ARTICLE IN PRESS

0021-9290/$ - se

doi:10.1016/j.jb

Abbreviations

PCL, polycapro

tricalciumphosp�Correspond

E-mail addr

Journal of Biomechanics 40 (2007) 750–765

www.elsevier.com/locate/jbiomech

Review

Osteochondral tissue engineering

Ivan Martin�, Sylvie Miot, Andrea Barbero, Marcel Jakob, David Wendt

Department of Research and Institute for Surgical Research and Hospital Management, University Hospital of Basel,

Hebelstrasse 20, 4031 Basel, Switzerland

Accepted 13 March 2006

www.JBiomech.com

Abstract

Osteochondral defects (i.e., defects which affect both the articular cartilage and underlying subchondral bone) are often associated

with mechanical instability of the joint, and therefore with the risk of inducing osteoarthritic degenerative changes. Current surgical

limits in the treatment of complex joint lesions could be overcome by grafting osteochondral composite tissues, engineered by

combining the patient’s own cells with three-dimensional (3D) porous biomaterials of pre-defined size and shape. Various strategies

have been reported for the engineering of osteochondral composites, which result from the use of one or more cell types cultured

into single-component or composite scaffolds in a broad spectrum of compositions and biomechanical properties. The variety of

concepts and models proposed by different groups for the generation of osteochondral grafts reflects that understanding of the

requirements to restore a normal joint function is still poor. In order to introduce the use of engineered osteochondral composites in

the routine clinical practice, it will be necessary to comprehensively address a number of critical issues, including those related to the

size and shape of the graft to be generated, the cell type(s) and properties of the scaffold(s) to be used, the potential physical

conditioning to be applied, the degree of functionality required, and the strategy for a cost-effective manufacturing. The progress

made in material science, cell biology, mechanobiology and bioreactor technology will be key to support advances in this challenging

field.

r 2006 Elsevier Ltd. All rights reserved.

Keywords: Cartilage repair; Chondrocyte; Scaffold; Bioreactor; Mechanobiology

Contents

1. Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 751

2. Described approaches for the engineering of osteochondral grafts . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 751

2.1. Scaffold for bone component, scaffold-free for cartilage (Scaffold strategy A) . . . . . . . . . . . . . . . . . . . . . . . . . . 751

2.2. Different scaffolds for bone and cartilage components (Scaffold strategy B) . . . . . . . . . . . . . . . . . . . . . . . . . . . 754

2.3. One heterogeneous/bilayered scaffold (Scaffold strategy C) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 755

2.4. One homogenous/single-layer scaffold (Scaffold strategy D) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 755

3. Towards design principles for the engineering of osteochondral composites. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 756

3.1. Size, shape of the graft and surgical approach for implantation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 756

e front matter r 2006 Elsevier Ltd. All rights reserved.

iomech.2006.03.008

: FDM, fused deposition modeling; FGF-2, fibroblast growth factor-2; HA, hydroxyapatite; MPC, mesenchymal progenitor cells;

lactone; PEG, poly-ethylene glycol; PLA, poly-lactic acid; PGA, poly-glycolic acid; PLGA, poly-lactic-coglycolic acid; TCP,

hate

ing author. Tel.: +41 61 265 2384; fax: +4161 265 3990.

ess: imartin@uhbs.ch (I. Martin).

ARTICLE IN PRESSI. Martin et al. / Journal of Biomechanics 40 (2007) 750–765 751

3.2. Cell source. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 757

3.3. Physical conditioning of engineered osteochondral tissues . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 758

3.4. Required in vitro maturation stage of engineered cartilage . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 759

3.5. Manufacture of engineered osteochondral composites . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 760

3.6. Osteochondral repair by cell-free scaffolds?. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 761

4. Conclusions. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 761

Acknowledgments . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 762

References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 762

Fig. 1. Schematic diagram of strategies used for the fabrication of

osteochondral grafts. Specific approaches proposed by different

research groups can be classified based on a combination of a scaffold

strategy (A–D) and a cell strategy (I–IV). Under this classification

system, for example, the approach of scaffold-free culture of

chondrocytes on a cell-free osteoconductive scaffold would be

designated as ‘‘A/I’’.

1. Introduction

Osteochondral defects, typically derived by trau-matic injuries or osteochondritis dissecans, are oftenassociated with mechanical instability of the joint,and therefore with the risk of inducing osteoarth-ritic degenerative changes. Grafting of osteochondralunits consisting of a superficial cartilaginous layer(corresponding to articular cartilage) and an under-lying calcified tissue (corresponding to subchondralbone) represents a promising approach to restorethe biological and mechanical functionality of thejoint. Despite the encouraging results reported, theclinical use of autologous osteochondral grafts(i.e., mosaicplasty technique) suffers from severallimitations, namely: (i) the amount of material available,(ii) the donor site morbidity, and (iii) the difficultyto match the topology of the grafts with the injuredsite. Tissue engineering of osteochondral compositeshas the potential to overcome these limits. Three-dimensional (3D) tissue grafts of pre-defined sizeand shape can be engineered by combining the patient’sown cells with 3D porous biomaterials, which providethe template for tissue development and degrade atdefined rates. Tissue engineering approaches wouldallow the properties of the graft to be specificallytailored, in order to introduce in the affected joints thestructural, biological and biomechanical cues which arenecessary and sufficient for a reproducible and durablerepair.

The progress made in material science, cell biologyand bioreactor technology has been instrumentalfor the flourishing of a broad spectrum of approachesto the engineering of osteochondral grafts. The varietyof concepts and models so far proposed and investi-gated by different groups for the generation ofosteochondral grafts reflects that understanding of therequirements to restore a normal joint function is stillpoor. While in principle it would be feasible to developtechnical solutions to a well-defined design, the princi-ples of the design itself still have to be defined. Based onthese considerations, here we will first review theapproaches so far proposed for the fabrication ofosteochondral grafts. We will then discuss some openquestions that we feel should be addressed in order toestablish the design criteria for engineered osteochon-dral composites.

2. Described approaches for the engineering of

osteochondral grafts

The different approaches so far proposed for thefabrication of osteochondral composite constructs canbe classified based on the combination of specificstrategies for the selection of the scaffold and the cellsource, as schematically outlined in Fig. 1. According tothis classification, osteochondral constructs have beengenerated using: (A) a scaffold for the bone componentbut a scaffold-free approach for the cartilage compo-nent; (B) different scaffolds for the bone and thecartilage components combined at the time of implanta-tion; (C) a single but heterogeneous composite scaffold;or (D) a single homogenous scaffold for both compo-nents. These scaffolds have been (I) loaded with a singlecell source having chondrogenic capacity, (II) loadedwith two cell sources having either chondrogenic orosteogenic capacities, (III) loaded with a single cellsource having both chondrogenic and osteogenic differ-entiation capacity, or (IV) used in a cell-free approach.In this section, some of the most relevant studies inosteochondral tissue engineering will be presented,following the outlined classification (see Table 1).

2.1. Scaffold for bone component, scaffold-free for

cartilage (Scaffold strategy A)

Composite implants can be produced by seeding andculturing a high density of chondrogenic cells directly ontop of an osteoconductive biomaterial. Following this

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Table 1

Summary of osteochondral studies grouped according to the repair strategy described in Fig. 1

Strategy Chondral component Subchondral component Assembly of chondral and

subchondral components

Model system for

osteochondral development

Reference

A/I Porcine articular chondrocytes

(scaffold-free)

Cell-free PLA or collagen/HA Chondrocytes seeded &

cultured on the surface of

bone substitutes

In vitro culture of cylindrical

composite constructs

Wang et al. (2004)

A/I Bovine articular chondrocytes

(scaffold-free)

Cell-free porous calcium

polyphosphate

Chondrocytes seeded &

cultured on the surface of

osteoconductive scaffold

In vitro pre-culture of

cylindrical composites, with

subsequent mechanical

conditioning

Waldman et al. (2003)

A/I Sheep articular chondrocytes

(scaffold-free)

Cell-free porous calcium

polyphosphate

Chondrocytes seeded &

cultured on the surface of

osteoconductive scaffold

In vitro culture of cylindrical

composite constructs &

further implantation in

osteochondral defects in

sheeps

Kandel et al. (2006)

A/III Cell pellet of human MPC

press-coated onto PLA

scaffold and pre-cultured in

chondrogenic medium

Osteogenic induced human

MPC seeded onto opposite

face of pre-cultured cartilage

component

Osteogenic MPC seeded onto

opposite face of pre-cultured

PLA construct

In vitro culture of cubical

composite constructs in

medium supporting both

chondrogenesis and

osteogenesis

Tuli et al. (2004)

B/I Bovine articular chondrocytes

in agarose gel

Devitalized bovine trabecular

bone

Cell-agarose suspension added

to, & partially penetrating

into, the surface of the bone

In vitro culture of cylindrical

& patella-shaped

osteochondral constructs

Hung et al. (2003)

B/I Bovine articular chondrocytes

cultured in PGA/PLA fleeces

Natural or synthetic calcium

carbonate

Chondrocyte seeded fleeces

attached to ceramic with

fibrin-cell solution

In vitro culture of cylindrical

composites in a perfused

bioreactor system

Kreklau et al. (1999)

B/I Porcine chondrocytes cultured

in gelatin scaffold

Calcium phosphate derived

from bovine cancellous bone

(cell-free)

Gelatin region of composite

partially infiltrated into pores

of bone component;

composite assembled prior to

cell seeding.

In vitro culture of cylindrical

composites in a double-

chamber bioreactor

Chang et al. (2004)

B/I & B/II Rabbit articular chondrocytes

cultured in PGA meshes

Collagen-hydroxyapatite

sponge: cell-free or with

absorbed bone marrow

Engineered cartilage sutured

to bone substitute

Composite blocks implanted

into critical-sized

osteochondral defects in adult

rabbits

Schaefer et al. (2002)

B/II Bovine articular chondrocytes

cultured in PGA meshes

Bovine periosteal cells

cultured in PLGA/PEG

sponges

Engineered components

sutured together following

different pre-culture times

In vitro culture of cylindrical

composites in osteogenic

medium

Schaefer et al. (2000)

B/III Chondrogenic rat MPC

seeded in hyaluronic acid

sponges

Osteogenic rat MPC seeded in

porous calcium phosphate

ceramics

Freshly seeded components

sealed together with fibrin

Composite construct blocks

implanted subcutaneously in

mice

Gao et al. (2001)

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SC/I Ovine articular chondrocytes

cultured in PLGA/PLA region

of composite scaffold

PLGA/TCP region of

composite scaffold

A single but heterogenous

scaffold with gradient of

materials and porosity

between compartments

In vitro culture of cylindrical

composites

Sherwood et al. (2002)

C/II Primary porcine articular

chondrocytes seeded into PLA

sponge

Human fibroblasts transduced

with adenovirus expressing

BMP-7, suspended in fibrin &

seeded into HA

Components of composite

scaffold combined with PLA

prior to cell seeding

Freshly seeded cylindrical

composites implanted

subcutaneously in mice

Schek et al. (2004)

C and D I

and IV

PLGA based scaffold (w/ or

w/o a PGA fiber

reinforcement), cell-free or

seeded with goat rib

chondrocytes

PLGA based scaffold

(w/ different additives)

A single homogenous or

heterogenous scaffold;

heterogeneous composites

‘‘glued’’ with solvent prior to

cell seeding

Cylindrical composites

implanted into both high &

low weight bearing sites in

goats

Niederauer et al. (2000)

D/II Human rib chondrocytes

cultured in one half of PCL

scaffold (manufactured by

FDM technique)

Human MPC pre-cultured in

one half of PCL scaffold

(manufactured by FDM

technique)

Homogeneous PCL scaffold

with gap between the two

compartments. Chondrocytes

seeded into empty half of pre-

cultured MPC-PCL construct

In vitro culture of composite

construct blocks

Cao et al. (2003)

D/III Chondrogenic rat MPC

embedded in

photopolymerizable PEG

hydrogel

Ostoegenic rat MPC

embedded in

photopolymerizable PEG

hydrogel

Sequential loading &

photopolymerization of

chondrogenic & osteogenic

hydrogel suspensions within

condyle-shaped mold

Human articular condyle-

shaped osteochondral

composites implanted

ectopically in mice

Alhadlaq et al. (2004)

D/IV Polyethylene fiber scaffold

coated with HA, impregnated

with type I collagen gel &

FGF-2 (cell-free)

Polyethylene fiber scaffold

coated with HA, impregnated

with type I collagen gel &

FGF-2 (cell-free)

A single & homogenous cell-

free scaffold

Scaffold blocks implanted into

patellar groove of rabbits

Fukuda et al. (2005)

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Fig. 2. In vitro engineered osteochondral composites. Histological appearance of osteochondral composites generated by seeding and culturing a

high density of porcine articular chondrocytes directly on top of a porous osteoconductive scaffold made of either poly-L-lactid (A) or collagen-HA

[Col-HA] (B) for 11 weeks (strategy A/I according to Fig. 1) (Wang et al., 2004). A cartilaginous matrix positively stained for Alcian blue was

obtained on top of both scaffolds, but neo-cartilage integrated more extensively into the Col-HA scaffold. Images have been kindly provided by Dr.

P. Mainil-Varlet, University of Bern, Switzerland (scale bar ¼ 500mm).

I. Martin et al. / Journal of Biomechanics 40 (2007) 750–765754

approach, Wang et al. seeded primary porcine articularchondrocytes directly onto the top surface of threecommonly used biomaterials: poly-L-lactide, poly-D,L-lactide or collagen-hydroxyapatite (HA) (Fig. 1: A/I)(Wang et al., 2004). After 7 weeks of in vitro culture,chondrocytes cultured on each of the biomaterialsproduced extracellular matrix (ECM) containing alarge proportion of collagen type II and glycosamino-glycans (GAG) that was partially integrated withthe subchondral base (Fig. 2). Constructs producedusing collagen-HA were superior in terms of cellviability, construct shape and cellular integration,although the authors pointed out that the outcomeon the stability of the osteochondral compositescould be different under physiological loading in anorthotopic model. In this regard, promising results wererecently reported in a sheep model using cartilage tissuesgrown on top of porous ceramic scaffolds (Kandel et al.,2006).

Using human trabecular bone-derived mesenchymalprogenitor cells (MPC) in conjunction with poly-D,L-lactic acid (PLA) scaffolds, Tuli et al. followed a similarapproach to generate osteochondral grafts (Fig. 1: A/III) (Tuli et al., 2004). A pellet of MPC was press-coatedon a PLA scaffold and cultured in a defined mediumsupplemented with transforming growth factor beta-1(TGF-b1) to support chondrogenesis. In parallel, MPCfrom the same patient were expanded and cultured inmonolayer with factors inducing osteogenesis (b-glycer-ophosphate and dexamethasone) and then seeded on theopposite face of the cartilage-PLA construct. Thebilayered constructs were then cultured in a cocktailmedium containing insulin and osteogenic factors tosupport/promote chondrogenesis and osteogenesis si-multaneously. The resulting composites consisted ofboth a hyaline cartilage-like layer and a dense bone-likecomponent, with a well developed transition zonebetween the two compartments.

2.2. Different scaffolds for bone and cartilage

components (Scaffold strategy B)

In an alternative approach, different scaffolds can beused for the cartilage and bone components. Cartilagi-nous and/or bone-like tissues can then be engineered invitro within the respective scaffold layer and combinedinto a single composite graft by suturing or adheringtogether the two layers.

Schaefer et al. seeded and cultured differentiatedbovine articular chondrocytes into poly-glycolic acid(PGA) meshes and periosteal-derived cells into poly-lactic-coglycolic acid/poly-ethylene glycol foams (PLGA/PEG) to independently generate the cartilage and bonelayers (Fig. 1: B/II) (Schaefer et al., 2000). Following 1–4weeks of independent culture, the generated cartilaginousand bone-like tissues were then sutured together andcultured for an additional period in osteogenic medium.The authors showed that chondrocytes within the PGAmeshes produced extracellular matrix containing highamounts of GAG and collagen type II during isolatedculture and had maintained their specific phenotypeduring the subsequent composite culture. Furthermore,periosteal cells within the PLGA/PEG foams depositedmineralized matrix and bone specific proteins (osteocalcinand osteopontin). This study demonstrated the possibilityof generating composites of cartilaginous and bone-liketissues in vitro, and pointed out that the maturation andintegration of the two components can be modulated bythe cultivation time. In addition, this was the firstevidence that articular chondrocytes can remain pheno-typically stable when cultured with factors typically usedto promote osteogenesis and adjacent to osteogenic cellsproducing mineralized matrix.

Using a single cell source, Gao et al. generatedcomposite osteochondral grafts from hyaluronic acid-based sponges and porous ceramics (Fig. 1: B/III) (Gaoet al., 2001). In this study, rat bone marrow-derived

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Fig. 3. Heterogeneous composite scaffold for osteochondral repair.

Example of a composite scaffold, consisting of two distinct but

integrated layers corresponding to the cartilage and bone components

(Scaffold strategy C according to Fig. 1). The implant is comprised of:

(i) a top layer of the elastomeric copolymer poly(ethylene glycol)-

terephthalate/poly(butylene)-terephthalate (PolyActive), having me-

chanical properties comparable to native cartilage (Miot et al., 2005),

and (ii) a bottom layer composed of HA and tricalcium phosphate,

displaying both osteoconductive and osteoinductive properties (Yuan

et al., 2002). The image has been kindly provided by Dr. J. Pieper,

IsoTis B.V., Bilthoven, The Netherlands.

I. Martin et al. / Journal of Biomechanics 40 (2007) 750–765 755

MPC were first ‘‘committed’’ in monolayers withchondrogenic (e.g., TGF-b1) or osteogenic (i.e., b-glycerophosphate and dexamethasone) culture supple-ments and then seeded, respectively, on the sponges orporous ceramic. The two parts were sealed togetherusing fibrin glue and implanted subcutaneously intonude mice. While lamellar bone was abundant in theceramic, fibrocartilage, and not hyaline cartilage, filledthe pores of the hyaluronic acid sponge.

Osteochondral composites based on different scaffoldswere first used to repair a critical-sized osteochondraldefects (7� 5� 5mm) in a rabbit model (Schaefer et al.,2002). In this study, engineered cartilage, generated invitro from autologous articular chondrocytes cultured inPGA meshes, was sutured to a collagen-HA sponge,loaded or not with autologous MPC at the time ofimplantation (Fig. 1: B/II and B/I, respectively). Theresulting composite was then press-fit into a surgicallycreated femoropatellar groove defect. Engineered carti-lage withstood physiologic loading and remodeled over 6months into osteochondral tissue with characteristicarchitectural features and Young’s moduli close to nativetissue (engineered cartilage: 0.68–0.80MPa; explanttissue: 0.84MPa). Constructs integrated well with hostbone but not with adjacent host cartilage. The initialloading of MPC in the osteoconductive material had noadvantage over the cell-free bone layer, likely due to thepresence of osteoprogenitor cells in the bleeding sub-chondral bone during implantation. This study demon-strated for the first time that functional engineeredcartilage grafts combined with an osteoconductive sup-port, even if the latter is not loaded with osteogenic cells,provide a template permitting the orderly repair ofcritical-sized osteochondral defects in adult rabbits.

2.3. One heterogeneous/bilayered scaffold (Scaffold

strategy C)

As opposed to generating composite osteochondralgrafts by combining independent cartilaginous andbone-like components, heterogeneous scaffolds havebeen proposed, composed of two distinct but integratedlayers for the cartilage and bone regions (Fig. 3).

Following this approach, Schek et al. seeded porcinearticular chondrocytes and human gingival fibroblaststransduced with adenovirus expressing BMP-7 into thetwo regions of a poly-L-lactic acid–HA (PLA–HA)composite scaffold (Fig. 1: C/II) (Schek et al., 2004). Athin film of PGA was deposited at the interface betweenthe two biomaterials to prevent cell migration betweenthe components. The scaffold shape and pore architec-ture were defined using an image-based design method(which could also be used to design patient-specificimplant geometries) and was manufactured using asolid-free-form fabrication technique. Following ectopicimplantation in mice, the cartilage layer contained ECM

rich in GAG, the bone layer contained regions withblood vessels, marrow stroma and adipose tissue, and amineralized interface was often present at the junction.

Using the TheriFormTM 3D printing process, Sher-wood et al. developed an innovative osteochondralcomposite scaffold, promoting preferential cell attach-ment and matrix deposition within the cartilage portion(Fig. 1: C/I) (Sherwood et al., 2002). The compositeconsisted of a cartilage region of 90% porous D,L-PLGA/L-PLA with macroscopic staggered channels, a boneregion of 55% porous L-PLGA/tricalciumphosphate(TCP), and a transition zone between the two regionscontaining a gradient of materials and porosity to preventdelamination at the interface. The bone region had tensileand compressive strengths (1.6 and 2.5MPa, respectively)similar in magnitude to fresh cancellous human bone (�8and 10–20MPa, respectively) and was fabricated in acloverleaf shape, with channels to increase the surfacecontact with bone marrow and to facilitate cell migration.

2.4. One homogenous/single-layer scaffold (Scaffold

strategy D)

The use of a single homogeneous scaffold and two celltypes having chondrogenic and osteogenic capacity(Fig. 1: D/II) was adopted by Cao et al. to engineer

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osteochondral composites (Cao et al., 2003). Using a fuseddeposition modeling (FDM) technique, polycaprolactone(PCL) scaffolds were fabricated with a honeycomb-likegeometry (in a 01/601/1201 lay-down pattern), with a welldefined pore volume, structure and porosity, partitionedvertically into two halves. Human bone marrow derivedMPC were first seeded in one-half of the partitionedscaffold and pre-cultured. Human rib chondrocytes weresubsequently seeded into the other half, and the compositeconstructs were co-cultured in vitro in medium containingosteogenic supplements. Using scanning electron micro-scopy, the authors observed different extracellular matricesin each compartment, however, no further characterizationof matrix composition or cell phenotype was performed.The authors concluded that the use of slow-degrading PCLin conjunction with the FDM technique allowed for thefabrication of a scaffold with sufficient mechanical strengthto endure initial in vivo loading.

Alhadlaq et al. also reported the use of a singlescaffold, a PEG hydrogel, but seeded with a single cellsource (Fig. 1: D/III) in order to engineer a humanshaped articular condyle (Alhadlaq et al., 2004). Rat bonemarrow-derived MPC were expanded, induced separatelyto chondrogenic or osteogenic differentiation using specificculture supplements and then loaded in two stratified andintegrated hydrogel layers that were photopolymerized in ahuman condylar mold. After 4 weeks implantation inimmunodeficient mice, most of the differentiated chon-drogenic and osteogenic cells had synthesized correspond-ing cartilaginous and bone-like matrices.

A cell-free approach was undertaken by Fukuda et al.who used a single 3D fabric (3DF) for the repair of a largefull-thickness osteochondral defect (6� 6� 3mm) in thepatellar groove of rabbits (Fig. 1: D/IV) (Fukuda et al.,2005). The scaffold, consisting of an ultra-high molecularweight polyethylene fiber and an unsintered HA coating,was impregnated with a type I collagen gel with or withoutfibroblast growth factor-2 (FGF-2). After 48 weeks in vivo,FGF-2 promoted biological resurfacing with hyaline-likecartilage and induced subchondral bone formation withinand around the implanted graft. The approach of using acell-free scaffold impregnated with growth factors is apromising strategy for the treatment of large osteochondraldefects, even if a future clinical application will necessitateto define optimal concentrations of released growth factorand to use more controllable delivery systems.

3. Towards design principles for the engineering of

osteochondral composites

3.1. Size, shape of the graft and surgical approach for

implantation

Resurfacing of an irregular defect with multiplecylindrical plugs, as in the mosaicplasty technique, is

associated with the generation of cartilage-to-cartilageinterfaces, known to have a limited capacity to integrate,and of incongruities in the articular surface, known tolead to significantly increased contact pressures (Kohet al., 2004). As an alternative to cylindrical plugs, thegeneration of anatomically shaped constructs thatreproduce the contour of the articular surfaces has thepotential to reduce the overall interfacial area betweenopposing cartilage surfaces and to restore normal loaddistributions across the joint when implanted. The needfor a custom-shaped graft is even more relevant whenthe defect is unconfined, as described in a recent casereport where the lesion was located on the edge of thelateral femoral condyle (Adachi et al., 2004). In thisstudy, the authors first restored the normal contour ofthe joint by implanting a cortical bone block, trimmedto adjust to the shape of the condyle bone, and thengrafted a layer of engineered cartilage. The feasibility toculture anatomically shaped grafts aimed at replacingthe entire articular surface of a diarthroidal joint wasdemonstrated by the generation of bilayered constructsconsisting of chondrocyte-seeded agarose on naturaltrabecular bone (Hung et al., 2003). In that study, thegeometry of the articular cartilage layer, previouslyacquired from human cadaver joints, was used inconjunction with computer-aided design and manufac-turing technology to create anatomically accuratemolds. As previously described in Section 2.4, ahuman-shaped condyle has also been generated throughthe photopolymerization of two stratified and integratedhydogel layers within a negative mold of a condyle(Alhadlaq et al., 2004). Based on the development ofrapid prototyping techniques to fabricate not onlyplastic molds, but also porous scaffolds directly fromcomputer-aided design (Woodfield et al., 2004; Yeonget al., 2004), it was proposed that 3D computedtomography, coupled with 3D computer aided designand rapid prototyping, could be used to design andengineer customized femoral and tibial cartilaginousimplants made of synthetic polymers in defined porearchitectures (Fig. 4). The concept of customizedosteochondral composites is in principle extremelyinteresting, but likely difficult to be implemented inpractice. Challenges to be overcome include the need ofa precise fit into a defect whose shape may change overtime and the definition of a surgical method for graftfixation. Also to be considered are the estimated highcosts of the procedure and the fact that it would notallow to be performed under minimally invasive condi-tions.

A promising approach which might allow arthro-scopic implantation of anatomically shaped grafts isbased on the use of shape memory materials, whichwould be fabricated in a condensed state, delivered viaminimally invasive surgery and would then acquire the‘‘programmed’’ shape in situ (Lendlein and Langer,

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Fig. 4. Rapid prototyping for anatomically shaped scaffolds. Using a

3D model generated by computed tomography scans of a rabbit knee

joint (A), porous scaffolds made of poly(ethylene glycol)-terephtha-

late/poly(butylene)-terephthalate (PolyActive) were produced by 3D

fiber deposition (B) (Woodfield et al., 2004). The images have been

kindly provided by Dr. T. Woodfield, University of Twente, The

Netherlands (scale bar ¼ 5mm).

I. Martin et al. / Journal of Biomechanics 40 (2007) 750–765 757

2002). Shape memory alginate hydrogels were recentlyused to demonstrate the possibility of cartilage tissueformation in predefined geometries following ectopicdelivery through a small catheter (Thornton et al., 2004).The scaffolds, initially reduced to less than 5% of thepredefined volume, returned to their programmed sizeupon rehydration with a solution containing chondrocytes,which resulted in efficient cell seeding and in cartilagestructures maintaining the predefined geometry up to 6months. Although a number of different classes of shapememory materials has already been described, theapproach has not yet been tailored for the generation ofosteochondral composites or the repair of joint defects.

From a surgical standpoint, the ideal consistency of agraft would be that of a paste, allowing the embeddingof chondrocytes and/or MPC, which might be deliveredarthroscopically and shaped in situ. Several promisingmaterials have been proposed to fill gaps in cartilagetissue (Sittinger et al., 2004), including chitosan-basedgels (Chenite et al., 2000) and chondroitin sulfatephotopolymerizing hydrogels (Li et al., 2004). However,none of these has yet been demonstrated to provide asubstrate with the appropriate biomechanical propertiesand the capacity to resorb/remodel, which couldpromote immediate joint loading and durable regenera-tion of osteochondral defects.

3.2. Cell source

The variety of concepts on the cell types proposed forthe repair of osteochondral defects has already been

introduced earlier in this review. Considering thepresence of osteoprogenitor cells in bleeding subchon-dral bone, paralleled by a generally high efficiency ofbone regeneration in the subchondral regions, it is likelythat fabrication of osteochondral composites willrequire only one cell type, to generate their cartilaginouspart. Articular chondrocytes have been the mostpopular source of cells to engineer cartilage grafts, butin the majority of the studies cells were used immediatelyafter isolation from the native cartilage of differentanimal species. For the clinical implementation of theapproaches proposed, instead, human articular chon-drocytes would have to be first expanded in culture,which is typically associated with cell de-differentiation,leading to the downregulation of cartilage-specific genes(Benya and Shaffer, 1982; Binette et al., 1998). Thereduced capacity of expanded chondrocytes, particularlyif of human origin, to re-differentiate and producecartilage-specific extracellular matrix, undermines thesuccess to engineer functional cartilage tissues.

One possible strategy that could be undertaken toovercome this limitation consists in expanding chon-drocytes in the presence of specific growth factors. Thepresence of FGF-2, TGF-b1 and platelet-derivedgrowth factor-bb during the expansion of human

articular chondrocytes was shown not only to repro-ducibly increase cell proliferation rate, but also toenhance the cell capacity to re-differentiate followingexpansion (Barbero et al., 2003; Jakob et al., 2001),particularly if the cells were from young adults below 40years of age (Barbero et al., 2004), and to respond todifferentiating factors present in the synovial fluid(Jakob et al., 2004).

A critical issue associated with the use of autologousarticular chondrocytes is the procurement of the biopsyfrom the patient. A cartilage biopsy in the joint, even ifharvested from a non-load bearing site, represents anadditional injury to the cartilage surface, and has beenreported to be detrimental to the surrounding healthyarticular cartilage (Lee et al., 2000). To overcome thisproblem, one alternative approach would be based onthe use of chondrocytes obtained from non-articularcartilage tissues. For example, biopsies of nasal or ribcartilage can be harvested under local anesthetic and bya less invasive procedure than removing tissue fromspecific areas of the joint. Morbidity is also reduced bythe fact that the donor site is not subjected to high levelsof physical forces, as in the joint. Several studies haveshown that chondrocytes derived from human nasalseptum or ear cartilage proliferate and generate cartila-ginous tissue after monolayer expansion with similar orsuperior capacity to those derived from articularcartilage (Kafienah et al., 2002; Tay et al., 2004; VanOsch et al., 2004). However, to demonstrate whether thetissue generated by non-articular chondrocytes isadequate for cartilage repair at articular sites, extensive

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data from in vivo orthotopic experimental studies and/or from in vitro loaded models will be needed.

To avoid the limitations of de-differentiated chon-drocytes, several groups have proposed the use of MPCto generate cartilaginous tissues. MPC have been foundin a variety of human adult tissues, including bonemarrow, periosteum, trabecular bone, synovial mem-brane, skeletal muscle, dermis, blood, and adipose tissue(Tuan et al., 2003). The most extensively studied MPCare those isolated from bone marrow, also called bonemarrow stromal cells (BMSC). However, BMSC differ-entiated towards the chondrogenic lineage were shownto express markers specific of hypertrophic chondro-cytes (e.g.: type X collagen and MMP-13) (Mackay etal., 1998; Winter et al., 2003) and to further mineralizethe deposited matrix when exposed to osteogenic stimuli(Mackay et al., 1998; Muraglia et al., 1998), thusindicating a potential instability of the acquiredchondrocytic phenotype. Despite a series of recentstudies reporting the use of BMSC for osteochondraldefect repair in different animal models (Gao et al.,2001; Oshima et al., 2004; Uematsu et al., 2005) and in afew clinical cases (Wakitani et al., 2004), the long-termefficacy of BMSC and their contribution to theregeneration of hyaline cartilage which does not remodelinto bone in the long term, still has to be demonstrated.

Yet another approach to bypass the difficultiesassociated with chondrocyte expansion would be theuse of a limited number of freshly isolated articularchondrocytes. Recent studies indicate that non-ex-panded chondrocytes can induce chondrogenic differ-entiation of other cell types (Hendriks et al., 2005) andthat even undigested cartilage tissue, minced into smallparticles, can be used to repair experimental cartilagedefects in goats and horses (Frisbie et al., 2005; Lu et al.,2005). These preliminary results are particularly inter-esting, since they would open the possibility of an intra-operative fabrication of osteochondral grafts, withobvious logistic and economic advantages.

3.3. Physical conditioning of engineered osteochondral

tissues

It is well known that physical stimuli can modulatethe metabolism of chondrocytes and osteoblasts, andwhen applied with specific magnitudes and frequencies,may upregulate the production of extracellular matrixcomponents. Consequently, to simulate specific physio-logical forces during joint loading, numerous tissueengineering groups have developed bioreactors to applymechanical stimuli to cell-seeded scaffolds/hydrogels inan effort to enhance cell differentiation and/or tissuedevelopment. A number of studies have indeed reportedstimulated chondrocyte metabolism and/or enhancedcartilage ECM production in response to dynamiccompression, although these responses were greatly

dependent upon the specific magnitude and/or fre-quency applied (Buschmann et al., 1995; Davisson etal., 2002a; Lee et al., 2003; Lee and Bader, 1997; Kisidayet al., 2004). Likewise, enhanced osteogenic differentia-tion and mineralized matrix deposition was reportedwhen human BMSC, pre-cultured within partiallydemineralized bone matrix, were subjected to physiolo-gical strain magnitudes in a four-point bending bior-eactor (Mauney et al., 2004). In addition to thesestudies, which focused on engineering either chondral orbone tissues, mechanical conditioning has also beenapplied to composite osteochondral constructs. Imple-menting intermittent regimes of either dynamic com-pression or dynamic shear to composite constructs (i.e.,chondrocytes pre-cultured on top of calcium polypho-sphate cylinders), Waldman et al. found that with aslittle as 6min of mechanical stimulation every other dayduring 4 weeks of culture, the accumulation of ECMand the equilibrium moduli were increased as comparedto unstimulated constructs (compression stimula-ted ¼ 80 kPa, shear stimulated ¼ 112 kPa, unstimula-ted ¼ 20 kPa) (Waldman et al., 2003). Applying adifferent regime of intermittent dynamic compressionto composite constructs, Hung et al. also reported anincrease in the GAG content and Young’s modulusthroughout 4 weeks of culture, although to a lowerextent than in chondral constructs stimulated in theabsence of the bone substrate (Hung et al., 2004). Thiscould be explained by the fact that the underlying bonesubstrate significantly altered the strains, hydrostaticpressure, and flow fields within the chondral region, asassessed by computational modeling (Lima et al., 2004).

While mechanical conditioning appears to have thepotential to improve the structural and functionalproperties of engineered tissues, little is currently knownabout which specific mechanical forces, or regimes ofapplication (i.e. magnitude, frequency, continuous orintermittent, duty cycle), are most stimulatory. Indeed,since strain transfer to the cells is initially dominated bythe scaffold system and is progressively modulated atincreasingly important degrees by the ECM beingdeposited and organized (Hung et al., 2004), differentregimes of conditioning might be required by tissues atdifferent stages of development (see Section 3.4; a morein-depth discussion of mechanobiology can be found inthe review of Van der Meulen and Huiskes, 2002).

Mass transfer limitations of nutrients and metabolicwaste products are often considered one of the greatestchallenges in the engineering of cartilage and bonetissues of clinically relevant sizes. Although whichparticular species is/are limiting is not decisively known,insufficient oxygen transport has been associated, bothexperimentally and computationally, with inhomoge-neous development of cartilaginous tissues (Malda et al.,2004; Lewis et al., 2005). External mass transferlimitations of oxygen and other nutrients can be reduced

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by culturing constructs in a stirred-flask, where turbu-lent eddies are associated with high shear stresses, orwithin rotating wall vessels, reported to enhance themass transfer of nutrients and wastes while generatinglow levels of turbulence and shear (Martin et al., 1999;Vunjak-Novakovic et al., 1999). Recently, Chang et al.designed an innovative double-chamber stirred bioreac-tor to engineer osteochondral constructs (Chang et al.,2004). The ultimate goal of the bioreactor was tofacilitate the co-culturing of chondrocytes and osteo-blasts simultaneously within respective sections of asingle-unit composite scaffold. The bioreactor wasseparated by a silicone membrane into two compart-ments, connected only through the pore structure of thecomposite scaffold, which traversed the membrane.Since each chamber contained independent mediarecirculation systems and media stirring mechanisms,different cocktails of factors could be used for theculture of each specific cell type within their respectivecompartment of the composite scaffold.

While widely used convective systems such as stirredflasks and rotating vessels may improve mass transferto/from the surface of the construct during prolongedtissue culture, direct perfusion of culture mediumthrough the scaffold pores can mitigate mass transferlimitations within the construct as well. Perfusion ofchondrocyte-seeded scaffolds was shown to supportelevated GAG synthesis and deposition (Davisson et al.,2002b; Pazzano et al., 2000), as well as a uniformdistribution of viable human chondrocytes and extra-cellular matrix (Wendt et al., 2006). Perfusion flow hasalso been associated with enhanced growth, differentia-tion and mineralized matrix deposition by rat marrowstromal cells (Bancroft et al., 2002; Goldstein et al.,2001; Sikavitsas et al., 2005) and with the generation ofhighly osteoinductive grafts starting from human bonemarrow nucleated cells, without prior monolayer isola-tion and expansion (Braccini et al., 2005). A perfusionsystem has also been used for the culture of achondrocyte seeded composite scaffold (PGA/PLLAfleece adhered to calcium carbonate), resulting in theformation of new ECM which adhered to the underlyingceramic biomaterial (Kreklau et al., 1999).

Optimizing a convective system for the engineering ofcartilaginous and osteoinductive grafts will have toaddress the balance between the mass transfer ofnutrients and waste products to and from cells, theretention of newly synthesized extracellular matrixcomponents within the construct, and the fluid-inducedshear stresses within the scaffold pores. The optimalflow conditions of a bioreactor should not be deter-mined through a trial-and-error approach, but rathershould be supported by computational fluid dynamics(CFD) models to simulate the velocity and shear profileswithin bioreactors and at the surface of engineeredconstructs (Sucosky et al., 2004; Williams et al., 2002),

as well as within the scaffold pores to better characterizethe local environment seen by the cells (Cioffi et al.,2006; Porter et al., 2005; Raimondi et al., 2002, 2004).

3.4. Required in vitro maturation stage of engineered

cartilage

A key question to be addressed in tissue engineering ishow developed an engineered graft should be prior toimplantation to support an optimal repair. Whileparticular biomechanical parameters have been pro-posed as design criteria for functional cartilage grafts(Hung et al., 2004), including local mechanical proper-ties reflecting cartilage heterogeneity (Kelly et al., inpress), these have yet to be validated, and existing datado not allow to reach a consensus.

In the most basic cell/scaffold-based approach tocartilage repair, cells are seeded onto the scaffold andimmediately, or soon thereafter, implanted. When usingthis method, however, the number of cells actuallyretained within the scaffold can be significantly lowerthan when pre-culturing constructs in vitro prior toimplantation, possibly due to enhanced cell-scaffoldadhesion or to the ECM stabilizing and protecting thecells (Ball et al., 2004). Yet, only minor differences in theectopic in vivo development of engineered septalcartilage were reported when scaffolds were eitherimmediately implanted after seeding or pre-culturedfor 3 weeks (Rotter et al., 2002). On the contrary,human articular chondrocyte-based constructs pre-cultured in a medium containing factors supportingchondrogenic differentiation were reported to containmore cartilaginous ECM and have higher equilibriummoduli following ectopic implantation than thoseimplanted directly after seeding (Moretti et al., 2005a).

Given the physiological load bearing requirementsthat an osteochondral graft must support once im-planted, the role of biomechanical signals will likelydetermine the key design criteria of a cartilage graft. Awell-developed cartilaginous ECM could allow for earlypost-operative rehabilitation and enhance the successrate of the graft survival when subjected to physiologicmechanical loads. In addition, considering the role ofECM in modulating the chondrocyte response tomechanical forces (Buschmann et al., 1995; Demarteauet al., 2003), a more developed ECM could provide amore physiological environment that supports further invivo development of the graft. Using an in vitro modelsystem, Demarteau et al. showed a correlation betweenthe effect of dynamic compression and the amount ofGAG contained within the engineered construct at theonset of compression (Fig. 5) (Demarteau et al., 2003).With the immature constructs, dynamic compressioninduced a down regulation in the GAG metabolism (i.e.,GAG synthesis and accumulation), while only in themost mature constructs was the loading advantageous

ARTICLE IN PRESS

0.4

0.6

0.8

1.0

1.2

1.4

GAG content before compression

Fo

ld d

iffe

ren

ce

fro

m u

nst

imu

late

d

r=0.783

0.0

0.5

1.0

1.5

2.0

2.5

0.0 10.0 15.0

GAG content before compression

Fo

ld d

iffe

ren

ce

fro

m u

nst

imu

late

d

GAG Synthesized

GAG Accumulated

5.0 10.0 15.00.0

p=0.002

r=0.912

p=0.035

5.0

(A)

(B)

Fig. 5. Correlation plots of the effect of dynamic compression and

maturation stage of engineered cartilage. Dynamic compression was

applied to pre-cultivated engineered cartilage constructs containing

different amounts of GAG at the onset of stimulation. In relation to

unstimulated constructs under freeswelling conditions, the effect of

dynamic compression on GAG synthesis (A) and GAG accumulation

(B) correlated with the initial amount of GAG within the constructs.

Only in the most mature constructs did compression result in an

upregulation in the GAG synthesized and accumulated as compared to

unstimulated constructs (i.e., points lying above the dashed line)

(quantities of GAG are reported following normalization to corre-

sponding DNA amounts (mg GAG/mg DNA)) (Demarteau et al.,

2003).

I. Martin et al. / Journal of Biomechanics 40 (2007) 750–765760

to the construct development. Bioreactors such as theone described by Demarteau et al. could be used ascontrolled model systems to assess when a construct isbiomechanically ready to be implanted, and could thushelp to answer the question ‘‘how good is goodenough’’.

While a more mature cartilaginous graft may impartgreater biomechanical functionality, less mature graftsmay have a greater capacity to integrate with theadjacent bone-substitute and the surrounding nativecartilage tissue. Using an in vitro model system,Obradovic et al. showed that while cartilaginousconstructs pre-cultured for 5 days had lower compres-sive stiffness than those pre-cultured for 5 weeks, theimmature constructs integrated better with adjacentcartilage explants during in vitro culture (Obradovicet al., 2001). Likewise, integration at the cartilage/boneinterface of osteochondral composites was better whenless mature cartilaginous constructs were sutured to thebone-substitute and cultured together in vitro (Schaeferet al., 2000). The inferior integrative capacity seen with

more mature constructs is consistent with the limitedintegrative capacity of native cartilage and could be theresult of low cell proliferation, cells being entrappedwithin their dense ECM, or the anti-adhesive propertiesof GAG (Obradovic et al., 2001). Nevertheless, given thepotential adverse consequences of physiological me-chanical loading on an immature cartilaginous graft,methods to improve the integrative capacity of maturecartilaginous grafts should be investigated. The use ofreproducible and controlled model systems to investi-gate specific mechanisms of integrative cartilage repairwill be essential in this endeavor (Moretti et al., 2005b).

3.5. Manufacture of engineered osteochondral composites

One of the major challenges to bring an autologouscell-based osteochondral product into routine clinicalpractice would be to translate research-scale productionmodels into clinically applicable manufacturing designsthat are reproducible, clinically effective, and economic-ally acceptable while complying with good manufactur-ing practice (GMP) requirements (Ratcliffe andNiklason, 2002). Production techniques for tissueengineering products currently rely on manual cellculture procedures, which suffer from high costs,complicated logistics and little standardization: thus,innovative and low-cost bioreactor systems might play akey role in the successful exploitation of an engineeredosteochondral product for wide-spread clinical use(Martin et al., 2004; Kino-Oka et al., 2005). A promisingmanufacturing concept is based on a de-centralized andclosed bioreactor system, such as the on-site hospitalbased ACTESTM (Autologous Clinical Tissue Engineer-ing System), under development by Millenium Biologix(www.millenium-biologix.com). As a fully automatedbioreactor system, ACTESTM will digest a patient’scartilage biopsy, expand the chondrocytes, seed andculture the cells onto the surface of an osteoconductiveporous scaffold, thereby generating the CartiGraftTM

osteochondral graft within a single closed bioreactorsystem. Bioreactor systems such as ACTESTM wouldeliminate logistical issues of transferring specimensbetween locations, eliminate the need for large andexpensive GMP tissue engineering facilities, and mini-mize operator handling, with the likely final result ofreducing the cost of engineered osteochondral grafts.

An interesting option to bypass the cultivation of cellsin the manufacture of osteochondral grafts is based onthe concept of an ‘‘in vivo bioreactor’’. The approachlies in the deliberate creation and manipulation of aspace between the tibia and the periosteum, a mesench-ymal layer rich in MPC, in such a way that the body’shealing mechanism is leveraged in the engineering ofneo-tissue (Emans et al., 2005). In a recent study,Stevens et al. demonstrated that by controlling angio-genesis it is possible to determine the formation of

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cartilage or bone tissues, thus opening the way to theectopic engineering of osteochondral composites, whichcan then be transplanted into joint defects (Stevens etal., 2005).

3.6. Osteochondral repair by cell-free scaffolds?

From a surgical and commercial standpoint, an idealgraft for osteochondral defect repair would be an off-the-shelf product, possibly free of the component whichtypically introduces the largest variability, namely cells.Is it possible for a cell-free implant to be sufficiently‘‘intelligent’’ to bring into the joint the appropriate cuesto induce orderly and durable tissue regeneration? Thedifferent approaches so far investigated to fabricate cell-free osteochondral composites have been developedbased on a variety of fundamental principles, but withthe common goal to provide the correct ‘‘regenerative’’signals to local cells. Mesenchymal cells from thesynovial membrane have been shown to be recruitedinto partial-thickness cartilage defects by local treatmentwith chondroitinase ABC, followed by filling of thedefect with a fibrin clot containing TGF-b1 (Hunzikerand Rosenberg, 1996). In order to extend the approachto osteochondral defects, it was proven to be necessaryto include an additional component to prevent rapidupgrowth of osseous tissue and vascular buds in thecartilage compartment, namely a blood vessel-excludingmembrane (Hunziker et al., 2001) or an anti-angiogenicfactor (Hunziker and Driesang, 2003). MPC from thesubchondral bone have been considered as the targetcells in an approach where microfracture was combinedwith the delivery of a chitosan-based gel, whichstabilized clot formation and resulted in improvedquantity and quality of repair cartilage in sheep(Hoemann et al., 2005). In order to selectively differ-entiate MPC recruited from the bone marrow, the use ofspecific scaffold compositions has also been explored. Inthis direction, a short-term study performed in rabbitsindicated that type I collagen-based cell-free matricesefficiently recruited osteoprogenitor cells from thesubchondral bone, whereas type II collagen-basedmatrices were more efficient to direct cells into achondrogenic phenotype (Buma et al., 2003). Based onthese observations, the authors proposed the concept ofa composite cell-free matrix consisting of a deep layer ofcollagen type I and a superficial layer of collagen type II.Another study reported the promising potential ofmultiphase implants based on polylactide-co-glycolideand various additives, used to vary the stiffness of thesuperficial and deep regions, for treatment of osteo-chondral defects (Niederauer et al., 2000). Interestingly,when implanted into a high weight-bearing region (goatcondyle), scaffolds containing a cartilage layer withhigher compressive stiffness (32 vs. 12MPa) rankedhigher in terms of overall repair, leading the authors to

speculate that it will be important to consider not onlythe composition, but also the mechanical properties ofthe scaffold.

The latter concept is in line with theories formechano-regulation of tissue differentiation, mostlydeveloped to model and predict patterns of fracturehealing (Carter et al., 1988; Lacroix and Prendergast,2002), and recently applied to the repair of osteochon-dral defects (Duda et al., 2005). In particular, amechano-regulation model for tissue differentiationwas recently used to determine the optimal mechanicalproperties of a construct to properly recruit, proliferateand differentiate MPCs from the bone marrow withinosteochondral defects, in order to induce orderlyformation of hyaline and bone tissues rather than offibrous tissue (Kelly and Prendergast, 2004). Thedeveloped criteria indicate that the bone region of thecomposite scaffold should be rigid enough to stabilizethe defect (e.g., Young’s modulus of 50MPa) and have apermeability sufficiently low (e.g., 2E�15m4/N s) toreduce large magnitudes of fluid flow that may preventosteogenesis at between the graft and native bone. Inaddition, the chondral region should possess a super-ficial layer (i) sufficiently rigid and impermeable toprevent cell death and fibrous tissue formation at thearticular surface, and (ii) with a gradual reduction instiffness (e.g., Young’s moduli from 60 to 10MPa) andincrease in permeability (e.g., from 1E�16 to 2E�15m4/N s) from the surface to the base of the chondral phaseto avoid inducing the dedifferentiation of chondrocytesin other regions of the defect (Kelly, 2003). While thedeveloped model includes several and sometimes arbi-trary assumptions, the study is a key milestone in thefield, since it represents a serious engineering effort todevelop quantitative principles of design for engineeredosteochondral composites, which should direct thesetting up of hypothesis-driven experimental studies.

4. Conclusions

By describing the wide variety of approaches cur-rently under investigation for the engineering ofosteochondral grafts, we have here provided an evidenceof the fact that an outcome-driven consensus overdesign principles is still far to be reached in thisrelatively new field. It should also be emphasized thatthe variability of the type and size of the defect inclinical cases, often not considered in standardizedanimal models, is expected to influence the suitabilityof the design. For example, it is possible that for smalland confined osteochondral lesions a cell-free approachwith appropriate scaffolds is sufficient, whereas for moreextended injuries the delivery of growth factors isnecessary for local cell recruitment, and where thewound bed is further compromised the use of a cell-

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based approach becomes mandatory. In the selection ofspecific strategies for the fabrication of osteochondralgrafts, it is important to consider that animal studiesmay be important to evaluate the safety and surgicalfeasibility of a procedure, or to test a general paradigmof tissue repair. However, due to the highly differentbiochemical and biomechanical milieu in animal andhuman joints, efficacy-driven guidelines could only bederived from prospective, randomized clinical trials.Finally, it is important to highlight that the patientpopulation targeted for treatment with engineeredosteochondral grafts is currently restricted to youngindividuals affected by traumatic injuries or by osteo-chondritis dissecans: future studies in the field shouldaddress if and how the same paradigm could beextended to treatment of degenerative joint pathologiesin the ageing population.

Acknowledgments

This work was partially supported by the SixthEuropean Framework Program (Integrated Project‘‘STEPS’’, Grant no. NMP3-CT-2005-500465).

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