WO2012021885A1 - Three-dimensional tissue engineering devices and uses thereof - Google Patents

Abstract

Three-dimensional tissue engineering devices are provided. These devices are useful for integration of soft tissue grafts with bone. Also provided are methods for their use as osteochondral interfaces.

Description

Three-dimensional Tissue Engineering Devices and Uses

Thereof

This patent application claims the benefit of priori from U.S. Provisional Application Serial No. 61/401,465, filed August 13, 2010, the entirety of the disclosure of which is explicitly incorporated by reference herein.

Statement Regarding Federally Sponsored Research or

Development

This invention was made with government support under

Grant Numbers NIH-NIA S 5R01AR0545280-02, NIH-NIA S

AR056459-01, and NIH-NIAMS 5R01AR055280-02 awarded by the National Institutes of Health. The government has certain rights in the invention.

BACKGROUND

Osteoarthritis is one of the leading causes of

disability among U.S. adults with 17.4 million adults experiencing activities limitation attributed to arthritis and 1 in 20 working-age adults experiencing limitations. The disease affects over 26.9 million adults in the U.S., affecting 13.9% of adults aged 25 and older, and 33.6% of those over 65 years of age.

Cartilage has limited capacity for self-repair as aneural and avascular tissue with restricted access to reparative cells and humoral factors. Options for cartilage repair include lavage, periosteal grafts, subchondral drilling, and mosaicplasty . However, there are limitations to these procedures including donor-site morbidity,

fibrocartilage formation, and poor graft-bone integration. SUMMARY

An aspect of this application relates to three- dimensional tissue engineering devices comprising a bottom portion and a side wall portion which are adjoined to form a hollow center portion. The bottom, side wall and hollow center portions function collectively to form a vessel or cup .

In one embodiment, the bottom portion and side wall portion comprise a mesh configured to permit cell migration from one side of the mesh through to an opposite side of said mesh. In one embodiment, this mesh comprises polymer nanofibers . The bottom and side wall portions may comprise the same mesh to promote growth and/or maintenance of the same selected tissue type or different mesh to promote growth and/or maintenance of different selected tissue types. In one embodiment, the bottom portion comprises a mesh which promotes growth and/or maintenance of cartilage and/or chondrocytes and the side wall portion comprises a mesh which promotes growth and/or maintenance of bone and/or osteoblasts. In this embodiment, the bottom of the cup promotes calcified cartilage formation and/or

osteointegration or binding with bone. In other

embodiments, the side wall and/or bottom portions may promote bone formation.

Devices of the present invention may further comprise one or more flaps at the top of the side wall portion for fixation of the cup or vessel.

Devices of the present invention may further comprise a cap for the cup or vessel.

In any of these embodiments, one or more agents promoting growth and/or maintenance of a selected tissue type can be added to the side wall and or bottom portions of the three dimensional tissue engineering device. Another aspect of this application relates to the above-referenced three dimensional tissue engineering device further comprising an insert member fitted into the hollow center portion of the tissue engineering device.

In one embodiment, the insert member may comprise a hydrogel or a plurality of hydrogel layers. In this embodiment, the hydrogel or hydrogel layer may be seeded with a selected cell type.

Alternatively, the insert member may comprise a tissue allograft or autograft.

The insert member may further comprise one or more agents which promote growth and/or maintenance of a selected tissue type.

Another aspect of this application relates to use of these three-dimensional tissue engineering devices as osteochondral interface tissue engineering devices.

Brief Description of the Drawings

Figure 1 shows fabrication of one embodiment of a three-dimensional tissue engineering device of this

application. In this embodiment, different meshes are used for the bottom and side wall portions and are adjoined via sintering to form the hollow center portion.

Figures 2A and 2B show fabrication of another

embodiment of a three-dimensional tissue engineering device of this application. In this embodiment, the bottom and side wall portions comprise a plurality of multi-phased, mesh sheets (Figure 2A) . As shown in Figure 2B, the plurality of sheets are folded to form a vessel wherein the side portion of mesh comprises aligned polymer nanofibers and the bottom portion comprises aligned polymer nanofibers and hydroxyapatite (HA) . Figure 2B further depicts

insertion of the three-dimensional tissue engineering device into a surgically relevant site in need of osteochondral interface regeneration.

Figures 3A through 3C show various nonlimiting examples of multi-phased meshes which can be used for the bottom and side wall portions of the three-dimensional tissue

engineering devices of this application.

Figure 4A through 4D show various nonlimiting examples of hydrogel insert members fitted into the hollow center portion of the tissue engineering device of this

application.

Figure 5 shows an embodiment of a tissue engineering device of this application wherein the insert member comprises an allograft fitted into the hollow center portion of the tissue engineering device.

Figures 6A and 6B show characterization of the

acellular alginate scaffolds of Example 2. In Figure 6Ά, environmental SEM was used to image acellular alginate (Alg) and alginate+HA (Alg+HA) scaffolds (Day 0, 15 kV, lOOx, bar=250 mm) . Corresponding EDAX analysis indicated the presence of both Ca and P peaks. Figure 6B shows the swelling ratio of acellular Alg scaffolds to be

significantly higher than Alg+HA scaffolds for all time points (n=4, *p<0.05). While both Alg and Alg+HA scaffolds experience an initial increase in swelling ratio on Day 2, their swelling behavior differs dramatically after Day 7 (n=4, #p<0.05) .

Figures 7A and 7B show characterization of DZC-seeded alginate scaffolds of Example 2. Figure 7A shows day 10 chondrocytes maintaining a rounded morphology within the alginate scaffolds, and intact, individual particles of hydroxyapatite are also seen in environmental scanning electron micrographs (15kV, lOOx, bar=500 mm; lOOOx, bar=50 mm) . Live/dead staining shows comparable cell viability of DZCs seeded in both alginate and alginate+HA scaffolds on Day 28 (n=2, lOx, bar=200 μπι) . Figure 7B shows the addition of hydroxyapatite to the alginate scaffold significantly decreases cell number on Day 14, 21, and 28 (n=6, *p<0.05) . There is also a significant increase in cell number until Day 21 for the alginate scaffold, whereas an increase in cell number only occurs from Day 1 to Day 7 for the

alginate+HA group (n=6, #p<0.05) .

Figures 8A and 3B show collagen and glycosaminoglycan deposition in alginate scaffolds of Example 2. As shown in Figure 8A, the addition of hydroxyapatite to the alginate scaffold significantly increases collagen deposition on Day 28 (n=6, *p<0.05), and there is a significant increase for both scaffolds from Day 14 to Day 28 (n=6, #p<0.05). Figure 8B shows Picrosirius Red and Alcian Blue staining of DZC- seeded alginate scaffolds on Day 28 indicative of collagen and GAG deposition, respectively (n=2, lOx, bar=200 μπι) .

Figures 9A through 9C show mechanical properties of alginate scaffolds of Example 2. The compressive modulus (Figure 9A) , magnitude of the complex shear modulus (Figure 9B) , and phase shift angle (Figure 9C) of both acellular (Acell) and DZC-seeded alginate (Alg) and alginate+HA

(Alg+HA) scaffolds are graphed. Relative to their

corresponding acellular controls, the addition of

hydroxyapatite to the alginate scaffold significantly increases all three mechanical properties on Day 28 (n=3, *p<0.05). There is also a significant increase in all three mechanical properties from Day 14 to Day 28 for the

alginate+HA scaffolds and a significant decrease from Day 14 to Day 28 for the alginate scaffolds (n=3, #p<0.05).

Figures 10A through IOC show mineralization potential and mineral deposition in alginate scaffolds of Example 2. As shown in Figure 10A, the addition of hydroxyapatite to the alginate scaffolds significantly decreases ALP activity on Day 5, 6, and 7 (n=6, *p<0.05). There is also a

significant decrease from Day 5 to Day 6 for both scaffolds (n=6, #p<0.05). Figure 10B shows Von kossa staining of DZC- seeded alginate scaffolds indicative of both the presence of scaffold-HA and matrix mineralization (n=2, lOx, bar=100 mm) . As shown in Figure IOC, the addition of hydroxyapatite to the alginate scaffolds also significantly decreases calcium concentrations (n=6, *p<0.05).

Figures 11A through 11C show expression of hypertrophic markers in the scaffolds of Example 2. The expression of SOX9 (Figure 11A) , MMP13 (Figure 11B) , and Collagen X

(Figure 11C) by deep zone chondrocytes in alginate (Alg) and alginate+HA (Alg+HA) scaffolds were compared. The addition of hydroxyapatite significantly increases P13 and Col X expression on Day 14 (n=3, *p<0.05) . Furthermore, there is a significant decrease in MMP13 expression from Day 14 to Day 28 in the alginate+HA group, while there is an increase in Collagen X (n=3, #p<0.05).

Figures 12A through 12D show comparisons between DZC and FTC response in the scaffolds of Example 2. DZC-seeded alginate scaffolds support cell proliferation, whereas FTC- seeded alginate scaffolds do not (Figure 12A) . The FTC- seeded scaffolds have a significantly lower ALP activity as compared to corresponding DZC-seeded scaffolds (Figure 12B) . The increase in collagen deposition for the DZC-seeded alginate+HA scaffold is not seen for the FTC-seeded

alginate+HA scaffolds (Figure 12C) . The FTC-seeded

scaffolds also have significantly decreased levels of GAG deposition for both alginate and alginate+HA scaffolds (Figure 12D) (n=6, *p<0.05).

Figures 13A and 13B show characterization of a biphasic alginate+HA scaffold with continuous yet distinct phases of gel and gel+mineral of Example 1. Mineral presence

identified by backscattered SEM is shown in Figure 13A.

Figure 13B shows ALP activity of DZCs cultured on mineral- free alginate (C) , alginate+HA and alginate+BG scaffolds (p<0.05) .

Figures 14A through 14C show GAG production over time (Figure 14A) , collagen content on Day 28 (Figure 14B) , normalized by wet weight (p<0.05), as well as histology for Day 14 alginate+BG constructs (Figure 14C) indicative of biphasic distribution of mineral, collagen and GAG (lOx, bar=200 m) in scaffolds of Example 1.

Figures 15A and 15B show mechanical properties of the scaffold of Example 1. Scaffold shear modulus (Figure 15A) and phase shift angle (Figure 15B) were compared to

acellular controls for each time point (p<0.05) . 1 Hz data is compared in these figures.

Figure 16 shows SEM images (lOOx) and corresponding EDX analysis of a fabricated scaffold from Example 3.

Figures 17A and 17B show cell proliferation on a scaffold of Example 3 in T3-free (Figure 17A) and T3-treated (Figure 17B) groups (*:p<0.05 between groups, #:p<0.05 over time) .

Figures 18A and 18B show matrix deposition in a scaffold of Example 3. Figure 18A shows GAG content and corresponding Alcian Blue staining (lOx, day 14) . Figure 18B shows collagen content and corresponding Picrosirxus Red staining (lOx, day 14) (*:p<0.05 between groups, #:p<0.05 over time) .

Figure 19A and B shows mineralization in a scaffold of Example 3 and ALP (Figure 19B) activity of T3-treated groups with corresponding Alizarin Red staining (Figure 19A;10x, day 7) (*:p<0.05 between groups, #:p<0.05 over time). Figure 20 shows mechanical properties of a scaffold of Example 3. In particular, phase shift angle of both

acellular and DZE-seeded scaffolds (*:p<0.05 between cell- seeded groups, #p<0.05 between acellular and cell-seeded groups) is shown.

Figures 21A and 21B show increased cell proliferation (Figure 21A) and media calcium concentration (Figure 21B) in 2D culture with addition of 10 mg BG (*p<0.05) in scaffolds of Example 4.

Figures 22A and 22B show a comparison of GAG deposition in 3D culture when BG is added both inside (Figure 22A) and outside (Figure 22B) to the scaffold of Example 4 (*p<0.05).

Figures 23A and 23B show ALP activity (Figure 23A) and media calcium concentration (Figure 23B) in 3D culture with addition of BG to the scaffold of Example 4 (*p<0.05).

Figures 24A through 24C provide a schematic of a procedure for culturing cells on scaffold as described in Example 5. As shown in Figure 24A, the scaffold is first immersed in 10 g/mL fibronectin for 12 hours. As shown in Figure 24B, bovine DZC is then seeded at 90,000 cells/cm² and allowed to attach for 15 minutes. As shown in Figure 24C, the scaffold is then maintained in ITS+media and 50 g/mL of ascorbic acid.

Figure 25 shows the effects of HA on DZC proliferation on a scaffold of Example 5. Deep zone chondrocytes remain viable and increased in number over time on both types of scaffolds. In addition, cell number decreased on PLGA scaffolds from day 7 to day 14 (*:p<0.05 over time.

Figures 26A and 26B show effects of HA on matrix deposition in a scaffold of Example 5. Collagen (Figure 26A) and proteoglycan (Figure 26B) production increased over time on both scaffold types (*:p<0.05 over time) . In addition, there is significantly higher proteoglycan content on PLGA scaffolds by day 14 (*:p<0.05 between groups).

Figure 27 shows the effects of HA on DZC mineralization in a scaffold of Example 5. Alkaline phosphatase (ALP) activity is significantly higher on PLGA at day 1 and decreased thereafter (*:p<0.05) . In addition, ALP activity is maintained in PLGA+HA scaffolds over time.

Figure 28 shows the effects of HA on DZC hypertrophy on a scaffold of Example 5. Hypertrophic markers were expressed on both scaffold types. In addition, there is no significant difference observed between PLGA and PLGA+HA scaffolds

Figures 29A and 29B show a comparison of ALP activity on Agarose and Agarose+HA hydrogel scaffolds (Figure 29A) and PLGA and PLGA nanofiber+HA scaffolds (Figure 29B) .

Figures 30A and 30B show normalized GAG content on PLGA and PLGA+HA nanofibers seeded with chondrocytes.

Figures 31A and 31B show normalized ALP content on PLGA and PLGA+HA nanofibers seeded with chondrocytes.

Figures 32A through 32D show a schematic of various tissue engineering device embodiments of this application. Figure 32A depicts an embodiment with a cup or vessel formed from PLGA mesh bottom and side wall portions and an agarose insert member serving as a hydrogel-based cartilage graft. Figure 32B depicts an embodiment in which the bottom portion of the cup or vessel further comprises HA. Figure 32C depicts an embodiment in which the agarose insert member further comprises HA. Figure 32D depicts an embodiment in this the bottom portion of the cup or vessel further comprises HA and the agarose insert member further comprises HA.

Figures 33A through 33H provide photographs and show characterization of the embodiment of the tissue engineering device depicted in Figure 32B. Figure 33A is photograph of the cup or vessel without the insert member from the top. Figures 33B through D are photographs depicting the side view (Figure 33B) , the bottom view (Figure 33C) and top view (Figure 33D) of the cup and vessel with the insert member. In Figure 33E and 33F, environmental SEM was used to image the side portion (Figure 33E) and bottom portion (Figure 33F) of the cup or vessel. Corresponding EDAX analyses of the side wall and bottom portions are depicted in Figures 33G and 33H, respectively.

Figures 34A and 34B show proteoglycan deposition in the device embodiments depicted in Figures 32A through 32D. The PLGA and PLGA+HA cups or vessels increased proteoglycan deposition in the hydrogel-based cartilage graft. Addition of HA to the hydrogel graft also increased proteoglycan deposition.

Figures 35A and 35B show collagen deposition in the device embodiments depicted in Figures 32A through 32D. The PLGA and PLGA+HA cups or vessels increased collagen deposition in the hydrogel-based cartilage graft. Addition of HA to the hydrogel graft also increased collagen deposition .

Figures 36A and 36B show mineralization potential in the device embodiments depicted in Figures 32A through 32D. The PLGA and PLGA+HA cups or vessels increased mineralization potential in the hydrogel-based cartilage graft .

DETAILED DESCRIPTION

Definitions

In order to facilitate an understanding of the material which follows, one may refer to Freshney, R. Ian. Culture of Animal Cells - A Manual of Basic Technique (New York: Wiley- Liss, 2000) for certain frequently occurring methodologies and/or terms which are described therein.

Unless defined otherwise, all technical and scientific terms used herein have the meaning commonly understood by a person skilled in the art to which this application belongs. However, except as otherwise expressly provided herein, each of the following terms, as used in this application, shall have the meaning set forth below.

As used herein, "agent" shall mean a component

incorporated into the bottom portion, side portion and/or insert member of a three-dimensional tissue engineering device of this application which, when released over time, supports alignment, proliferation and matrix deposition of a selected musculoskeletal cell and/or promotes cell migration and integration of a selected tissue such as cartilage or bone. Examples of such agents include, but are in no way limited to ceramics, in particular calcium phosphate (CaP) ceramics, bioactive glasses and glass ceramics and

hydroxyapaptite (HA) , growth factors such as platelet- derived growth factor (PDGF), transforming growth factor- beta 3(TGF- 3), fibroblast growth factor (FGF), basic fibroblast growth factor (bGF) , growth/differentiation factor-5 (gdf-5) and insulin derived growth factors, bone morphogenetic proteins, platelet-rich plasma, and anti- angiogenesis factors. By "agent" it is also meant to be inclusive of components which promote differentiation of stem cells to a selected cell type. A single agent or a combination of agents may be incorporated into the tissue engineering scaffolds of this application.

As used herein, "aligned fibers" shall mean groups of fibers which are oriented along the same directional axis. Examples of aligned fibers include, but are not limited to, groups of parallel fibers. As used herein, "ALP activity" shall mean alkaline phosphatase activity.

As used herein, "bioactive" shall include a quality of a material such that the material has an osteointegrative potential, or in other words the ability to bond with bone. Generally, materials that are bioactive develop an adherent interface with tissues that resist substantial mechanical forces .

As used herein, a "biocompatible" material is a synthetic or natural material used to replace part of a living system or to function in intimate contact with living tissue. Biocompatible materials are intended to interface with biological systems to evaluate, treat, augment or replace any tissue, organ or function of the body. The biocompatible material has the ability to perform with an appropriate host response in a specific application and does not have toxic or injurious effects on biological systems. Nonldmiting examples of biocompatible materials include a biocompatible ceramic, a biocompatible polymer or a

biocompatible hydrogel.

As used herein, "biodegradable" means that the

material, once implanted into a host, will begin to degrade.

As used herein, "biomimetic" shall mean a resemblance of a synthesized material to a substance that occurs naturally in a human body and which is not substantially rejected by (e.g., does not cause an unacceptable adverse reaction in) the human body. When used in connection with the tissue scaffolds, biomimetic means that the scaffold is substantially biologically inert (i.e., will not cause an unacceptable immune response/rejection) and is designed to resemble a structure (e.g., soft tissue anatomy) that occurs naturally in a mammalian, e.g., human, body and that promotes healing when implanted into the body. As used herein, "chondrocyte" shall mean a

differentiated cell responsible for secretion of

extracellular matrix of cartilage.

As used herein, "chondrogenesis" shall mean the

formation of cartilage tissue.

As used herein, "fibroblast" shall mean a cell of connective tissue, mesodermally derived, that secretes proteins and molecular collagen including fibrillar

procollagen, fibronectin and collagenase, from which an extracellular fibrillar matrix of connective tissue may be formed .

As used herein, "hydrogel" shall mean any colloid in which the particles are in the external or dispersion phase and water is in the internal or dispersed phase.

As used herein, "matrix" shall mean a three-dimensional structure fabricated from biomaterials . The biomaterials can be biologically derived or synthetic.

As used herein, "mesh" means a network of material. The mesh may be woven synthetic fibers, non-woven synthetic fibers, microfibers and nanofibers suitable for implantation into a mammal, e.g., a human. The woven and non-woven fibers may be made according to well known techniques . The mesh may be made according to techniques known in the art and those disclosed in, e.g., co-owned international application no. PCT/US2008/001889 filed on February 12, 2008 to Lu et al . , which application is incorporated by reference as if recited in full herein.

As used herein, "microfiber" shall mean a fiber with a diameter no more than 1000 micrometers.

As used herein, "nanofiber" shall mean a fiber with a diameter no more than 1000 nanometers.

In one embodiment, the microfibers and/or or nanofibers are comprised of a biodegradable polymer that is electrospun into a fiber. The microfibers and/or nanofibers of the scaffold are oriented in such a way (i.e., aligned or unaligned) so as to mimic the natural architecture of the soft tissue to be repaired. Moreover, the microfibers and/or nanofibers and the subsequently formed microfiber and/or nanofiber scaffold are controlled with respect to their physical properties, such as for example, fiber diameter, pore diameter, and porosity so that the mechanical properties of the microfibers and/or nanofibers and

microfiber and/or nanofiber scaffold are similar to the native tissue to be repaired, augmented or replaced.

As used herein, "microspheres", mean microbeads, which are suitable, e.g., for cell attachment and adhesion.

Microspheres of a tissue scaffold may be made from polymers such as aliphatic polyesters, poly(amino acids),

copoly (ether-esters ) , polyalkylenes oxalates, polyamides, poly ( iminocarbonates ) , polyorthoesters , polyoxaesters , polyamidoesters , poly ( ε-caprolactone ) s , polyanhydrides , polyarylates , polyphosphazenes , polyhydroxyalkanoates , polysaccharides, or biopolymers, or a blend of two or more of the preceding polymers. Preferably, the polymer comprises at least one of the following materials: poly ( lactide-co- glycolide), poly ( lactide ) or poly (glycolide) . More

preferably, the polymer is poly ( lactide-co-glycolide )

(PLGA) .

As used herein, "osteoblast" shall mean a bone-forming cell that is derived from mesenchymal osteoprogenitor cells and forms an osseous matrix in which it becomes enclosed as an osteocyte. The term is also used broadly to encompass osteoblast-like, and related, cells, such as osteocytes and osteoclasts .

As used herein, "osteogenesis" shall mean the

production of bone tissue. As used herein, "osteointegrative" shall mean having the ability to chemically bond to bone.

As used herein, "polymer" shall mean a chemical compound or mixture of compounds formed by polymerization and including repeating structural units. Polymers may be constructed in multiple forms and compositions or

combinations of compositions.

As used herein, "porous" shall mean having an

interconnected pore network.

As used herein, "stem cell" means any unspecialized cell that has the potential to develop into many different cell types in the body, such as mesenchymal osteoprogenitor cells, osteoblasts, osteocytes, osteoclasts, chondrocytes, chondrocyte progenitor cells, fibrochondrocytes , fibroblasts and fibroblast progenitor cells. Nonlimiting examples of

"stem cells" include mesenchymal stem cells, embryonic stem cells and induced pluripotent cells.

Osteochondral grafts have emerged as a promising alternative for treating osteoarthritis (Gao et al . Tissue Eng 2002 8:827-37; Gao et al . Tissue Eng 2001 7:363-71;

Sherwood et al . Biomaterials 2002 23:4739-51; Schaefer et al. Arthritis Rheum 2002 46:2524-34; Cao et al . Tissue Eng 2003 9 (Suppl 1) :S103-S112; Tuli et al . Tissue Eng 2004 10:1169-79; Alhadlag et al . J Bone Joint Surg Am 2005

87:936-44). However, there remains a challenge in consistent formation of a stable osteochondral interface. Currently, the formation of a stable interface between the cartilage and bone portions of the osteochondral graft poses a significant challenge (Hunziker et al. Clin Orthop Relat Res 2001 S182-S189) . Articular cartilage integrates with bone via the osteochondral interface, which consists of an interdigitated calcified cartilage region that separates the uncalcified, deep-zone cartilage region from the subchondral bone. The regeneration of a biomimetic soft tissue-to-bone interface will be critical for promoting the functional integration of osteochondral grafts, and thereby extending their clinical functionality (Hunziker et al. Clin Orthop Relat Res 2001 S182-S189) .

Exemplary Embodiments

This disclosure relates to three-dimensional tissue engineering devices. Nonlimiting embodiments of devices of this application and methodologies and components used to form such devices are depicted in Figures 1 through 5 as well as Figures 32A through D.

As shown in, for example Figure 1, in simplest form the three dimensional tissue engineering device of this

application comprises a bottom portion 2 and a side wall portion 3 which is adjoined to the bottom portion to form a hollow center portion 4 . The bottom portion 2 , the side wall portion 3 and the hollow center portion 4 collectively form a vessel or cup 5.

Devices of the present invention may further comprise one or more flaps 7 at the top of the side wall portion 3 for fixation of the cup or vessel at a defect site. In one embodiment, the one or more flaps are formed by a portion of the side wall which extends higher than the defect into which the device is surgically implanted.

Devices of the present invention may further comprise a cap 8 covering the opening at the top of the cup or vessel (See Figure 4D) . The cap 8 functions to close off the wound and/or promotes tissue growth. In one embodiment, the cap is formed by a portion of the side wall which extends higher than the defect into which the device is surgically

implanted . The bottom portion 2, side wall portion 3, flaps 7 and/or cap 8 comprise material or materials for promoting growth and maintenance of a selected tissue type or selected tissue types and/or integration with underlying tissues.

Examples of materials which can be used to form the cup or vessel include, but are in no way limited to fiber mesh, microspheres and hydrogel.

In one embodiment, the material comprises mesh configured to permit cell migration from one side of the mesh through to an opposite side of the mesh. As such, when inserted into the site requiring tissue formation, the device would allow cells from the natural tissue of the subject to migrate through the vessel made from mesh into the insert member, thereby promoting tissue integration. (See, e.g., Figure 5)

In one embodiment, the mesh comprises polymer fibers, microfibers and/or nanofibers . Examples of polymers which can be selected for the polymer fiber, microfiber and/or nanofiber mesh include, but are not limited to,

biodegradable polymers selected from the group consisting of aliphatic polyesters, poly(amino acids), modified proteins, polydepsipeptides , copoly ( ether-esters ) , polyurethanes , polyalkylenes oxalates, polyamides, poly ( iminocarbonates ) , polyorthoesters , polyoxaesters , polyamidoesters , poly(e- caprolactone ) s , polyanhydrides , polyarylates ,

polyphosphazenes , polyhydroxyalkanoates , polysaccharides, modified polysaccharides, polycarbonates,

polytyrosinecarbonates , polyorthocarbonates ,

poly (trimethylene carbonate), poly (phosphoester) s ,

polyglycolide, polylactides , polyhydroxybutyrates ,

polyhydroxyvalerates , polydioxanones, polyalkylene oxalates, polyalkylene succinates, poly (malic acid), poly(maleic anhydride), polyvinylalcohol , polyesteramides , polycyanoacrylates , polyfumarates , poly (ethylene glycol), polyoxaesters containing amine groups, poly ( lactide-co- glycolides) , poly(lactic acid)s, poly (glycolic acid)s, poly (dioxanone ) s , poly (alkylene alkylate) s, biopolymers, collagen, silk, chitosan, alginate, and a blend of two or more of the preceding polymers. In one embodiment, the polymer comprises poly ( lactide-co-glycolide ) .

Fibers, microfibers and/or nanofibers of the mesh can be aligned or unaligned, or a combination of both. Alignment may be selected to mimic the anatomy of the site requiring tissue formation. For example, the side wall portion may comprise aligned nanofibers while the bottom portion may comprise unaligned nanofibers. See Figure 3C .

In one embodiment, the bottom portion and/ side wall portion are formed by a plurality of sheets of mesh. The plurality of sheets can be arranged and combined in a number of ways.

For example, the plurality of sheets of mesh can overlap in at least one area. See, e.g., Figure 2A and Figure 2B.

The plurality of sheets of mesh can also be sintered together. For example, the bottom portion can be sintered to the side wall portion. See, e.g., Figure 1.

As will be understood by the skilled artisan upon reading this disclosure, alternative methods for joining the bottom and side wall portions such as, but not limited to adhesives, methods of solvent evaporation and polymer cross linking can be used.

In one embodiment, the bottom portion can be formed from one or more sheets of mesh, and the side wall portion can be formed from one or more sheets of mesh.

In another embodiment, the bottom portion and the side wall portion can be formed from a continuous mesh sheet. Accordingly, in one embodiment, the bottom portion and side wall portion may comprise the same mesh which promotes growth and/or maintenance of the same selected tissue type and/or integration with the same underlying tissue type. In another embodiment, the bottom portion may comprise a mesh which promotes growth and/or maintenance of a first selected tissue type and/or integration with a first selected underlying tissue type and the side wall portion may comprise a mesh which promotes growth and/or maintenance of a second selected tissue type and/or integration with a second selected underlying tissue type. For example, in this embodiment, the bottom portion may comprise a mesh which promotes growth and/or maintenance of cartilage and/or chondrocytes and/or integration with underlying cartilage tissue and the side wall portion may comprise a mesh which promotes growth and/or maintenance of bone and/or

osteoblasts and/or integration with underlying bone tissue, or vice versa.

The bottom portion 2 , side wall portion 3 , or both the bottom portion and side wall portion may further comprise an agent or agents which promote growth and/or maintenance of a selected tissue type. In embodiments wherein both the bottom and side wall portions further comprise an agent which promotes growth and/or maintenance of a selected tissue type, the agent may be the same in the bottom portion and side wall portion to promote growth and/or maintenance of the same selected tissue type. Alternatively, the agent in the bottom portion may be different from the side wall portion to promote growth and/or maintenance of two

different selected tissue types.

Any agent which, when released over time, supports alignment, proliferation and matrix deposition of a selected musculoskeletal cell and/or promotes cell migration and integration of a selected tissue such as cartilage can be incorporated into the bottom portion and/or side wall portion of the device of this application.

In one embodiment, the agent of the bottom portion, side wall portion, or both the bottom portion and side wall portion is a ceramic. The ceramic can be present in the apparatus as micro-sized particles, nano-sized particles or a mixture of both. In one embodiment, the ceramic comprises calcium phosphate. In another embodiment, the agent is hydroxyapatite (HA) or bioactive glass.

Thus, in one embodiment, the bottom portion can comprise a ceramic and a polymer, e.g., polymer nanofibers and ceramic particles, configured to support the growth and maintenance of the first tissue type, e.g., bone.

Additionally, the polymer nanofiber alignment and

orientation can be chosen to mimic the anatomy of the first tissue type, e.g., bone. Further, the side wall portion may comprise a polymer, e.g., polymer nanofibers, configured to support the growth and maintenance of the second tissue type, e.g., cartilage. The polymer nanofiber alignment and orientation can be chosen to mimic the anatomy of the second tissue type, e.g., cartilage.

In one embodiment, the agent is a growth factor such as platelet-derived growth factor (PDGF) incorporated into the side wall portion of the device to promote cell migration or functionalize with chondroitin sulfate or collagen peptide to promote cartilage-cartilage integration.

In one embodiment, the agent is a bone morphogenetic protein incorporated into the bottom portion of the device to promote osteointegration or the formation of calcified cartilage . In one embodiment, the agent is an anti-angiogenesis factor incorporated into the bottom portion of the device to prevent bony ingrowth.

Figure 2B shows a nonlimiting example of three polymer nanofiber and ceramic particle mesh sheets overlapping in one area and folded to form a cup-shaped three-dimensional tissue engineering device of this application.

Figures 3A-3C show three other embodiments with varying mesh alignment designs.

In one embodiment, the three dimensional tissue

engineering device further comprises an insert member 6 fitted into the hollow center portion of the tissue

engineering device. Figures 4A-4D, Figure 5 and Figures 32A through 32D depict embodiments of the device of this

application further comprising the insert member 6 and photographs of the device of Figure 32D are provided in Figures 33A through 33D. As shown therein, the insert member may also comprise, as a constituent thereof one or more agents which, when released over time, supports alignment, proliferation and matrix deposition of a selected musculoskeletal cell and/or promotes cell migration and integration of a selected tissue such as cartilage. The insert member can be synthetic or biologically-derived and can be seeded with cells relevant to the tissue to be engineered. In the case of device for osteochondral

interface tissue engineering, the insert member can be seeded with at least some of the cells for chondrogenesis , e.g., chondrocytes, and in particular, deep zone

chondrocytes or stem cells capable of differentiating into chondrocytes . Figure 4A shows one embodiment of the device wherein the insert member is a cylindrical hydrogel scaffold for cartilage generation. Figure 4B shows another

embodiment of the device wherein the insert member is a biphasic cylindrical scaffold comprising one layer of hydrogel material and another layer of hydrogel and

hydroxyapatite composite. Figure 4C shows another

embodiment of the device wherein the insert member is a triphasic cylindrical scaffold comprising three layers of hydrogel scaffold with surface zone chondrocytes in one layer, middle zone chondrocytes in a second layer adjoining the first layer, and deep zone chondrocytes in a third layer adjoining the second layer and positioned next to the subchondral bone. Figure 4D shows an embodiment similar to Figure 4A wherein the insert member is a cylindrical hydrogel scaffold for cartilage generation of the device with a cap covering the top of the cup or vessel. Figure 5 shows another embodiment of the device wherein the insert member is a cartilage allograft.

In one embodiment, the vessel 5 and the insert member 6 are used in combination to facilitate the formation of a stable interface which mimics that found in nature. Thus, the shape, composition, orientation, positioning, and other characteristics of the vessel 5 and the insert member 6 should be chosen for this end.

In one embodiment, the device is designed for

osteochondral interface tissue engineering. In this

embodiment, the various characteristics of the device can be selected to facilitate the growth and maintenance of cartilage tissue, bone tissue, and the cartilage-to-bone osteochondral interface. According to this embodiment, the side wall portion can be configured to facilitate cartilage growth and regeneration and the bottom portion can be configured to facilitate cartilage-to-bone interface/bone tissue growth and regeneration. To this end, the first material can include at least one of cells and agents for promoting bone tissue growth. Similarly, the second material can include at least one of cells and agents for promoting cartilage tissue growth. In this embodiment, the device further comprises an insert member comprising a hydrogel with a ceramic dispersed in at least a portion of the hydrogel or a cartilage allograft. In one embodiment, the insert member is seeded with chondrocytes or stem cells capable of differentiating into chondrocytes.

In each of the various embodiments described herein, one or more active pharmaceutical ingredients selected from a group comprising the following are introduced into the device: anti-infectives ; hormones, analgesics; antiinflammatory agents; growth factors; chemotherapeutic agents; anti-rejection agents; and RGD peptides.

All combinations of the various elements are within the scope of the invention.

This disclosure also provides a method for engineering tissue, for example, osteochondral interface tissue, in a subject in need thereof, the method comprising affixing the apparatus described herein to a site requiring tissue formation/a surgically relevant site, thereby engineering the tissue (e.g., osteochondral interface tissue) in the subject. A discussion of scaffolds for osteochondral repair is also provided in U.S. Application Publication US 2006- 0036331 Al, the entire contents of which are incorporated herein by reference. Further, PTHrP treatment is a

therapeutic aspect that can be incorporated either to the three-dimensional tissue engineering devices of this application and/or the insert member. Use of PTHrP for musculoskeletal tissue engineering is discussed in PCT International Application No. PCT/US2008/007323, filed June 11, 2008, the entire contents of which are herby

incorporated herein by reference. The ability of various tissue engineering devices of this application as depicted in Figures 32A through 32D to increase proteoglycan and collagen deposition and increase mineralization potential in a hydrogel-based cartilage graft was demonstrated.

Figures 34A and 34B show proteoglycan deposition in the device embodiments depicted in Figures 32A through 32D. As shown therein, the PLGA and PLGA+HA cup or vessel increased proteoglycan deposition in the hydrogel-based cartilage graft. Further, addition of agent HA to the insert member also increased proteoglycan deposition.

Figures 35A and 35B show collagen deposition in the device embodiments depicted in Figures 32A through 32D.

Again, addition of the PLGA and PLGA+HA cup or vessel increased collagen deposition in the hydrogel-based

cartilage graft. Addition of the agent HA to the insert member also increased collagen deposition.

Figures 36A and 36B show mineralization potential in the device embodiments depicted in Figures 32A through 32D. Addition of the PLGA and PLGA+HA cup or vessel also

increased mineralization potential in the hydrogel-based cartilage graft.

These results are indicative of the usefulness of the devices of this application as tissue engineering devices, particularly in applications wherein a universal interface is reguired such as, but not limited to, the osteochondral interface .

Advantages of the devices of this application include, but are in no way limited to, flexibility in use of either biological or synthetic grafts, biodegradability, biomimetic organization, osteointegrative ability, cartilage-cartilage integration, easy application during surgery, ease in coupling with hydrogel+HA grafts, use with adhesives to promote cartilage integration, acellular versus cellular flexibility, stem cell versus chondrocyte flexibility, use with or without pre-incorporated agents which, when released over time, support alignment, proliferation and matrix deposition of a selected musculoskeletal cell and/or promote cell migration and integration of a selected tissue such as cartilage, and the ability to serve as a scaffold in all types of cartilage grafts reguiring integration with cone and/or calcified cartilage formation.

The ability of other scaffolds to function as insert members and/or meshes for the bottom and side wall portions of the devices of the instant application was also

demonstrated .

For example, a biphasic scaffold of calcium phosphate ceramic and alginate hydrogel as described in Example 1 was evaluated for use as an insert member of this tissue engineering device of this application.

Alginate has been extensively used for chondrocyte culture and cartilage tissue engineering. It is a well characterized biopolymer with the advantage of being biocompatible, nonimmunogenic, and biodegradable. The polymer structure is comprised of linear block copolymer and unbranched polysaccharide chains based on D-mannuronic acid and L-guluronic acid that are crosslinked by divalent ion binding (Grant et al . FEBS Lett 1973 32:195-8; Marttinen et al. Biotechnol Bioeng 1989 33:79-89). The ambient gelation preserves the bioactivity of cells and factors . In addition to its mild gelation conditions, alginate does not have toxic byproducts or residues (Augst et al . acromol Biosci 2006 6:623-33). In addition, chondrocytes have been shown to maintain their morphology and phenotypic stability in alginate culture, as well as deposit a matrix rich in proteoglycans and collagen (Guo et al. Connect Tissue Res 1989; 19 277-97; Hauselmann et al . J Cell Sci 1994; 107 ( Pt 1) 17-27; Heywood et al . Tissue Eng 2004; 10 1467-79; Lee et al. Biomaterials 2007; 28 2987-93; Masuda et al. J Orthop Res 2003; 21 139-48; Wong et al . Tissue Eng 2002; 8 979-87).

Deep zone chondrocyte (DZC) growth and biosynthesis was evaluated in the hybrid scaffold.

Deep zone chondrocytes proliferated and remained viable in the biphasic scaffold, with an initial period of cell growth during first two-week of culture and stabilization thereafter. At day 28, the alginate+BG group supported a lower cell number than the control but was not significantly different from the alginate+HA group.

DZC ALP activity (Fig. 13B) was suppressed in the alginate+HA group on Day 5 relative to control, while no difference were found for cells in alginate+BG. Positive von Kossa staining was seen in the alginate+mineral phase of the scaffold (Fig. 14C) .

No significant difference in GAG production was found between groups (Fig 14A) . In contrast, the presence of HA or BG both significantly increased DZC collagen deposition by Day 28. Histological analysis revealed significant cell aggregation and matrix production in the alginate+mineral phase, with positive staining for GAG, collagen, as well as mineralization.

For day 28, both alginate+HA and alginate+BG groups measured significantly higher Young's modulus, shear modulus, and phase shift angle compared to the mineral-free control (Fig. 15)

Thus, as shown by these experiments, DZC culture on the biphasic scaffold resulted in the formation of non-calcified cartilage continuous with a calcified cartilage region, mimicking the organization of the osteochondral interface. The presence of the mineral phase significantly enhanced collagen production by DZCs and promoted the formation of a dense GAG and collagen-rich matrix. Moreover, the increased biosynthesis translated into significant increases in compressive and shear moduli, as well as phase shift angle, relative to the mineral-free control on Day 28. These observations are indicative of the formation of a stiffer and more viscoelastic matrix that is consistent with the

increased matrix collagen content (Mow et al . J Biomechanics 1984 17:377-394). DZC matrix production also compensated for alginate degradation, with significant temporal increases in both compressive and shear moduli over acellular controls.

Accordingly, results from these experiments showed that HA suppressed DZC mineralization potential resulting in significant lower matrix properties when compared to DZC cultured with a dynamic mineral phase such as BG. Moreover, this difference becomes more pronounced if the HA to cell ratio increased. Thus, results from this study collectively suggest that the biphasic scaffold design enabled spatial- control of DZC response, with both mineral presence and chemistry regulating DZC biosynthesis and mineralization potential .

The potential of a hybrid scaffold of hydroxyapatite (HA) and alginate hydrogel as described in Example 2 as the insert member in a device of this application for

osteochondral interface engineering was also evaluated.

Specifically, the effects of HA on matrix production and mineralization, as well as scaffold mechanical properties were determined. It was found that the HA phase of the hybrid scaffold enhanced DZC collagen production and promoted the formation of a dense proteoglycan and collagen- rich matrix. Over time, enhanced biosynthesis translated into significant increases in both compressive and shear moduli and a higher phase shift angle relative to the mineral-free control. Presence of HA also promoted the expression of classic hypertrophic markers by DZC, further demonstrating the potential for DZC-mediated formation of a calcified interface region in the hybrid scaffold system.

More specifically, acellular alginate and alginate+HA scaffolds were imaged using SEM (See Fig. 6A) . While the alginate scaffold was homogeneous and uniform in appearance, the alginate+HA scaffold exhibited a heterogeneous mineral distribution with increased HA content in the bottom third of the scaffold. Energy-dispersive X-ray analysis of the alginate scaffold confirmed the presence of sodium (Na) , chlorine (CI) , and calcium (Ca) as a result of the sodium- alginic salt and calcium chloride (CaCl₂) cross-linking reaction. The presence of the phosphorus (P) peak was only detected for the alginate+HA scaffolds, accompanied by increased peak intensity for Ca .

Both alginate and alginate+HA scaffolds exhibited a time-dependent swelling behavior (See Fig. 6B) .

Furthermore, the swelling ratio of the alginate group remained significantly higher than that of the alginate+HA scaffold for all time points examined and is due to the increased dry weight of HA within the alginate+HA group.

While there was an initial decrease in thickness for both acellular alginate and alginate+HA scaffolds, a significant increase was evident for all groups after 14 days.

Furthermore, the presence of HA resulted in increased scaffold thickness at all time points (See Table 2) .

The cell-seeded alginate and alginate+HA scaffolds measured significantly higher wet weight and thickness as compared to their corresponding acellular controls at 14 and 28 days. Furthermore, a significant increase in wet weight was observed for the cell-seeded alginate and alginate+HA scaffolds from 1 to 14 days. While there was a significant decrease in scaffold thickness for the cell-seeded alginate group, a significant increase in thickness was observed for the cell-seeded alginate+HA group at 28 days.

Chondrocyte clusters and matrix elaboration were evident throughout the alginate and alginate+HA scaffolds. Scanning electron microscopy confirmed the spherical

morphology of individual chondrocytes as well as their viability over time (See Fig. 7A) . Cells seeded in

alginate+HA scaffolds proliferated over the first week of culture, with cell number staggering thereafter. In

contrast, cell number increased in the control over time, with a significantly higher cell number at 14, 21, and 28 days as compared to the alginate+ΗΔ group (p<0.05, Fig. 7B) .

Matrix deposition was observed over time in both alginate and alginate+HA scaffolds. While collagen

production increased in both alginate and alginate+HA scaffolds during culture, the alginate+HA group measured a significantly higher collagen content as compared to the alginate group at 28 days (See Fig. 8A) . These results were confirmed with histological staining (See Fig. 8B) . While proteoglycan deposition also increased over time for all scaffold types, no significant difference between the alginate and alginate+HA groups was observed. Histological analysis revealed that matrix deposition was well

distributed throughout the scaffolds with localization of proteoglycan deposits surrounding chondrocyte clusters (See Fig. 8B) .

The mechanical properties of alginate and alginate+HA scaffolds were assessed both in compression and shear (See Fig 9A-9C) . The cell-seeded alginate scaffolds exhibited significantly higher compressive modulus, shear modulus, and phase shift angle as compared to the acellular controls at Day 14. By Day 28, there was a significant decrease in the relative increase between cell-seeded and acellular scaffold mechanical properties. Similarly, the cell-seeded

alginate+HA scaffolds had significantly higher compressive modulus, shear modulus, and phase shift angle as compared to the acellular alginate+HA scaffold at 14 days. When the cell-seeded groups were normalized relative to their corresponding acellular controls, it was evident that by 28 days, there was a significant increase in the relative mechanical properties of those scaffolds with HA.

Correlation analysis was performed in order to relate Day 28 matrix deposition to scaffold mechanical properties. Specifically, both proteoglycan and collagen content were normalized by wet weight and correlated with both

compressive modulus and shear modulus (See Fig. 9A-9C) .

Proteoglycan content correlated more highly with compressive modulus (R=0.838), whereas collagen content correlated more highly with the magnitude of the complex shear modulus (R=0.972). Combining total matrix content increased the degree of correlation for the compressive modulus (R=0.838 to R=0.868) .

Cells exhibited ALP activity in both the alginate and alginate+HA scaffolds, although ALP activity decreased over time for both groups. While there was no significant difference in ALP activity at time points after 14 days, there was a significant decrease in the ALP activity of DZC cells in the alginate+HA scaffold as compared to those seeded in the alginate scaffold at early time points (See Fig. 10A) . Mineral deposition visualized by von Kossa confirmed the presence of HA particles in the alginate+HA scaffold as well as cell-mediated matrix mineralization (See Fig. 10B) . The solution calcium concentration measured a significant increase for the alginate scaffolds as compared to plain media, whereas the alginate+HA scaffold media was not significantly different (See Fig. IOC).

The effects of the HA phase of the composite were examined on the expression of hypertrophic markers.

Specifically, expression of SOX9 was detected in both groups at 14 and 28 days, with no significant difference in levels between the alginate and alginate+HA groups (See Fig. 11A) . Similarly, MMP13 expression was evident in both the alginate and alginate+HA groups (See Fig. 11B) . While there was no change in MMP13 expression over time for the alginate group, a significant decrease was detected in the alginate+HA group. Furthermore, MMP13 expression was significantly higher in the alginate+HA group as compared to the alginate group at Day 14. While minimal collagen X expression was observed in the alginate group (See Fig. 11C) , significantly higher collagen X expression was evident in the alginate+HA group at both 14 and 28 days.

The response of DZC was also compared to those of FTC which consist of surface, middle, and deep zone

chondrocytes. Specifically, alginate and alginate+HA scaffolds were analyzed in terms of cell proliferation, mineralization potential, and matrix deposition. The DZC- seeded alginate scaffolds measured increased cell number as compared to the FTC-seeded alginate scaffolds at all time points, whereas increased cell number was increased for the alginate+HA scaffolds only on Day 7 (See Fig. 12A) . In terms of mineralization potential, the DZC-seeded scaffolds had increased ALP activity as compared to the FTC-seeded scaffolds (See Fig. 12B) .

No significant difference in collagen deposition was observed between the DZC-seeded and FTC-seeded alginate scaffolds. However, the increase in collagen deposition observed with the addition of HA in the DZC-seeded groups was not seen in the FTC-seeded groups (See Fig. 12C) . On the other hand, there was increased proteoglycan deposition for DZC-seeded alginate and alginate+HA scaffolds as

compared to the FTC-seeded scaffolds (See Fig. 12D) .

Results from these studies are further depicted in

Tables 1 and 2.

Table 1:

Table 1 shows physical properties of acellular and DZC- seeded alginate scaffolds. While the wet weight of the acellular scaffolds decreases from Day 1 to Day 28, the wet weight of DZC-seeded scaffolds increases (n=4, #p<0.05) . While there is an initial decrease in thickness for the acellular scaffolds, there is a significant increase from Day 14 to Day 28 (N=4, #p<0.05) . Furthermore, acellular alginate+HA scaffolds are significantly thicker than alginate scaffolds on Day 28 (n=4, *p<0.05) . There is also a significant difference in thickness between alginate and alginate+HA scaffolds on Day 28 (n=4, *p<0.05).

Table 2:

Table 2 shows correlation between matrix deposition and scaffold mechanical properties. On Day 28, both GAG and collagen deposition are correlated with compressive and shear modulus using a linear regression fit (n=6) . GAG correlates more closely with compressive modulus, and collage correlates more closely with shear modulus. The interaction of GAG and collagen improves the quality of both fits .

As shown by these experiments, the HA phase of the hybrid scaffold enhanced DZC collagen production and promoted the formation of a dense proteoglycan and collagen- rich matrix. Over time, enhanced biosynthesis translated into significant increases in both compressive and shear moduli and a higher phase shift angle relative to the expression of classic hypertrophic markers by DZC, further demonstrating the potential for DZC-mediated formation of a calcified interface region in this hybrid scaffold system.

A composite hydrogel scaffold of hydroxyapatite (HA) and agarose as described in Example 3 was also evaluated. Agarose has been successfully utilized for functional cartilage tissue engineering both in vitro and in vivo (Mauck et al . Osteoarthritis Cartilage 2006 14 (2 ): 179-89) . Inclusion of HA as a substrate phase for interface

formation is advantageous as it resembles the mineral of the calcified cartilage layer. At the interface, aggregates of HA platelets instead of individual crystals are observed (Arsenault et al . Calcif Tissue Int. 1988 43 ( 4 ) : 219-25 ) . In this study, the effect of HA presence and particle size was examined as well as the effects of triiodothyronine (T3) , a known promoter of hypertrophy (Cook, 1993) on deep zone chondrocyte (DZC) proliferation, matrix deposition, and mineralization in the hydrogel+HA scaffold.

A significant increase in cell number was observed by day 7 for all groups. The addition of micro-HA increased cell number, while T3 reduced cell number by day 14 in the absence of HA (See Fig. 17A and 17B) .

Extensive matrix deposition was observed over time for all groups. No significant difference in GAG and collagen deposition is observed in all groups without T3. In

contrast, significantly higher GAG and collagen deposition are found by day 14 in the micro-HA group with T3

stimulation (See Fig. 18A-D) .

In both untreated and treated T3 groups, the presence of micro-HA initially suppresses ALP activity at day 7. As expected, the addition of T3 enhances mineralization potential (See Fig. 19) . Alizarin Red staining revealed the presence of both scaffold-HA and cell-mediated mineralization .

Similar to the suppression of ALP activity by the presence of HA, there was a down-regulation of chondrocyte hypertrophic markers. This suppression was again reversed with the addition of T3. The cell-seeded scaffolds exhibited higher shear mechanical properties than the acellular controls due to matrix elaboration (See Fig. 20) . While the addition of ΗΆ at the current concentration (2%) had little effect on the compressive properties of the scaffolds, it resulted in increased shear modulus and phase shift, an indicator of matrix viscoelasticity, as compared to DZC grown in hydrogel without HA (See Fig. 20) .

Thus, overall in this study, T3 stimulated DZC

hypertrophy and promoted mineralization. Previous studies have shown that the presence of HA modulates DZC

biosynthesis by suppressing ALP activity and increasing collagen deposition (Rosenthal et al. J. Rheumatol. 1999 26 ( 2 ): 395-401) . Results herein indicate that these changes are also dependent on HA particle size, with the micro-HA stimulating the formation of a calcified matrix rich in GAG and collagen that resembles the calcified cartilage

interface .

The potential of 45S5 bioactive glass (BG) as an agent to be added to devices of the instant application in accordance with Example 4 for interface formation was also examined as BG has been shown to exhibit osteointegration potential (Jones et al . J. Miomed. Mater. Res. 2001

58(6):720-6) and promote the differentiation and

mineralization of nasal chondrocytes (Asselin et al .

Biomaterials 2004 25 (25) : 5621-5630) . In this study articular chondrocyte response to BG was evaluated in both 3D hydrogel and 2D direct cultures. Chondrocytes cultured with BG exhibited increased proliferation in 2D culture, with the highest cell number found in the lOmg BG group at day 7. In contrast, no significant change in cell number over the day 1 control was observed in 3-D culture. The presence of BG within the hydrogel significantly enhanced matrix production (Fig. 22A) , with a dose- dependent increase in GAG observed in 3D culture. Similar to the 2D control group, no significant change in GAG synthesis was observed when compared to the hydrogel group cultured with BG outside of the hydrogel.

The addition of 2 mg BG significantly increased

chondrocyte ALP activity in both 2D and 3D culture. However this increase is not dose-dependent. No significant change in solution [Ca] was observed for the 2mg BG group over time. In contrast, a significantly increase in [Ca] was found at day 1 for the lOmg group, which decreased over time in both 2D (Fig. 22B) and 3D culture (Fig. 23B) . These changes may be indicative of Ca precipitation and

mineralization.

These results are indicative of BG enhancing

chondrocyte proliferation, biosynthesis and mineralization potential, although these effects are modulated by

culturing conditions (2D vs. 3D) . Moreover, increased interaction of BG with chondrocytes in 2-D culture promotes cell proliferation, while both BG and 3D culture are

required for enhanced GAG synthesis. Overall these

observations demonstrate the potential of incorporation of BG as an agent in the tissue engineering device of this application.

A biomimetic interface scaffold comprising a biodegradable polymer and hydroxyapatite composite nanofiber scaffold as described in Example 6 was also evaluated. Specifically, the effects of hydroxyapatite (HA) nanoparticles on deep zone chondrocyte (DZC) biosynthesis and mineralization on polylactide-co-glycolide (PLGA) -based nanofiber scaffolds ( DZC on PLGA vs. DZC on PLGA+HA) were examined . Figure 25 shows the effects of HA on DZC proliferation. Deep zone chondrocytes remained viable and increased in number over time on both types of scaffolds . In addition, cell number decreased on PLGA scaffolds from day 7 to day 14 (*:p<0.05 over time).

Figures 26A and 26B show the effects of HA on matrix deposition. Proteoglycan and collagen production increased over time on both scaffold types (*:p<0.05 over time). In addition, there was significantly higher proteoglycan content on PLGA scaffolds by day 14 (^Λ:ρ<0.05 between groups) .

Figure 27 shows the effects of HA on DZC mineralization. Alkaline phosphatase (ALP) activity was significantly higher on PLGA at day 1 and decreased thereafter (*:p<0.05). In addition, ALP activity was maintained in PLGA+HA scaffolds over time.

Figure 28 shows the effects of HA on DZC hypertrophy. Hypertrophic markers were expressed on both scaffold types. In addition, there is no significant difference observed between PLGA and PLGA+HA scaffolds.

Thus, this study demonstrated that both PLGA and PLGA+HA scaffolds supported chondrocyte viability and growth. In addition, both supported the formation of collagen-proteoglycan rich matrix.

Experiments were also performed to compare PLGA nanofibers with and without HA. See Figures 29-31.

Chondrocyte mineralization potential was analyzed by alkaline phosphatase assay. It was found that incorporation of HA in PLGA decreased DZC ALP activity or mineralization potential. This was similar to the trend of DZC cultured in agarose+HA hydrogel. ALP activity of DZC in the presence of HA was shown to decrease by day 14 (Khanarian et al. Transactions of the 56th Annual Meeting of the Orthopaedic Research Society 2010 page 36) . These results were confirmed with von Kossa staining, which stains black in the presence of phosphorus, showing positive staining for mineral in the deposited EC .

As compared to full thickness chondrocytes

(SZC+MZC+DZC) on PLGA+HA nanofibers and cultured in ITS+10% FBS (Moffat, there was higher GAG production (See Figure 30A and B) and higher ALP activity (see Figure 29A and B) on PLGA+HA scaffold. This suggests that the observed response is both cell type and media dependent.

Thus, as shown herein, PLGA+HA scaffolds supported DZC growth and production of a collagen-proteoglycan rich matrix. PLGA+HA scaffolds further provided the mineralized matrix for the formation of the calcified cartilage interface.

Incorporation of HA in the PLGA nanofiber scaffolds modulated biosynthesis and mineralization potential of DZCs.

Accordingly, PLGA-HA scaffolds are also expected to provide a useful scaffold in the tissue engineering devices of this application .

Throughout this application, certain publications are referenced. Full citations for these publications, as well as additional related references, may be found immediately preceding the claims. The disclosures of these publications are hereby incorporated by reference into this application in order to more fully describe the state of the art as of the date of the invention described and claimed herein.

The following section provides further illustration of devices of this applications and methodologies used to demonstrate their usefulness as tissue engineering devices and in particular as osteochondral interfaces. These examples are illustrative only and are not intended to limit the scope of the invention in any way. EXAMPLES

EXAMPLE 1 : BI -PHASIC SCAFFOLD OF CALCIUM PHOSPHATE CERAMIC AND ALGINATE HYDROGEL

Biphasic scaffold Fabrication and Characterization:

Deep zone chondrocytes (DZC, 1 week-old calves) were isolated via enzymatic digestion (Jiang et al.

Osteoarthritis Cartilage 2008 16:70-82) and embedded in a biphasic scaffold of alginate hydrogel and ceramic (Fig.

8A) . The scaffold was formed by first mixing the

cell+hydrogel phase (20M/mL) with particulates (1.5wt%) of either hydroxyapatite (HA) or 45S5 bioactive glass (BG) .

The construct was cross-linked with 50 mM CaCl2. Mineral presence and distribution were confirmed by SEM in the backscattered mode (Fig. 8A) Control groups included DZC in alginate hydrogel and acellular scaffolds of alginate, alginate+HA and alginate+BG. DZC+scaffold were cultured over time in ITS media and 50μg/mL ascorbic acid.

Biochemical Analysis:

Cell growth (n=5) was quantified by Picogreen assay; collagen and glycosaminoglycan (GAG) synthesis (n=5) were measured via Sircol and DMMB assays. Alkaline phosphatase (ALP) activity (n=5) was measured by enzymatic assay (Jiang et al. Osteoarthritis Cartilage 2008 16:70-82). Collagen, GAG and mineral distribution were visualized by Picrosirius Red, Alcian Blue and von Kossa staining (with neutral red counterstaining) , respectively.

Mechanical Testing:

Shear and compressive moduli (n=3) were determined under static compression and dynamic torsional shear using a rheometer (TA Instruments). After a 0.025N tare load, 15% compression was applied and the normal equilibrium force was recorded. The dynamic shear test was performed at a shear strain of 0.01 radians on a logarithmic frequency sweep (0.01-20Hz), and the complex shear modulus and phase shift angle (n=3) were calculated (Zhu et al. J Orthop Res 1993 11: 771-81) .

EXAMPLE 2 : HYDROGEL-MINERAL COMPOSITE SCAFFOLD

Cells and Cell Culture:

Primary bovine articular chondrocytes used in this study were isolated from neonatal calf knees purchased from a local abattoir (Green Village, NJ) . Deep zone chondrocytes (DZC) were obtained from the bottom 30% of the articular cartilage adjacent to the subchondral bone (Jiang et al. Biochem Biophys Res Commun 2005 338:762-70). For comparison, chondrocytes from the full thickness of cartilage (FTC) were also isolated. The cells were enzymatically digested from their respective regions. Briefly, the tissue was incubated for 16 hours with 0.1 w/v% collagenase (Sigma, St. Louis, MO) in Dulbecco' s modified Eagle's medium (DMEM) supplemented with 5% fetal bovine serum ( FBS , Atlanta Biologicals, Atlanta, GA) , 2% antibiotics (10,000 U/ml penicillin, 10 mg/ml streptomycin), and 0.1% antifungal (amphotericin B) . The isolated chondrocytes were maintained in high-density culture (4xl0⁵ cells/cm²) in fully-supplemented DMEM supplemented with 10% FBS, 1% non-essential amino acids, 1% antibiotics, and 0.1% antifungal for 48 hours before seeding. All media supplements were purchased from Cellgro-Mediatech unless otherwise specified.

Alginate Scaffold Fabrication, Characterization, and Culture : Medium viscosity sodium salt alginic acid (Sigma) was used in this study. A 2% alginate solution was prepared and sterile filtered (0.22 μπι, Nalgene, Rochester, NY). A custom mold was used to crosslink the alginate scaffolds (10 mm diameter x 1.6 mm height) with 50 mM CaCl2 (Sigma) and 150 mM NaCl (Sigma) for 30 minutes (Wan, 2005) . Both mineral- free alginate scaffolds and composite scaffolds with 1.5% hydroxyapatite (HA, Sigma) with and without cells were fabricated. The experimental group consisted of cell-seeded alginate scaffolds with hydroxyapatite (Alg+HA) , whereas the control group consisted of cell-seeded alginate scaffolds (Alg) . Corresponding acellular scaffolds were fabricated for each group. For cell-seeded scaffolds, the cell pellet was mixed with the alginate solution, resulting in a final seeding density of 10x10^s cells/ml. All samples were cultured under humidified conditions at 37°C and 5% CO2 and maintained in ITS media, consisting of DMEM supplemented with 1% ITS+ Premix (BD Biosciences, San Jose, CA) , 1% antibiotics, 0.1% antifungal, and 40 pg/ml proline (Sigma). The media was changed every other day and supplemented with fresh 50 pg/mL ascorbic acid (Sigma) . Chondrocyte response in the alginate+HA composite scaffolds was evaluated over a four-week culturing period. Specifically, cell viability, proliferation, mineralization, hypertrophy, collagen and glycosaminoglycan (GAG) deposition, and scaffold mechanical properties were evaluated over time.

Mineral distribution (n=2) in the as-fabricated acellular and cell-seeded alginate+HA scaffolds was characterized using environmental scanning electron microscopy (ESEM, 15 kV, JEOL 5600LV, Tokyo, Japan) on 1 and 10 days, and elemental composition (n=2) was determined using energy dispersive x-ray analysis (EDAX, 15kV, FEI Quanta 600, FEI Co., Hillsboro, OR). Scaffold water content and swelling ratio (n=6) were determined on 1, 2, 3, 14, 21, and 28 days. Briefly, samples were weighed and desiccated for 24 hours (CentriVap Concentrator, Labconco Co., Kansas City, MO) , after which time the swelling ratio (wet weight/dry weight) and water content (water weight/wet weight) were calculated.

Cell Proliferation and Viability:

Cell viability (n=2) throughout the scaffold was assessed by the Live/Dead Cytotoxicity Kit (Invitrogen) at 1, 14, and 28 days and imaged using the Axiovert 35 microscope (Zeiss, Oberkochen, Germany) with fluorescence. Cell proliferation (n=6) was determined at 1, 7, 14, 21, and 28 days using the PicoGreen^® total DNA assay (Molecular Probes, Eugene, OR), following the manufacturer's suggested protocol. Briefly, the samples were first rinsed with phosphate buffered saline (PBS, Sigma) and the cells were lysed in 300 μΐ of 0.1% Triton X solution (Sigma). An aliquot of the sample (25 μΐ) was then added to 175 μΐ of the PicoGreen^® working solution. Fluorescence was measured with a microplate reader (Tecan, Research Triangle Park, NC) , at the excitation and emission wavelengths of 485 and 535 nm, respectively. Total cell number in the sample was obtained by converting the amount of DNA per sample to cell number using the conversion factor of 7.7 pg DNA/cell (Jiang et al. Biochem Biophys Res Commun 2005 338:762-70) .

Matrix Deposition:

Collagen deposition (n=6) was quantified on 1, 14, and 28 days using the Sircol assay (Biocolor, Belfast, UK) according to the manufacturer's suggested protocol. The samples were first desiccated for 24 hours and then digested for 16 hours at 60° C with 20 μΐ/ml papain in 0.1M sodium acetate (Sigma) , 10 mM cysteine HC1 (Sigma) , and 50 mM ethylenediaminetetraacetate (EDTA, Sigma) . Absorbance was measured at 555 nm using a microplate reader (Tecan) . Additionally, collagen distribution (n=2) was visualized by Picrosirius red staining at 14 and 28 days. Samples were first fixed with neutral buffered formalin and 1% cetylpyridinium chloride (CPC, Sigma) for 24 hours, followed by dehydration with an ethanol series. The dehydrated samples were then embedded in paraffin (Type 9, Richard- Allan Scientific, Kalamazoo, MI) , sectioned (7 pm) , and mounted on microscope slides prior to staining prior to imaging (Axiovert 35 microscope, Zeiss) .

Total proteoglycan deposition (n=6) was measured using a modified 1 , 9-dimethylmethylene blue (DMMB) dye-binding assay with chondroitin-6-sulfate (Sigma) as a standard. The pH of the DMMB dye was adjusted to 1.5 with formic acid (Sigma) . Absorbance was measured at both 540 nm and 595 nm for improved signal detection. Distribution (n=2) was visualized by Alcian blue histology staining.

Scaffold Mechanical Properties :

Scaffold mechanical properties (n=3) were determined on 1, 14, and 28 days following published protocols (Wan et al . Cell Mol. Bioeng. 2008 1 ( 1 ): 93-102 ) . Briefly, the sample was placed between two porous platens of a rheometer (ARES- LSI, TA instruments, New Castle, DE) and immersed in DMEM. A normal tare load of 0.025N was first applied, and sample thickness was determined from the axial position readings of the rheometer. After 5 minutes of stress relaxation, the normal force (Fl) was recorded. A 15% normal strain was then applied, and the normal force (F2) was determined after another 25-minute relaxation period. The compressive modulus (E_eq) was calculated under static compression with the following equation:

(1) E_eq = σ/ε where o=F2-Fl ^~ / (nd²/4 ) where d is the sample diameter. Sample shear modulus was determined under a dynamic torsional test at a shear strain of 0.01 radians on a logarithmic frequency sweep (0.01-20 Hz). The complex shear modulus (G*) was calculated with the following formula:

(2) G* = Td/2I_PY

where I_p is the polar moment of inertia of a cylinder and given by I_p = nd⁴/32, γ is a sinusoidal shear strain applied on the sample, and T is the torque response. The complex shear modulus can be expressed as G* = G' + iG", where G' is the storage modulus and G' ' is the loss modulus. The magnitude of the shear modulus and phase shift angle (δ) were then calculated based on the following equations:

(3) G* = sqrt(G'P + G"²) I ( 4 ) δ = tan^"1 (G"/G)

Mineralization :

Mineralization was determined by measuring ALP activity and mineral deposition at 1, 5, 6, 7, 14, and 28 days. Quantitative ALP activity (n=6) was measured using an enzymatic assay based on the hydrolysis of p-nitrophenyl phosphate (pNP-PO ) to p-nitrophenol (pNP) (Lu et al. Biomaterials 2005 26 ( 32 ) : 6323-34 ) . The samples were lysed in 0.1% Triton-X solution, then added to pNP-P0₄ solution (Sigma) and allowed to react for 30 min at 3 °C. The reaction was terminated with 0.1 N NaOH (Sigma), and sample absorbance was measured at 415 nm using a microplate reader (Tecan) . Mineral distribution (n=2) was evaluated by von Kossa staining with 5% silver nitrate and 30 minutes of UV exposure (Wang et al . J Orthop Res 2007 25:1609-20). Additionally, media calcium concentrations (n=6) were quantified using Arsenazo III dye (Pointe Scientific, Lincoln Park, MI), and absorbance was measured at 620 nm using a microplate reader (Tecan) . Gene Expression:

The expression of SOX9, MMP13, Ihh, and collagen X (n=3) were measured at 14 and 28 days using reverse transcription followed by polymerase chain reaction (RT- PCR) . The custom-designed oligonucleotides are listed in Jiang et al . (Osteoarthritis Cartilage 2008 16:70-82). Total RNA was isolated using the TRIzol reagent (Invitrogen) extraction method. The isolated RNA was reverse-transcribed into cDNA using the Superscript III First-Strand Synthesis System (Invitrogen) following the manufacturer's suggested protocol, and the cDNA product was amplified with recombinant Platinum Taq DNA polymerase (Invitrogen) . Expression band intensities of relevant genes were analyzed semiquantitatively and normalized against that of the housekeeping gene GAPDH. The primer sequences utilized in PCR analysis are also disclosed in Jiang et al. (Osteoarthritis Cartilage 2008 16:70-82), the entirety of which is incorporated herein by reference.

Statistical Analysis :

Results are presented in the form of mean ± standard deviation, with n equal to the number of samples analyzed. A two-way analysis of variance (ANOVA) was performed to determine the effects of mineral presence, culturing time, cell type on cell parameters (cell proliferation, matrix deposition, mineralization potential, gene expression) , as well as scaffold parameters (weight, thickness, swelling ratio, mechanical properties) . The Tukey-Kramer post hoc test was used for all pair-wise comparisons, and significance was attained at p<0.05. All statistical analyses were performed using the JMP software (SAS Institute, Cary, NC) . EXAMPLE 3 : COMPOSITE HA-AGAROSE SCAFFOLD FOR INTERFACE REGENERATION

Cells and Cell Culture:

DZC were isolated from the bottom 30% of cartilage tissue of neonatal calf knees (Jiang et al. Osteoarthritis Cartilage 2008 16:70-82) and were maintained in ITS media with 50 fig/mL ascorbic acid.

Scaffold Fabrication/Characterization :

The cell suspension was first mixed with 2% w/v nano- sized (Nanocerox, 100 nm) or micro-sized (Sigma, 20 urn) HA and then with 4% agarose for a final seeding density of 10 million cells/ml. The presence and distribution of HA within the hydrogel were confirmed by SEM (backscattered mode, 15 keV) (Fig. 12) and FTIR. Experimental groups were stimulated with 50 nM T3 for the first 3 days of culture. Control groups included DZC in HA-free scaffolds, corresponding acellular scaffolds, and T3-untreated samples.

End-point Analyses:

Samples were collected on days 1, 7, and 14, with total

DNA (n=6) determined by Picogreen assay and alkaline phosphatase (ALP, n=6) activity evaluated by both enzymatic assay and histology. Matrix production of glycosaminoglycans (GAG, n=6) and collagen (n=6) were determined by the DMMB and Sircol assays, respectively, and assessed by histology. Mineral distribution (n=2) was visualized by Alizarin Red staining. Additionally, gene expression (n=3) for collagen I, collagen II, collagen X, and MP13 was determined by RT- PCR, with β-actin as housekeeping gene.

Mechanical Testing:

Shear and compressive moduli (n=3) were determined under static unconfined compression (0.025 N tare load, 15% compression) and dynamic torsional shear using a rheometer (TA Instruments) (Arsenault, et al . , 1988). The dynamic shear test was performed at a shear strain of 0.01 radians on a logarithmic frequency sweep (0.01-20 Hz), and the magnitude of the complex shear modulus and phase shift angle (n=3) were calculated.

Statistical Analysis:

ANOVA and the Tukey-Kramer post-hoc tests were used for all pair-wise comparisons (p<0.05).

EXAMPLE 4: EFFECTS OF 45S5 BIOACTIVE GLASS PARTICLES ON CHONDROCYTE BIOSYNTHESIS AND MINERALIZATION

Cells/Cell Culture:

Articular chondrocytes were isolated via enzymatic digestion from the knee joints of neonatal calves. All cultures were maintained in ITS media supplemented with 50 ug/mL ascorbic acid.

2D Culture with BG:

Chondrocytes were seeded at 1 million cells/well in culture wells pre-coated with 2 mg or 10 mg of 45S5 bioactive glass (BG, 2 urn, MOSCI), while cells in monolayer without BG served as control.

3D Culture with BG: Chondrocytes (25million cells/mL) were loaded into 2% agarose hydrogel. containing 2mg or lOmg of BG (BG inside) . The control groups include cells in agarose disks without BG (control), and cell-laden disks cultured in wells pre-coated with 2mg or lOmg BG (BG outside) .

End-point Analyses:

Cell growth (n=5) was quantified by Picogreen assay, while glycosaminoglycan (GAG) synthesis (n=5) by DM B assay. Alkaline phosphatase (ALP) activity (n=5) was measured by an enzymatic assay, and calcium content in the media (n=5) was quantified with Arsenazo III dye.

EXAMPLE 5 : EVALUATION OF CHONDROCYTE RESPONSE ON PLGA NANOFIBER SCAFFOLDS WITH AND WITHOUT HA NANOPARTICLES

Scaffold fabrication:

Aligned PLGA and PLGA+HA nanofiber scaffolds were fabricated by electrospinning as described by Moffat et al . (Tissue Eng. Part A 2009 15 (1) : 115-26) :

1. PLGA 85:15 solution: 35% (v/v) PLGA (85:15), 55% dimethylformamide , 10% ethanol

2. PLGA+HA solution : PLGA solution, 3% hydroxyapatite nanoparticles (average size 100-150nm)

Grounded rotating collector.

Cell And Cell Culture:

DZC were isolated via enzymatic digestion of the bottom third of femoral cartilage in bovine knee joints (Jiang et al. Osteoarthritis Cartilage 2008 16:70-82) . The first step was to immerse scaffolds in 10 pg/mL fibronectin for 12 hours (fibronectin enhances cell attachment). Then, bovine DZC were seeded at 90,000 cells/cm² and allowed to attach for 15 minutes. Then, scaffolds were maintained in ITS+ media and summarized by Figure 24.

Experimental Design:

PLGA+HA versus PLGA control scaffolds were compared 1, 7, and 14 days.

End point Analysis :

Cell response was assessed at 1, 7, and 14 days.

The following were assessed: 1) cell proliferation (n=5) , 2) Matrix deposition - Glycosaminoglycan (GAG) by DMMB assay (n=5) and Collagen by Sircol assay (n=5) , 3) Hypertrophy - Gene expression of MMP13, PTHrp, Ihh, Col X (n=3) , and 4) Mineralization - Alkaline phosphatase (ALP) activity (n=5) In addition, statistical analysis is done by two-way ANOVA and Tukey-HSD (*,p<0.05).

Claims

What is claimed is :

1. A three-dimensional tissue engineering device comprising :

a bottom portion; and

a side wall portion with a top and a bottom adjoined at the bottom to said bottom portion to form a hollow center portion,

wherein said bottom, side wall and hollow center portions collectively form a cup or vessel.

2. The three-dimensional tissue engineering device of claim 1 further comprising one or more flaps at the top of the side wall portion for fixation of the cup or vessel.

3. The three-dimensional tissue engineering device of claims 1 or 2 further comprising a cap at the top of said side wall portion for closure of the cup or vessel.

4. The three-dimensional tissue engineering device of any of claims 1 through 3 wherein said bottom portion, said side wall portion, said one or more claims and/or said cap comprise a material or materials for promoting growth and maintenance of a selected tissue type or selected tissue types .

5. The three-dimensional tissue engineering device of claim 4 wherein said bottom portion and said side wall portion comprise mesh configured to permit cell migration from one: side of the mesh through to an opposite side of said mesh.

6. The three-dimensional tissue engineering device of claim 5 wherein said mesh comprises polymer nanofibers.

7. The three-dimensional tissue engineering device of claim 5 wherein said bottom portion and said side wall portion comprise a plurality of sheets of mesh.

8. The three-dimensional tissue engineering device of claim 7 wherein said plurality of sheets of mesh overlap in at least one area.

9. The three-dimensional tissue engineering device of claim 7 wherein said plurality of sheets of mesh are joined together .

10. The three-dimensional tissue engineering device of claim 5 wherein said bottom portion and said side wall portion comprise the same mesh which promotes growth and/or maintenance of a selected tissue type and/or integration with a selected underlying tissue type.

11. The three-dimensional tissue engineering device laim 10 wherein said bottom portion and said side wall ortion are continuous.

12. The three-dimensional tissue engineering device o: claim 5 wherein said bottom portion comprises a mesh which promotes growth and/or maintenance of a first selected tissue type and/or integration with a first selected underlying tissue type and said side wall portion comprises a mesh which promotes growth and/or maintenance of a second selected tissue type and/or integration with a second selected underlying tissue type.

13. The three-dimensional tissue engineering device of claim 12 wherein said bottom portion comprises a mesh which promotes growth and/or maintenance of cartilage and/or chondrocytes and/or integration with underlying cartilage tissue and said side wall portion comprises a mesh which promotes growth and/or maintenance of bone and/or

osteoblasts and/or integration with underlying bone tissue.

14. The three-dimensional tissue engineering device of any of claims 1 through 13 wherein said bottom portion comprises an agent which supports alignment, proliferation and/or matrix deposition of a selected musculoskeletal cell and/or promotes cell migration and integration and/or which promotes differentiation of stem cells to a selected cell type.

15. The three-dimensional tissue engineering device of any of claims 1 through 14 wherein said side wall portion comprises an agent which supports alignment, proliferation and/or matrix deposition of a selected musculoskeletal cell and/or promotes cell migration and integration and/or which promotes differentiation of stem cells to a selected cell type.

16. The three-dimensional tissue engineering device of claim 15 wherein said agent of said side wall is the same as said agent of said bottom portion.

17. The three-dimensional tissue engineering device of claim 15 wherein said agent of said side wall is different from said agent of said bottom portion .

18. The three-dimensional tissue engineering device of any of claims 14 through 17 wherein said agent in said bottom portion and/or said agent in said side walls

comprises a ceramic.

19. The three-dimensional tissue engineering device of claim 18 wherein said ceramic comprises calcium phosphate.

20. The three-dimensional tissue engineering device of any of claims 1 through 19 further comprising an insert member fitted into the hollow center portion of said tissue engineering device.

21. The three-dimensional tissue engineering device of claim 20 wherein the insert member has a shape and

orientation complementary to the hollow center portion.

22. The three-dimensional tissue engineering device of claim 20 wherein said insert member comprises a hydrogel .

23. The three-dimensional tissue engineering device of claim 20 wherein said insert member comprises a plurality of hydrogel layers .

24. The three-dimensional tissue engineering device of any of claims 20 through 23 wherein said insert member further comprises one or more agents which support

alignment, proliferation and/or matrix deposition of a selected musculoskeletal cell and/or promote cell migration and integration and/or which promote differentiation of stem cells to a selected cell type.

25. The three-dimensional tissue engineering device any of claims 20 through 24 wherein said hydrogel or hydrogel layer is seeded with a selected cell type.

26. The three-dimensional tissue engineering device claim 25 wherein said selected cell type comprises chondrocytes .

27. The three-dimensional tissue engineering device claim 20 wherein said insert member comprises a tissue allograft or autograft.

28. The three-dimensional tissue engineering device claim 27 wherein said tissue allograft or autograft comprises a cartilage allograft or autograft.

29. Use of the three-dimensional tissue engineering device of any of claims 1 through 28 as an osteochondral interface .