divalent calciumions with monovalent sodium or potassium ions.

Polymer Science 36 (2011) 238–268 251

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Chitosan also is preferred in articular tissue engineer-ing applications because of its structural similarity toGAG components in cartilage-stimulating chondrogenesis[105]. In a study characterizing cartilage repair in sheep,a chitosan-glycerol phosphate solution was mixed withautologous whole blood from sheep to facilitate the healingprocess in cartilage defects with blood clots [106]. Inter-estingly, the mixture of chitosan-glycerol phosphate andwhole blood showed increased adhesion to walls of thecartilage defect region and significant repair in hyaline car-tilage at a 6-month follow-up. In a subsequent study with arabbit model, it was reported that beneficial blood clot for-mation could be stabilized by chitosan-glycerol phosphatesolution via inhibition of clot retraction [107]. Since chi-tosan possesses primary amino groups at the C2 position ofeach repeating unit, which is only protonable below pH 6.5,this unique characteristic allows for convenient processingin artificial ECMs under mild conditions. The polycationicproperty of chitosan can be exploited to form a more stableionic complex with polyanionic macromolecules includ-ing alginate, GAGs, synthetic poly(acrylic acid), and smalleranionic molecules. Chitosan microspheres were fabricatedby drop-wise addition of anionic tripolyphosphate in achitosan emulsion containing TGF-�1 for the tissue engi-neering of articular cartilage [108]. Controlled release ofa bioactive growth factor was achieved from ionicallycrosslinked chitosan microspheres, and in vitro culture ofporcine chondrocytes on the microspheres demonstrateda significant increase in cell proliferation and ECM produc-tion.

Recently, a novel peptide amphiphile (PA) with a TGF-�1-binding peptide domain was explored for cartilageregeneration applications [109]. This self-assembling PAformed supramolecular nanofibers upon the addition ofcharge-shielding ions, which resulted in a hydrogel dis-playing a high density of exposed TGF-�1-binding peptidefor efficient capture and release of TGF-�1 growth fac-tor. When human mesenchymal stem cells (hMSCs) weresuspended in PA solution and injected in a full-thickness,articular cartilage rabbit model with microfracture, thedata demonstrated improved survival and increased chon-drogenic differentiation of hMSCs. Furthermore, PA itself,even without TGF-�1 growth factor, promoted the regen-eration of articular cartilage in a full-thickness chondraldefect, suggesting the possibility that TGF-�1-bindingdomains in PA might serve as a reservoir for attracting andreleasing the endogenous TGF-�1.

3.2. Synthetic material-based artificial ECMs

Synthetic materials including poly(lactic acid) (PLA),poly(glycolic acid) (PGA), and their copolymer, poly(lacticacid-co-glycolic acid) (PLGA), have been tested for cartilagetissue engineering potential (Table 6). They are thought tohave advantages including ease in molding and the abilityto design a degradation rate to match tissue growth into the
artificial ECM. However, these synthetic materials are notas preferred for cartilage tissue engineering as materialsof natural origin, because the acidic byproducts gener-ated during the degradation process cause an inflammationreaction, giant cell reaction, and acute chondrocyte death Ta
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ue to the abrupt drop in pH in the local microenviron-ent inside the engineered cartilage. The closed cartilage

nvironment deters rapid clearance of acidic byproductsrom the degrading synthetic artificial ECMs, eventuallynducing undesirable side reactions in cartilage. Anotherisadvantage of these materials is their poor cell attach-ent. Commonly, these materials were modified to possess

n anchorage site for cell adhesion or mixed with otheraturally-derived materials [110].

.2.1. Poly(ethylene glycol) (PEG) hydrogelsRecently, hydrogel-based artificial ECMs have gained

reater attention for cartilage tissue engineering applica-ions because of their similarity to the natural cartilagenvironment. Hydrogels contain high water content sim-lar to that of natural cartilage, which serves as auitable environment for chondrocytes. These materialsre composed of synthetic or natural-based hydrophiliciomaterials cross-linked by physical, ionic, or chemical

nteractions. They also can be injected transcutaneouslynto the defect region of the joint, which avoids invasiveurgery required for the implantation of prefabricated arti-cial ECMs. PEG is a popular biocompatible hydrophilicolymer approved by the FDA, and it has been exten-ively explored for the use of formulating hydrogels toncapsulate bioactive drugs or cells. It is thought thatrosslinked PEG hydrogel may provide a better environ-ent for culturing chondrocytes due to its high water

ontent and mechanical strength, because chondrocytesre surrounded by hydrophilic ECM components in highbundance. In one approach to minimize invasiveness dur-ng treatment of the cartilage defect while maintainingts mechanical and biological properties, photopolymeriza-ion was employed using a PEG macromer polymerizable atoth ends and a photoinitiator, followed by exposure of theolution to UV light [111]. Photopolymerized PEG hydrogels an attractive material for tissue engineering applicationsecause of its cell-friendly properties and in situ gelationpon exposure to UV. In this study, a PEG macromer wasixed with isolated allogenic chondrocytes in the pres-

nce of a photoinitiator. The photopolymerization reactionan be performed at physiological pH and temperature,hich allows for safe encapsulation of cells and proteins.hondrocyte viability in this hydrogel was maintainedith characteristic gene expression for up to three days

f in vitro culture. In another study, cell binding peptideas conjugated to one end of a bifunctional PEG macromer

nd mixed with PEG diacrylate in the presence of a pho-oinitiator for crosslinking. When BMSCs and embryonictem cell-derived cells (ESCs) were encapsulated in thisydrogel with a mechanical stimulus, gene expression ofartilage-related markers, such as Sox-9, type II collagen,nd aggrecan, was noticed [112]. However, it was deter-ined that highly crosslinked PEG hydrogel might hinder

he proliferation and proteoglycan synthesis of encapsu-ated chondrocytes [113,114]. As the seeded cells grow and
orm new tissue inside the hydrogel, the scaffold shouldegrade accordingly. The effect of the degradation rate ofydrogel on the engineered cartilage formation was evalu-ted using hydrogels of different degradation rates, whichere prepared by copolymerizing a degradable PLA-b-PEG-
Science 36 (2011) 238–268

b-PLA macromer with non-degradable PEG macromers atpre-determined mixing ratios [115]. It was found thatchondrocytes encapsulated in a PEG hydrogel with a higherdegradation rate produced more DNA content and type IIcollagen, indicating the important role of hydrogel degra-dation in controlling and influencing the deposition anddistribution of ECM molecules. The photopolymerized PEGhydrogel also can be functionalized for cellular anchorageby incorporating a cell adhesion peptide, such as Arg-Gly-Asp (RGD), at PEG chain terminals during polymerization[116].

3.2.2. Injectable hydrogelsIn situ injectable hydrogel systems have generated

increased interest for cartilage repair applications. Theycan be injected with encapsulated cells and/or bioactivematerials of interest into the cartilage defect in a minimallyinvasive manner and easily can fill the three-dimensionalshape of the defect for facilitated integration with the addi-tional mechanical property of temperature-dependent sol-gel transition. Poly(N-isopropylacrylamide) (PNIPAAm)exhibits reversible phase separation with a lower criticalsolution temperature (LCST) of approximately 32 ◦C. Thus,chondrocytes cells can be dispersed in PNIPAAm solutionat room temperature (RT) lower than LCST and injectedinto the cartilage defect for in situ gelation. When iso-lated bovine articular chondrocytes were seeded into theloosely crosslinked copolymer hydrogel of NIPAAm andacrylic acid, p(NIPAAm-co-AAc), and incubated at 37 ◦C invitro, the viability of seeded chondrocytes was maintainedfor up to four weeks, and cartilage-like tissue formationwas observed in the matrices [117]. In an in vivo studyusing this thermo-reversible hydrogel, a larger volume ofcartilage-associated ECM was observed with the addition ofdexamethasone, ascorbate, and TGF �-3 eight weeks aftersubcutaneous implantation in nude mice [118]. However,the temperature-dependent phase transition might hinderthe thermo-responsive gel from being injected through along needle or catheter to the desired site in the humanbody, because the warmed needle or catheter can even-tually cause premature gelation during injection beforereaching the target site. In another approach to inhibitthis pre-mature gelation, a p(NIPAAm-co- vinylimidazole)[p(NIPAAm-co-VI)] copolymer hydrogel that responds toboth temperature and ionic strength was used [119]. Rabbitchondrocytes were embedded in this composite hydrogelwith TGF �-1-loaded nanoparticles. Chondrocyte-specificECMs were observed eight weeks after subcutaneousimplantation of the hydrogel in nude mice.

Pluronic F-127 consists of approximately 70% ethyleneoxide and 30% propylene oxide by weight and is avail-able commercially [120]. It transforms to hydrogel at RTat concentrations of 20% or higher. After administration invivo, it can be slowly dissolved and cleared by renal andbiliary excretion. Although several trials have been con-ducted using Pluronic F-127, poor mechanical properties
hinder its application for tissue engineering due to a diffi-culty maintaining the constructed shape. Thus, compositeartificial ECMs that compensate for the poor mechanicalproperties of Pluronic F-127 have been preferred, whichwill be described later.

Polymer
3.3. Composite artificial ECMs

A few specific characteristics of single ECM componentsmight be insufficient to create the optimal environmentto mimic the natural proliferation and differentiationof chondrocytes in cartilage. Therefore, in most casesa combination of multiple components is desired toaddress various features required for culturing chondro-cytes. Most synthetic material-based artificial ECMs sufferfrom poor anchorage of chondrocytes or stem cells, andincorporation of natural ECMs including collagen, fib-rin, and hyaluronic acid in synthetic artificial ECMs forcell attachment has been a popular approach studiedextensively. Synthetic materials provide relatively highmechanical strength with a tunable degradation rate,whereas their hydrophobicity and lack of cellular anchor-age sites are drawbacks for their application in tissueengineering. Conversely, naturally-derived ECM polymerssupport excellent cellular adhesion and growth due totheir specific cell interaction peptides and hydrophilic-ity, even though their weak mechanical properties makeit difficult to use them in load-bearing regions, suchas cartilage. Thus, these characteristics of naturally- orsynthetically-originated materials drive us to attempt acombination of the two materials in order to affordhigher mechanical strength, tunable degradation, and cel-lular attachment [121]. When collagen microsponges wereembedded in a PLGA mesh and transplanted subcuta-neously into nude mice for chondrocyte tissue engineering,homogeneous cell distribution, natural chondrocyte mor-phology, and abundant cartilaginous ECM deposition wereobserved. As a biomimetic approach, the active cellanchorage domain of collagen (e.g. RGD) was conjugatedchemically onto the surface of synthetic material-basedartificial ECMs [122]. Chondrocytes seeded on RGD-modified PLGA/gelatin microspheres demonstrated thebest attachment, proliferation, viability, and sulfated GAGsecretion.

Another purpose of hybrid artificial ECMs is to incor-porate thermally or chemically responsive componentsin natural or synthetic material-based artificial ECMs toincrease the bioavailability of seeded cells while min-imizing their leakage out of the artificial ECMs. Fibrinpossesses a chemically active gelling property, which typ-ically occurs during blood coagulation in addition to theadvantages described previously. Several studies investi-gating a fibrin-based hybrid artificial ECM demonstratedits potential for the promotion of homogeneous cell dis-tribution and cartilaginous tissue formation [110,123,124].PNIPAAm also was used for this purpose, and it is asynthetic thermoresponsive polymer for tissue engineer-ing with versatile manipulation of its chemical structureand molecular weight as well as easy conjugation withother components. Chondrocytes could be seeded easilyinto the PLA-PNIPAAm hybrid artificial ECMs at RT, atwhich the hydrodynamic structure of PNIPAAm was shrunk
to create larger pore sizes for facile cell loading [125].Upon incubation at 37 ◦C, the PNIPAAm was rehydratedto provide a more hydrophilic and chondrocyte-friendlymicroenvironment, and its swelling property at physio-logical temperature in PBS solution was maintained for
Science 36 (2011) 238–268 253

4 weeks. Chondrocytes cultured in this hybrid artificialECM could be harvested easily from the artificial ECM bysimply lowering the temperature to RT without any addi-tional treatment by proteolytic enzymes, and the harvestedchondrocytes exhibited a round morphology, indicative ofmaintenance of the chondrocyte phenotype during culturein the hybrid artificial ECM.

4. Bone

4.1. Artificial ECMs for bone tissue engineering

4.1.1. Design criteria of artificial ECMs for bone tissueengineering

A key component in bone tissue engineering is an arti-ficial ECM that serves as a template for cell interactionsand the formation of bone ECM to provide structural sup-port for the newly formed tissue. The artificial ECMs forbone tissue engineering should meet several criteria toserve this function including possession of mechanicalproperties similar to those of the bone repair site, biocom-patibility, biodegradability at a rate commensurate withremodeling, and porosity of artificial ECMs that allowsmigration and proliferation of osteoblasts and mesenchy-mal cells as well as vascularization. In addition to thesecriteria, artificial ECMs should be able to regulate cellu-lar functions even though several reviews have focusedon artificial ECM properties including mechanical proper-ties [126,127], degradation [128,129], porosity [130], andosteoconductive activity of the artificial ECM [131,132] forbone tissue engineering. In this section, we reviewed stud-ies that focused on the regulation of cellular function byartificial ECMs after the introduction of materials used inbone tissue engineering. In addition to polymeric materi-als, non-polymeric materials were also used to fabricateartificial ECMs for bone tissue engineering because bone isa hard tissue.

4.1.2. Non-polymeric materialsNon-polymeric materials including metallic and

ceramic materials have been widely used for bone tis-sue engineering. Generally, the mechanical propertiesof metallic materials (e.g. stainless steel, cobalt-chromebased alloys, titanium (Ti) alloys) are superior to those ofpolymeric ones. However, there are several disadvantages,such as the lack of tissue adherence, low rate of degra-dation, toxicity due to accumulation of metal ions dueto corrosion, and mismatch of Young’s moduli betweenmetallic materials and bone [133–135]. Furthermore, theTi alloys are not osteoconductive [136]. Ceramic materialssuch as calcium phosphate also have been used exten-sively in bone tissue engineering because they are naturalcomponents of bone tissue [137–140]. Calcium phosphatecan be used in permanent implants or as biodegradableartificial ECMs due to its osteoconductive and osteoin-
ductive properties as well as mechanical compatibilitywith native bone. Although calcium phosphate has severalattractive properties, its mechanical properties are poor[126]. Therefore, its clinical use is limited to non-weightbearing sites.

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.1.3. Synthetic polymeric materialsCompared to metallic and ceramic materials, polymeric

aterials are materials in which it is easy to chemi-ally incorporate moieties that regulate cellular functions.lso, synthetic polymers can be produced under con-

rolled conditions and, therefore, exhibit predictable andeproducible mechanical and physical properties, suchs tensile strength, elastic modulus, and degradationate. Moreover, three-dimensional, biodegradable, syn-hetic polymeric systems are of particular interest becauseheir porosity, hydrophilicity, and degradation time can bearied with a high degree of reproducibility. For bone tissuengineering, there are several commonly used syntheticolymers, such as PLA, PGA, poly(�-caprolactone) (PCL),nd PLGA [100,140–143]. To fabricate these polymersnto three-dimensional artificial ECMs, there are severalechniques including salt-leaching, gas foaming, electro-pinning, and emulsion polymerization [130]. PLGA-basedrtificial ECMs fabricated via consolidation by pressurerop have been used as teeth implants [141]. PLGA/polyvinyl alcohol) blend fabricated via a salt-leaching tech-ique were used as an artificial ECM for skull bone defectreatment [142] even though they have been applied inreas of low mechanical stress in vivo because of their lowechanical properties [143].

.1.4. Naturally-derived polymersNaturally-derived polymers (e.g. collagen, hyaluronic

cid, glycosaminoglycans) have the advantages of biocom-atibility and biodegradability because they are composedf the structural materials of tissues. Kim et al. usedyaluronic acid-based hydrogels containing bone morpho-enetic protein-2 (BMP-2) and human mesenchymal stemells to treat a rat calvarial defect model [144]. Collagen-ased artificial ECMs also have been used for bone tissuengineering [145,146]. Type I collagen is a major com-onents of in vivo bone ECM and leads to new bonerom stem/progenitor cells via the developmental cascade.herefore, type I collagen might be a good candidate mate-ial for a biomimetic approach to design artificial ECMs forone tissue engineering. In addition to hyaluronic acid andype I collagen, silk fibroin also has been used as the arti-cial ECM in bone tissue engineering [147]. Salt-leachingechniques and freeze-drying techniques are commonlysed to fabricate these materials. Since they are generallyechanically weak, it might be difficult for natural materi-

ls to be used as the bone tissue engineering artificial ECMsy themselves even though they have high biocompatibil-

ty and biodegradability.

.1.5. Composite materialsEven though each individual material has advantages

or bone tissue engineering applications, each materialiscussed has disadvantages in various properties includ-

ng the brittleness of calcium phosphate and the inferiorechanical properties of natural and synthetic polymers.

he combination of different materials to form compos-tes has led to overcoming the disadvantages of any onearticular material. For example, to develop Ti-based arti-cial ECMs with osteoconductive activity, artificial ECMurface was coated with a hydroxyapatite layer which

Science 36 (2011) 238–268

improved interactions with osteoblastic cells and pro-moted osteogenic activity (e.g. osteocalcin expression) invitro compared to that of a non-coated Ti-based artificialECM [148]. Brittleness of calcium phosphate is a majorproblem for its implementation as an artificial ECM forbone tissue engineering. A hydroxyapatite-based artifi-cial ECM was coated with a hydroxyapatite/PCL compositeand demonstrated increased compressive strength andelastic modulus compared to an uncoated hydroxyapatite-based ECM [149]. In addition, a collagen/hydroxyapatitecomposite artificial ECM was developed to support theattachment and proliferation of rabbit periosteal cells[150]. Furthermore, a PLGA sponge hybridized with col-lagen microsponges on which apatite particulates weredeposited is an example of a three-component artificialECM for bone tissue engineering [151].

4.2. Regulation of cellular functions by artificial ECMs

4.2.1. Importance of regulation of cellular function byartificial ECMs

When materials were used for bone regeneration, thesole criterion was “to achieve a suitable combination ofphysical properties to match those of the replaced bone tis-sue with a minimal toxic response of the host” [152,153].However, all bone tissue functions cannot be mimicked bythe materials currently used for bone regeneration becausethey are unstable and cannot work well in vivo after longperiods of implantation. To reconstruct new bone tissue,regulation of cellular function is necessary. There are sev-eral approaches, such as the application of growth factorsand mechanical stimuli that in turn regulate cellular func-tion. In addition, artificial ECMs play a central role in theregulation of cellular function because the artificial ECMsrespond to growth factors and transmit mechanical stim-uli. Therefore, the design of artificial ECMs will becomea primary issue for the regulation of cellular function forbone tissue engineering applications in the future. For bonetissue engineering applications using cells, osteoblastsand mesenchymal stem cells (MSCs) are important cellsources. In this section, we primarily discussed materi-als for the regulation of osteoblast function. The materialsfor bone tissue engineering using MSCs are reviewed inSection 6.

4.2.2. Regulation of osteoblast functionsTo regulate cellular function using artificial matri-

ces, mimicking native ECM is a well established, reliableapproach. To mimic native ECM in bone, ECM proteinssecreted from osteoblasts or MSCs are used for theregulation of cell function [154–158]. Using these mate-rials, it is possible to regulate cellular functions, such asproliferation and differentiation. To mimic native boneECM more efficiently, fragments of the ECM proteins areincorporated within polymeric materials. Type I collagenhas been used widely as an artificial ECM, mimicking
in vivo ECM in bone tissue engineering applications. Theosteoblasts cultured on type I collagen matrix expressedan osteoblastic phenotype at higher levels than those cellson plastic surfaces [159]. These cells recognize type I col-lagen by the cell surface receptor, integrin �2�1 [160],

Polymer

as a barrier or substratum for skin cells. The artificial ECM


where the integrin receptor activated intracellular sig-naling pathways to express the osteoblastic phenotype.To confirm this mechanism, synthetic peptides also havebeen used. When integrins bind to the ECM molecules,integrins recognize the biologically significant peptideregion, Arg-Gly-Asp (RGD). Hu et al. developed an RGDpeptide-modified PLA film [161] on which they demon-strated enhanced osteoblast attachment. Also, alkalinephosphatase activity and calcium deposition increased onthis film, indicating that the osteogenic phenotype wasexpressed. In addition to the RGD peptide, a collagenpeptide motif also was applied to the osteoblast culture[162]. The ECM also regulated the activity of cytokinesby increasing their accessibility to receptors [155,156].Therefore, the incorporation of cytokines is one of themost useful approaches for regulating cellular function[163,164].

4.3. Artificial ECMs for osteochondral defects

Osteochondral defects are often associated withmechanical instability of the joint and, therefore, withthe risk of causing osteoarthritic degenerative changes.Grafting of osteochondral units with both chondral andosseous layers represents a promising approach to restorethe biological and mechanical functionalities of the joint.To transplant osteochondral units, autologous osteochon-dral grafts (i.e. mosaicplasty technique) are currently usedclinically; however, this technique suffers from severallimitations, such as lack of availability of materials, donorsite morbidity, and difficulty in matching the topology ofthe grafts with the injured site. Therefore, osteochondralcomposites have been developed to overcome theselimitations. Although homogenous artificial ECMs wereoften fabricated for the reconstruction of osteochondralcomposites [165,166], the materials were designed to pro-duce cartilaginous and bone parts separately [167–170].Because cartilage and bone are different tissues, the mate-rial design should be optimized accordingly for each tissue.To fabricate osteochondral composites, two approacheshave been utilized: (1) cartilage and bone tissues areindependently reconstructed and combined into a singlecomposite graft by suturing or adhering them together,and (2) cartilage and bone tissues are reconstructed inmulti-phasic artificial ECMs.

For fabricating a single phase of each tissue layerfollowed by reconstruction in osteochondral composites,several approaches that were optimized for each tissuehave been utilized. These materials have been discussedin detail previously. For example, periosteal-derived cellswere cultured in a PLGA and PEG hybrid artificial ECM togenerate reconstructed bone. The cells in the PLGA/PEGartificial ECM deposited mineralized matrix and bonespecific proteins (e.g. osteocalcin, osteopontin). The engi-neered bone was sutured together with reconstructed
PGA-based engineered cartilage seeded with articularchondrocytes [167]. Similarly, collagen-hydroxyapatitehybrid sponges were used as the bone component of osteo-chondral composites, and the artificial ECMs were suturedto the engineered cartilage [168].
Science 36 (2011) 238–268 255

In addition to the physical combination of indepen-dently engineered cartilage and bone tissues, reconstruc-tion of osteochondral composites in a multi-phasic artificialECM has been explored. Scheck et al. developed a biphasicartificial ECM composed of PLA and HA. At the interfacebetween PLA and HA, a thin PGA film was deposited to pre-vent cell migration between the components [169]. Porcinearticular chondrocytes were seeded in the PLA region ofthis artificial ECM (cartilage layer), while human gingi-val fibroblasts transduced with an adenovirus expressingBMP-7 were seeded in the HA region of the artificial ECM(bone layer). Following ectopic implantation in mice, thecartilage layer contained ECM rich in glycosaminoglycans.In contrast, blood vessels were observed in bone layer.Moreover, a mineralized interface was often present at thejunction. Kon et al. also investigated multilayered gradientartificial ECMs composed of HA and type I collagen [170].They divided the artificial ECM into three parts (cartilage,transition, and bone regions). According to their results,the compositions of type I collagen and HA varied in eachregion (cartilage region: 100% type I collagen; transitionregion: 40% HA and 60% type I collagen; and bone region:70% HA and 30% type I collagen). In this case, chondrocyteswere seeded only in the cartilage region and were trans-planted in the osteochondral defect model. The regenera-tion of bone and cartilage was significantly enhanced withthe artificial ECMs compared to that of the control group.

In summary, the incorporation of bioactive molecules isa good approach for regulating cellular function. The opti-mization of physical properties (e.g. elasticity, porosity ofthree-dimensional artificial ECMs) also is required to regu-late cellular function precisely for bone tissue engineeringapplications.

5. Skin

Several tissue-engineered skin products are commer-cially available. The application of tissue-engineered skin isexpanding from laboratory use (e.g. skin biology research,drug evaluation) to clinical use [171]. Although tissue-engineered skins are commercially available, skin tissueengineering research should progress to include additionalapplications. Several types of artificial ECMs composedof naturally-derived ECM proteins and synthetic poly-meric materials have been used to develop engineered skin[172–179].

Skin tissue is divided into two parts: epidermis anddermis. In the epidermis, keratinocytes are the most com-mon cell type and form a surface barrier layer. In thedermis, fibroblasts exist in skin ECM and provide strengthand resilience. Most tissue-engineered skins are created byexpanding skin cells and using them to restore barrier func-tion or to initiate wound healing. Many tissue engineeredskin dermal replacement products contain artificial ECMs

must eventually be discarded or replaced by live skin cellsfor long-term healing, but artificial ECM can be used tem-porarily to provide a barrier or a substratum for the cells.Here, we summarized artificial ECMs used in the skin tissueengineering field.


Table 7Polymers used in skin tissue engineering.

Polymers Name in market Application Reference

Natural-derived polymers Collagen Integra Dermal replacement [172]Collagen Apligraf Epidermal/dermal replacement [173]Collagen Cincinnati skin substrate Epidermal/dermal replacement [174]Fibrin Bioseed-S Epithelial cover [175]Hyaluronic acid Hyalograft-3D Dermal replacement [175]Acellular matrix Alloderm Dermal replacement [176]

y availaby availab

5

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5

s(csbei

5

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6

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Polyglactin910 Dermagraft

Synthetic polymers PCL Not commerciallPLGA Not commerciall

.1. Naturally-derived materials

For dermal replacement, several naturally-derived ECMroteins have been used as artificial ECMs for skin tis-ue engineering. Type I collagen has been used extensivelyn this application since is a major component of der-

is ECM. Integra (produced by Integra Life Sciences) is anxample of type I collagen-based artificial ECM for skinissue engineering [172]. Type I collagen-based artificialCM also has been developed to replace epidermis andermis [173,174]. Apligraf (produced by Organogenesis)

s composed of keratinocytes and fibroblasts with type Iollagen-based artificial ECM. In addition to type I colla-en, fibrin (BioSeed-S) and hyaluronic acid (Hyalograft-3D)ave been used as an artificial ECM for skin tissue engineer-

ng [175]. In addition, acellular dermis from cadaveric skinAlloderm) has been used as an artificial ECM to mimic then vivo ECM of skin tissue [176].

.2. Synthetic polymeric materials

Synthetic polymeric materials have also been used forkin tissue engineering applications. Bioresorbable PLGApolyglactin 910) is used in Dermagraft which is a commer-ially available skin substitute [177]. Furthermore, otherynthetic polymeric materials, such as PCL and PLGA, haveeen used for a long time as artificial ECM in skin tissuengineering [178,179]. Table 7 summarizes the character-stics of polymers used in skin tissue engineering.

.3. Future challenges in skin tissue engineering

Although tissue-engineered skins are commerciallyvailable, there are several challenging issues includinghe improvement of safety and angiogenesis. In additiono these issues, reconstruction of more complex structuresn skin is expected because in addition to epidermal andermal tissues there are hairs, sweat glands, and toucheceptors in native skin tissues. Therefore, it is necessary toevelop engineered skin with these structures in the future.

. Stem cells

Artificial ECMs can modulate stem cells’ behavior,ncluding adhesion, proliferation and differentiation. Sol-ble factors, cell adhesion signals and mechanical signalsan be incorporated into artificial ECMs for the control oftem cell behavior. The chemical and physical properties

Dermal replacement [177]

le Mechanical support to type I collagen scaffold [178]le Dermal replacement [179]

of artificial ECMs, such as material chemistry, architectureand mechanical properties can also modulate stem cells’behavior, which is summarized in Table 8.

6.1. Delivery of soluble factors using artificial ECMs

Soluble factors, including growth factors and hormones,play a significant role in the modulation of stem cell behav-ior. Incorporation or local delivery of certain growth factorsusing artificial ECMs may increase the accessibility of saidfactors to seeded stem cells as well as prevent growth fac-tor loss due to diffusion, which would be an effective wayof directing stem cell behavior. Artificial ECMs can retainand present growth factors via direct entrapment, affinityand covalent tethering.

Growth factors can be directly entrapped in artificialECMs. Moreover, the release of growth factors directlyentrapped in artificial ECMs or delivery systems is generallycontrolled via diffusion through or degradation of deliverysystems, or even a combination of these two mechanisms[180,181]. One example of stem cell control by the growthfactor delivery is that vascular endothelial growth factor(VEGF) delivery using a hydrogel can induce vascular differ-entiation of human embryonic stem cells (ESCs) entrappedin the hydrogel [182]. Another example is that insulin-like growth factor-1 and transforming growth factor-�1(TGF-�1) released from gelatin microparticles entrapped inoligo[poly(ethylene glycol)fumarate] hydrogel can inducechondrogenic differentiation of mesenchymal stem cells(MSCs) entrapped in the hydrogel [183].

An affinity-based delivery system can release growthfactors retained in a delivery system via non-covalentinteractions. Heparin-conjugated artificial ECMs have beendeveloped for the local release of growth factors withaffinity to heparin. The delivery system is based on theinherent capacity of heparin to bind to growth factors suchas fibroblast growth factor 2 (FGF2), VEGF, platelet-derivedgrowth factor (PDGF), TGF-� and bone morphogeneticprotein (BMP) via electrostatic interactions between thenegatively-charged sulfate groups of heparin and thepositively-charged amino acid residues of the growth fac-tor proteins [184,185]. Such a system can deliver FGF2,which allows long-term self-renewal [186] and pluripo-
tency maintenance [187] of human ESCs. Heparin was alsoconjugated to poly(lactic-co-glycolic acid) (PLGA) scaffoldsfor BMP-2 delivery, creating a system capable of potentiat-ing the osteogenic efficacy of BMP-2 [188]. Local deliveryof BMP-2 using this system induced in situ osteogenic dif-


Table 8Modulation of stem cells’ behavior with interactions with signals from artificial ECMs.

Parameter Artificial ECM Stem cell Stem cell’s behavior References

Soluble factorVEGF Dextran hydrogel Human ESCs Vascular cell differentiation [182]IGF-1/TGF-�1 Gelatin/PEG-fumarate

hydrogelRabbit MSCs Chondrogenic differentiation [183]

BMP-2 Heparin-conjugated PLGAscaffold

Human MSCs Osteogenic differentiation [189]

EGF PMMA-g-PEO Human MSCs Promotion of survival [190]

Cell adhesion factorRGD peptide PEG diacrylate hydrogel Human MSCs Osteogenic differentiation [192]RGD peptide PEG-fumarate hydrogel Rat MSCs Osteogenic differentiation [193]RGD peptide PEG diacrylate hydrogel Human ESCs Chondrogenic differentiation [195]RGD peptide P(NIPAAm-co-AAc) Human ESCs Maintenance of pluripotency [197]IKVAV peptide (laminin epitope) Self-assembling peptide

nanofibersMouse neuralprogenitor cells

Neuronal differentiation [215]

Fibronectin PEG diacrylate hydrogel Mouse MSCs Osteogenic differentiation [198]Fibronectin PET Human cord blood

CD34+ cellsPromotion of proliferation [199]

Material chemistryHyaluronic acid Hyaluronic acid/PEG diacrylate

hydrogelGoat MSCs Chondrogenic differentiation [200]

Hyaluronic acid Hyaluronic acid hydrogel Sheep MSCs Ligament protein expression [201]Hyaluronic acid Hyaluronic acid hydrogel Human ESCs Maintenance of pluripotency [202]Collagen type II Collagen type II Human MSCs Chondrogenic differentiation [203]Apatite Apatite-coated PLGA scaffold Human ESCs Osteogenic differentiation [205]Apatite Apatite-coated PLGA scaffold Human MSCs Osteogenic differentiation [208]Hydroxypatite Hydroxypatite/Self-assembling

peptide hydrogelMouse ESCs Osteogenic differentiation [209]

Material architectureNanofiber Polyamide nanofiber scaffold Mouse ESCs Promotion of proliferation [213]Nanofiber Titanium oxide nanotube

scaffoldHuman MSCs Osteogenic differentiation [214]

3D hydrogel PEG hydrogel Mouse ESCs Chondrogenic differentiation [216]

Material mechanicsSoft matrix Polyacrylamide Human MSCs Neuronal differentiation [229]Rigid matrix Polyacrylamide Human MSCs Osteogenic differentiation [229]Soft gel PEG-silica gel Human MSCs Neuronal differentiation [236]Rigid gel PEG-silica gel Human MSCs Osteogenic differentiation [236]

Cell shape and sizeSpread and large Fibronectin-coated polystyrene Human MSCs Osteogenic differentiation [240]Round and small Fibronectin-coated polystyrene Human MSCs Adipogenic differentiation [240]Round Tissue adhesive-coated

polystyreneMouse embryonicmesenchymal cells

Undifferentiation [241]

Elongated Tissue adhesive-coatedpolystyrene

Mouse embryonicmesenchymal cells

Smooth muscle cell differentiation [241]

Flower shape Fibronectin-coated surface Human MSCs Adipogenic differentiation [242]Star shape Fibronectin-coated surface Human MSCs Osteogenic differentiation [242]

Mechanical signalShear force Polyurethane Mouse ESCs Vascular cell differentiation [245]Cyclic strain Collagen type I-coated surface Mouse embryonic

mesenchymal cellsSmooth muscle cell differentiation [246]

BovinGoat MMouse

Multidimensional strain Collagen type I gelCyclic compression PEG diacrylate hydrogelCyclic strain Poly(lactide-co-caprolactone)

scaffold

ferentiation of human bone marrow MSCs transplanted inundifferentiated state [189].

Growth factors can be retained on artificial ECMs viacovalent tethering. Recently, a study showed that a growthfactor tethered to an artificial ECM exhibits higher activity
than soluble growth factor [190]. The authors hypothesizedthat the cellular response to tethered growth factor wouldbe different due to cellular binding to the growth factoron a culture surface, which would inhibit internalization ofthe growth factor into the cytoplasm. MSCs were cultured
e MSCs Ligament cell differentiation [248]SCs Chondrogenic differentiation [112]ESCs Cardiomyogenic differentiation [249]

on an epidermal growth factor (EGF)-tethered poly(methylmethacrylate)-graft-poly(ethylene oxide) (PMMA-g-PEO)surface or a PMMA-g-PEO surface with EGF supplementedto the culture medium. EGF is known to promote MSCproliferation and migration without affecting pluripotency
[191]. Surprisingly, tethered EGF was superior to EGF insolution for the promotion of MSCs spreading and survivalin the presence of Fas ligand, a potent cell death factor. Theauthors concluded that EGF-tethered artificial ECM mayoffer a protective advantage over MSCs seeded on artificial

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CM in vivo if the artificial ECM causes an acute inflamma-ory response detrimental to the MSCs.

.2. Cell adhesion signal presentation by artificial ECMs

Conjugation of cell adhesion peptides to artificial ECMsan promote stem cell differentiation. Previously, RGDeptides were covalently conjugated to PEG diacrylateydrogels using conventional N-hydroxysuccinimidyl-ster (NHS) chemistry. It was found that functionalizationf PEG diacrylate hydrogels promotes the osteogenic dif-erentiation of human MSCs [192]. Another study showedhat RGD peptide incorporation into oligo(PEG-fumarate)ydrogels promotes osteogenic differentiation of MSCs193]. The osteogenic differentiation of MSCs on peptide-

odified oligo(PEG-fumarate) hydrogels is dependent onhe concentration of incorporated peptide: the ALP activi-ies of MSCs in hydrogels modified with 2.0 and 1.0 �moleptide/g were significantly greater compared to MSCs inydrogels modified with 0.1 �mol peptides/g or no pep-ide. The behavior of bone marrow stromal osteoblastsultured in hydrogels modified with two types of pep-ides, rat osteopontin-derived peptide and RGD, wereompared previously [194]. The results showed that hydro-el modified with osteopontin-derived peptide exhibitsignificantly higher cell migration than RGD-modifiedydrogels. RGD peptide-conjugated PEG diacrylate hydro-els also promote the chondrogenic differentiation ofuman ESCs [195]. Compared to human ESCs in PEGiacrylate hydrogels, cells in RGD peptide-conjugated PEGiacrylate hydrogels exhibit enhanced gene expression ofartilage-specific markers (i.e., aggrecan and collagen typeI).

The incorporation of cell adhesion peptide into arti-cial ECMs may favorably influence stem cell behavior.rtificial ECM comprising a semi-interpenetrating poly-er network, or polymer hydrogel, was designed to mimic

he native ECM in terms of mechanical properties andell adhesion properties. Poly(N-isopropylacrylamide-co-crylic acid) [p(NIPAAm-co-AAc)] was loosely cross-linkedith an acrylated peptide, Gln-Pro-Gln-Gly-Leu-Ala-Lys-H2 (QPQGLAK-NH2), the sequence of which was designed

o be cleaved by matrix metalloproteinase-13 and otherollagenases [196]. The mechanical properties of the(NIPAAm-co-AAc) hydrogel were controlled by vary-

ng the density of the network cross-linker. To provideell adhesion properties to the hydrogel, the hydrogelas further interpenetrated by polyacrylic acid-graft-Ac-GGNGEPRGDTYRAY-NH2 [p(AAc)-g-RGD] linear polymerhains, which provide an active RGD motif [197]. HumanSCs cultured in this hydrogel showed increased adherencend viability, as well as the expression of markers of undif-erentiated human ESCs. In contrast, human ESCs culturedn gelatin-adsorbed polystyrene exhibited morphologiesf spontaneously differentiating cells.

In addition to cell adhesion peptide conjugation, cell
dhesion proteins such as fibronectin have been incor-orated into artificial ECMs made of synthetic polymers.he surfaces of photo-crosslinkable PEG diacrylate scaf-olds were modified by covalently conjugating fibronectinor the promotion of cell adhesion properties [198]. Mouse
Science 36 (2011) 238–268

MSCs cultured on the fibronectin-conjugated PEG diacry-late hydrogel also adhered to the hydrogel and underwentosteogenic differentiation, as indicated by matrix miner-alization [198]. Compared to free fibronectin, fibronectinconjugated to artificial ECM promotes cell proliferation[199]. On the other hand, conjugation of fibronectin topolyethylene terephthalate scaffolds results in higher pro-liferation of human umbilical cord blood CD34+ cellscompared to free, soluble fibronectin. This result could havebeen due to the prevention of fibronectin detachment fromthe scaffolds by covalent conjugation, compared to the highfrequency of fibronectin desorption that occurs with free,soluble fibronectin.

6.3. Chemistry of artificial ECMs

Hyaluronic acid is a polysaccharide found in manytissues and is known to influence stem cells’ behavior.Hyaluronic acid has been shown to induce chondrogenicdifferentiation of MSCs. Entrapment of hyaluronic acidin photo-crosslinkable PEG diacrylate hydrogel containingcultured MSCs enhances cartilage-specific gene expres-sion [200]. In addition, MSCs cultured in hyaluronic acidhydrogels express a variety of ligament proteins [201].Hyaluronic acid also influences ESC behavior; human ESCsencapsulated in hyaluronic acid hydrogels maintain theirpluripotency and undifferentiated state, whereas humanESCs encapsulated in hydrogels of dextran, a similarlystructured polysaccharide, undergo spontaneous differ-entiation [202]. Furthermore, simple addition of solublegrowth factors to these hydrogels triggers lineage-specificdifferentiation of the entrapped ESCs.

Collagen was also shown to favorably modulate stemcells’ behavior. Human MSCs were mixed with collagentype II isolated from bovine cartilage and then cultured inpellet form [203]. The cells on the collagen type II scaffoldunderwent chondrogenic differentiation into cartilage-specific ECMs. Chondrogenesis of MSCs is most likelymediated by interaction with collagen type II throughintegrin �1, since blocking of this integrin impeded chon-drogenesis. Implantation of type II collagen-cell constructsresulted in maintenance of chondrogenesis, inhibition ofendochondral ossification and prevention of vessel inva-sion in vivo.

Hydroxyapatite can promote bone formation by stemcells. Hydroxyapatite is known to enhance osteoblastic dif-ferentiation as well as osteoblast growth [204]. A studypreviously demonstrated that PLGA scaffolds incorporat-ing apatite on their surfaces can stimulate the regenerationof bone tissue in vivo by osteogenic cells derived fromhuman ESCs [205]. Efficient apatite incorporation intoPLGA scaffolds can be achieved through either gas foam-ing of PLGA-hydroxyapatite scaffolds [206] or biomimeticapatite-coating [207]. PLGA-hydroxyapatite compositescaffold exposed to high amounts of hydroxyapatite canbe fabricated by the gas-forming or particulate-leaching
method without the use of organic solvents. The sur-face of the PLGA-hydroxyapatite composite scaffold canfurther be coated with bone-like apatite by incubationin simulated body fluid [207]. The biomimetic apatite-coating process was accelerated by the introduction of

Polymer
hydroxyapatite nanoparticle nucleation sites to the PLGA-hydroxyapatite composite scaffold. Bone-like apatite onthe scaffold surface significantly enhances the in vivoosteogenic potential of the scaffold due to improved osteo-conductivity [207]. Apatite-coated PLGA films enhanceproliferation, alkaline phosphatase activity and osteocalcinproduction in human MSCs [208]. Further, the inclusion ofhydroxyapatite particles in self-assembling peptide hydro-gel promotes osteogenic differentiation of mouse ESCs[209].

6.4. Architecture of artificial ECMs

The macroscale, microscale and nanoscale structuresof artificial ECMs can determine the fate of stem cells.The macroscale and microscale ECM structures affectcell migration, nutrient and waste diffusion and bulkmechanical properties. Generally, artificial ECMs aremacroscopically designed to have physical properties(mechanical properties, pore size, porosity, and degrada-tion rate) for the regeneration of specific tissues or organs.The resultant artificial ECMs may lack the nanoscale fea-tures required for precise control of cell behavior and tissueformation. Therefore, the nanoscale design of artificialECMs is critical, since the nanoscale structure of the arti-ficial ECMs can provide a substantial cellular interface foreffective interactions with cells [210] as well as the modu-lation of cell behavior, including cell adhesion, biophysicalforce transmission, cell shape, cytoskeletal structure anddifferentiation [211,212].

Stem cells interact typically with nanoscaled ECMsin vivo, and this interaction may influence stem cell behav-iors. Natural ECMs are composed of a complex networkof ECM molecules with nanoscale and microscale dimen-sions, such as a nanofibrillar collagen fibril network. Toinvestigate whether or not the nanostructure of the artifi-cial ECMs influences stem cells’ behavior, 3D nanofibrillar,artificial ECMs were produced by depositing electrospunnanofibers composed of polyamide onto the surfaces ofglass coverslips [213]. Culture on the nanofibrillar matrixgreatly enhanced the proliferation and self-renewal ofmouse ESCs compared to culture without nanofibers [213].The dimensions of the nanotubular matrix significantlyinfluence stem cell behavior. Human MSCs were cul-tured on nanotubular-shaped titanium oxide matrices, andosteogenic differentiation of human MSCs was induced bycontrolling the diameter of the nanotubes in the absenceof osteogenic soluble factors [214]. As a result, nanotubeswith large diameters (70–100 nm) were found to promotecell elongation and osteogenic differentiation of MSCs,whereas nanotubes with small diameters (30 nm) elicit celladhesion only without noticeable differentiation.

Artificial nanofiber ECMs can be functionalized usingbioactive factors to enhance stem cell differentiation. Theaddition of hydroxyapatite particles into self-assemblingpeptide nanofiber matrices promotes osteogenic dif-
ferentiation of cultured mouse ESCs [214]. Further,incorporation of specific peptide sequences into self-assembling peptide nanofibers can promote neuronaldifferentiation of progenitor cells. Finally, incorporationof laminin epitope IKVAV into self-assembling peptide
Science 36 (2011) 238–268 259

nanofibers promote neuronal differentiation of neural pro-genitor cells [215].

Stem cell differentiation is also significantly affected bythe dimensions of the culture system, which may be dueto differences in cellular interactions between cells cul-tured in 2D and 3D systems. In 2D culture systems, cellsadhere to one side of the culture surface. In contrast, cellscultured in 3D hydrogels or scaffolds may attach to allsides of the hydrogels or scaffolds, to other cells, or evento secreted ECMs. ESC-derived EBs cultured in monolayerin the presence of chondrogenic soluble factors show lim-ited expression of chondrogenic markers (aggrecan and Sox9) [195,216]. However, when the EBs were cultured in 3DPEG hydrogels, the soluble factors dramatically enhancedchondrogenic marker expression. Therefore, a 3D cultureenvironment more closely mimics chondrogenesis in vivoand may be responsible for enhanced chondrogenic differ-entiation.

Nanofabrication techniques have been used to studystem cells’ behavior in response to nanopatterned artifi-cial ECMs. The nanopatterned surface of artificial ECMs canmodulate diverse cell behaviors in terms of cell adhesion,cell orientation, cell migration, cytoskeletal change andgene expression [210]. Nanoscale topographical featuresthat modulate cell behavior include scale (5 nm to microm-eter scale), order type (e.g., ridge, step, groove, pillar andpit) and symmetry (e.g., orthogonal or hexagonal packing ofnanopit) [217–223]. Increasing the nanoscale roughness ofpore walls of artificial ECMs has been shown to enhance celladhesion, proliferation and expression of ECM components[224]. Modulation of the nanoscale topography of cells maybe mediated by changes in the adsorption and conforma-tion of cell adhesion proteins, the type of integrin that bindsto cell adhesion proteins and, consequently, integrin signal-ing [222]. Nanoscale topography may also modulate focaladhesion of the cells as well as cytoskeletal organization,both of which affect intracellular signaling [217].

6.5. Mechanical properties of artificial ECMs

The cellular microenvironments within tissues ororgans can be physically diverse. The mechanical prop-erties of the artificial ECMs on which cells reside canalso vary as much as 300-fold from soft brain tissueto rigid, calcified bone [225]. Such variations in themechanical properties of ECMs within various tissuesare obvious during development [226,227], and suchproperties may play important roles in tissue or organdevelopment [228]. Muscle cells prefer a compliant matrixfor optimal contraction and relaxation, whereas bonecells prefer a rigid matrix on which they can miner-alize. While the effects of growth factors on stem cellcommitment have been well recognized, the importanceof mechanical properties of matrices as a regulator ofstem cell differentiation is just now being appreciated[229].

MSCs can sense the mechanical properties of an arti-ficial ECM and transduce this information as changes incytoskeletal organization or as stem cell commitment.Cells can feel the resistance to deformation of the artifi-cial ECM by adhering to and pulling against the artificial

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CM. Physical forces are generated by cells adhered to theatrix due to the sliding of myosin bundles along actin

laments. Actin structures are linked to focal adhesions,hich provide a pathway for the transmission of force

rom the cell to the matrix inside [230]. The tension gen-rated from the cell-matrix interaction induces changesn cytoskeletal organization [231] as well as cellular con-raction [232]. The biophysical cues are then convertednto biochemical signals via signaling molecules that act as

echano-transducers [233]. Finally, the biochemical sig-als commit the cell to a specific lineage [229].

The mechanical properties of a matrix or culture sur-ace can modulate stem cell differentiation. Engler et al.ound that matrix elasticity directs the differentiation ofuman MSCs into cells the tissue type of which matcheshe stiffness of the culture environment [229]. Specifically,

atrix elasticity was varied by controlling the cross-linkingf polyacrylamide gels [234]. To promote cell-adhesiveroperties, the gels were coated with collagen type I,hich is known to support both myogenic and osteogenicifferentiation [235]. MSC differentiation correlated withhe in vivo matrix elasticity. Soft matrices that mimicrain induced neurogenic differentiation, stiffer matriceshat mimic muscle induced myogenic differentiation, andomparatively rigid matrices that mimic bone inducedsteogenic differentiation (Fig. 7). Various mechanical sig-als were generated by the cells pulling on the matricesith various levels of elasticity. Moreover, the inhibition

f nonmuscle myosin II, a cytoskeletal motor, blocks theodulation of cell differentiation by variable matrix elas-

icity.The relationship between MSC differentiation and

atrix elasticity in 2D culture systems [229] has also beenemonstrated in 3D culture systems [236]. 2D culture sys-ems for the modulation of stem cell differentiation mayy incompatible with 3D tissue fabrication methods thatse stem cells. Pek et al. demonstrated that mechanicalroperties modulate the differentiation of MSCs cultured

n 3D gels [236]. They previously developed a thixotropicEG-silica nanocomposite gel for 3D cell culturing [237].his gel was used as a 3D matrix to investigate the effectsf mechanical signaling on MSC differentiation in a 3Datrix. The stiffness of the gel was varied by controlling

he amount of fumed silica [238]. When human MSCs wereultured in the gels with varying degrees of mechanicaltiffness, the highest expression levels of neural (ENO2),yogenic (MYOG) and osteogenic (Runx2, OC) transcrip-

ion factors were obtained using gels with low, mediumnd high degrees of stiffness, respectively [236]. In addi-ion, the immobilization of RDG ligand to the gel furthernhanced osteogenic factor expression, due possibly to theGD ligand-integrin complex acting as a mechanical signaleceptor that allows the transmission of signals through aignaling pathway [239].

.6. Control of cell shape and size with artificial ECMs

Soft lithography techniques allow for spatial patterningf cell adhesion molecules, which can be utilized to controlell size and shape. Chen and colleagues showed that cellhape and size direct stem cell fate [240]. Micropatterned

Science 36 (2011) 238–268

substrates containing fibronectin-coated islands of varioussizes were fabricated using polydimethylsiloxane (PDMS)stamps. Human MSCs cultured on larger islands tendedto flatten, spread and then differentiate into osteoblasts,whereas cells cultured on smaller islands tended to beround and unspread and differentiate into adipocytes. Theshape-dependent control of differentiation was mediatedby RhoA activity and cytoskeletal tension. Surprisingly,controlling RhoA activity did not require the use of solubledifferentiation factors for adipogenic or osteogenic differ-entiation.

Cell shape also modulates differentiation of embry-onic mesenchymal cells [241]. The role of embryonicmesenchymal cell shape in smooth muscle differentia-tion has been investigated by plating undifferentiatedembryonic mesenchymal cells on microsurfaces. Micro-surfaces with diameters of 10 and 20 �m were preparedby coating multi-perforated polycarbonate membraneswith holes of identical diameter using tissue adhesive.Embryonic MSCs cultured on 10 �m diameter microsur-faces conserved their original round cell shape, whereascells cultured on 20 �m diameter microsurfaces were elon-gated. The round-shaped cells remained undifferentiatedand expressed �-fetoprotein, which is produced by fetalcells but not by differentiated cells. In comparison, the elon-gated cells differentiated into smooth muscle, as indicatedby the expression of smooth muscle-specific proteins andmembrane potential-dependent calcium ion currents. Thecell shape was the critical factor in directing differentiationinto smooth muscle, regardless of the supplementation ofTGF-�1.

When cells are of identical size, cell shape becomes sig-nificant in modulating stem cell differentiation. A study byChen et al. demonstrated that cell shape and size directstem cell fate [240] and that cell shape also affects cellsize. However, their study did not isolate the effect ofeither cell size or shape from each other. Kilian et al.investigated the effect of cell shape only on stem cell dif-ferentiation by varying cell shape and keeping cell sizeconstant [242]. A microcontact printing technique usingPDMS stamps was used to produce a patterned culture ofhuman MSCs with various cell shapes and with identicalcell sizes in the presence of both adipogenic and osteogenicsoluble factors. Cells that were flower-shaped differenti-ated into adipocytes, whereas star-shaped cells resultedin osteogenic differentiation, possibly due to a differencein cytoskeletal contractility. A previous study showed thatcontractile cytoskeletal movement changes the cell froma flower shape to a star shape [243]. The concave regionsbetween the points of the star are highly contractile regionswhere the cell spans across a non-adhesive area [242]. Star-shaped cells have larger focal adhesion and cytoskeletaltension than cells that are flower-shaped. It has been sug-gested that local curvatures that increase the cytoskeletaltension of the cell also promote osteogenic differentiationrather than adipogenic differentiation [242]. It was also
suggested that cells with enhanced contractility undergoosteogenic differentiation due to the increased activationof c-Jun N-terminal kinase (JNK) and extracellular relatedkinase (ERK1/2) in conjunction with elevated wingless type(Wnt) signaling.


ix leads), and coier Inc.,

Fig. 7. Elasticity of artificial ECM directs MSC differentiation. Soft matrmatrix leads to myogenic differentiation (mid-size arrow: MyoD-positiveCBF�1-positive). The scale bars denote 5 �m. [229] Copyright 2006, Elsev

6.7. Mechanical signal loading on artificial ECMs

Stem cell differentiation into certain types of cellsis influenced by the mechanical microenvironment. Forexample, cells in tissues or organs in vivo are subjected to

to neuronal differentiation (�3 small arrows: Tubulin-positive), stiffermparatively rigid matrix leads to osteogenic differentiation (large arrow:Oxford, UK.

various forms of mechanical forces. Cells of the heart, arter-ies, bladder and intestines are subjected to cyclic strain.Further, chondrocytes are subjected to cyclic compressionin vivo. It is thought that the same forces that govern tis-sue development in vivo also enhance tissue development

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n vitro [244]. Thus, the mechanical signals could play sig-ificant roles in the lineage commitment of stem cells. Forxample, shear force [245] or cyclic strain [246] inducesascular cell differentiation of ESCs. It is speculated thatechanical signals are transmitted to the cells via cell

urface receptors at focal adhesion sites, which serve asechanical transducers between the cell-adhesion surface

nd cytoskeleton of the cell. The transmitted mechanicalignals are converted into biochemical responses through aechanism termed as mechanotransduction [247], result-

ng in activation of specific genes that modulate stem cellifferentiation.

Mechanical stimulation alone, without exogenousrowth factors, can induce the differentiation of bone mar-ow mesenchymal progenitor cells into a ligament cellineage [248]. Collagen gel was used as an artificial ECM forhe transduction of mechanical signals to cells in vitro. Thepplication of ligament-like multidimensional mechanicaltrains (translational and rotational strain) to undifferen-iated cells embedded in a collagen gel over a period of 21ays resulted in increased expression of ligament fibroblastarkers, including collagen types I and III and tenascin-C,

s well as induced cell alignment and collagen fiber orien-ation, all features characteristic of ligament cells.

Mechanical compression could promote the chondro-enic differentiation of stem cells. Hydrogels can be useds an artificial ECM for the transduction of compressiveechanical signals to cells. Exposure of bone marrow-

erived MSCs cultured in PEG diacrylate hydrogels to cyclicompressive strain in vitro increases the gene expres-ion of cartilage-related markers such as Sox-9, type IIollagen and aggrecan regardless of TGF-�1 supplemen-ation [112]. On the other hand, mechanical compres-ion increases expression of cartilage-specific genes inuman ESC-derived cells encapsulated in tyrosinearginine-lycine-aspartate-serine (YRGDS)-modified PEG-acrylateYRGDS-PEG-acrylate) hydrogels only in the presence ofGF-�1 [112].

Elastic, artificial ECMs permit the transduction ofechanical cyclic strain signals for cardiomyogenesis. Car-

iomyocytes in the body are subjected to cyclic strainnduced by the heart beating. A study has demonstratedhat cyclic strain promotes cardiomyogenesis of embryonictem cell-derived cardiomyocytes (ESCCs) [249]. The appli-ation of cyclic strain to ESCCs cultured on elastic polymerpoly(lactide-co-caprolactone), PLCL] scaffolds in vitro ele-ated cardiac gene expression compared to unstrainedontrols. Elastic PLCL scaffolds were also used as a vehi-le for cell transplantation into infarcted rat myocardium.LCL scaffolds were seeded with ESCCs and implanted onhe surface of infracted rat heart. As the heart beats, cyclictrain signals were transmitted to ESCCs seeded on thelastic scaffolds. Importantly, cardiac gene expression waspregulated in the elastic patches compared to unstrainedontrol patches (non-elastic PLGA scaffolds seeded withSCCs), suggesting for the presence of cardiomyocyte-
pecific microstructures including myofibrillar bundlesnd Z-lines.
Mechanical signals in combination with soluble sig-als synergistically stimulate the differentiation of stemells. Either cyclic strain [249] or TGF-�1 [250] alone has

Science 36 (2011) 238–268

been shown to stimulate the cardiomyogenic differentia-tion of stem cells. Bhang et al. showed that combining amechanical signal (cyclic strain) and soluble signal (TGF-�1) promotes cardiomyogenic marker expression in MSCsmore effectively than soluble signal alone [251]. Theexpression of cardiac-specific markers (cardiac �-actin,tropomyosin, cardiac troponin-I and cardiac myosin heavychain) was more upregulated in MSCs cultured with TGF-�1 and subjected to cyclic strain than in MSCs culturedstatically with TGF-�1. Incomplete cardiomyogenic dif-ferentiation of MSCs could result in the persistence ofundifferentiated stem cells that undergo unanticipateddifferentiation upon implantation. On the other hand, acombination of two or more types of signals could enhancethe efficiency of stem cell differentiation.

7. Concluding remarks

Artificial ECMs can regulate the cellular function to engi-neer functional tissues. For engineering functional tissues,signals and parameters that can modulate cell behav-ior can be engineered into artificial ECMs. The signalsand parameters include soluble signals, cell-adhesion sig-nals, mechanical signals, and chemistry, architecture, andmechanical properties of artificial ECMs. Design criteria forartificial ECMs vary considerably depending on the tissueof interest. Artificial ECMs that can present appropriatesignals and parameters can coordinate the interactionsbetween the artificial ECMs and cells to regenerate adesired tissue. The artificial ECMs developed for tissueengineering may in turn provide a novel experimental sys-tem to elucidate the mechanisms by which native ECMsregulate tissue development and stem cell behavior.

Acknowledgements

This work was supported by a grant (ROA-2010-0018428) from NRL program of Korea Science andEngineering Foundation (KOSEF), the Ministry of Educa-tion, Science and Technology, Republic of Korea and a grant(2010-0020352) from the National Research Foundation ofKorea.

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