IMPLANTS FOR TISSUE REPAIR

The invention relates to implants for the repair or replacement of defective biological tissue parts, in particular for the repair or replacement of damaged load-bearing cartilaginous tissue, such as the meniscus and articular cartilage.

Popular current procedures for the treatment of articular cartilage defects include high tibial osteotomy (primarily performed in

Europe) and muscle release to alter joint biomechanics and loading; lavage and debridement to remove osteophytes and fibrillated areas of cartilage and perforation and penetration of the subchondral bone to induce bleeding and clot formation. It has been established, however, that none of these techniques lead to successful regeneration of the tissue that duplicate the structure, composition, mechanical properties or durability of articular cartilage.

The common technique is that of subchondral drilling which results in the creation of a fibrin clot and a fibrous tissue. In about

75% of patients there is a satisfactory result, but two years after the operation, only 12% of patients remain free of symptoms. Limitations of this approach are the difficulty in predicting the quality and duration of the clinical outcome and, in the long term, this procedure often leaves the patient in a worse condition. On the other hand, subchondral drilling and debridement are generally successful in alleviating symptoms and in postponing the requirement for total knee replacement by between one and three years.

A cartilage repair product which can predictably alleviate pain, restore function, achieve good fixation and which has a durability enabling it to last at least ten to fifteen years is seen as an important objective by most surgeons.

It has long been known that cartilage cells grown on porous, 3 dimensional scaffolds are able to form cartilage-like tissue in vitro. In principal, this "neocartilage" could be surgically implanted as a graft to promote the repair of otherwise untreatable full thickness articular cartilage defects. This approach to joint repair is known as tissue engineering. Tissue engineered scaffolds must be (i) porous, (to allow the ingress of cells and nutrients during graft production); (ii) mechanically robust, (to allow surgical handling and minimise the shear arising from modulus mismatches between the implant and surrounding tissue when the graft is loaded) and (iii) conformable (to allow the device to follow the contours of the joint surface). The scaffolds should also be biocompatible and, preferably, biogradable. Examples of scaffolds known to the art include the non-woven polyglycolate felts described by Vacanti 5,041 ,138.

In order to attach a tissue engineered cartilage implant into the defect site, an implant could either be (i) press-fitted into a full thickness chondral lesion, (created by surgically debriding the remaining cartilage from the defect site and exposing the underlying bone; (ii) glued into a full thickness cartilage lesion prepared as above, or (iii) placed into a deep, cylindrical "pocket" formed by drilling out the residual cartilage, subchondral bone plate and underlying trabecular bone. Unfortunately, none of these approaches to graft fixation yields fully satisfactory results. When the implant is loaded during normal locomotion, the graft is sequentially (I) compressed as the two joint surfaces come into opposition, (ii) exposed to shear as the joint surfaces roll across each other and then (iii) "sucked" from the defect site as the two joint surfaces separate and create a partial vacuum in the joint space (Triffit ICRS 2002). The combined effect of these forces either produces small, repeated displacements of the implant within the defect site, (preventing the implant from integrating with

the surrounding cartilage or underlying bone), or detaches the implant from the defect site altogether, (leaving it as a free body within the joint space). Resolving the problem of graft fixation is therefore a key step toward the overall goal of using tissue engineered cartilage to promote joint repair.

Numerous attempts have been made to create a tissue engineered cartilage scaffold with (i) the necessary biological properties to allow cell-seeding, growth and matrix deposition and (ii) the required physical properties for attachment, conformability and mechanical strength. None of these earlier approaches has been entirely successful.

The scaffold component of a tissue engineered cartilage device should facilitate the biological processes required for graft formation, (e.g. homogeneous cell-seeding; adequate nutrient and O2 availability during cell growth and matrix deposition; waste product egress). Many variations in scaffold design have been proposed, (e.g. meshes, foams, perforated materials), which seek to address these biological issues without considering the question of graft fixation. Methods based upon these approaches to scaffold design have succeeded in producing neo-cartilage but the resulting implants have failed to integrate with the host tissues during in vivo trials.

The simplest scaffold designs, (e.g. Vacanti 5,041 ,138), envisage a non-woven scaffold that is either placed into a chondral defect or an osteochondral defect. Unfortunately, press fitted chondral implants detach rapidly and are therefore inappropriate.

Implants that are fitted into deep osteochondral pockets, (e.g. Athanasiou, 5,875,452), provide better graft fixation, but destroy the subchondral bone plate and trabecular bone beneath the implant,

disrupting the mechanical integrity of the joint. Cylindrical osteochondral pockets of the diameter and depth required to attach a tissue engineered cartilage implant are associated with excessive bone resorption immediately below the implant, bone cyst formation beneath the implant and sclerotic changes to the bone at the periphery of the joint. These degenerative changes tend to undermine the implant, causing it to subside from alignment with the joint surface, thereby rendering it functionally useless. Excessive bone resorption may be due to inadequate conformability of the implant and/or compression of the implant.

In an attempt to improve mechanical attachment, ingenious cutters have been designed to produce undercut osteochondral "pockets" to receive an implant (Smith 5,817,095). Whilst these cutters may create a "pocket" that gives better mechanical fixation to the implant than those created by a cylindrical drill, such cutters still destroy the subchondral bone plate and underlying trabecular bone below the implant and may therefore cause failure by giving rise to bone-cyst formation, sclerotic bone changes at the periphery of the joint and excessive bone resorption immediately below the implant.

It has been suggested that a tissue engineered cartilage graft maybe held in a full thickness chondral defect by suturing it in place or holding it in place by a membrane that is sutured over the implant and defect site. It is however very difficult and time consuming to suture into cartilage and the resulting bond is very weak. Furthermore, the needle holes formed in the surrounding host cartilage do not heal and therefore constitute additional defects in the joint structure. Consequently, surgeons do not like to suture into cartilage. Nevertheless suture-based devices have been proposed for anchoring and manipulating cartilage implants, (Hayhurst 4,741 ,330; Pachence 5,713,374; Naughton 5,842,477).

Various attempts have been made to use glues, gels, enzymes or active exogenous factors to promote the attachment of chondral tissue engineered implants to a defect site. Examples include fibrin glues, (Asculai, 6,569,172); transglutaminase, (Fraji, 6,190,896); biological gels (Zaleske, 6,183,737); angiotensins (Rodgers, 6,258,778); in situ curable gels, (Felt, 5,795,353); proteoglycanases (Berlowitz-Tarrant, 5,916,557). Results show that fibrin glue does improve the attachment of an implant to a chondral defect over that observed with simple press fitting alone (table 1) but the improvement is small and manual inspection shows that the bond strength is weak. It is therefore likely that glue-based approaches will only be of value if used in conjunction with improved scaffold design and graft site preparation techniques.

Alternatively, again, a neo-cartilage implant could be held in place by means of anchors or pins that penetrate through the graft into the underlying bone. Examples include Kay, 5,662,683. The use of anchors would be expected to have several undesirable consequences, e.g. making holes in the neo-cartilage graft, the likelihood of the graft tearing around the anchors and the danger of the head of an anchor protruding into the joint space and scoring the cartilage on the opposing surface of the joint.

A small number of groups have attempted to produce scaffolds for cartilage tissue engineering that specially address the need to improve graft fixation whilst minimising damage to the subchondral bone. These devices resemble 'thumb tacks", (a.k.a. drawing pins), in that they combine a central shank that penetrates into the subchondral bone to provide attachment, with a planar structure that occupies the space previously filled with the articular cartilage. Examples include Schwartz, 6,468,314, Overaker, 6,371 ,958 and Schwartz, 5,769,899. Thumb-tack-based approaches

share several common disadvantages. Firstly, any device with the general form of a "thumbtack" will tend to flex at the junction between the "shaft" and the "top" when the "top" is exposed to the shear forces and cyclical, off-axial compression/tensional loads that occur within a joint. This cyclical flexure of the junction between the shaft and top is likely to lead to fatigue fracture and consequent failure of the device. Stress concentration at the junction between the shaft and top is also likely to lead to catastrophic failure under peak loads, (e.g. jumping or twisting whilst running). As a consequence the central shaft of thumb-tack-like device will need to be more massive (>2.5mm diameter; essentially solid) and therefore create a greater barrier to boney ingrowth and osseous healing than the thin and porous peripheral outer shell wall (0.5mm thick; >90% porosity), envisaged in the present device.

Secondly, it is now appreciated (Triffit 2002) that one of the causes of implants detachment from a defect site is the combination of (i) compression and shear at the centre of the implant with (ii) a "suction force" at the trailing edge of implant (that occurs as the two opposing joint surfaces separate during locomotion). This "suction force" lifts the trailing edge of the implant out of the defect site, removing any barrier to lateral displacement and allowing the implant to detach. The "suction force" upon the periphery of the device is not resisted by the central attachment of a "thumb-tack-like" device described in Schwartz, 6,468,314, Qveraker, 6,371 ,958 and Schwartz, 5,769,899 etc. but is resisted by the peripheral attachment of the scaffold described in this proposal. Furthermore, because the force required to detach a graft from an implant site is proportional to the area of the contacting surface between the graft and the surrounding tissue, (because friction is proportional to the contact area between two surfaces), then it can be shown that it will require 2.4 fold more force to pull out a device of the present invention as

hereinafter described than an equivalent sized "thumb-tack" shaped device.

Thirdly, the 'thumb-tack" approach to scaffold design envisaged in the earlier patents envisages the use of multiple, essentially rigid implants to form a mosaic across the defect site. This is especially true if the top of the device is rigid and perpendicular to the central shank. It is known to be surgically difficult to position and align such multiple implants so that they recreate the 3 dimensionally curved contours of the cartilage in a healthy joint.

Some known scaffold designs have featured honeycombs of closely packed and/or laterally bonded cells. The problem with these designs is that they are rigid planar structures that cannot easily conform to the contours of an individual patient's joint defect. These devices would therefore have to be produced on a one-off basis for individual patients. This would have very adverse consequences for their clinical and commercial acceptability. As with deep ostechondral pocket type implants, these tightly packed rigid structures have also suffered from the problems of causing destruction of the subchondral bone plate and trabecular bone beneath the implant, disrupting the mechanical integrity of the joint. Furthermore, if they were manufacture in a curved conformation, they might be difficult to seed evenly with cells and maintain in culture prior to implantation.

It is a primary object of the present invention to provide an implant for tissue repair, adapted to receive, support the growth of and anchor cellular tissue, particularly cartilage tissue, which addresses the above problems.

In accordance with a first aspect of the present invention, there is presented a device for the repair of cartilage defects comprising a flexible material and a plurality of elongate elements attached substantially perpendicular to the flexible material, wherein the elongate elements provide compression resistance.

As the flexible material may not be even or flat, the elongate elements may not, indeed need not, be strictly perpendicular to the flexible material.

Preferably the flexible material and elongate elements are biocompatible and suitable for implantation into the human body. Suitable cartilage defects include partial replacements of tissue, total replacement of tissue or reinforcement of tissue.

When implanted in a site with a curved surface, for example articular cartilage, the flexible material forms an architrave that will conform to the shape of the curved surface whilst the elongate elements give strength in their longitudinal direction.

The present invention may have an outer shell wall surrounding the elongate elements. Other embodiments of the present invention may not have this outer shell wall.

Preferably the device of the present invention will have an outer shell wall as this provides improved fixation of the device into the tissue. This outer shell wall need not be continuous. It may, for example, comprise at least one slot, slit or opening. It is envisaged that the outer shell wall in particular embodiments of the invention will act to anchor the device into the defect site with minimal disruption to the tissue e.g. the subchondral bone plate and underlying trabecular bone. Typically the elongate elements are

shorter than the outer shell wall.

The design offers better attachment than press-fitted full thickness or partial thickness chondral implants which rapidly fall out. It offers better attachment than sutured or fibrin glued press-fitted full thickness or partial thickness chondral implants which are surgically difficult and complicated to fit and which still fall out. It offers better attachment than periosteal flaps, which are traumatic to harvest, surgically difficult and time consuming to fit and which often rapidly fall out. It is better than osteochondral implants which stay in place but which require the surgeon to breach the underlying subchondral bone plate and remove the trabecular bone beneath it, thereby placing abnormal loads upon the cartilage surrounding the defect — which causes it to degenerate — and upon the bone, which becomes sclerotic away from the defect and excessively resorbed under the implant — which causes the joint/implant to fail.

In addition, the shell wall is a better approach to fixation from a theoretical mechanical perspective than previous "thumb-tack" designs because (a) it does not concentrate the stresses at a single point, (i.e. the junction between "shaft" and "top") and (b) because the greater contact area between the implant and host tissue offered by a thin peripheral wall compared to a central spike of equal mass leads to greater friction and therefore a 2.4 — 9.7 fold greater "pull out force".

The outer shell wall and flexible material can be produced from a single disk of reinforced non-woven fabric, thereby minimising the likelihood that the top of the device will shear of the underlying supports.

Inner elongate elements resist compression. When arranged

as cylinders below the "architrave" surface formed by the flexible material when the device is implanted, the elongate elements provide a high degree of compression resistance, (approximately 0.1 — 0.8 fold that of normal cartilage), whilst leaving a very high percentage void space to facilitate cell-seeding and nutrient access during the in vitro phase of tissue engineered graft formation. The compressive modulus is typically between 0.1MPa and 2MPa and is preferably approximately 1MPa. The elongate elements may either be open-ended cylinders or, preferably, inverted cups i.e. cylinders closed at one end. The advantage of using cylinders closed at one end is that they are (a) stronger and (b) they can be assembled from disks of reinforced fabric and then adhered onto the underside of the flexible material with the required spacing to give the desired radius of curvature.

There is no need for the elongate elements to have a uniform diameter or spacing. These variables can be modified so that the device has differing properties across its surface.

The device has a plurality of elongate elements. This is defined as at least two elongate elements, although the actual number will vary depending on the size of the defect to be repaired and the dimensions of the elongate elements themselves. The elongate elements are typically between 1mm and 5mm long, preferably approximately 2mm long. Furthermore they are typically between 0.1mm and 1mm thick, preferably approximately 0.5mm thick.

Spaces between elongate elements enable the device to be handled as a planar structure in vitro but to conform to the contours of the defect site following implantation. Previous designs of scaffolds (e.g. honeycombs with closely packed, laterally bonded

cells or "spoked wheel" scaffolds) give rise to devices with high strengths and high porosity but they are rigid planar structures that cannot conform to the specific contours of an individual patient's joint surface during surgical implantation. The only way of making these structures conformable is to make them weaker so that they can deform, but this is counter productive because it reduces their compression resistance. Alternatively, these devices could be made on a one-off basis for individual patients but this would be expensive, inconvenient, require very accurate data on the precise dimensions and curvature of the defect site and necessitate very accurate preparation of the defect site.

In contrast, the spaces between the elongate elements allow the flexible material of the device to flex thereby allowing it to follow the curvature of the joint surface. This flexure does not require any weakening of the elongate elements themselves; indeed it maintains them perpendicular to the subchondral bone plate, optimizing their strength.

The elongate elements of the present invention may mimic the overall orientation of collagen bundles found in connective tissue, such as cartilage. This property allows the device to respond to tension, compression and torsion in a way similar to natural tissue resulting in a reduced incidence of failure. As the cartilage and bone tissue underneath the device of the present invention is subject to tension, compression and torsion in a more similar way to the natural material tissue than previous implants there is less likelihood of eroding or natural resorption of the tissue under the device of the present invention.

Reference herein to any material being "biocompatible" means that the material gives rise to essentially no acute reaction when

implanted into a patient.

Reference below to a "support structure" means connective tissue, injured or healthy, into which the device can be implanted.

The flexible material employed in the device according to the invention may comprise, woven, non-woven (fibrous material), knitted, braided or crocheted material, foam, sponge, dendritic material, a polymeric film or membrane or a mixture of two or more of these materials. The flexible material employed in the device according to the invention may comprise a porous or a non-porous structure. Preferably, it comprises an at least partially porous structure. This has the advantage of allowing tissue ingrowth, helping to reinforce the structure and avoid mechanical failure. In the event that the flexible material comprises a porous structure, the percentage open volume of the tape may be in the range 30-99%. The optimum value will depend upon the application and may be a compromise between attaining a high open volume for rapid and efficient penetration of tissue, and good initial mechanical properties, such as tensile or compressive modulus. Typically, the percentage open volume will be in the range 65-90%. In a preferred embodiment the open volume will be 70%.

The elongate elements according to the present invention may be integrally formed with the flexible material or attached in some way to the surface of the flexible material, for example they may be adhered.

Advantageously, each elongate element is maintained spaced apart from the other elongate element or elements in order to aid overall flexibility of the device. However some elongate elements may not be spaced apart provided sufficient flexibility of the device is

maintained.

The elongate elements may be generally cylindrical, but need not have a circular cross-section: the cross-section may be essentially circular, but may also be rectangular, triangular, hexagonal, octagonal or have an irregular shape or may vary along the length of the elongate element. The elongate elements need not be uniform in shape and therefore more than one shape of elongate elements can be used in one implant. Indeed the cross-section of the elongate element need not be a closed shape, like a circle or square but could be any shape even if that shape is not closed e.g. folds or furrows. Preferably the cross-section shape of the elongate elements will be a shape that facilitates cell growth on their longitudinal surface but also gives longitudinal rigidity like corrugated shaped material e.g. with alternate ridges and grooves. The elongate elements, or indeed the flexible material as well may also comprise woven, non-woven (fibrous material), knitted, braided or crocheted material, foam, sponge, dendritic material, a polymeric film or membrane or a mixture of two or more of these materials. Part or all of the device may comprise a porous or a non-porous structure. Preferably, it comprises an at least partially porous structure, since this permits tissue ingrowth into the structure, helping to reinforce the structure and avoid mechanical failure. When the elongate elements are cylindrical they typically have a diameter between 2mm and 5mm, although preferably the diameter is approximately 3mm. When the elongate elements are hollow they typically have wall thicknesses between 0.25mm and 1mm, but preferably have wall thicknesses of approximately 0.5mm.

Use of bonded non-woven fabric to form a mechanically robust device matched to the mechanical environment of the joint. The use of a bonded non-woven fabric (e.g. PCI re-enforced PET or 5050

DLA re-enforced PGA) gives a strong but highly porous material (>80%) with a compressive modulus of - 1 MPa. The aim of

producing a device that has a stiffness only slightly less than the surrounding cartilage is that it minimizes the shear that occurs at the interface when two adjacent materials of differing stiffness are compressed without "stress shielding" the repair tissue. Less shearing at the junction between the repair tissue and the cartilage at the edge of the defect should maximize the chance for lateral integration between the device and surrounding host tissue. (This is an improvement over traditional, un-reinforced non-woven scaffolds which have little intrinsic strength).

The device according to the invention may comprise bioresorbable or non-bioresorbable material or a mixture of the two.

Reference herein to a material being bioresorbable means that it breaks down over time due to the chemical/biological action of the body and the terms "resorption" and "resorb" are to be interpreted accordingly. Preferably, complete resorption occurs within about 5 years of implantation, more preferably within about 3 years. An advantage of using bioresorbable materials is that further. Surgery to remove them is not necessary, since they are absorbed back into the body.

A wide range of bioresorbable materials is known, with differing in vivo resorption times. Not only does the resorption time vary according to the material, but the resorption time of a single material itself can also vary significantly with molecular weight. Finally, it can readily be appreciated that by blending and/or copolymerizing different bioresorbable materials and/or by modifying the molecular weights of the components, it is possible precisely to

tailor the resorption time of the bioresorbable material to the requirement at hand.

With the above in mind, the device may comprise bioresorbable polymers or copolymers comprising the following monomers or mixtures of polymers and/or copolymers formed thereby: hydroxy acids, particularly lactic acid, glycolic acid; caprolactone; hydroxybutyrate; dioxanone; orthoesters; orthocarbonates; aminocarbonates.

The device may also comprise natural materials such as collagen, cellulose, fibrin, hyaluronic acid, fibronectin, chitosan or mixtures of two or more of these materials. The bioresorbable materials may also comprise devitalized xenograft and/or devitalized allograft.

Preferred bioresorbable materials comprise poly(lactic acid), poly(glycolic acid), polydioxanone, polycaprolactone, polyhydroxybutyrate and poly(trimethylene carbonate) or mixtures thereof.

It is particularly preferred that the device comprise poly(lactic acid) . This material has the advantage that it has good mechanical strength and does not resorb too quickly, thus allowing its mechanical properties to be retained for a sufficient time for tissue repair to occur at which point the repaired tissue can take over load- bearing functions reference is made to A.G.A. Coombes and M. C. Meikle, "Resorbable Synthetic Polymers as Replacements for Bone Graft" Clinical Materials 17, (1994), pp 35-67.

Appropriate non-bioresorbable materials include polyesters,

particularly aromatic polyesters, such as polyalkylene terephthalates, like polyethylene terephthalate and polybutylene terephthalates; polyamides; polyalkenes such as polyethylene and polypropylene; polyvinyl fluoride), polytetrafluoroethylene, carbon fibres, silk

(natural or synthetic), carbon fibre, glass and mixtures of these materials. An advantage of non-bioresorbable materials is that they essentially retain their initial mechanical properties — i.e. properties such as strength do not reduce over time.

All components of the devices according to the invention may comprise the same materials. Alternatively, some components may comprise different materials or each component of the device may comprise a different material from the other components: in the case where the device comprises warp threads, weft threads and elongate elements, each of these three components may comprise a different material. Alternatively, each may comprise the same material, for example poly(lactic acid).

The device may comprise biocompatible fibres cross-linked with a bonding agent that is a different material from the biocompatible fibre. Suitable biocompatible fibres include polyethylene terephthalate (PET), polyglycolic acid (PGA), polylactic acid (PLA), polyhydroxybutyrate (PHB), poly (lactic-glycolic) acid (PLGA), polyesters and polyamides. Suitable bonding agents include polycaprolactone (PCL), PLA, PLGA, polyesters and polyamides.

In a further form of the present invention, the device may be loaded with cells. Incorporation of cells may be carried out either before or after implantation, but is preferably carried out prior to implantation. The cells may be incorporated within the flexible

material, within the shell walls, within the elongate elements, within the void spaces or a combination of one or more of the above.

The cells are generally incorporated by means of a carrier medium. The carrier medium may be a medium which is retained by

the device, for example a gel such as a hydrogel, or one which substantially passes through the device such that, after seeding, it is

substantially no longer present therein - the cells remaining within the device. Examples of this type of carrier medium are cell culture media, like DMEM (Dulbeco"s Modified Eagle"s Medium).

If the carrier medium is a gel, such as a hydrogel, it may be incorporated within and/or on and/or around the device. In one preferred form, the carrier medium is incorporated within the implantable material, since this efficiently utilizes the available open volume for cellular growth. More preferably, the carrier medium occupies the entire open volume of the device. Alternatively, the carrier gel may be incorporated by overlaying a confluent/sub- confluent cell layer onto the device.

Hydrogels which may be used as carrier media according to the invention comprise positively charged, negatively charged and neutral hydrogels which may be saturated or unsaturated. Examples of hydrogels, which may be used according to the invention are collagen (particularly Type I), fibrin, TETRONICS™ and

POLOXAMINES™, which are poly(oxyethylene)poly(oxypropylene) block copolymers of ethylene diamine; polysaccharides, chitosan, polyvinyl amines), polyvinyl pyridine), polyvinyl imidazole), polyethylenimine, poly-L-lysine, growth factor binding or cell adhesion molecule binding derivatives, derivatised versions of the

above, e.g. polyanions, polycations, peptides, polysaccharides, lipids, nucleic acids or blends, block copolymers or combinations of the above or copolymers of the corresponding monomers; agarose, methylcellulose, hydroxyproylmethylcellulose, xyloglucan, acetan, carrageenan, xanthan gum/locust beangum, gelatine, collagen (particularly Type 1), PLURONICS™, POLOXAMERS™, POLY(N

isopropylacrylmide) and N-isopropylacrylmide copolymers.

The device may be seeded with cells that are either terminally

differentiated or capable of undergoing phenotypic change e.g. stem cells, pluripotent cells and other precursor cells. Suitable cells include stem cells, cartilage precursor cells, cartilage forming cells, chondrocytes, cells that can differentiate into cartilage forming cells or cells that can induce other cells to differentiate into cartilage forming cells or a mixture thereof. Preferably, the cells used according to the present invention are autologous or allogenic, although xenogenic cells may also be used. The cells are advantageously chosen from the list comprising stem cells, cartilage precursor cells, cartilage forming cells, chondrocytes, cells that can differentiate into cartilage forming cells or cells that can induce other cells to differentiate into cartilage forming cells or a mixture thereof.

The cell-loaded device may be incubated under standard cell culturing techniques known to those in the art. Furthermore, the device may be incubated under mechanical strain, as disclosed in our patent application PCT/GB94/01455, the entire contents of which are incorporated herein by reference.

Cells and/or actives to be seeded into the device to form a layered structure, as occurs in native growing cartilage, rather than

as the homogeneous mass that usually occurred with earlier tissue engineered scaffolds. Previous tissue engineered cartilage devices have tended to be seeded with a homogeneous distribution of mixed chondrocytes with the intention of producing a uniform piece of cartilage tissue. Cartilage is however a layered structure with a distinct superficial zone of chondroprogenitor cells, a mid zone of chondrocytes and a deep zone which interacts with the underlying

mineralized cartilage, sub-chondral bone plate and trabecular bone. The design of the device of the present invention facilitates the production of an implant that recreates this layered structure. The flexible material, shell wall and elongate elements can be seeded

with one type of cell, e.g. a chondroprogenitor cell, whilst the void spaces between the flexible surface, shell wall and elongate elements can be seeded with a second type of cell, e.g. mature chondrocytes. The whole assembly can be embedded onto the subchondral bone plate with a "grout" that contains active agents to promote osseous integration, (e.g. BMPs and bisphosphonates).

The device may be sold with cells already pre-loaded or may be sold as a kit of parts.

In contrast to the prior art the device of the present invention conforms to the contours of the defect. The formula Gap=((2πrφ/360j - (2π(r- H)φ/360) - D predicts the theoretical distance between the elongate elements required to form a device that will conform to a surface with a particular radius of curvature (see Fig 5 and Fig 6). A device can therefore be mass-produced with elongate elements distributed so that it will conform to an "average" defect, (e.g. on a medial femoral condyl), but which retains the flexibility to conform to the particular curvature of an individual patients joint defect. In

preferred embodiments the gap between elongate elements is between 0.5mm and 6mm. In other embodiments the gap between elongate elements is approximately 2.5mm.

Although the formula Gap=((2πrφ/360j - (2π(r- H)φ/360) - D predicts the theoretical spacing, in reality the elongate elements flex slightly, so the actual spacing required is 70-90% of the theoretical figure. The exact number required depends upon the material used

to form the scaffold and can be determined experimentally.

It is this combination of (i) improved fixation compared with earlier chondral implants, (see table 1); (ii) minimal damage to the sub-chondral bone plate and underlying bone compared with earlier

osteochondral defects (see figure 1); (iii) improved mechanical properties over earlier non woven scaffolds (whilst retaining > 90% porosity in the structure of the scaffold); and (iv) the ability to change from a planar conformation to the contours of an individual joint defect that marks this design as an improvement over earlier tissue engineered cartilage scaffold designs.

The device of the present invention may be used for the partial or total replacement of connective tissue such as cartilage and / or bone in mammalian organisms.

In accordance with a second aspect of the present invention there is provided a method for the total or partial replacement of connective tissue in a mammalian patient comprising the step of implanting a device as hereinbefore described.

According to an aspect of the present invention there is

provided a method of repairing a cartilage lesion by (i) debriding the defective articular surface down to the sub-chondral bone; (ii) cutting a circumferential groove approximately 1mm wide and 5 mm deep around the perimeter of the defect into the subchondral bone; (iii) implanting a device according to the present invention such that the shell wall enters the circumferential groove and the flexible material conforms to the curved joint surface.

According to an aspect of the present invention there is

provided a method of treating cartilage defects comprising the steps of debriding the damaged cartilage back to, but not into, the subchondral bone; cutting a peripheral groove into the subchondral bone; inserting the device such that the shell wall enters the peripheral groove, the elongate elements abut onto the subchondral bone and the flexible material is positioned substantially flush with

the cartilage surface.

Reference is made, by way of example, to the following figures and examples:

Figure 1 shows an example of a device with a flexible surface, outer shell wall and cylindrical elongate elements.

Figure 2 shows a transverse section of the device shown in Figure 1.

Figure 3 shows a comparison between devices of the prior art and a device according to the present invention.

Figure 4 is a diagrammatic representation of a device after implantation into a cartilage defect.

Figure 5 is a diagrammatic representation of how a device curves to align with a curved joint surface.

Figure 6 shows the definition of parameters used in the gap formula.

Figure 7 shows the prepared cartilage defect site and an example of the device partially in place.

Figure 8 shows an example of cartilage tissue formed within the

device in vivo.

Detailed description

In Figure 1 flexible material 100 is in a circular configuration. Outer shell wall 120 is integrally formed with flexible material 100 and extends continuously around the circumference of, and longitudinally

away from, flexible material 100. A plurality of elongate elements 110 are bonded to flexible material 100. The elongate elements are hollow circular cylinders that are open at one end.

Figure 2 shows a transverse section of Figure 1. Outer shell wall 220 and elongate elements 210 are attached substantially perpendicular to flexible material 200. Outer shell wall 220 is longer than elongate elements 210. There are void spaces 230 within the hollow of the elongate elements and void spaces 240 between the elongate elements and within the outer shell wall.

Figure 3 shows a comparison between devices of the prior art and a device according to the present invention. In this diagrammatic representation there is a cartilage layer 300 and bone layer 310.

Prior art device 320 sits on the osteochondral interface. Prior art device 330 is stuck to the osteochondral interface by an adhering agent. Prior art device 340 is an example of an osteochondral device that is inserted into a deep socket cut into the subchondral bone. In contrast, in device 350 the elongate elements rest on the osteochondral interface. The outer shell wall penetrates into bone 310 anchoring the implant in the defect site.

Figure 4 is a diagrammatic representation of a device after implantation into a cartilage defect. The device is anchored into the

cartilage 400 with the flexible material 430 substantially flush with the cartilage surface. Elongate elements 440 rest on osteochondral interface region 410. Outer shell wall 450 penetrates into bone 420 to anchor the device in place.

Figure 5 is a diagrammatic representation of how a device curves to align with a curved joint surface. The ostechondral interface region 530 has a curved surface. Elongate elements 510 abut substantially

perpendicular to this region but are attached substantially perpendicular to flexible material 500. This geometry results in the flexible material curving to follow the contour of the joint and the elongate elements become angled relative to each other with angle 550. As the flexible material curves the gap between elongate elements 540 will narrow at a point distal to the flexible material. The gap proximal to the flexible material will remain constant. Outer shell wall 520 anchors the implant into the subchondral bone

Figure 6 shows the definition of parameters used in the gap formula to calculate the optimal distance S between adjacent elongate elements. D is the diameter of a cylindrical elongate element; H is

the length of a cylindrical elongate element; φ is the angle between two elongate elements when resting on the osteochondral interface; r is the radius of curvature for the joint surface at the defect site.

Figure 7 shows the prepared cartilage defect site and an example of the device partially in place. Panel A shows a cartilage defect site prepared for device implantation. Cartilage 700 has been debrided back to subchondral bone 720. A circumferential groove 710 has been cut into the subchondral bone. Panel B shows a device 750 partially inserted into the prepared defect site. Once fully inserted the device will be substantially flush with cartilage surface 740.

Figure 8 shows an example of cartilage tissue formed within the device in vivo. Prior to implantation chondroprogenitor cells were seeded into flexible material 810 and mature chondrocyte aggregates were seeded into the void spaces of the device.

Following implantation in vivo the chondroprogenitor cells formed a layer 800 upon the surface of the flexible material reproducing the superficial layer of progenitor cells seen in normal growing cartilage.

The mature chondrocytes secreted extracellular matrix, fusing the

aggregates into continuous cartilage tissue 820.

Examples

Example 1. Production of cartilage repair scaffold

Cartilage repair scaffolds were prepared from PCL re- enforced PET fabric. Powdered PCL (CAPA686, Solway) was dissolved in Chloroform (GPC Grade) to form a 6% w/v solution. Strips of PET felt, (0.5mm thick, 100mg/cc, 2.5dTex filaments) were bonded by dipping them into the PCL solution and drying them flat on release paper. The coating weights were in the range 0.88 to

0.93 gPCL/gPET. The re-enforced PET felts were then cut into discs with a die cutter. To form the scaffolds, the PCL re-enforced PET felt discs were heated on a release-paper covered hot plate (~110°C) and pressed into shape over a metal former. Samples were inspected for correct shape and absence of viable defects. Incorrectly shaped samples were either reworked (a maximum of 2 times) or discarded. On final inspection all the scaffolds in the batch were assessed to be correctly formed and free from viable defects.

Example 2. Mechanical testing of scaffold fixation

The medial and lateral femoral condyls of ovine stifle joints were implanted with either (a) a 0.5 mm x 8mm diameter PET non

woven disk press fitted into an 8mm full thickness chondral defect, 'Implant A'; (b) a 0.5 mm x 8mm diameter PET non woven disk glued into an 8mm full thickness chondral defect with fibrin glue, 'Implant B'; (c) a cylindrical osteochondral implant consisting of a PET 3 dimensional knitted fabric bonded to a non-woven PET fabric (5mm x 8mm diameter) press-fitted into an 8mm diameter x 5mm deep osteochondral defect created with a guide wire and flat bottomed drill, 'Implant C; or (d) an implant of the type described in this disclosure formed from PCL re-enforced PET (5mm deep shell wall x

8mm diameter) press fitted into an 8mm full thickness chondral defect with a 5mm deep x 0.5 mm wide peripheral gutter cut into the underlying bone using an Acufex powered trephine, 'Implant D'.

The ovine stifle joints with the implants were mounted in a test rig and cyclically loaded in a simulated walking action at 0.5 Hz for a total of 84,000 cycles, (46.7 hours). Specimens were inspected after sequential loading periods of 5h (9000 cycles), 16 h (30000 cycles), 24 hours (45000 cycles) and at the completion of the experiment.

The implants were scored at each time point as either (1) in place and intact; (2) in place but damaged or (3) detached.

Results are as recorded in table 1. Results showed that biomaterial disks (Implant A) that had been press fitted into chondral defects were inadequately fixed into the defect site. All the implants became detached within 5 hours. This finding is in accordance with our previous experience with early cell-seeded tissue engineered cartilage devices which had similar dimensions to the samples tested here and which also showed a high tendency to detach shortly after implantation. Implants from group B (biomaterial disks glued in place with fibrin glue) were likewise shown to be inadequately fixed into the defect site. All the implants became detached during the first 24 hours of cyclic loading. Again this result was in accordance with previous findings that suggested that whilst

fibrin glue could be used to hold an implant in position within an unloaded chondral defect site, it was not strong enough to withstand the forces applied to a joint during locomotion. This finding suggests that a mechanical fixation method will be needed to secure ah implant in the defect site. The results from implant group C, (cylindrical devices press fitted into deep sockets cut into the subchondral bone), showed that this type of implant is adequately fixed into the defect site. All the samples remained in place for the duration of the experiment. This is in accordance with previous in vivo results in which showed that biomaterial implants of this type

could be securely fixed in place by locating them in deep osteochondral sockets. Unfortunately, previous results also showed that this method of fixation leads to severe disruption of the joints physiology and that it is associated with bone cyst formation, widespread osteosclerotic changes within the joint and the

subsequent collapse of the implant. Overall these results suggest that a fixation method is required that gives the mechanical resilience associated with group C without causing unnecessary damage to the underlying bone. Results from group D, (implants of the type disclosed in this patent), remained intact and in place for the duration of the experiment suggesting that they were adequately attached to the implant site. Furthermore, because the implant fixation method does not require major damage to the subchondral bone below the implant, this type of device should not initiate bone cyst formation or sclerotic changes within the joint. It therefore constitutes an improved scaffold design.

Table 1. Experimental investigation of the effect of graft- fixation and design upon the stability of attachment to implant site in an ex vivo ovine stifle joint model.

Example 3. Cartilage tissue formation in cell-seeded PET/ PCL scaffolds

Combinations of chondroprogenitor cells (either superficial zone chondroprogenitor cells or marrow derived mesenchymal stem cells) and/ or mature chondrocytes were seeded into PET /PCL scaffolds and implanted sub-cutaneously in nude mice to investigate the ability of cells to form macroscopic pieces of continuous cartilage within this type of scaffold.

(i) Isolation of superficial zone chondroprogenitor cells. The wells of a 24 well plate were coated with bovine plasma fibronectin (pFN; 10μg ml ^"1 in Dulbecco's PBS with 1mM MgCI ₂ and ImMCaCI ₂ ;

Sigma; 16 h; 4 ⁰ C), the supernatant discarded and the plates blocked with 1% BSA.

Superficial zone cartilage was isolated from the surface of 2-3 week old bovine metatarsophalangeal joints by fine dissection. The cartilage was digested with 0.1% pronase (Merck, 4x10 ⁶ units/g in DMEM/ 5% FCS; 37°C for 3 hours), rinsed in PBS and the cells released by incubation in 0.04% collagenase (Worthington, 237U/mg; in DMEM/ 5% FCS; 16 hours at 37°C with gentle shaking). The tissue digests were filtered (70μm; Falcon), centrifuged (1000 rpm; 5 minutes), washed with serum free DMEM (10 ml), re- centrifuged and resuspended in serum free DMEM (10ml). Cell number was then counted using a haemocytometer and the volume adjusted to a final concentration of 700 cells ml ^"1 . After isolation, 700 superficial zone chondrocytes in serum free DMEM were seeded into wells of coated 24 well plates and incubated at 37 ⁰ C for 20 minutes. The media was gently swirled,

removed and replaced with fresh DMEM containing 5% FCS. Plates were incubated at 37 ⁰ C in a humidified atmosphere of 5% CO ₂ .

(W) Isolation of bone marrow stromal cells Subchondral trabecular bone (~1 cm ³ ) was harvested from 2-3 week old bovine metatarsophalangeal joints using a 3.5 mm Acufex mosaicplasty punch. The bone was fragmented into 0.1% collagenase (Worthington type 2 in DMEM/ 5%FCS; 20 ml) and incubated at 37 ⁰ C for 90 minutes. The bone fragments were washed twice with DMEM/ 10% FCS (20 ml), a further twice with vortexing and the residual bone fragments discarded. The digests and washings were combined, centrifuged (lOOOrpm; 5 minutes), the cell pellets resuspended in DMEM/FSC, re-centrifuged, resuspended in DMEM/FCS (30ml) and cultured in 2 x T175 flasks. After 5-7 days

the cultures were rinsed with PBS to remove the non-adherent cells, trypsinised, the released cells centrifuged (lOOOrpm; 5 minutes), the cell pellets resuspended in DMEM/FCS (35ml) and transferred to 2 x fresh T175 flasks. The cultures were incubated at 37°C in a humidified atmosphere of 5% CO ₂ .

(iii) Isolation of bovine chondrocytes. Pairs of full thickness cartilage strips (~5 x 35 mm) were obtained from 2-3 week old bovine metatarsophalangeal joints, finely sliced and digested in collagenase, (Worthington type 2; 0.1% in DMEM/5%FCS; 16 hours; 37 ⁰ C). The tissue digests were filtered (70μm; Falcon), centrifuged (lOOOrpm; 10 minutes; Megafuge), the supernatant discarded and the cell pellet washed twice with DMEM/FCS.

(M Sterilisation and FCS pre-treatment of PET/PCL scaffolds. PET/PCL scaffolds were immersed in 70% ethanol/H ₂ O and gently agitated until all the bubbles had been removed. The scaffolds were

then transferred to 70% ethanol/ acetone for 10 minutes and pre- treated with foetal calf serum by immersion in FCS (16 hours; 4 ⁰ C).

(v) Seeding scaffolds with superficial zone chondroproqenitor cells. FCS pre-treated PET/PCL scaffolds were placed into a 50 ml Falcon tube containing 5 x10 ⁶ superficial zone chondroprogenitor cells (P3; see section ii) in DMEM/FCS/ASCP (30ml) and the tubes sealed with vented caps (Sartorius 0.2μm). The cultures were incubated on the modified Spiramixer for 5 days and then incubated statically for the remainder of the 2 week period at 37°C in a humidified atmosphere of 5% CO ₂ .

(vi) Seeding scaffolds with bone marrow stromal cells An FCS pre- treated PET/PCL scaffold was placed into a 50 ml Falcon tube containing 5 x10 ⁶ bone marrow stromal cells (P3; see section ii) in DMEM/FCS/ASCP (30ml) and the tube sealed with a vented cap (Sartorius 0.2μm). The culture was incubated o/n on the modified Spiramixer and then incubated statically for 2 weeks at 37°C in a humidified atmosphere of 5% CO ₂ .

(vii) Seeding scaffolds with chondrocytes. An FCS pre-treated PET/PCL scaffold was placed into a 50 ml Falcon tube containing 5 x10 ⁶ freshly isolated chondrocytes (see section iii) in DMEM/FCS/ASCP (30ml) and the tube sealed with a vented cap (Sartorius 0.2μm). The culture was incubated o/n on the modified spiramixer and then incubated statically for 2 weeks at 37°C in a humidified atmosphere of 5% CO ₂ .

(viii) Seeding superficial zone chondroproαenitor cells into pots to form SZC aggregates. Superficial zone chondroprogenitor cells (7 x 10 ⁶ cells; P2; see section i) were added to 60 ml sample pots

(Sterilin) that had been fitted with vented caps (0.2μm; Sartorius). The cells were allowed to settle and aggregate for 16 hours and then gently agitated in an attempt to break the cell layer into clumps. The cultures were then incubated for 2 weeks at 37 ⁰ C in a humidified atmosphere of 5% CO2.

(ix) Seeding chondrocytes into pots to form chondrocyte aggregates. Freshly isolated chondrocytes (7 x 10 ⁶ cells; see section iii) were added to 60 ml sample pots (Sterilin) that had been fitted with vented caps (0.2μm; Sartorius). The cells were allowed to settle and aggregate for 16 hours and then gently agitated to break the cell

layer into clumps. The cultures were then incubated for 2 weeks at 37°C in a humidified atmosphere of 5% CO ₂ .

M Preparation of vented 60ml pots with agarose underlays. Agarose, (2%, Promega V3125; 1 litre in DMEM, Sigma D6429), was sterilised by autoclaving (121°C; 15 minutes), and melted by incubation in a near-boiling water bath. The melted sterile agarose (2%; 5ml) was added to the base of 60 ml sample pots (Sterilin) and the culture underlays set at 4 ⁰ C. Vented caps were sterilised by sequential immersion in 1% Virkon, water, 70% ethanol/H ₂ O and 70% ethanol/ acetone and allowed to air dry. A sterile vented cap was fitted to each of the 60 ml pots with the agarose underlays and closed with a 0.2 μm filter (Sartorius).

(xi) Preparation of centrifuge tubes with agarose underlays. Agarose, (4%, Sigma A-9045; 80ml in DMEM, Sigma D6429), was sterilised by autoclaving (121 °C; 15 minutes), and melted by incubation in a near-boiling water bath. The melted sterile agarose (4%; 0.4ml) was added to the base of centrifuge tubes (Sterilin, 144AS) and the culture underlays set at 4 ⁰ C.

(xii) Addition of chondrocyte clumps into cell-seeded scaffolds. DMEM/FCS/ASCP (5ml) was added to the centrifuge tubes with agarose bases prepared in section (xi) and either (a) a superficial zone chondroprogenitor-seeded PET/PCL scaffold (see section v); (b) a bone marrow stromal cell-seeded PET/PCL scaffold (see section vi); or (c) a chondrocyte-seeded PET/PCL scaffold (see section vii) placed into each tube. In order to add 100μl of chondrocyte 'clumps' to each cell-seeded PET/PCL scaffold, the contents of the 60 ml pots of chondrocyte clumps, (section ix), were centrifuged, (2000rpm; 10 min), the pellets combined in a single 10

ml aliquot of serum free DMEM, re-centrifuged (1500 rpm; 5 min) and the supernatant discarded. The volume of the combined chondrocyte 'clump' pellet was estimated by adding known volumes of liquid to a second centrifuge tube. The chondrocyte 'clump' pellet was then resuspended in sufficient DMEM/FCS/ASCP to give a final concentration of 100μl of chondrocyte 'clumps' in 5 ml of medium. Aliquots, (5ml) of this chondrocyte 'clump' suspension were then added to each of the centrifuge tubes containing the PET/PCL scaffolds and the tubes placed vertically into a rack and incubated at 37°C for 48 hours. At the end of this incubation period the cell- seeded scaffolds were recovered from the centrifuge tubes and placed into vented 60ml pots with agarose underlays (see section x)

containing 30 ml DMEM/FCS/ASCP and incubated at 37 ⁰ C in a humidified atmosphere of 5% CO ₂ .

(xiii) Addition of superficial zone chondrocyte clumps into SZC seeded scaffolds. In order to seed groups A, B, D and E with 100μl of chondrocyte clumps, the products of 7 x 10 ⁶ chondrocytes, (see section ix), were diluted to 8.6 ml and 5 ml of that suspension added to each scaffold. Thus each scaffold received the equivalent of 0.58 of a 'pot' i.e. the products of 4 x 10 ⁶ cells.

Therefore, to seed the scaffolds from group C with SZC clumps, the SZC clumps were recovered from the 60 ml pots (see section viii) by sequential vigorous pipetting and recovery of the culture medium, washing the culture surface with PBS (10 ml) and trypsinisation (5 ml). These washings (containing the SZC aggregates) were pooled, centrifuged, (2000rpm; 10 minutes), the supernatant discarded and the SZC clumps from each pot resuspended in 8.6 ml of DMEM/FCS/ASCP. 5 ml of this suspension was then added to each of the centrifuge tubes containing the SZC seeded PET/PCL scaffolds from group C and the tubes placed vertically into a rack

and incubated at 37 ⁰ C for 48 hours. At the end of this incubation period the cell-seeded scaffolds were recovered from the centrifuge tubes and placed into vented 60ml pots with agarose underlays (see section x) containing 30 ml DMEM/FCS/ASCP and incubated at 37 ⁰ C in a humidified atmosphere of 5% CO ₂ .

(xiv) Addition of BMX layer to cell-seeded devices. Bone matrix extract/ fibrin paste was prepared as follows. Neonatal bovine phalanges were dissected free from adherent soft tissue and periosteum, cut across their epiphesis with a hack-saw, the diaphyses cut longitudinally and the marrow removed by washing in

warm water. The resulting cleaned cortical bone (59g) was powdered in a liquid nitrogen cooled mill (3 x 2 minutes; 10Hz; Spex Mill), and washed with 70% ethanol (total volume 150ml; 2 x 100 ml changes at 4 ⁰ C over a total of 16 hours). The supernatant (100ml) was removed and the bone powder defatted in 70% ethanol/ 30% acetone (1 hour; room temperature). The supernatant was completely removed and the bone dried under aseptic conditions at room temperature. The resulting bone fragments were sieved (710μm; Endecotts; 8.3g discarded) and extracted twice more with 70% ethanol/ 30% acetone (200ml) for a total of 16 hours at 4 ⁰ C. The defatted bone powder was rinsed with PBS, (150ml), the

supernatant discarded, and demineralised with HCI (0.6 M; 400 ml; 5 hours with occasional mixing). The material was centrifuged (1000 rpm; 5min; Megafuge 1.0), resuspended in 150 ml PBS and NaOH added (1 N; 20 ml). The material was allowed to equilibrate for 16 hours at 4 ⁰ C after which time it was re-neutralised with an additional

10 ml of 1 N NaOH. The demineralised bone powder was then centrifuged (lOOOrpm; 5 minutes; Megafuge), resuspended in PBS

(200ml) and stored at 4 ⁰ C. The resulting material was described as

BMX suspension. BMX suspension (10 ml) was centrifuged (1000

rpm; 5 minutes; Megafuge), the supernatant discarded and the pellet resuspended in a sterile fibrinogen solution (1 ml; 5mg/ml in PBS;

Sigma F8630). The fibrinogen / BMX mixture was aliquoted (150μl) to the wells of a 24 well plate and fibrinogen (400μl; 5mg/ml) and thrombin (60μl; 1000U/ml; Sigma T4648) added to each well. The contents of the wells were allowed to coagulate for > 5 minutes before their contents were gently mixed and stirred with a small spatula. The insoluble material in the wells separated from the liquid to form a 'plug' of fibrin and BMX that was lifted from the well and

pressed against the side of an empty well to remove the residual liquid. The resulting BMX/fibrin paste was compressed into a flat cylinder with the plunger of a 2ml syringe and a spatula and carefully placed onto the underside of the seeded PET/PCL scaffolds from experimental group A (seeded with SZC and chondrocytes). The resulting constructs were transferred to vented 60 ml pots with 'agarose underlays ¹ containing 30 ml of UMEM/IO^oFCS/ascorbate phosphate and incubated at 37°C in a humidified atmosphere of 5% CO ₂ .

Results showed that when chondrocytes were seeded into the PET/ PCL scaffolds and the cell-seeded devices implanted in vivo, then the cells formed a continuous piece of cartilage within the scaffold. When chondroprogenitor cells were seeded into the

flexible material, shell wall and elongate elements of the PET/ PCL scaffolds and mature chondrocyte aggregates seeded into the void spaces contained within the scaffolds and the resulting construct implanted in vivo, then cells formed a continuous piece of tissue with two distinct layers, an upper layer of flattened progenitor cells with a fibroblastic morphology (see figure 8, item 800) and a lower layer consisting of a continuous piece of cartilaginous tissue in which the cells have a mature chondrocyte morphology (see figure 8, item

820). Addition of bioactive molecules to the mature chondrocyte layer (e.g. demineralised bone particles) improved the quality of the tissue formed, indicating that the scaffold described within the patent could be used to deliver combinations of cells and/ or biological molecules to a cartilage defect to improve repair.

Example 4 Repair of full thickness cartilage lesions or early stage QA in humans.

A device is seeded with a superficial layer of

chondroprogenitor cells, a middle layer of "mature chondrocyte clumps", (derived from in vitro pre-differentiated chondroprogenitor cells) and a bone facing layer of "grout" consisting of demineralised bole matrix/fibrin/bisphosphonate paste.

The surgeon exposes the patient's defect site, (generally a circumscribed articular cartilage lesion), and debrides any remaining cartilage/fibrous repair tissue to create a circular, full thickness chondral defect with sound cartilaginous margins. This debridement exposes, but does not significantly damage, the underlying subchondral bone plate. A circular, peripheral groove (5 mm deep x 0.5 mm wide) is then cut around the circumference of the debrided defect site through the sub chondral bone plate and into the underlying trabecular bone using a powered, toothed trephine.

The cell-seeded device is then inserted into the defect site such that the lower portion of the peripheral, shell-wall becomes located in the peripheral groove of the prepared defect site. Pressing upon the circumference of the implant aligns the margins of the device with the surrounding cartilage, draws the flexible material of the device tight, locates the elongate elements upon the subchondral bone and therefore brings the surface contour of the device

into line with the healthy contour of the joints cartilage.

The patient's joint is then closed and the defect allowed to heal. During healing the patient's bone rapidly grows into and across the 0.5 mm thick porous wall of the device shell-wall, firmly anchoring the device in place. Meanwhile, the implanted mature chondrocytes proliferate, deposit extracellular matrix and agglomerate to form a continuous cartilaginous mass within the

defect site. The chondroprogenitor cells near the articular surface form a superficial fibrous layer and potentially act as a reserve of proliferative cartilage forming cells that could contribute to the long term viability of the implant. The BMX/fibrin releases growth factors (e.g. TGFβ superfamily members) into the developing implant, promoting chondrocyte differentiation, increasing cartilage matrix synthesis and potentially promoting osseous integration. The limited

mobility of the periphery of the device relative to the surrounding host tissue and the presence of chondroprogenitor cells at the periphery and top of the device maximizes the chance of osteointegration.