WO2008070186A2

WO2008070186A2 - Scaffold apparatus for promoting tendon-to-bone fixation

Info

Publication number: WO2008070186A2
Application number: PCT/US2007/025127
Authority: WO
Inventors: Helen H. Lu; Jeffrey Spalazzi
Original assignee: The Trustees Of Columbia University In The City Of New York
Priority date: 2006-12-06
Filing date: 2007-12-06
Publication date: 2008-06-12
Also published as: WO2008070186A3

Abstract

This invention provides a graft collar for fixing tendon to bone. In one embodiment, the graft collar comprises a sheet of biopolymer mesh. In another embodiment, the graft collar comprises a polymer-fiber mesh. In another embodiment, the graft collar comprises (a) a first region comprising a biopolymer mesh and hydrogel and (b) a second region adjoining the first region and comprising a biopolymer mesh. In another embodiment, the graft collar comprises (a) a first region comprising a polymer-fiber mesh and hydrogel and (b) a second region adjoining the first region and comprising a polymer-fiber mesh.

Description

SCAFFOLD APPARATUS FOR PROMOTING TENDON-TO-BONE FIXATION

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 methods and apparatuses described and claimed herein.

Background

This application relates to musculoskeletal tissue engineering, and more particularly, to techniques for tendon-to-bone fixation. For example, a graft collar for fixing tendon to bone in a subject is discussed below. Some exemplary embodiments which include a soft tissue- bone interface are discussed.

As an example of a soft tissue-bone interface, the human anterior cruciate ligament (ACL) is described below. The ACL and ACL-bone interface are used in the following discussion as an example and to aid in understanding the description of the methods and apparatuses of this application. This discussion, however, is not intended to, and should not be construed to, limit the claims of this application.

The ACL consists of a band of regularly oriented, dense connective tissue that spans the junction between the femur and tibia. It participates in knee motion control and acts as a joint stabilizer, serving as the primary restraint to anterior tibial translation. The natural ACL-bone interface consists of three regions: ligament, fibrocartilage (non-mineralized and mineralized) and bone. The natural ligament to bone interface is arranged linearly from ligament to fibrocartilage and to bone. The transition results in varying cellular, chemical, and mechanical properties across the interface, and acts to minimize stress concentrations from soft tissue to bone.

The ACL is the most often injured ligament of the knee. Due to its inherently poor healing potential and limited vascularization, ACL ruptures do not heal effectively upon injury, and surgical intervention is typically needed to restore normal function to the knee.

Clinically, autogenous grafts based on either bone- patellar tendon-bone (BPTB) or hamstring-tendon (HST) grafts are often a preferred grafting system for ACL reconstruction, primarily due to a lack of alternative grafting solutions. Current ACL grafts are limited by donor site morbidity, tendonitis and arthritis. Synthetic grafts may exhibit good short term results but encounter clinical failure in long-term follow-ups, since they are unable to duplicate the mechanical strength and structural properties of human ACL tissue. ACL tears and ruptures are currently commonly repaired using semitendinous grafts. Although semitendinous autografts are superior, they often fail at the insertion site between the graft and the bone tunnel. One of the major causes of failure in this type of reconstruction grafts is its inability to regenerate the soft-tissue to bone interface . Despite their distinct advantages over synthetic substitutes, autogenous grafts have a relatively high failure rate. A primary cause for the high failure rate is the lack of consistent graft integration with the subchondral bone within bone tunnels. The site of graft contact in femoral or tibial tunnels represents the weakest point mechanically in the early post-operative healing period. Therefore, success of ACL reconstructive surgery depends heavily on the extent of graft integration with bone.

ACL reconstruction based on autografts often results in loss of functional strength from an initial implantation time, followed by a gradual increase in strength that does not typically reach the original magnitude. Despite its clinical success, long term performance of autogenous ligament substitutes is dependent on a variety of factors, including structural and material properties of the graft, initial graft tension, intrarticular position of the graft, as well as fixation of the graft. These grafts typically do not achieve normal restoration of ACL morphology and knee stability.

There is often a lack of graft integration with host tissue, in particular at bony tunnels, which contributes to suboptimal clinical outcome of these grafts. The fixation sites at the tibial and femoral tunnels, instead of the isolated strength of the graft material, have been identified as mechanically weak points in the reconstructed ACL. Poor graft integration may lead to enlargement of the bone tunnels, and in turn may compromise the long term stability of the graft. Increased emphasis has been placed on graft fixation, as post surgery rehabilitation protocols require the immediate ability to exercise full range of motion, reestablish neuromuscular function and weight bearing. During ACL reconstruction, the bone-patellar tendon-bone or hamstring-tendon graft is fixed into the tibial and femoral tunnels using a variety of fixation techniques. Fixation devices include, for example, staples, screw and washer, press fit EndoButton® devices, and interference screws. In many instances, EndoButton® devices or Mitek® Anchor devices are utilized for fixation of femoral insertions. Staples, interference screws, or interference screws combined with washers can be used to fix the graft to the tibial region.

Recently, interference screws have emerged as a standard device for graft fixation. The interference screw, about 9 mm in diameter and at least 20 mm in length, is used routinely to secure tendon to bone and bone to bone in ligament reconstruction. Surgically, the knee is flexed and the screw is inserted from the para-patellar incision into the tibial socket, and the tibial screw is inserted just underneath the joint surface. After tension is applied to the femoral graft and the knee is fully flexed, the femoral tunnel screw is inserted. This procedure has been reported to result in stiffness and fixation strength levels which are adequate for daily activities and progressive rehabilitation programs.

While the use of interference screws have improved the fixation of ACL grafts, mechanical considerations and biomaterial-related issues associated with existing screw systems have limited the long term functionality of the ligament substitutes. Screw-related laceration of either the ligament substitute or bone plug suture has been reported. In some cases, tibial screw removal was necessary to reduce the pain suffered by the patient. Stress relaxation, distortion of magnetic resonance imaging, and corrosion of metallic screws have provided motivation for development of biodegradable screws based on poly-α-hydroxy acids. While lower incidence of graft laceration was reported for biodegradable screws, the highest interference fixation strength of the grafts to bone is reported to be 475 N, which is significantly lower than the attachment strength of ACL to bone. When tendon-to-bone fixation with polylactic acid-based interference screws was examined in a sheep model, intraligamentous failure was reported by 6 weeks. In addition, fixation strength is dependent on quality of bone (mineral density) and bone compression.

Two insertion zones can be found in the ACL, one at the femoral end and another located at the tibial attachment site. The ACL can attach to mineralized tissue through insertion of collagen fibrils, and there exists a gradual transition from soft tissue to bone. The femoral attachment area in the human ACL was measured to be 113 ± 27 mm² and 136 ± 33 mm² for the tibia insertion. With the exception of the mode of collagen insertion into the subchondral bone, the transition from ACL to bone is histologically similar for the femoral and tibial insertion sites.

The insertion site is comprised of four different zones: ligament, non-mineralized fibrocartilage, mineralized fibrocartilage, and bone. The first zone, which is the ligament proper, is composed of solitary, spindle-shaped fibroblasts aligned in rows, and embedded in parallel collagen fibril bundles of 70-150 μm in diameter. Primarily type I collagen makes up the extracellular matrix, and type III collagen, which are small reticular fibers, are located between the collagen I fibril bundles. The second zone, which is fibro-cartilaginous in nature, is composed of ovoid-shaped chondrocyte-like cells. The cells do not lie solitarily, but are aligned in rows of 3-15 cells per row. Collagen fibril bundles are not strictly parallel and much larger than those found in zone 1. Type II collagen is now found within the pericellular matrix of the chondrocytes, with the matrix still made up predominantly of type I collagen. This zone is primarily avascular, and the primary sulfated proteoglycan is aggrecan. The next zone is mineralized fibrocartilage . In this zone, chondrocytes appear more circular and hypertrophic, surrounded by larger pericellular matrix distal from the ACL. Type X collagen, a specific marker for hypertrophic chondrocytes and subsequent mineralization, is detected and found only within this zone. The interface between mineralized fibrocartilage and subjacent bone is characterized by deep inter-digitations . Increasing number of deep inter- digitations is positively correlated to increased resistance to shear and tensile forces during development of rabbit ligament insertions. The last zone is the subchondral bone and the cells present are osteoblasts, osteocytes and osteoclasts. The predominant collagen is type I and fibrocartilage-specific markers such as type II collagen are no longer present. Progressing through the four different zones which make up the native ACL insertion zone, several cell types are identified: ligament fibroblasts, chondrocytes, hypertrophic chondrocytes and osteoblasts, osteoclasts, and osteocytes. The development of in vitro multi-cell type culture systems facilitates the formation of the transition zones.

For bone-patellar tendon-bone grafts, bone-to-bone integration with the aid of interference screws is the primary mechanism facilitating graft fixation. Several groups have examined the process of tendon-to-bone healing.

Blickenstaff et al. (1997) evaluated the histological and biomechanical changes during the healing of a semitendinous autograft for ACL reconstruction in a rabbit model. Graft integration occurred by the formation of an indirect tendon insertion to bone at 26 weeks. However, large differences in graft strength and stiffness remained between the normal semi-tendinous tendon and anterior cruciate ligament after 52 weeks of implantation.

In a similar model, Grana et al. (1994) reported that graft integration within the bone tunnel occurs by an intertwining of graft and connective tissue and anchoring of connective tissue to bone by collagenous fibers and bone formation in the tunnels. The collagenous fibers have the appearance of Sharpey's fibers seen in an indirect tendon insertion. Rodeo et al. (1993) examined tendon-to-bone healing in a canine model by transplanting digital extensor tendon into a bone tunnel within the proximal tibial metaphysis. A layer of cellular, fibrous tissue was found between the tendon and bone, and this fibrous layer matured and reorganized during the healing process. As the tendon integrated with bone through Sharpey's fibers, the strength of the interface increased between the second and the twelfth week after surgery. The progressive increase in strength was correlated with the degree of bone in growth, mineralization, and maturation of the healing tissue.

In most cases, tendon-to-bone healing with and without interference fixation does not result in the complete re- establishment of the normal transition zones of the native ACL-bone insertions. This inability to fully reproduce these structurally and functionally different regions at the junction between graft and bone is detrimental to the ability of the graft to transmit mechanical stress across the graft proper and leads to sites of stress concentration at the junction between soft tissue and bone.

Zonal variations from soft to hard tissue at the interface facilitate a gradual change in stiffness and can prevent build up of stress concentrations at the attachment sites.

The insertion zone is dominated by non-mineralized and mineralized fibrocartilage, which are tissues adept at transmitting compressive loads. Mechanical factors may be responsible for the development and maintenance of the fibrocartilagenous zone found at many of the interfaces between soft tissue and bone. The fibrocartilage zone with its expected gradual increase in stiffness appears less prone to failure.

Benjamin et al. (1991) suggested that the amount of calcified tissue in the insertion may be positively correlated to the force transmitted across the calcified zone .

Using simple histomorphometry techniques, Gao et al. determined that the thickness of the calcified fibrocartilage zone was 0.22 ± 0.7 mm and that this was not statistically different from the tibial insertion zone. While the ligament proper is primarily subjected to tensile and torsional loads, the load profile and stress distribution at the insertion zone is more complex.

Matyas et al. (1995) combined histomorphometry with a finite element model (FEM) to correlate tissue phenotype with stress state at the medial collateral ligament (MCL) femoral insertion zone. The FEM model predicted that when the MCL is under tension, the MCL midsubstance is subjected to tension and the highest principal compressive stress is found at the interface between ligament and bone.

Calcium phosphates have been shown to modulate cell morphology, proliferation and differentiation. Calcium ions can serve as a substrate for Ca²⁺-binding proteins, and modulate the function of cytoskeleton proteins involved in cell shape maintenance. Gregiore et al. (1987) examined human gingival fibroblasts and osteoblasts and reported that these cells underwent changes in morphology, cellular activity, and proliferation as a function of hydroxyapatite particle sizes. Culture distribution varied from a homogenous confluent monolayer to dense, asymmetric, and multilayers as particle size varied from less than 5 μm to greater than 50 μm, and proliferation changes correlated with hydroxyapatite particles size.

Cheung et al. (1985) further observed that fibroblast mitosis is stimulated with various types of calcium- containing complexes in a concentration-dependent fashion.

Chondrocytes are also dependent on both calcium and phosphates for their function and matrix mineralization. Wuthier et al. (1993) reported that matrix vesicles in fibrocartilage consist of calcium-acidic phospholipids- phosphate complex, which are formed from actively acquired calcium ions and an elevated cytosolic phosphate concentration.

Phosphate ions have been reported to enhance matrix mineralization without regulation of protein production or cell proliferation, likely because phosphate concentration is often the limiting step in mineralization. It has been demonstrated that human foreskin fibroblasts when grown in micromass cultures and under the stimulation of lactic acid can dedifferentiate into chondrocytes and produce type II collagen. Cheung et al. (1985) found a direct relationship between β-glycerophosphate concentrations and mineralization by both osteoblasts and fibroblasts. Increased mineralization by ligament fibroblasts is observed with increasing concentration of β-glycerophosphate, a media additive commonly used in osteoblast cultures. These reports strongly suggest the plasticity of the fibroblast response and that the de-differentiation of ligament fibroblasts is a function of mineral content in vitro.

Goulet et al . (2000) developed a bio-engineered ligament model, where ACL fibroblasts were added to the structure and bone plugs were used to anchor the bioengineered tissue. Fibroblasts isolated from human ACL were grown on bovine type I collagen, and the bony plugs were used to promote the anchoring of the implant within the bone tunnels.

Cooper et al. (2000) and Lu et al. (2001) developed a tissue engineered ACL scaffold using biodegradable polymer fibers braided into a 3-D scaffold. This scaffold has been shown to promote the attachment and growth of rabbit ACL cells in vitro and in vivo.

There remains a need, however, for improved techniques for tendon-to-bone fixation. Summary

This application provides improved scaffold apparatuses for promoting tendon-to-bone fixation. According to some exemplary embodiments (discussed infra) , a scaffold apparatus for promoting tendon-to-bone fixation can include (or take on the form of) a graft collar.

A graft collar, according to one embodiment, comprises a sheet of collagen mesh.

According to another embodiment, a graft collar comprises a sheet of polymer-fiber mesh.

This application further provides a graft collar for fixing tendon to bone in a subject, wherein the graft collar comprises, according to yet another embodiment, (a) a first region comprising a hydrogel and (b) a second region adjoining the first region and comprising a collagen mesh.

According to yet another embodiment, a graft collar for fixing tendon to bone in a subject comprises (a) a first region comprising a hydrogel and (b) a second region adjoining the first region and comprising a polymer-fiber mesh.

This application further provides a graft collar for fixing tendon to bone in a subject, wherein said graft collar comprises a sheet of mesh comprising fibers aligned substantially perpendicular in relation to a longitudinal axis of said tendon, wherein said mesh applies compression to the graft. This application further provides a graft collar for fixing tendon to bone in a subject, wherein said graft collar comprises a sheet of mesh comprising fibers aligned substantially parallel in relation to a longitudinal axis of said tendon, wherein said mesh applies lateral tension to the graft.

Brief Description of the Figures

Figure IA: A schematic diagram of a graft collar, wherein the graft collar comprises a sheet of biopolymer mesh or polymer-fiber mesh, according to one embodiment. Figure IB: A schematic diagram of a graft collar, wherein the graft collar comprises 2 regions wherein (i) region 1 comprises a biopolymer mesh or a polymer-fiber mesh and (ii) region 2 comprises a biopolymer mesh or a polymer-fiber mesh and a hydrogel, according to one embodiment .

Figure 2: A schematic diagram of a graft collar, wherein the graft collar comprises 2 regions wherein (i) region A comprises a biopolymer mesh or a polymer-fiber mesh and (ii) region B comprises a biopolymer mesh or a polymer- fiber mesh and a hydrogel, according to one embodiment. As indicated, additional substances can be added to regions A and B.

Figure 3A: Posterior view of an intact bovine anterior cruciate ligament (ACL) connecting the femur to the tibia (left). Figure 3B: Environmental scanning electron microscope (ESEM) image of transition from ligament (L) to fibrocartilage (FC) to bone (B) at the ACL insertion (upper right). Figure 3C: Histological micrograph of similar ACL to bone interface additionally showing mineralized fibrocartilage (MFC) zone (lower right).

Figures 4A and 4B show Bovine tibial-femoral joint after ACL and insertion site extraction (right) , ACL and insertion sites after excision. Figure 5A shows FTIR Spectra of BG immersed in SBF for up to 7 days. Presence of an amorphous Ca-P layer at 1 day, and of a crystalline layer at 3 days.

Figure 5B: SEM image of Ca-P nodules on BG surface (3 days in SBF) . Nodules are ~ 1 μm in size initially, and grew as immersion continued (15,00Ox). Figure 5C: EDXA spectrum of BG surfaces immersed in SBF for 3 days. The relative Ca/P ratio is « 1.67.

Figures 6A and 6B show environmental SEM images of Bovine ACL insertion Site (1 and 2), including a cross section of the ACL-femur insertion site, ACL fiber (L) left, fibrocartilage region (FC) middle, and sectioned bone (B) right (Figure 6A: 250X; Figure 6B: 500X) .

Figure 7A: SEM of the cross section of the femoral insertion zone, 100OX; Figure 7B: EDAX of the femoral insertion zone. The peak intensities of Ca, P are higher compared to those in ligament region.

Figure 8 shows apparent modulus versus indentation X- position across sample.

Figures 9A and 9B show X-Ray CT scans of discs made of poly-lactide-co-glycolide (PLAGA) 50:50 and bioactive glass (BG) submerged in SBF for 0 days (Figure 9A) and 28 days; Figure 9B shows the formation of Ca-P over time.

Figure 1OA: SEM image; Figure 1OB: EDAX of PLAGA-BG immersed in SBF for 14 days.

Figure 11 shows osteoblast grown on PLAGA-BG, 3 weeks. Figure 12 shows higher type I collagen type synthesis on PLAGA-BG.

Figure 13A: ALZ stain, ACL fibroblasts 14 days, 2Ox; Figure 13B: ALZ stain, interface, ACL 14 days, 2Ox; Figure 13C: ALZ stain, osteoblasts, ACL 14 days, 2Ox.

Figure 14A: ALP stain, ACL fibroblasts, 7 days, 32x; Figure 14B: ALP+DAPI stain, co-culture, 7 days, 32x; Figure 14C: ALP stain, osteoblasts, 7 days, 32x.

Figures 15A-15F show images of multiphase scaffold (Figures 15A-15C) and blow-ups of respective sections (Figures 15D-15F) .

Figures 16A-16C show multiphasic scaffold for co-culture of ligament fibroblasts and osteoblasts; Figure 16A and Figure 16B: images of a sample scaffold; Figure 16C: schematic of scaffold design depicting the three layers.

Figures 17A-17D show Micromass co-culture samples after 14 days. Figure 17A: H&E stain; Figure 17B: Alcian blue; Figure 17C: Type I collagen (green); Figure 17D: Type II collagen (green) + Nucleic stain (red) .

Figures 18A and 18B show RT-PCR gel for day 7 micromass samples. Figure 18A: Type X collagen expression. Figure 18B: Type II collagen expression. (C: control micromass sample; E: experimental co-culture sample) .

Figures 19A and 19B show SEM image of cellular attachment to PLAGA-BG scaffold after 30 min; Figure 19A: chondrocyte control (2000X) ; Figure 19B: co-culture (1500X) .

Figures 20A-20C show Cellular attachment to PLAGA-BG scaffold; Figure 2OA: chondrocyte control, day 1 (500X) ; Figure 2OB: co-culture, day 1 (500X) . Figure 2OC: co- culture, day 7 (750X) .

Fig. 21-1 shows a table of porosimetry data, including intrusion volume, porosity, and pore diameter data, in another set of experiments.

Figs. 21-2A through 21-2C show fluorescence microscopy images (day 28, xlO) for Phases A through C, respectively.

Figs. 21-3A and 21-3B are images showing extracellular matrix production for Phases B and C, respectively.

Fig. 22-1 shows a schematic of the experimental design, in another set of experiments, for in vitro evaluations of human osteoblasts and fibroblasts co-cultured on multi-phased scaffolds.

Fig. 22-2 shows a graph which demonstrates cell proliferation in Phases A, B, and C during 35 days of human hamstring tendon fibroblast and osteoblast co- culture on multiphased scaffolds.

Figs. 22-3A and 22-3B graphically show Mechanical testing data for multiphased scaffolds seeded with human hamstring tendon fibroblasts and human osteoblasts over 35 days of culture (n=4) . Fig. 23-1 schematically shows a method for producing multi-phasic scaffolds, in another set of experiments. First, Ethicon PLAGA mesh is cut into small pieces and inserted into a mold. By applying compression force (F) and heating (H) at 150°C for time (t) = 20 hours, the mesh segments are sintered into a mesh scaffold, which is removed from the mold. Next, PLAGA microspheres are inserted into the mold, sintered, then removed as a second scaffold. The same process is performed for the PLAGA-BG microspheres. Finally, Phases A and B are joined by solvent evaporation, then all three scaffolds are inserted into the mold and sintered together, forming the final multi-phasic scaffold.

Fig. 23-2 shows a schematic of a co-culture experimental design.

Fig. 23-3 shows a table summarizing mercury porosimetry data.

Figs. 23-4A and 23-4B show graphically scaffold phase thicknesses and diameters, in the experiments of Fig. 23- 1 through Fig. 23-3.

Fig. 23-5 shows graphically a comparison of microsphere initial mass and final mass after undergoing a sintering process .

Figs. 23-6A and 23-6B show graphically mechanical testing data for multiphased scaffolds seeded with human hamstring tendon fibroblasts and human osteoblasts over 35 days of culture (n=4) . Scaffolds were tested in uniaxial compression. Compressive modulus (A) and yield strength (B) were calculated from the resulting stress- strain curves. Both cell seeded (C) and acellular (AC) scaffolds were examined at days 0, I₁ 21, and 35. Scaffold compressive modulus was significantly greater at day 0 than for all subsequent time points and groups (p<0.05) .

Fig. 24-1 shows a table illustrating the compositions of polymer solutions tested, in another set of experiments.

Fig. 24-2 shows a table illustrating drum rotational velocity (rpm) and surface velocity (m/s) for each gear.

Figs. 25-3A and 25-3D show SEMs of electrospun meshes spun at:

A) 1^st gear, 7.4 m/s;

B) 2^nd gear, 9.4 m/s;

C) 3^rd gear, 15 m/s; and

D) 4^th gear, 20 m/s.

Fig. 25-4A and 25-4B show scanning electron microscopy (SEM) images of another embodiment of multi-phased scaffold, with 85:15 PLAGA electrospun mesh joined with PLAGA: BG composite microspheres.

Fig. 26 schematically shows one exemplary embodiment of multi-phased scaffold as a hamstring tendon graft collar which can be implemented during ACL reconstruction surgery to assist with hamstring tendon-to-bone healing.

Fig. 27A shows an exemplary embodiment of a graft collar (A)_comprising a mesh, wherein the fibers of the mesh are aligned substantially parallel to a longitudinal axis of the tendon (B) . Figure 27B shows an exemplary embodiment of a graft collar (C) comprising a mesh, wherein the fibers of the mesh are aligned substantially perpendicular to a longitudinal axis of the tendon (D) .

Fig. 28. Characterization of Nanofiber Mesh Contraction. A) As-fabricated nanofiber mesh with preferential fiber alignment at low (left) and high (right) magnification as shown by scanning electron microscopy (low: x500, high: x2000) . B) Percent contraction of the aligned nanofiber mesh in the direction along (y-axis) and normal to (x- axis) fiber alignment (*p<0.05). Significant mesh contraction was the greatest along the direction of fiber alignment, and contraction stabilized after 24 hours.

Fig. 29. Compression of Graft Collar Scaffold with Nanofiber Mesh. A) Microsphere scaffold wrapped with nanofiber mesh before (top) and after (bottom) 24 hours of mesh contraction. B) Changes in scaffold inner diameter due to compression induced by the nanofiber mesh. While the scaffold-only control swelled (4%), nanofiber mesh contraction induced over 15% decrease in scaffold diameter after 24 hours.

Fig. 30. Compression of Tendon Graft with Nanofiber Mesh. A) Nanofiber mesh wrapped around a patellar tendon sample before (top, day 0) and after (bottom, day 1) mesh contraction. B) Effects of mesh contraction on tendon matrix organization. After five days of culture, the compressed tendon matrix exhibited greater cell density and is morphologically distinct from the unloaded control. After 14 days, however, no difference was observed between the groups. (H&E, xlO, arrows denote the direction of compressive loading applied by the mesh) .

Fig. 31. Compression of Tendon Graft with Graft Collar Scaffold and Nanofiber Mesh. A) Wrapping of the tendon graft with graft collar scaffold and mesh (top) and the tendon graft with mesh+scaffold complex after 24 hours

(bottom) . B) Effects of compression on tendon matrix organization. Within 24 hours of loading, the tendon matrix no longer exhibits the crimp pattern evident in the unloaded control. In addition, local cell density increased and there is evidence of matrix remodeling, and this organization is maintained after two weeks of static compression. (H&E, x20, arrows denote the direction of compressive loading applied by the scaffold) .

Fig. 32. Effects of Compression on Tendon Cellularity and Matrix Composition. A) Cells proliferated in the unloaded group and cell number was significantly higher in the control tendons compared to the compressed tendons after 24 hours of loading (p<0.05). B) Glycosaminoglycan content in the mesh was significantly higher in the compressed group after 24 hours of loading ( *p<0.05) .

Fig. 33. Effects of Compression on the Expression of Fibrocartilage-Related Markers. Scaffold-induced compression of the tendon graft resulted in significant up-regulation of type II collagen, aggrecan, and TGF-β3 after 24 hours (*p<0.05) .

Fig. 34. Effects of wrapping tendon with a PLGA electrospun mesh wherein fibers are either perpendicular or parallel to the longitudinal axis of the tendon. A) Tendons before mesh wrapping: Control Group, no mesh wrapping (left column) ; Tendons wrapped with mesh having fibers perpendicular to longitudinal axis of tendon

(center column) ; Tendons wrapped with mesh having fibers

) parallel to longitudinal axis of tendon (right column) .

B) Tendons 24 hours after mesh wrapping.

Detailed Description

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.

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, "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, "allogenic", in regards to a biopolymer mesh, shall mean a biopolymer mesh derived from a material originating from the same species as the subject receiving the biopolymer mesh.

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, "biomimetic" shall mean a resemblance of a synthesized material to a substance that occurs naturally in a human body and which is not rejected by (e.g., does not cause an adverse reaction in) the human body.

As used herein, "biopolymer mesh" shall mean any material derived from a biological source. Examples of a biopolymer mesh include, but are limited to, collagen, chitosan, silk and alginate.

As used herein, "BMP" shall mean bone morphogenetic protein.

As used herein, "BMSC" shall mean bone marrow-derived stem cells.

As used herein, "chondrocyte" shall mean a differentiated cell responsible for secretion of extracellular matrix of cartilage.

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, "GDF" shall mean growth differentiation factor.

As used herein, "matrix" shall mean a three-dimensional structure fabricated from biomaterials . The biomaterials can be biologically-derived or synthetic. 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. For example, a chondrocyte-embedded agarose hydrogel may be used in some instances. As another example, the hydrogel may be formed from hyaluronic acid, chitosan, alginate, collagen, glycosaminoglycan and polyethylene glycol

(degradable and non-degradable) , which can be modified to be light-sensitive. It should be appreciated, however, that other biomimetic hydrogels may be used instead.

As used herein, "lyophilized", in regards to a graft collar, shall mean a graft collar that has been rapidly frozen and dehydrated.

As used herein, "osteoblast" shall mean a bone-forming cell that is derived from mesenchymal osteoprognitor 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, "osteointegrative" shall mean ability to chemically bond to bone.

As used herein, "PDGF" shall mean platelet-derived growth factor.

As used herein, "photopolymerized" shall mean using light (e.g. visible or ultraviolet light) to convert a liquid monomer or macromer into a hydrogel by free radical polymerization. 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, "porosity" shall mean the ratio of the volume of interstices of a material to a volume of a mass of the material.

As used herein, "PTHrP" shall mean parathyroid hormone- related protein.

As used herein, "sinter" or "sintering" shall mean densification of a particulate polymer compact involving a removal of pores between particles (which may be accompanied by equivalent shrinkage) combined with coalescence and strong bonding between adjacent particles. The particles may include particles of varying size and composition, or a combination of sizes and compositions. For example, sintering a polymer would involve heating the polymer above the glass transition temperature, wherein the polymer chains rearrange and link together to form sintering necks.

As used herein, "TGF" shall mean transforming growth factor.

As used herein, "VEGF" shall mean vascular endothelial growth factor.

As used herein, "xenogenic", in regards to a biopolymer mesh, shall mean a biopolymer mesh derived from a raaterial originating from a species other than that of the subject receiving the biopolymer mesh.

The following exemplary embodiments and experimental details sections are set forth to aid in an understanding of the subject matter of this disclosure but are not intended to, and should not be construed to, limit in any way the invention as set forth in the claims which follow thereafter.

Apparatuses for promoting tendon-to-bone fixation, according to some embodiments of the techniques of this application, can include a graft collar for fixing tendon to bone in a subject. For example, the graft collar may be adapted for hamstring tendon-to-bone healing.

In one exemplary embodiment, a graft collar comprising a sheet of biopolymer mesh is provided for fixing tendon to bone in a subject.

The biopolymer mesh may comprise aligned fibers. Depending on the alignment of the fibers of the biopolymer mesh, the contraction of the fibers of the biopolymer mesh can be used to exert lateral tension or shear on the tendon (i.e., when fibers are aligned substantially parallel in relation to a longitudinal axis of the tendon) or vertical compression on the graft (i.e., when fibers are aligned substantially perpendicular in relation to a longitudinal axis of the tendon) .

Examples of biopolymer meshes include, but are not limited to, meshes derived from at least one of collagen, chitosan, silk and alginate. The biopolymer mesh can also be allogenic or xenogenic. For example, the graft collar may optionally be sutured around a tendon graft. The subject may be a mammal. In another embodiment, the mammal is a human. In a preferred embodiment, the graft collar promotes integration of the tendon graft to bone.

The graft collar may optionally include at least one of the following substances: anti-infectives, antibiotics, bisphosphonate, hormones, analgesics, anti-inflammatory agents, growth factors, angiogenic factors, chemotherapeutic agents, anti-rejection agents, and RGD peptides. In another embodiment, the growth factors are selected from the group consisting of TGFs, BMPs, IGFs, PTHrP, GDFs, VEGFs and PDGFs. In one embodiment, the TGF is TGF-β. In another embodiment, the TGF-β is TGF-β3. In another embodiment, the BMP is BMP-2. In another embodiment, the GDF is GDF-5 or GDF-7.

In one embodiment, the graft collar may include one or more of the following types of cells: chondrocytes, osteoblasts, osteoblast-like cells and stem cells. In another embodiment, the graft collar includes at least one of the following: osteogenic agents, osteogenic materials, osteoinductive agents, osteoinductive materials, osteoconductive agents, osteoconductive materials and chemical factors. In one embodiment, the graft collar promotes regeneration of an interfacial region between tendon and bone.

The graft collar may optionally be lyophilized. In another embodiment, the graft collar is biodegradable. In another embodiment, the graft collar is osteointegrative .

A graft collar for fixing tendon to bone in a subject, according to another exemplary embodiment, comprises a sheet of polymer-fiber mesh.

The polymer-fiber mesh preferably comprises aligned fibers. For example, the graft collar may optionally be sutured around a tendon graft. Depending on the alignment of the fibers of the polymer-fiber mesh, the contraction of the fibers of the polymer-fiber mesh can be used to exert lateral tension or shear on the graft

(i.e., when fibers are aligned substantially parallel in relation to a longitudinal axis of the tendon) or vertical compression on the graft (i.e., when fibers are aligned substantially perpendicular in relation to a longitudinal axis of the tendon) .

The subject may be a mammal. In another embodiment, the mammal is a human. In a preferred embodiment, the graft collar promotes integration of the tendon graft to bone.

The graft collar may optionally include at least one of the following substances: anti-infectives, antibiotics, bisphosphonate, hormones, analgesics, anti-inflammatory agents, growth factors, angiogenic factors, chemotherapeutic agents, anti-rejection agents, and RGD peptides. In another embodiment, the growth factors are selected from the group consisting of TGFs, BMPs, IGFs,

PTHrP, GDFs, VEGFs and PDGFs. In one embodiment, the TGF is TGF-β. In another embodiment, the TGF-β is TGF-β3. In another embodiment, the BMP is BMP-2. In another embodiment, the GDF is GDF-5 or GDF-7.

The graft collar may optionally include one or more of the following types of cells: chondrocytes, osteoblasts, osteoblast-like cells and stem cells. In another embodiment, the graft collar includes at least one of the following: osteogenic agents, osteogenic materials, osteoinductive agents, osteoinductive materials, osteoconductive agents, osteoconductive materials and chemical factors.

The polymer-fiber mesh can be selected from the group comprising 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, and biopolymers, and a blend of two or more of the preceding polymers. In another embodiment, the polymer-fiber mesh comprises at least one of the poly (lactide-co-glycolide) , poly (lactide) and poly (glycolide) .

In a preferred embodiment, the graft collar promotes regeneration of an interfacial region between tendon and bone.

In another embodiment, graft collar may optionally be lyophilized. In another embodiment, the graft collar is biodegradable. In another embodiment, the graft collar is osteointegrative . According to another exemplary embodiment, a graft collar for fixing tendon to bone in a subject comprises (a) a first region comprising a biopolymer mesh and hydrogel and (b) a second region adjoining the first region and comprising a biopolymer mesh.

The subject may be a mammal. In another embodiment, the mammal is a human.

The first region preferably supports the growth and maintenance of an interfacial zone between tendon and bone, and the second region supports the growth and maintenance of bone tissue.

The graft collar can include at least one of the following substances: anti-infectives, antibiotics, bisphosphonate, hormones, analgesics, anti-inflammatory agents, growth factors, angiogenic factors, chemotherapeutic agents, anti-rejections agents, and RGD peptides. In another embodiment, the hydrogel is photopolymerized, thermoset or chemically cross-linked. In one embodiment, the hydrogel is polyethylene glycol. In one embodiment, the biopolymer mesh comprises aligned fibers .

Depending on the alignment of the fibers of the biopolymer mesh, the contraction of the fibers of the biopolymer mesh can be used to exert lateral tension or shear on the graft (i.e., when fibers are aligned substantially parallel in relation to a longitudinal axis of the tendon) or vertical compression on the graft (i.e., when fibers are aligned substantially perpendicular in relation to a longitudinal axis of the tendon) .

The first region may optionally contain TGF. In one embodiment, the TGF is TGF-β. In another embodiment, the TGF-β is TGF-β3. In another embodiment, the first region may optionally contain PTHrp or GDF. In another embodiment, the GDF is GDF-5 or GDF-7. In another embodiment, the first region contains chondrocytes. In another embodiment, the chondrocytes are BMSC-derived. In another embodiment, the first region contains stem cells. In another embodiment, the stem cells are BMSCs.

Examples of biopolymer meshes include, but are not limited to, meshes is derived from at least one of collagen, chitosan, silk and alginate. In another embodiment, the biopolymer mesh is allogenic or xenogenic.

In another embodiment, the second region contains at least one of the following growth factors: BMP, IGF, PTHrP, GDF, VEGF and PDGF. In one embodiment, the BMP is BMP-2. In another embodiment, the GDF is GDF-5 or GDF-7. In another embodiment, the second region includes osteoblasts and/or osteoblast-like cells. In another embodiment, the osteoblasts and/or osteoblast like cells are BMSC-derived.

The second region can include at least one of the following: osteogenic agents, osteogenic materials, osteoinductive agents, osteoinductive materials, osteoconductive agents, osteoconductive materials and chemical factors. In another embodiment, the second region contains nanoparticles of calcium phosphate. In another embodiment, the calcium phosphate is selected from the group comprising tricalcium phosphate, hydroxyapatite and a combination thereof. In another embodiment, the second region contains nanoparticles of bioglass .

The graft collar is preferably biodegradable. In another embodiment, the graft collar is osteointegrative.

A graft collar for fixing tendon to bone in a subject, according to another embodiment, comprises (a) a first region comprising a polymer-fiber mesh and hydrogel and (b) a second region adjoining the first region and comprising a polymer-fiber mesh.

The subject can be a mammal. In another embodiment, the mammal is a human.

The graft collar can include at least one of the following substances: anti-infectives, antibiotics, bisphophonate, hormones, analgesics, anti-inflammatory agents, growth factors, angiogenic ^''Xfactors, chemotherapeutic agents, anti-rejections agents, and RGD peptides. In another embodiment, the hydrogel is photopolymerized, thermoset or chemically cross-linked. In one embodiment, the hydrogel is polyethylene glycol. In another embodiment, the polymer-fiber mesh comprises aligned fibers.

The first region may optionally contain TGF. In one embodiment, the TGF is TGF-β. In another embodiment, the

TGF-β is TGF-β3. In another embodiment, the first region may optionally contain PTHrp or GDF. In another embodiment, the GDF is GDF-5 or GDF-7. In another embodiment, the first region contains chondrocytes. In another embodiment, the chondrocytes are BMSC-derived.

In another embodiment, the first region contains stem cells. In one embodiment, the stem cells are BMSCs.

In one embodiment, the second region contains at least one of the following growth factors: BMP, IGF, PTHrP,

GDF, VEGF and PDGF. In one embodiment, the BMP is BMP-2.

In another embodiment, the GDF is GDF-5 or GDF-7. In another embodiment, the second region includes osteoblasts and/or osteoblast-like cells. In another embodiment, the osteoblasts and/or osteoblast like cells are BMSC-derived.

The second region can include at least one of the following: osteogenic agents, osteogenic materials, osteoinductive agents, osteoinductive materials, osteoconductive agents, osteoconductive materials and chemical factors. In another embodiment, the second region contains nanoparticles of calcium phosphate. In another embodiment, the calcium phosphate is selected from the group comprising tricalcium phosphate, hydroxyapatite and a combination thereof. In another embodiment, the second region contains nanoparticles of bioglass. The polymer-fiber mesh in the second region can be selected from the group comprising 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, and biopolymers, and a blend of two or more of the preceding polymers. In another embodiment, the polymer- fiber mesh comprises at least one of the poly (lactide-co- glycolide) , poly (lactide) and poly (glycolide) .

Depending on the alignment of the fibers of the polymer- fiber mesh, the contraction of the fibers of the polymer- fiber mesh can be used to exert lateral tension or shear on the graft (i.e., when fibers are aligned substantially parallel in relation to a longitudinal axis of the tendon) or vertical compression on the graft (i.e., when fibers are aligned substantially perpendicular in relation to a longitudinal axis of the tendon) .

The graft collar is preferably biodegradable. In another embodiment, the graft collar is osteointegrative .

The specific embodiments described herein are illustrative, and many variations can be introduced on these embodiments without departing from the spirit of the disclosure or from the scope of the appended claims. Elements and/or features of different illustrative embodiments may be combined with each other and/or substituted for each other within the scope of this disclosure and appended claims. This application claims the benefit of U.S. Provisional Application No. 60/873,518, filed December 6, 2006, the entire contents of which are incorporated herein by reference.

Further non-limiting details are described in the following Experimental Details section which is set forth to aid in an understanding of the invention but is not intended to, and should not be construed to, limit in any way the claims which follow thereafter.

Experimental Details

First Set of Experiments

To address the challenge of graft fixation to subchondral bone, a normal and functional interface may be engineered between the ligament and bone. This interface, according to one exemplary embodiment, was developed from the co- culture of osteoblasts and ligament fibroblasts on a multi-phased scaffold system with a gradient of structural and functional properties mimicking those of the native insertion zones to result in the formation of a fibrocartilage-like interfacial zone on the scaffold. Variations in mineral content from the ligament proper to the subchondral bone were examined to identify design parameters significant in the development of the multi- phased scaffold. Mineral content (Ca-P distribution, Ca/P ratio) across the tissue-bone interface was characterized. A multi-phased scaffold with a biomimetic compositional variation of Ca-P was developed and effects of osteoblast-ligament fibroblast co-culture on the development of interfacial zone specific markers (proteoglycan, types II & X collagen) on the scaffold were examined.

The insertion sites of bovine ACL to bone (see Figs. 3A- 3C) were examined by SEM. Pre-skinned bovine tibial- femoral joints were obtained. The intact ACL and attached insertion sites were excised with a scalpel and transferred to 60mm tissue culturing dishes filled with Dulbecco's Modified Eagle Medium (DMEM) (see Figures 4A and 4B) . After isolation, the samples were fixed in neutral formalin overnight, and imaged by environmental SEM (FEI Quanta Environmental SEM) at an incident energy of 15 keV. ACL attachment to the femur exhibited an abrupt insertion of the collagen bundle into the cartilage/subchondral bone matrix. Examination of collagen bundle revealed that the surface was ruffled and small collagen fibrils can be seen. When a cross section was imaged, three distinct zones at the insertion site were evident: ligament (L), fibrocartilage (FC), and subchondral bone (B) . The interface region spans proximally 200μm. These cross section views showed the insertion of Sharpey fiber into the fibrocartilage (see Figures 5A, 5B and 5C) . Mineralized fibrocartilage was not distinguishable with regular cartilage from these images .

The insertion sites of bovine ACL to bone were examined by scanning electron microscopy (SEM) . Bovine tibial- femoral joints were obtained. The intact ACL and attached insertion sites were excised with a scalpel and transferred to 60 mm tissue culturing dishes filled with Dulbecco' s Modified Eagle Medium (DMEM). After isolation, the samples were fixed in neutral formalin overnight, and imaged by environmental SEM (FEI Quanta Environmental SEM) at 15 keV.

ACL attachment to the femur exhibited an abrupt insertion of the collagen bundle into subchondral bone. When a cross section was imaged (see Figures 6A and 6B) , three distinct zones at the insertion site were evident: ligament (L), fibrocartilage (FC), and subchondral bone (B) . Sharpey fiber insertion into the fibrocartilage (see Figure 6A) was observed. The bovine interface region spans proximally 600 μm. Examination of the interface using energy dispersive X-ray analysis (EDAX, FEI Company) enable the mineralized and non-mineralized FC zones to be distinguished. A zonal difference in Ca and P content was measured between the ligament proper and the ACL-femoral insertion (see Table I) .

TABLE I

At the insertion zone (see Figures 7A and 7B), higher Ca and P peak intensities were observed, accompanied by an increase in Ca/P ratio as compared to the ligament region. Higher sulfur content due to the presence of sulfated proteoglycans at the FC region was also detected. The zonal difference in Ca-P content was correlated with changes in stiffness across the interface. Nanoindentation measurements were performed using atomic force microscopy (AFM, Digital Instruments). An increasing apparent modulus was measured as the indentation testing position moved from the ligament region into the transition zone (see Figure 8).

Ca-P distribution on polylactide-co-glycolide (50:50) and 45S5 bioactive glass composite disc (PLAGA-BG) after incubation in a simulated body fluid (SBF) was evaluated using μCT (μCT 20, Scanco Medical, Bassersdorf, Switzerland) following the methods of Lin et al. The sample was loaded into the system, scanned at 20 mm voxel resolution and an integration time of 120 ms . Figures 9A and 9B compare the amount of calcified region (dark areas) observed on the PLAGA-BG disc as a function of incubation time in SBF (from day 0 to day 28) . Using custom image analysis software, it was determined that at day 0, the mineralized region corresponded to 0.768% of the total disc (quartered) area, and at day 28, the mineralized region corresponded to 12.9% of the total area. Results demonstrate the Ca-P distribution on scaffolds measured by μCT analysis.

The scaffold system developed for the experiments was based on a 3-D composite scaffold of ceramic and biodegradable polymers. A composite system has been developed by combining poly-lactide-co-glycolide (PLAGA) 50:50 and bioactive glass (BG) to engineer a degradable, three-dimensional composite (PLAGA-BG) scaffold with improved mechanical properties. This composite was selected as the bony phase of the multi-phased scaffold as it has unique properties suitable as a bone graft.

A significant feature of the composite was that it was osteointegrative, i.e., able to bond to bone tissue. No such calcium phosphate layer was detected on PLAGA alone, and currently, osteointegration was deemed a significant factor in facilitating the chemical fixation of a biomaterial to bone tissue. A second feature of the scaffold was that the addition of bioactive glass granules to the PLAGA matrix results in a structure with a higher compressive modulus than PLAGA alone. The compressive properties of the composite approach those of trabecular bone. In addition to being bioactive, the PLAGA-BG lends greater functionality in vivo compared to the PLAGA matrix alone. Moreover, the combination of the two phases serves to neutralize both the acidic byproducts produced during polymer degradation and the alkalinity due to the formation of the calcium phosphate layer. The composite supports the growth and differentiation of human osteoblast-like cells in vitro.

The polymer-bioactive glass composite developed for the experiments was a novel, three-dimensional, polymer- bioactive biodegradable and osteointegrative glass composite scaffold. The morphology, porosity and mechanical properties of the PLAGA-BG construct have been characterized. BG particle reinforcement of the PLAGA structure resulted in an approximately two-fold increase in compressive modulus (p < 0.05). PLAGA-BG scaffold formed a surface Ca-P layer when immersed in an electrolyte solution (see Figure 10A), and a surface Ca-P layer was formed. No such layer was detected on PLAGA controls. EDXA spectra confirmed the presence of Ca and P (see Figure 10B) on the surface. The Ca, P peaks were not evident in the spectra of PLAGA controls.

In vitro formation of a surface Ca-P layer indicates PLAGA-BG composite's osteointegrative potential in vivo. The growth and differentiation of human osteoblast-like cells on the PLAGA-BG scaffolds were also examined. The composite promoted osteoblast-like morphology and stained positive for alkaline phosphatase, and promoted synthesis to a greater extent of Type I collagen synthesis than tissue culture polystyrene controls. The porous, interconnected network of the scaffold was maintained after 3 weeks of culture (see Figure 11). Mercury porosimetry (Micromeritics Autopore III, Micromeritics, Norcross, GA) was used to quantify the porosity, average pore diameter and total surface area of the composite construct. The construct porosity was determined by measuring the volume of mercury infused into the structure during analysis. In addition, the construct (n = 6) was tested under compression. BG particle reinforcement of the PLAGA structure resulted in approximately two-fold increase in compressive modulus (see Table II, p < 0.05).

TABLE II

Porosity, pore diameter, and mechanical properties of the scaffold may be variable as a function of microsphere diameter and BG content. The growth and differentiation of human osteoblast-like cells on the PLAGA-BG scaffolds were also examined. The composite supported osteoblast- like morphology and stained positive for alkaline phosphatase .

The porous, interconnected network of the scaffold was maintained after 3 weeks of culture (see Figure 11) . The synthesis of type I collagen was found to be the highest on the composite, as compared to the PLAGA and tissue culture polystyrene (TCPS) controls (n = 3, p < 0.05) (see Figure 12) .

The effects of bovine osteoblast and fibroblast co- culture on their individual phenotypes were examined. The cells were isolated using primary explant culture. The co-culture was established by first dividing the surfaces of each well in a multi-well plate into three parallel sections using sterile agarose inserts. ACL cells and osteoblasts were seeded on the left and right surfaces respectively, with the middle section left empty. Cells were seeded at 50,000 cells/section and left to attach for 30 minutes prior to rinsing with PBS. The agarose inserts were removed at day 7, and cell migration into the interface was monitored. Control groups were fibroblasts alone and osteoblasts alone.

In time, both ACL fibroblasts and osteoblasts proliferated and expanded beyond the initial seeding areas. These cells continued to grow into the interfacial zone, and a contiguous, confluent culture was observed. All three cultures expressed type I collagen over time. The co-culture group expressed type II collagen at day 14, while the control fibroblast did not.

Type X collagen was not expressed in these cultures, likely due to the low concentration of b-GP used.

Alizarin Red S stain intensity was the highest for the osteoblast control, (see Figure 13C) followed by the co- cultured group (see Figure 13B). Positive ALP staining was also observed for osteoblast control and co-culture groups (see Figures 14C and 14B, respectively) . Scaffold of four continuous, graded layers with different sizes of microspheres was formulated (see Figures 15A- 15F) . Layered inhomogeneity was pre-designed into the scaffold. Due to differences in packing efficiency between different sizes of microspheres, the porosity of the scaffold decreases from layers of large microsphere to those consisting of small microspheres. PLAGA-BG composite microspheres were produced via the emulsion method. Three layers of PLAGA-BG microspheres of different diameters (250-300, 300-355, 355-500 μm, from top to bottom) were used, shown in Figures 15A-15F. Microsphere layers were sintered at 70⁰C for 20 hours.

Image analysis confirmed that pore size increased from bottom to top of scaffold. For the growth of ACL fibroblasts on the scaffold, another type of multi-phased scaffold was fabricated using a PLAGA mesh (Ethicon, NJ) and two layers of PLAGA-BG microspheres. The layers were sintered in three stages in a Teflon mold. First the mesh was cut into small pieces and sintered in the mold for more than 20 hours at 55⁰C. A layer of PLAGA-BG microspheres with diameter of 425-500 μm was then added to the mold. This layer was sintered for more than 20 hours at 75°C. The final layer consisted of PLAGA-BG microspheres with diameter greater than 300 μm. The scaffolds and three distinct regions were readily observed (see Figures 16A-16C) .

Kinetics of Ca-P layer formation on BG surfaces was related to changes in surface zeta potential in a simulated body fluid (SBF) . The chemical and structural changes in BG surface Ca-P layer were characterized using Fourier transform infrared spectroscopy (FTIR) , SEM and energy dispersive x-ray analysis (EDXA) . FTIR provides information on the degree of crystallinity (amorphous vs. crystalline) of the Ca-P layer formed (see Figure 4) as well as the functional groups present on BG surface (carbonated Ca-P layer versus non-carbonated, protein adsorption, etc.). FTIR is much more surface sensitive than X-ray diffraction in detecting the Ca-P crystalline structures when the surface layer is only several microns in thickness. SEM combined with EDXA is a powerful tool in relating elemental composition to specific surface morphology and distributions (see Figures 5B and 5C) . EDXA provides a direct calculation of Ca/P ratio (Ca/P = 1.67 for bone mineral and crystalline Ca-P layer) when appropriate standards are used. FTIR, SEM, and EDXA are complimentary techniques which together provide quantitative data on the crystallinity, composition of and functional groups pertaining to the Ca-P layer.

Evaluation of the effects of co-culturing on the growth and phenotypic expression of osteoblasts and chondrocytes. Osteoblasts were seeded directly on high density chondrocyte micromasses. Specific effects of co- culture on the expression of chondrogenic markers were observed primarily at the top surface interaction zone instead of within the micromass. Alcian blue staining (see Figure 17B) revealed characteristic peri-cellular sulfated GAG deposition by chondrocytes. GAG deposition was found largely within the micromass, instead of at the co-culture zone where elongated osteoblasts and chondrocytes were located. Sulfated GAG was not detected in the predominantly osteoblast monolayer surrounding the micromass. Surface chondrocytes may have dedifferentiated due to co-culturing with osteoblasts. The expression of type I collagen was observed to be distributed mainly on the top surface of the co-cultured mass (Figure 17C), where osteoblasts were located. Type I was also found at the primarily osteoblastic monolayer surrounding the micromass (see Figure 17C, left) . No type I collagen expression was observed in the chondrocyte-dominated center and bottom surface of the micromass. High expression of type II collagen was observed within the micromass (see Figure 17D) .

As types I and II collagen were detected at the surface, it is possible that due to co-culture, chondrocytes and osteoblasts were forming an osteochondral-like interface at the surface interaction zone. Alizarin Red (ALZ) staining revealed that there was limited mineralization in the co-cultured group, while the osteoblast control stained increasingly positive for calcium. It is likely that co-culture with chondrocytes may have delayed osteoblast mineralization. Preliminary PCR results (see Figures 18A and 18B) showed that the 7 day co-culture group expressed types II and X collagen, as detected by RT-PCR.

Effects of media additives on the growth and mineralization of osteoblasts and human ACL fibroblasts

(hACL) were examined. During mineralization, ALP reacted with β-glycerophosphate (βGP) and the phosphate product was utilized for mineralization. Concentrations (0, 1.0,

3.0, 5.0 mM) effects were examined over time. No significant change in cell number was observed for the

[βGP] investigated. At 1.0 mM, a significant difference between 1-day & 7-day samples (p < 0.05) was observed.

No differences were found between 1.0 mM and 3.0 mM cultures. ALZ stains for the osteoblast cultures were more intense for 3.0 mM than for 1.0 mM. Ectopic mineralization was observed for hACL cultures at 3.0 mM suggesting a potential change in cell phenotype. Interaction of osteoblasts and chondrocytes on a 3-D composite scaffold during co-culture was examined. Scaffolds seeded with only osteoblasts or chondrocytes at the same densities served as controls. Both short-term and long-term co-culture experiments were conducted. Extensive SEM analysis revealed that significant interactions occurred between osteoblasts and chondrocytes during co-culture. Differences in cellular attachment were observed between the chondrocyte control scaffolds and the co-cultured scaffolds. On the co- cultured scaffolds, focal adhesions were evident between the spherical chondrocytes and the surface, indicated by the arrow in Figure 19B.

No comparable focal adhesions were observed on the chondrocyte controls at the same time point. Chondrocyte morphology changed over time as it assumed a spherical morphology in the first 8 hours, and then spread on the surface of the microspheres (see Figure 20A) . The nodules on the surface of the microspheres correspond to the flattened chondrocytes. These nodules were likely chondrocytes instead of calcium phosphate nodules, since calcium phosphate nodules were approximately 1-5 μm in diameter at the culture duration observed and these nodules were ~10 μm, approximately the diameter of an ovoid cell. After 7 days of culture, the co-culture group exhibited extensive matrix production (see Figure

20C) and expansion on the scaffold. Examination of the ACL-bone interface confirmed existence of a mineral gradient across the insertion zone and correlation to changes in material properties. Multi- phased scaffolds with controlled morphology and porosity were fabricated. The osteochondral graft developed from co-culture on PLAGA-BG and hydrogel scaffold supported growth of multiple matrix zones with varied GAG and mineral content. BMSCs differentiated into ligament fibroblast and produced a functional extracellular matrix when cultured with growth factors on a fiber-based scaffold. Mineral content, distribution, and chemistry at the interface and on the scaffold were quantifiable using a complimentary set of surface analysis techniques (FTIR, SEM, EDAX, μCT) . Electron microscopy examination of the ACL-bone interface revealed insertion zone including three different regions: ligament, fibrocartilage-like zone, and bone. Co-culture of osteoblasts and ligament fibroblasts on 2-D and 3-D scaffolds resulted in changes in cell morphology and phenotype. Type X collagen, an interfacial zone marker, was expressed during co-culture. Multi-phased scaffold with layered morphology and inhomogenous properties were designed and fabricated. FTIR, SEM and EDXA are complimentary techniques which collectively provided qualitative and quantitative information on the Ca-P layer and composition of the calcium phosphate surface.

These experiments illustrate, in relevant part, the interaction between osteoblasts and chondrocytes on a scaffold apparatus during co-culture.

Second Set of Experiments The degree of graft integration is a significant factor governing clinical success and it is believed that interface regeneration significantly improves the long term outcome. The approach of this set of experiments was to regenerate the ACL-bone interface through biomimetic scaffold design and the co-culture of osteoblasts and fibroblasts. The interface exhibits varying cellular, chemical, and mechanical properties across the tissue zones, which can be explored as scaffold design parameters. This study describes the design and testing of a multi-phased, continuous scaffold with controlled heterogeneity for the formation of multiple tissues. The continuous scaffold consists of three phases: Phase A for soft tissue, Phase C for bone, and Phase B for interface development. Each phase was designed with optimal composition and geometry suitable for the tissue type to be regenerated. Fibroblasts were seeded on Phase A and osteoblasts were seeded on Phase C, and the interactions of osteoblasts and fibroblasts (ACL and hamstring tendon) during co-cultures on the scaffolds were examined in vitro.

Phases A, B and C consist of poly (lactide-co-glycolide) (PLAGA, 10: 90) woven mesh, PLAGA (85:15) microspheres, and PLAGA (85: 15) /Bioactive Glass (45S5,BG) composite microspheres, respectively. The microspheres were formed via a double emulsion method, and the continuous multi- phased scaffolds were formed by sintering above the polymer T_g. Scaffold porosity and pore diameter were determined by porosimetry (Micromeritics, n=3) and the samples were tested under uniaxial compression (MTS 810, n=5) at 1.3 mm/min up to 5% strain with 10 N preload. Bovine and human osteoblasts (bOB and hOB) , and bovine ACL fibroblasts (bFB) and human hamstring tendon fibroblasts (hFB) were obtained through explant culture. In experiment I, bOB and bFB (5xlO⁵ cells each/scaffold) were co-cultured on the scaffold, and cell viability, attachment, migration and growth were evaluated by electron and fluorescence microscopy. The bOB were pre- labeled with CM-DiI, and both cell types were labeled with calcein AM (Molecular Probes) prior to imaging. Matrix production and mineralization were determined by histology. After ascertaining cell viability on the scaffolds, a more extensive experiment using hOB and hFB was conducted in which cell proliferation and differentiation and above analyses were investigated. The mechanical properties of the seeded scaffolds were also measured as a function of culture time.

Compression testing of scaffolds indicated an average modulus of 120120 MPa and yield strength of 2.3 MPa. The intrusion volume, porosity and pore diameter data are summarized in the table shown in FIG. 21-1.

The fibroblasts and osteoblasts were localized primarily at the two ends of the scaffolds after initial seeding, with few cells found in Phase B. After 28 days, both cell types migrated into Phase B (Fig. 21-2B), and extensive cell growth was observed in Phases A and C (Figs. 21-2A and 21-2C) .

Extensive collagen-rich matrix production was found throughout the three phases at day 28 (Figs. 21-3A and 21-3B) . The biomimetic, multi-phased scaffolds supported the growth and ECM production of both osteoblasts and fibroblasts. After 28 days of culture, collagen production was evident in all three phases and mineralized matrix was found in the bone and interface regions. Osteoblast and fibroblast interaction at the interface (Phase B) suggests that these cells may play a significant role in the development of a functional insertion site.

These findings demonstrate, in relevant part, that this novel scaffold is capable of simultaneously supporting the growth of multiple cell types and can be used as a model system to regenerate the soft tissue to bone interface.

Third Set of Experiments

This set of experiments was directed to in vitro evaluations of human osteoblasts and fibroblasts co- cultured on multi-phased scaffolds. A schematic of the experimental design for the in vitro study is shown in Fig. 22-1. Phase A (mesh) was seeded with human hamstring tendon fibroblast cell suspension. Phase C was seeded with osteoblasts. Cell interaction in the interfacial Phase B was monitored over time. Acellular scaffolds served as controls.

Cell proliferation in Phases A, B, and C during 35 days of human hamstring tendon fibroblast and osteoblast co- culture on multiphased scaffolds is shown in Fig. 22-2. A general trend of increasing cell number was observed in each phase over time. Data demonstrates that all three phases of the scaffold support cellular viability and proliferation. A higher number of cells were seeded on phase A due to its inherently larger surface area compared to phase C.

Mechanical testing data for multiphased scaffolds seeded with human hamstring tendon fibroblasts and human osteoblasts over 35 days of culture (n=4) is graphically shown in Figs. 22-3A and 22-3B. Scaffolds were tested in uniaxial compression. Compressive modulus (Fig. 22-3A) and yield strength (Fig. 22-3B) were calculated from the resulting stress-strain curves. Both cell seeded (C) and acellular (AC) scaffolds were examined at days 0, 7, 21, and 35.

Compared to the acellular controls, the cell seeded scaffolds degraded slower and better maintained their structural integrity over time. The yield strength of the acellular scaffold decreased over 35 days, while the seeded scaffolds maintained its yield strength.

These experiments, in relevant part, illustrate the interaction between osteoblasts and human hamstring tendon fibroblasts on a multi-phase scaffold.

Fourth Set of Experiments The scaffold designed for this study consisted of three phases and were fabricated in four stages (Figure 23-1). First, Phase A was formed from polyglactin 10:90 PLGA mesh sheets (Vicryl VKML, Ethicon) . Mesh sheets were cut into small segments (approximately 5 mm x 5 mm) and inserted into cylindrical molds (7.44 mm diameter). Molds were heated to 150°C for 20 hours to sinter the segments together to form a cylindrical mesh scaffold. The next phase (Phase B) consisted of 100% 85:15- poly (DL-lactide- co-glycolide) (PLAGA, Alkermes Medisorb, M_w « 123.6 kDa) microspheres formed by a water/oil/water emulsion. Briefly, Ig PLAGA was dissolved in 10 mL methylene chloride (EM Science, Gibbstown, New Jersey) and poured into a mixing 1% PVA surfactant solution (Sigma Chemicals, St. Louis, MO) . Microspheres were mixed for 4 hours, recovered by filtration, allowed to dry in a fume hood overnight, then vacuum desiccated for 24 hours. To form the PLAGA microsphere phase, -0.075 g microspheres were inserted into the same molds as used previously, and sintered at 55°C for 5 hours. The last phase (Phase C) consisted of composite microspheres formed from an 80:20 ratio of PLAGA and 45S5 bioactive glass (BG, MO-SCI Corporation, Rolla, MD) . Again, microspheres were formed by emulsion, except with 0.25 g bioactive glass suspended in a solution of 1 g PLAGA in 10 mL methylene chloride. Microspheres (28-30 mg/scaffold) were sintered in the same molds at 55⁰C for five hours. After all three phases were sintered separately, Phases A and B were joined by methylene chloride solvent evaporation, and then sintered to Phase C for 10 hours at 55°C in the same molds. Subsequently, scaffolds were sterilized with ethylene oxide. Final scaffold dimensions are detailed in Figs. 23-4A and 23-4B.

Human osteoblast-like cells and hamstring tendon fibroblasts were obtained from explant culture of tissue isolated from humerus trabecular bone and hamstring tendon respectively. Trabecular bone was rinsed with PBS, then cultured in Dulbecco' s Modified Eagle's Medium (DMEM, Mediatech, Herndon, VA, USA) supplemented with 10% fetal bovine serum, 1% non essential amino acids, and 1% penicillin/streptomycin (Mediatech, Herndon, Virginia) , and incubated at 37°C in a 5% CO2 incubator to allow for cell migration. Hamstring tendon obtained from excess tissue utilized for hamstring tendon ACL reconstruction autografts was minced and cultured in similarly supplemented DMEM. The first migrations of cells were discarded to obtain a more uniform cell distribution. Second migration, passage 2 osteoblast-like cells and second and third migration, passage 5 hamstring tendon fibroblasts were utilized for the co-culture experiment.

Scaffold dimensions were measured prior to cell seeding and before and after EtO sterilization. Phase thickness was calculated by image analysis, while phase diameter was determined using a digital caliper. Scaffold porosity and pore diameter (Phases A and B: n = 3; Phase C: n = 1) were determined by mercury porosimetry (Micromeritics Autopore III and Autopore IV 9500, Micromeritics, Norcross, GA) . The porosity data were utilized to determine cell seeding densities and cell suspension volumes for Phases A and C, with the volumes calculated such that fibroblasts suspension remains in Phase A and osteoblasts suspension in Phase C.

Hamstring tendon fibroblasts were seeded at a density of 250,000 cells/scaffold in a volume of 40.7 μL/scaffold on Phase A (Fig. 23-2) . After allowing the fibroblasts to attach to the scaffolds for 20 minutes, the scaffolds were rotated upside down so that Phase C faced upwards. Subsequently, 75,000 osteoblast-like cells were seeded per scaffold in a volume of 12.5 μL. After allowing the osteoblasts to attach to the scaffold for 20 minutes, the scaffolds were covered with DMEM supplemented with 10% FBS, 1% NEAA, and 1% penicillin/streptomycin, and incubated at 37⁰C and 5% CO₂. Ascorbic acid at a concentration of 20 μg/mL was added beginning at day 7. Media was exchanged every two days. Scaffolds were cultured in 6-well plates and covered with 7 mL of supplemented media per scaffold to minimize pH fluctuations due to rapid poly (glycolic acid) degradation.

Cell attachment, migration, and proliferation on the multi-phased scaffolds were examined using SEM (5kV, JEOL

5600LV) at days 7, 21, and 35. The scaffolds were fixed with Karnovsky's glutaraldehyde fixative, and stored at

4⁰C for 24 hours. The samples were then rinsed with

Hank's buffered salt solution two times, and serially dehydrated with ethanol . Cross-sections of the scaffold phases were mounted on an aluminum post and gold-coated prior to analysis.

Extracellular matrix production and mineralization were determined via histology at day 35. Scaffolds were rinsed two times with room temperature PBS. The scaffolds were then covered with 10% neutral buffered formalin and stored at 4 degrees C. Samples were plastic embedded using a modification of a procedure developed by Erben. The scaffolds were first suspended in 2% agarose (low gelling temperature, cell culture grade, Sigma, St.

Louis, Missouri) , then serially dehydrated with ethanol and cleared with xylene substitute (Surgipath, Sub-X,

Richmond, Illinois). Following dehydration, samples were embedded in poly (methyl methacrylate) (Polysciences,

Inc., Warrington, Pennsylvania) and sectioned into 10 μm slices. The scaffold sections were stained with either hematoxylin and eosin, von Kossa or Picrosirius Red stains and imaged with light microscopy.

At days 1, 7, 21, and 35, scaffolds were rinsed twice with PBS and subseguently the three phases were separated. Each phase was then stored in 0.1% Triton-X at

-80⁰C. Cellular proliferation in each phase was determined by means of PicoGreen DNA quantitation assay.

In addition, cellular phenotype for mineralization was evaluated using a quantitative alkaline phosphatase (ALP) assay.

At days 0, 7, 21, and 35, seeded and acellular scaffolds were tested under uniaxial compression (MTS 810, n=4) . The crosshead speed was 1.3 mm/min, and the scaffolds were compressed up to 35-40% strain. A 10 N preload was applied prior to testing. The effects of scaffold degradation and extracellular matrix production on scaffold compressive modulus were examined.

Mercury porosimetry data for each phase are summarized in the table shown in Fig. 23-3. Scaffold dimensions are shown in Figs. 23-4A and 23-4B. The thickness of Phase C decreased significantly (p<0.05) due to contraction during the EtO sterilization (Figure 23-4A). In addition, the thicknesses of all phases were significantly different from each other after sterilization. Scaffold diameters also varied due to contraction during sintering, in the case of Phase A, and contraction of Phase C during sterilization. The diameters of Phases B and C decreased significantly after sterilization, and the diameters of all phases were significantly different from each other after sterilization (p<0.05). During the scaffold fabrication process, microspheres are lost between weighing and filling the molds. This loss is mainly due to static charge accumulation in one or more of the microspheres, weighing paper, or mold, which prevents a small percentage of the microspheres from entering the molds. PLAGA-BG microspheres for Phase C generally experience a 2.1 ± 1.4 % loss in mass, while the PLAGA microspheres for Phase B suffer a loss of 4.0 ± 1.8 % (Fig. 23-5). Composite microspheres are generally more statically charged than the PLAGA microspheres; however, the stainless steel mold, used more often for the composite microspheres, dissipates charge buildup more readily than the PTFE mold, which is used more often for the PLAGA microspheres, possibly explaining why there is a significant loss for Phase B (p<0.05). Mesh for Phase A is not susceptible to this loss.

Compressive modulus and yield strength were obtained for seeded and acellular control scaffolds at days 0, 7, 21, and 35 of culture. A rapid decrease in compressive modulus was observed following day 0, possibly due to rapid initial polymer degradation. By day 35, the seeded scaffolds exhibited a greater compressive modulus (Fig. 23-6A) and yield strength (Fig. 23-6B), possibly due to cellular extracellular matrix and mineralization compensating loss of scaffold strength due to polymer degradation.

In this experiment, the cell types were switched from bovine ACL fibroblasts and trabecular bone osteoblast- like cells to human hamstring tendon fibroblasts and trabecular bone osteoblasts due to the increased clinical relevance of these new cell types. This experiment aimed to acquire quantitative data about cell proliferation and migration throughout the three phases, as well as cellular alkaline phosphatase activity in each phase of the scaffold.

Based on the previous experiment performed with bovine cells, it is apparent that the biomimetic, multi-phased scaffolds support the growth and ECM production of both osteoblasts and fibroblasts. After 28 days of culture, collagen production was evident in all three phases and mineralized matrix was found in the bone and interface regions. Osteoblast and fibroblast interaction at the interface (Phase B) suggests that these cells may play a significant role in the development of a functional insertion site.

These findings demonstrate that this novel scaffold is capable of simultaneously supporting the growth of multiple cell types and can be used as a model system to regenerate the soft tissue to bone interface.

Fifth Set of Experiments

The objective of the set of experiments was to incorporate electrospun PLAGA meshes into the multi- phased scaffold design, substituting the Ethicon mesh phase, and allowing the entire scaffold to be made in- house .

Electrospinning, short for electrostatic spinning, is a relatively new term that describes a principle first th discovered in the first half of the 20 century (see, for example, U.S. Patents Nos. 1,975,504, 2,160,962, 2,187,306, 2,323,025 and 2,349,950 to Formhals, the entire contents of which are incorporated herein by reference) . Electrostatic spinning involves the fabrication of fibers by applying a high electric potential to a polymer solution. The material to be electrospun, or dissolved into a solution in the case of polymers, is loaded into a syringe or spoon, and a high potential is applied between the solution and a grounded substrate. As the potential is increased, the electrostatic force applied to the polymer solution overcomes surface tension, distorting the solution droplet into a Taylor cone from which a jet of solution is ejected toward the grounded plate. The jet splays into randomly oriented fibers, assuming that the solution has a high cohesive strength, linked to polymer chain molecular weight, to prevent droplets from forming instead of fibers in a process known as electrospraying. These fibers have diameters ranging from nanometer scale to greater than 1 μm and are deposited onto the grounded substrate or onto objects inserted into the electric field forming a non-woven mesh. Mesh characteristics can be customized by altering electrospinning parameters. For example, fiber diameter and morphology can be altered, including the formation of beads along the fibers, by controlling applied voltage and polymer solution surface tension and viscosity. Also, fiber orientation can be controlled by rotating the grounded substrate . This high degree of customizability and ability to use many different materials, such as biodegradable polymers and silks, grant this fabrication method a high potential in the development of materials for biomedical application. Management of fiber diameter allows surface area to be controlled, and polymers with different degradation rates can be combined in various ratios to control fiber degradation, both of which are significant in drug delivery applications. Also, controlling the orientation of fiber deposition grants a degree of control over cell attachment and migration. Moreover, the ability to electrospin fiber meshes onto non-metal objects placed in the electric field enables the fabrication of multiphasic scaffold systems.

Here, in order to obtain precise parameters for the mesh fibers, including fiber diameter, morphology, and alignment, the effects of processing parameters on fiber characteristics were studied. A variable-speed rotating drum was designed and constructed to serve as a substrate for aligned fibers, and rheological experiments were performed on the polymer solutions to determine the effect of polymer concentration on solution viscosity and the subsequent effect of solution viscosity on fiber diameter and morphology.

In addition to determining the speed of each gear, the effect of each speed on fiber alignment was determined qualitatively. A 30% v/v PLAGA solution was prepared with 60% dimethylformamide and 10% ethanol, and this solution was electrospun onto the rotating drum at each of the four speed settings. The resulting meshes were examined by scanning electron microscopy (JEOL 5600LV) .

The relationship between polymer concentration (Alkermes 85:15 PLAGA) and solution viscosity was determine by means of a rheological study. Three concentrations of polymer were tested - 20%, 30%, and 40% v/v - in dimethylformamide (DMF) and ethanol. The composition of each solution is listed in the table shown in Fig. 24-1. Solutions were analyzed using an Advanced Rheometer AR 200Ot. There was variability in the viscosity measurements (n=l) at different strain rates due to the evaporation of solvent during testing. The geometry used for the viscosity measurements was a 25 mm stainless steel disc. A solvent trap was not used since it is not designed to fit with this geometry and a prior trial using the solvent trap with another geometry resulted in poor results, possibly because water from the solvent trap seal interacted with the polymer solution. Additional trials can use a solvent trap to obtain consistent and reliable values for viscosity. For the present study, averages were taken of the viscosity measurements taken at strain rates tested after the equipment had equilibrated. As a result, there are standard deviations for the viscosity measurements even with an n of 1.

The surface velocity of the rotating drum was seen to increase with increased pulley positions from gear 1 to gear 4 (see the table shown in Fig. 24-2) . The degree of fiber alignment increased with increasing drum velocity, as seen in the SEMs of each mesh (see Figs. 25-3A through 25-3D) .

It was found that (as expected) the degree of fiber orientation increased with increasing drum rotational velocity. The image was analyzed and a histogram of fiber angles was generated against the horizontal axis of the image at regular interval across the image. Thus, the degree of alignment of the fibers can be quantified. It is desirable to control the degree of fiber alignment in the electrospun meshes so that the extracellular environment found at the interface can be mimicked. By producing biomimetic scaffolds, it was intended to direct cell growth to reproduce the tissue inhomogeneity found at the native ACL insertions. In addition to controlling the fiber alignment, it is desirable to control fiber diameter and morphology. It was previously determined that substituting 10% of the DMF in the polymer solutions with ethanol reduces the surface tension of the solution and results in a significant reduction in the number of beads formed along the fibers when electrospinning PLAGA.

This effect was also observed by Fong et al., who reduced the number of beads in electrospun poly (ethylene oxide)

(PEO) meshes by the addition of ethanol. Surface tension of the polymer solution acts to form spheres during the electrospinning process. By reducing the solution surface tension, the formation of spheres is less favorable and straighter fibers result. Fong et al. also determined that the addition of ethanol increased the viscosity of the PE0:water solutions, which also favors the formation of straight fibers, and results in increased fiber diameter. Deitzel et al. also have demonstrated a relationship between PEO: water solution viscosity and fiber diameter, with fiber diameter increasing with increasing viscosity according to a power law. A relationship between solution viscosity and concentration of polymer can be determined in order to understand how PLAGA: N,W-DMF viscosity affects fiber diameter and morphology. The effect of solution viscosity on fiber diameter and morphology can be determined by spinning the various solutions and examining the resulting meshes by SEM. Other variables can affect the fiber parameters. By changing the percentage of polymer, the surface tensions of the polymer solutions also change in addition to the viscosity. Therefore, in addition to testing the viscosities of each solution, the surface tension of each solution are measured. It is desirable to keep all variables constant except for viscosity in order to truly determine the effect of solution viscosity on fiber characteristics. However, the interrelation of many of the electrospinning parameters complicates the process.

A PLAGA mesh was electrospun directly onto a microsphere scaffold. This is one way to incorporate the mesh. In addition, the scaffolds can be secured to the drum and aligned fibers electrospin directly onto the scaffolds. However, because of the high rotational velocities, it is difficult to secure the scaffolds and prevent them from flying off the drum when it begins rotating. Alternatively, aligned fiber meshes can simply be spun separately, and then later sintered to the microsphere scaffolds. For example, aligned fiber meshes can be electrospun onto aluminum foil, then wrapped around a rod with multiple mesh sheets sintered together to obtain a hollow cylinder of aligned fibers.

These experiments illustrate one possible method, i.e. electrospinning, for the production of aligned fibers for use in PLAGA scaffold apparatuses.

Fig. 25-4A and 25-4B show scanning electron microscopy (SEM) images of another embodiment of multi-phased scaffold, with 85:15 PLAGA electrospun mesh joined with PLAGA: BG composite microspheres. References for Background and First Through Fifth Sets of Experiments

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Biological fixation of soft tissue-based grafts utilized for Anterior Cruciate Ligament (ACL) reconstruction poses a significant clinical challenge. The ACL integrates with subchondral bone through a fibrocartilage interface, which serves to minimize stress concentrations and facilitate load transfer between two distinct types of tissue. Functional integration thus requires the re- establishment of this fibrocartilage region on the reconstructed grafts. To this end, this study focuses on the design and evaluation of a mechano-active scaffold system based on a composite of poly-α-hydroxyester nanofiber mesh and sintered microspheres. Specifically, the effects of scaffold-induced compression on tendon matrix remodeling and the development of fibrocartilage- related markers are evaluated over a two-week period. Scaffold contraction resulted in over 15% compression of the patellar tendon graft and up-regulated the expression of fibrocartilage-related markers such as type II collagen, aggrecan and transforming growth factor-β3. In addition, proteoglycan content was significantly higher in the compressed tendon group after one day of loading. This is the first reported study describing scaffold- mediated mechanical loading, and the findings of this study demonstrate the potential of the mechano-active scaffold to promote the formation of an anatomic fibrocartilage transition on tendon-based ACL reconstruction grafts, which is critical for achieving biological fixation and extending graft functionality. Materials and Methods

Tendon Graft Isolation

Patellar tendon grafts were isolated from neonatal bovine tibiofemoral joints (1-7 days old) obtained from a local abattoir (Green Village Packing, Green Village, NJ) . Briefly, the joints were first cleaned in an antimicrobial bath. Under antiseptic conditions, midline longitudinal incisions were made through the subcutaneous fascia to expose the patellar tendon. The paratenon was removed, and the patellar tendon dissected from the underlying fat pad. Sharp incisions were made through the patellar tendon at the patellar and tibial insertions, and the insertions were completely removed from the graft.

Nanofiber Mesh Fabrication and Characterization

Aligned nanofiber meshes (Fig. 28A,B) were fabricated by electrospinning¹³. A viscous polymer solution consisting of 35% poly (DL-lactic-co-glycolic acid) 85:15 (PLGA, I. V. = 0.70 dL/g, Lakeshore Biomaterials, Birmingham, AL), 55% N, N-dimethylformamide (Sigma, St. Louis, MO), and 10% ethanol (Commercial Alcohol, Inc., Toronto, Ontario) was loaded into a syringe fitted with an 18-gauge needle (Becton Dickinson, Franklin Lakes, NJ) . Aligned fibers⁸² were obtained using an aluminum drum with an outer diameter of 10.2 cm rotating with a surface velocity of 20 m/s. A constant flow rate of 1 mL/hr was maintained using a syringe pump (Harvard Apparatus, Holliston, MA) , and an electrical potential was applied between the needle and the grounded substrate (distance=10 cm) using a high voltage DC power supply (Spellman, Hauppauge, NY, 8-1OkV) . Fiber morphology, diameter and alignment of the as-fabricated mesh samples were analyzed using scanning electron microscopy (SEM) . Briefly, the samples were sputter-coated with gold (LVC-76, Plasma Sciences, Lorton, VA) and subsequently imaged (JSM 5600LV, JEOL, Tokyo, Japan) at an accelerating voltage of 5 kV.

Graft Collar Scaffold Fabrication

A tendon graft collar based on a sintered microsphere scaffold was fabricated following published methods³⁸'^{' 69}. Specifically, the scaffold is composed of composite microspheres consisting of PLGA (85:15, I. V. = 3.42 dl/g, Purac, Lincolnshire, IL) and 45S5 bioactive glass (BG, 20 μm, MO-SCI Corporation, Rolla, MD) . The microspheres were formed following the methods of Lu et al.³⁸, where the polymer was first dissolved in dichloromethane (Acros Organics, Morris Plains, NJ) and then BG particles were added (20 wt%) . After vortexing, the suspension was poured into a 1% solution of polyvinyl alcohol (Sigma, St. Louis, MO) to form the microspheres. The microspheres were subsequently sintered at 70⁰C for 5 hours in a custom mold to form cylindrical scaffolds with an outer diameter of 0.7 cm and an inner diameter of 0.3 cm.

Characterization of Nanofiber Mesh Contraction

Mesh contraction was evaluated using digital image analysis. Briefly, the nanofiber meshes were cut into 10 mm x 10 mm squares and immersed in Dulbecco' s Modification of Eagle's Medium (DMEM, Mediatech, Inc., Herndon, VA) supplemented with 10% fetal bovine serum (FBS, Atlanta Biologicals, Norcross, GA) and incubated at 37⁰C and 5% CO₂. The meshes were imaged using stereomicroscopy at 0, 2, 24, and 72 hours. Mesh dimensions (n=5) were measured by image analysis (ImageJ 1.34s, NIH, Bethesda, MD), and contraction was calculated based on percent change in length both in the x-axis and along the direction of fiber alignment (y-axis) .

Compression of Graft Collar Scaffold with Nanofiber Mesh

In addition to mesh contraction, the nanofiber mesh- mediated compression of the microsphere-based graft collar was also evaluated in vitro. Briefly, strips of nanofiber mesh (15.5 cm x 1.5 cm) were wrapped around the graft collar scaffold, with the fibers aligned perpendicular to the scaffold long axis. The mesh+scaffold was then incubated in PBS at 37⁰C and 5% CO₂, and changes in scaffold diameter (n=6) due to mesh contraction were monitored over 24 hours using image analysis (ImageJ).

Compression of Tendon with Nanofiber Mesh

The potential of utilizing nanofiber mesh contraction to directly apply compression to the tendon graft was evaluated over time. Briefly, the aligned electrospun meshes were cut into 10 cm x 2 cm strips, with fiber alignment oriented along the long axis of the mesh. The patellar tendon graft was bisected along its long axis, and one half of the tendon was wrapped with the nanofiber mesh while the other half served as the unloaded control

(Fig. 30A) . The samples were cultured in DMEM supplemented with 1% non-essential amino acids, 1% antibiotics, and 0.1% antifungal (all from Mediatech) and 10% FBS (Atlanta Biologicals) . At days 5 and 14, the effects of compression on tissue morphology and cellularity were characterized by histology⁶⁸. The samples were rinsed with phosphate buffered saline (PBS, Sigma), fixed with 10% neutral buffered formalin (Fisher Scientic and Sigma) and embedded in paraffin (Fisher Scientific, Pittsburgh, PA) . The samples were then cut into 7-μm thick sections and stained with hematoxylin and eosin (H&E) .

Compression of Tendon Graft with the Graft Collar Scaffold and Nanofiber Mesh

The potential of the graft collar scaffold and nanofiber mesh complex to apply static compression to the patellar tendon graft was also evaluated in vitro. Specifically, the patellar tendon graft was dissected into 2 cm x 0.3 cm segments and the cylindrical scaffold was halved along its long axis. Each tendon segment was inserted between the two scaffold halves (Fig. 31A, top). For the experimental group, the tendon+graft collar was wrapped with the aligned nanofiber mesh (15.5 cm x 1.5 cm), while the control scaffolds were wrapped with pre-contracted electrospun mesh (n=2). In addition, to ensure static compression of the tendon graft, the experimental group was wrapped with new mesh strips on every other day during the two week culturing period. The scaffold+tendon graft complex (Fig. 31A, bottom) was cultured in fully supplemented media at 37⁰C and 5% CO₂. Effects of Compression on Tendon Graft Cellularity and Matrix Content

The effects of static compression on tendon matrix organization (n=2) were analyzed at 1 and 14 days via histology (H&E) . In addition, since most of the mesh compression occurs within the first 24 hours, total cell number (n=5) and proteoglycan content in the tendon graft were evaluated at day 1. For the biochemical assays²⁸'^{" 29;} ⁶⁹, both the wet and dry weights of the tendon samples were determined at day 0 and day 1, and the tissue was subsequently digested for 16 hours in 2% papain (Sigma) buffer at 60⁰C. Total DNA content of the digest was determined with the PicoGreen dsDNA assay (Molecular Probes), following the manufacturer's suggested protocol. Sample fluorescence was measured using a microplate reader (Tecan, Research Triangle Park, NC) , with excitation and emission wavelengths set at 485 and 535 nm, respectively. The total number of cells in the sample was calculated using the conversion factor of 8 pg DNA/cell⁴⁰.

Sulfated glycosaminoglycan (GAG) content was quantified using a colorimetric 1, 9-dimethylmethylene blue (DMMB) assay. Tissue digest from the cell quantitation assay was combined with DMMB dye, and the concentration of GAG-

DMMB complexes was determined using a plate reader at 540 and 595 nm and correlated to a standard prepared with chondroitin-6-sulfate. CeIl Phenotype

Gene expression for fibrocartilage markers (n=2) such as collagen I, II, aggrecan, and Transforming Growth Factor- Beta 3 (TGF-β3) was determined at day 1 using reverse- transcription polymerase chain reaction (RT-PCR) . Briefly, after removing the graft collar and nanofiber mesh, total RNA of the tendon graft was obtained using the Trizol extraction method (Invitrogen, Carlsbad, CA) . The isolated RNA was reverse-transcribed into cDNA using the Superscript III First-Strand Synthesis System (Invitrogen, Carlsbad, CA) and the cDNA product was amplified using recombinant Platinum Tag DNA polymerase (Invitrogen) . GAPDH was used as the housekeeping gene, and expression band intensities were measured (ImageJ) and normalized against GAPDH.

Statistical Analysis

Results are presented in the form of mean ± standard deviation, with n equal to the number of samples analyzed. Two-way analysis of variance (ANOVA) was first performed to assess if differences exist among the means. Fisher's LSD post-hoc test was subsequently performed for all pair-wise comparisons and statistical significance was attained at p<0.05. For gene expression, a one-way ANOVA and Fisher's LSD post-hoc test were performed to determine the effect of compression. All statistical analyses were performed using the JMP statistical software package (SAS Institute, Cavy, NC) . RESULTS

Nanofiber Characterization and Mesh Contraction

The nanofiber mesh exhibited a high degree of alignment with an average fiber diameter of 0.9 ± 0.4 μm (Fig. 28A) . Anisotropic mesh contractile behavior was observed in the mesh, with significantly higher contraction found in the direction of nanofiber alignment. Specifically, the mesh contracted over 57% along the aligned fiber direction (y-axis) by 2 hours, with less than 13% reduction in the x-axis (Fig. 28B). Mesh contraction continued over time, exhibiting over 70% contraction in the y-axis and 20% in the x-axis by 24 hours and stabilizing thereafter, with no significant differences found between the 24- and 72-hour groups.

Compression of Graft Collar Scaffold with Nanofiber Mesh

After the nanofiber mesh was wrapped around the graft collar scaffold, mesh contraction resulted in a significant decrease in scaffold inner diameter, averaging 15% strain within 24 hours (Fig. 29). In contrast, the control scaffold without mesh cultured under similar conditions expanded and measured an increase in inner diameter (4%), although the difference was not statistically significant (p<0.05).

Compression of Tendon with Nanofiber Mesh

When the nanofiber mesh was used to compress the tendon graft, mesh contraction resulted in an approximately 30% decrease in graft diameter by 24 hours (Fig. 30A) . After five days of explant culture, the compressed tendon exhibits less of the crimp structure evident in the control group, and remodeled into a dense matrix with high cellularity (Fig. 30B) . However, by day 14, the crimp pattern was restored in the compressed group, with ultrastructure and cellularity indistinguishable from the unloaded control group.

Compression of Tendon with the Graft Collar Scaffold and Nanofiber Mesh

In order to apply a physiological level of loading (10- 15%), the tendon graft was compressed by a complex of the graft collar scaffold and nanofiber mesh. It was observed that at 24 hours post-compression (Fig. 31B, top) , the tendon graft matrix organization was distinct from that of the unloaded control, with increased matrix density and less of the characteristic crimp of the tendon. After 14 days of compression by the scaffold+mesh complex, it was found that the matrix remodeling visible 24 hours following the onset of loading was maintained over time (Fig. 31B, bottom). In contrast, the control tendon retained its characteristic crimp, with evident disruption of the matrix ultrastructure. In addition to changes in tendon matrix organization, total cell number in the tendons remained relatively constant in the compressed group, with a significantly higher number of cells found in the control tendons by day 1 (Fig. 32A). Interestingly, matrix glycosaminoglycan (GAG) content was found to be significantly higher in compressed tendon group after one day of culture (Fig. 32B) . Effects of Compression on the Expression of Fibrocartilage-Related Markers

The expression of fibrocartilage markers such as types I and II collagen, aggrecan and TGF-β3 were evaluated after compression with the graft collar scaffold and nanofiber mesh. As shown in Fig. 33, after 24 hours of compression, gene expression of type II collagen, aggrecan and TGF-β3 were all up-regulated in the loaded group when compared to non-compressed tendons (Fig. 33), with significant differences found in aggrecan and TGF-β3 expression.

DISCUSSION

The long term goal is to achieve biological fixation by engineering a functional and anatomical fibrocartilage interface on biological and synthetic soft tissue grafts used in orthopaedic repair³⁹. To this end, the current study focuses on the design and evaluation of a novel graft collar scaffold system capable of applying mechanical loading and inducing fibrocartilage formation on tendon grafts. Specifically, scaffold-mediated compression of a patellar tendon graft was evaluated over time, focusing on the effects of loading on tendon matrix organization and cell response. It was found that the complex of the nanofiber mesh and graft collar was able to apply a physiological range of compressive loading to tendon grafts. Moreover, scaffold-mediated compression promoted matrix remodeling, maintained graft glycosaminoglycan content and, interestingly, induced gene expression for fibrocartilage markers, including type II collagen, aggrecan, and TGF-β3. These promising results demonstrate that compressive loading can be incorporated into scaffold design and used to promote fibrocartilage formation on tendon grafts.

Two scaffold-based loading systems were described in this study. The first design involved using a nanofiber mesh to directly load the tendon graft. The pre-designed alignment of the nanofiber mesh results in anisotropic mesh contractile behavior, effectively translating contractile force into compression, which has been utilized in this study to apply compressive loading to the tendon grafts. Histological analysis of the grafts revealed that the scaffold-mediated compression induced extensive remodeling of the tendon ultrastructure, with the compressed graft exhibiting a denser matrix with increased local cell density. This matrix modulation effect, however, diminished over time, with the control and loaded groups nearly indistinguishable by day 14. As mesh contraction stabilizes after 24 hours, it is likely that the tendon graft is no longer experiencing mechanical stimulation in long term cultures. These observations suggest that it is necessary to incorporate extended mechanical stimulation into scaffold design.

The short-term effect of mesh-induced compressive loading on graft matrix organization and the high magnitude of compression (approximately 30%) initiated the development of the second mechano-active scaffold system. Specifically, the nanofiber mesh was combined with a degradable microsphere-based graft collar system in order to achieve a physiological range of loading (15%) . Moreover, to maintain static compression, the tendon- scaffold complex was wrapped with new nanofiber mesh every other day. It was observed that under static compression, the remodeled tendon matrix with cells embedded in a dense matrix was maintained over time, with marked differences observed between control and the loaded groups. These observations demonstrate the potential of this scaffold system to provide continuous mechanical stimulation and promote sustained tissue remodeling. Proteoglycan content of the tendon matrix was also significantly higher in the compressed group compared to the control at day 1, further indicating that the scaffold-induced compression influences matrix maintenance and remodeling.

Scaffold-mediated compression also resulted in the up- regulation of fibrocartilage markers including type II collagen, aggrecan, and Transforming Growth Factor-β3 (TGF-β3) . It is well known that fibrocartilage in tendons is largely comprised of types I and II collagen, as well as proteoglycans⁵'^{' 15; 32; 47}. Moreover, compressive loading of fibrocartilaginous regions of tendons has been reported to increase the synthesis of Transforming Growth Factor-βl (TGF-βl)⁵⁸ and large proteoglycans, as well as enhancing aggrecan gene expression¹⁵'^{' 32}. Compression of the non-fibrocartilaginous regions of the deep flexor tendon has also been reported to promote proteoglycan synthesis¹⁵. The findings of this study are in agreement with these published studies on the effects of compressive loading, and demonstrate the feasibility of implementing a degradable scaffold system for fibrocartilage interface formation on tendon grafts. In addition to applying continuous compressive loading to the graft within a physiological range, it is anticipated that this novel scaffold system also can be used to deliver cells and growth factors. These design optimizations will be critical for allograft recellularization and exercising biochemical stimulation to direct cellular differentiation as well as transformation of the tendon matrix into fibrocartilage .

Contraction of PLGA meshes has been previously reported in the literature⁸⁴, although the phenomenon has been discredited as a shortcoming rather than promoted as an advantageous attribute of the system. Currently, the mechanism underlying mesh contraction is not known. Zong et al.^%i have observed that electrospun nanofiber mesh comprised of crystalline polyesters contract significantly less than amorphous polyester co-polymers such as PLGA 75:25. It was proposed that when nanofiber meshes comprised of crystalline polymers are incubated at 37⁰C, the polymer glass transition temperature is approached and crystallization rapidly occurs, resulting in a lamellar structure that constrains the relaxation of the polymer chains and in turn prevents contraction⁸⁴. The polyester co-polymer utilized in this study has a high D,L-lactide content (85%) and is non-crystalline, thus the above mechanism may explain the high degree of contraction observed. Although not the focus of the current study, fiber alignment-related scaffold anisotropy may be controlled to modulate mesh contraction, and consequently, the magnitude and direction of compressive loading on the graft may be controlled by customizing the degree of fiber alignment. Future studies will focus on elucidating the mechanism of mesh contraction as well as exploring methods to control this process for mechanical stimulation. This is the first study to incorporate mechanical loading into scaffold design and to demonstrate the potential of using this mechano-active scaffold system to induce fibrocartilage formation on soft tissue grafts. The mesh- collar system is intended to be applied clinically as a degradable graft collar, and will be used to initiate and direct regeneration of an anatomical fibrocartilage interface at the insertion of tendon-based ACL reconstruction grafts. In addition to providing a three- dimensional environment for matrix development and growth factors for guided cell differentiation, the innovative scaffold system described here can also apply physiologic mechanical stimulation crucial for directing cellular function and tissue remodeling. For utilization with viable autografts, it is envisioned that the graft would be inserted through the collars immediately prior to implantation, and compression of the graft and subsequent fibrocartilage formation would occur in vivo. Allografts, which do not contain viable cells necessary for remodeling the tendon matrix, would need to be repopulated with fibroblasts or stem cells delivered either from the scaffold in vitro prior to graft implantation. It has been reported that mesenchymal stem cell (MSC) -seeded type I collagen sponges inserted into excised sheep patellar tendons and loaded using an ex vivo wrap-around system results in an up-regulation of chondrogenic markers such as Sox9 and Fos³⁴. A similar response by a cell-populated tendon allograft is anticipated following scaffold-mediated compressive loading. Moreover, the mesh-scaffold system is based on degradable poly-α-hydroxyester polymers, thus it is expected that the mechano-active scaffold will be replaced by newly formed tissue after a functional fibrocartilage interface has been formed on the graft. Future studies will evaluate the potential of coupling the mechano-active scaffold with ACL reconstruction grafts using in vivo models.

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The objective of this experiment was to determine the effect of wrapping tendon with a PLGA electrospun mesh wherein the fibers of the mesh were either perpendicular or parallel to the longitudinal axis of the tendon.

Three groups (3 tendons per group) were examined: (1) A control group with no mesh wrapping; (2) a group wrapped with mesh, wherein the fibers of the mesh were perpendicular to the longitudinal axix of the tendon; and

(3) a group wrapped with mesh, wherein the fibers of the mesh were parallel to the longitudinal axis of the tendon. The meshes were allowed to contract for 48 hours. See Figures 34A and 34B.

Results

The control group exhibited a 13.3 ± 6.4 percentage change in tendon diameter and a -6.2 ± 5.2 percentage change in tendon length. The perpendicular fiber group exhibited a -40.0 ± 63.6 percentage change in tendon diameter and a 12.9 ± 2.2 percentage change in tendon length. The parallel fiber group exhibited a 5.6 ± 6.7 percentage change in tendon diameter and a -16.3 ± 5.6 percentage change in tendon length.

Discussion

This experiment indicates that wrapping a tendon with a PLGA electrospun mesh having fibers perpendicular to the longitudinal axis of the tendon results in decreased tendon diameter and increased tendon length (due to compression of the center of the tendon) . Wrapping a tendon with a PLGA electrospun mesh having fibers parallel to the longitudinal axis of the tendon results in decreased tendon length and no significant change in tendon diameter compared to the control.

Claims

What is claimed is:

1. A graft collar for fixing tendon to bone in a subject, wherein said graft collar comprises a sheet of biopolymer mesh.

2. The graft collar of claim 1, wherein the biopolymer mesh comprises aligned fibers.

3. The graft collar of claim 1, wherein the biopolymer mesh is derived from at least one of collagen, chitosan, silk and alginate.

4. The graft collar of claim 1, wherein the biopolymer mesh is allogenic or xenogenic.

5. The graft collar of claim 1, wherein the graft collar is sutured around a tendon graft.

6. The graft collar of claim 1, wherein the subject is a mammal.

7. The graft collar of claim 6, wherein the mammal is a human .

8. The graft collar of claim 1, wherein the graft collar promotes integration of the tendon graft to bone.

9. The graft collar of claim 1, wherein the graft collar includes at least one of the following substances: anti-infectives, antibiotics, bisphosphonate, hormones, analgesics, anti- -Ill-

inflammatory agents, growth factors, angiogenic factors, chemotherapeutic agents, anti-rejection agents, and RGD peptides.

10. The graft collar of claim 9, wherein the growth factors are selected from the group consisting of TGFs, BMPs, IGFs, VEGFs and PDGFs.

11. The graft collar of claim 10, wherein the TGF is TGF-β.

12. The graft collar of claim 10, wherein the BMP is BMP-2.

13. The graft collar of claim 1, wherein the graft collar includes one or more of the following types of cells: chondrocytes, osteoblasts, osteoblast- like cells and stem cells.

14. The graft collar of claim 1, wherein the graft collar includes at least one of the following: osteogenic agents, osteogenic materials, osteoinductive agents, osteoinductive materials, osteoconductive agents, osteoconductive materials and chemical factors.

15. The graft collar of claim 1, wherein the graft collar promotes regeneration of an interfacial region between tendon and bone.

16. The graft collar of claim 1, wherein the graft collar is lyophilized.

17. The graft collar of claim 1, wherein the graft collar is biodegradable.

18. The graft collar of claim 1, wherein the graft collar is osteointegrative .

19. A graft collar for fixing tendon to bone in a subject, wherein said graft collar comprises a sheet of polymer-fiber mesh.

20. The graft collar of claim 19, wherein the polymer- fiber mesh comprises aligned fibers.

21. The graft collar of claim 19, wherein the graft collar is sutured around a tendon graft.

22. The graft collar of claim 19, wherein the subject is a mammal.

23. The graft collar of claim 22, wherein the mammal is a human .

24. The graft collar of claim 19, wherein the graft collar promotes integration of the tendon graft to bone .

25. The graft collar of claim 19, wherein the graft collar includes at least one of the following substances: anti-infectives, antibiotics, bisphosphonate, hormones, analgesics, antiinflammatory agents, growth factors, angiogenic factors, chemotherapeutic agents, anti-rejection agents, and RGD peptides.

26. The graft collar of claim 25, wherein the growth factors are selected from the group consisting of TGFs, BMPs, IGFs, VEGFs and PDGFs.

27. The graft collar of claim 26, wherein the TGF is TGF-β.

28. The graft collar of claim 26, wherein the BMP is BMP-2.

29. The graft collar of claim 19, wherein the graft collar includes one or more of the following types of cells: chondrocytes, osteoblasts, osteoblast- like cells and stem cells.

30. The graft collar of claim 19, wherein the graft collar includes at least one of the following: osteogenic agents, osteogenic materials, osteoinductive agents, osteoinductive materials, osteoconductive agents, osteoconductive materials and chemical factors.

31. The graft collar of claim 19, wherein the polymer- fiber mesh in the second region is selected from the group comprising 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, and biopolymers, and a blend of two or more of the preceding polymers.

32. The graft collar of claim 31, wherein the polymer comprises at least one of the poly (lactide-co- glycolide) , poly (lactide) and poly (glycolide) .

33. The graft collar of claim 19, wherein the graft collar promotes regeneration of an interfacial region between tendon and bone.

34. The graft collar of claim 19, wherein the graft collar is lyophilized.

35. The graft collar of claim 19, wherein the graft collar is biodegradable.

36. The graft collar of claim 19, wherein the graft collar is osteointegrative .

37. A graft collar for fixing tendon to bone in a subject, wherein the graft collar comprises:

(a) a first region comprising a biopolymer mesh and hydrogel; and

(b) a second region adjoining the first region and comprising a biopolymer mesh.

38. The graft collar of claim 37, wherein the subject is a mammal.

39. The graft collar of claim 38, wherein the mammal is a human .

40. The graft collar of claim 37, wherein the first region supports the growth and maintenance of an interfacial zone between tendon and bone, and the second region supports the growth and maintenance of bone tissue.

41. The graft collar of claim 37, wherein the graft collar includes at least one of the following substances: anti-infectives, antibiotics, bisphophonate, hormones, analgesics, antiinflammatory agents, growth factors, angiogenic factors, chemotherapeutic agents, anti-rejections agents, and RGD peptides.

42. The graft collar of claim 37, wherein the hydrogel is photopolymerized, thermoset or chemically cross- linked.

43. The graft collar of claim 37, wherein the hydrogel is polyethylene glycol.

44. The graft collar of claim 37, wherein the biopolymer mesh comprises aligned fibers.

45. The graft collar of claim 37, wherein the first region contains TGF.

46. The graft collar of claim 45, wherein the TGF is TGF-β.

47. The graft collar of claim 37, wherein the first region contains chondrocytes.

48. The graft collar of claim 47, wherein the chondrocytes are BMSC-derived.

49. The graft collar of claim 37, wherein the first region contains stem cells.

50. The graft collar of claim 49, wherein the stem cells are BMSCs.

51. The graft collar of claim 37, wherein the biopolymer mesh is derived from at least one of collagen, chitosan, silk and alginate.

52. The graft collar of claim 37, wherein the biopolymer mesh is allogenic or xenogenic.

53. The graft collar of claim 37, wherein the second region contains at least one of the following growth factors: BMP, IGF, VEGF and PDGF.

54. The graft collar of claim 53, wherein the BMP is BMP-2.

55. The graft collar of claim 37, wherein the second region includes osteoblasts and/or osteoblast-like cells .

56. The graft collar of claim 55, wherein the osteoblasts and/or osteoblast like cells are BMSC- derived.

57. The graft collar of claim 37, wherein the second region includes at least one of the following: osteogenic agents, osteogenic materials, osteoinductive agents, osteoinductive materials, osteoconductive agents, osteoconductive materials and chemical factors.

58. The graft collar of claim 37, wherein the second region contains nanoparticles of calcium phosphate.

59. The graft collar of claim 58, wherein the calcium phosphate is selected from the group comprising tricalcium phosphate, hydroxyapatite and a combination thereof.

60. The graft collar of claim 37, wherein the second region contains nanoparticles of bioglass.

61. The graft collar of claim 37, wherein the graft collar is biodegradable.

62. The graft collar of claim 37, wherein the graft collar is osteointegrative.

63. A graft collar for fixing tendon to bone in a subject, wherein the apparatus comprises:

(a) a first region comprising a polymer-fiber mesh and hydrogel; and

(b) a second region adjoining the first region and comprising a polymer-fiber mesh.

64. The graft collar of claim 63, wherein the subject is a mammal.

65. The graft collar of claim 64, wherein the mammal is a human.

66. The graft collar of claim 63, wherein the first region supports the growth and maintenance of an interfacial zone between tendon and bone, and the second region supports the growth and maintenance of bone tissue.

67. The graft collar of claim 63, wherein the graft collar includes at least one of the following substances: anti-infectives, hormones, analgesics, anti-inflammatory agents, growth factors, angiogenic factors, chemotherapeutic agents, anti-rejections agents, and RGD peptides.

68. The graft collar of claim 63, wherein the hydrogel is photopolymerized, thermoset or chemically cross- linked.

69. The graft collar of claim 63, wherein the hydrogel is polyethylene glycol.

70. The graft collar of claim 63, wherein the polymer- fiber mesh comprises aligned fibers.

71. The graft collar of claim 63, wherein the first region contains TGF.

72. The graft collar of claim 71, wherein the TGF is TGF-β

73. The graft collar of claim 63, wherein the first region contains chondrocytes.

74. The graft collar of claim 73, wherein the chondrocytes are BMSC-derived.

75. The graft collar of claim 63, wherein the first region contains stem cells.

76. The graft collar of claim 75, where the stem cells are BMSCs.

77. The graft collar of claim 63, wherein the second region contains at least one of the following growth factors: BMP, IGF, VEGF and PDGF.

78. The graft collar of claim 77, wherein the BMP is BMP-2.

79. The graft collar of claim 63, wherein the second region includes osteoblasts and/or osteoblast-like cells .

80. The graft collar of claim 79, wherein the osteoblasts and/or osteoblast like cells are BMSC- derived.

81. The graft collar of claim 63, wherein the second region includes at least one of the following: osteogenic agents, osteogenic materials, osteoinductive agents, osteoinductive materials, osteoconductive agents, osteoconductive materials and chemical factors.

82. The graft collar of claim 63, wherein the second region contains nanoparticles of calcium phosphate.

83. The graft collar of claim 82, wherein the calcium phosphate is selected from the group comprising tricalcium phosphate, hydroxyapatite and a combination thereof.

84. The graft collar of claim 63, wherein the second region contains nanoparticles of bioglass.

85. The graft collar of claim 63, wherein the polymer- fiber mesh in the second region is selected from the group comprising 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, and biopolymers, and a blend of two or more of the preceding polymers.

86. The graft collar of claim 85, wherein the polymer comprises at least one of the poly (lactide-co- glycolide) , poly (lactide) and poly (glycolide) .

87. The graft collar of claim 63, wherein the apparatus is biodegradable.

88. The graft collar of claim 63, wherein the apparatus is osteointegrative .

89. A graft collar for fixing tendon to bone in a subject, wherein said graft collar comprises a sheet of mesh comprising fibers aligned substantially perpendicular in relation to a longitudinal axis of said tendon, wherein said mesh applies compression to the graft.

90. The graft collar of claim 89, wherein the mesh comprises a biopolymer.

91. The graft collar of claim 89, wherein the mesh comprises a polymer-fiber.

92. The graft collar of claim 89, wherein the graft collar comprises:

(a) a first region comprising a mesh and hydrogel; and

(b) a second region adjoining the first region and comprising a mesh.

93. A graft collar for fixing tendon to bone in a subject, wherein said graft collar comprises a sheet of mesh comprising fibers aligned substantially parallel in relation to a longitudinal axis of said tendon, wherein said mesh applies lateral tension to the graft.

94. The graft collar of claim 93, wherein the mesh comprises a biopolymer.

95. The graft collar of claim 93, wherein the mesh comprises a polymer-fiber.

96. The graft collar of claim 93, wherein the graft collar comprises:

(a) a first region comprising a mesh and hydrogel; and

(b) a second region adjoining the first region and comprising a mesh.