RELATED APPLICATIONS

[01] This application claims priority to U.S. Provisional Application No. 62/064,748, filed on October 16, 2014, and U.S. Provisional Application No. 62/205,488, filed on August 14, 2015. The entire contents of each of the foregoing applications are incorporated herein by reference.

FIELD OF THE INVENTION

[02] The present invention relates to engineered constructs for tissue repair comprising collagen and alginate.

BACKGROUND OF THE INVENTION

[03] Multiple applications in tissue repair and regeneration depend on availability of suitable artificial substitutes that can replace a damaged natural tissue and successfully integrate with the host tissue and that promote vascular network formation. Many types of materials have been used as implants in tissue engineering, including natural and synthetic materials. Natural materials can include, for example, extracellular matrices (ECMs) obtained from allografts (obtained from a human donor other than a patient) or xenografts (obtained from an animal donor). Although natural materials typically have good

biocompatibility, they can have insufficient mechanical stability and can cause adverse immune reactions in a host. Synthetic materials provide improved and controllable mechanical properties, but their biocompatibility can be low. Some clinical applications include the use of combinations of natural and synthetic materials.

[04] However, despite a variety of existing approaches, many challenges remain in creating sufficiently strong, safe, and efficient artificial tissues that would repair a damaged natural tissue, assist in healing process, and facilitate host integration without causing undesirable complications. For example, some existing techniques for abdominal wall hernia reconstruction may be associated with seroma formation, infection or even reherniation.

[05] Accordingly, there remains a need for engineered substitutes to restore, replace or regenerate damaged tissues. SUMMARY OF THE INVENTION

[06] The invention addresses these issues and provides methods and systems for tissue repair and regeneration using collagen-alginate constructs comprising stacked layers of collagen sheets having an alginate incorporated therein. Various patterns of pores are formed in the collagen sheets. A pore pattern determines a mechanical strength of the construct and promotes vascular formation.

[07] The described constructs are configured to be sufficiently strong and at the same time flexible for a number of clinical applications. The structure of the constructs improves healing and promotes a strength of integration between the construct and host tissue. The collagen component of the constructs provides structural support for cell attachment and subsequent tissue development, whereas the alginate promotes vascularization and tissue ingrowth.

[08] In some embodiments, the alginate comprises cells and/or bioactive molecules (e.g. , growth factors and cytokines) and serves as an efficient vehicle for cell incorporation and release. When the collagen-alginate construct seeded with cells or bioactive molecules is implanted into a subject, the alginate acts as a temporary bioactive interface between the construct and a host environment and as carrier of cells and bioactive molecules to promote a beneficial host response.

[09] Constructs generated in accordance with some embodiments may have tunable mechanical characteristics, cytocompatibility, biocompatibility, and implant/host integration that make them suitable for use in various tissue repair and regeneration, drug delivery, and other applications.

[010] Accordingly, in some embodiments, the present invention provides a construct having a top side and a bottom side and comprising a plurality of collagen sheets and a plurality of pores. The construct further comprises alginate at least partially incorporated into both the top side and the bottom side of the construct.

[011] In some aspects, the construct further comprises at least one component selected from the group consisting of a cell, a growth factor and a cytokine.

[012] In some embodiments, the alginate comprises a mixture of high molecular weight alginate and low molecular weight alginate. The degradation rate of the alginate may depend on a ratio of the high molecular weight alginate to the low molecular weight alginate. In a further aspect, the ratio of the high molecular weight alginate to the low molecular weight alginate is a ratio selected from the group consisting of 90: 1, 90: 10, 75:25, 50:50, 25:75 and 10:90.

[013] In one embodiment, each collagen sheet comprised in the plurality of collagen sheets has a thickness of between about 10 micrometers and about 40 micrometers. In another embodiment, each collagen sheet comprised in the plurality of collagen sheets has a thickness of between about 20 micrometers and about 100 micrometers. In a further embodiment, each collagen sheet has a thickness of about 65 micrometers. In another further embodiment, each collagen sheet has a thickness of about 40 micrometers.

[014] In one aspect, each collagen sheet comprised in the plurality of collagen sheets has a generally rectangular shape and has a length of about 25 mm and a width of about 15 mm.

[015] In some embodiments, the construct of the invention comprises about 15 to about 150 collagen sheets. In another embodiment, the construct of the invention comprises about 10 to about 100 collagen sheets. In a further embodiment, the construct of the invention comprises about 10 to about 20 collagen sheets. In yet another further embodiment, the construct of the invention comprises 9 collagen sheets.

[016] In one aspect, the plurality of pores in the construct of the invention are adapted to facilitate cell infiltration and vascularization to form new tissue within the construct upon implantation of said construct into a subject.

[017] In one embodiment, each pore comprised in the plurality of pores in the construct of the invention has a hexagonal shape. In another embodiment, each pore has a round shape.

[018] In some aspects, the plurality of pores in the construct of the invention form a uniform pattern in the construct. In other aspects, the plurality of pores form a non-uniform pattern in the construct.

[019] In some embodiments, the plurality of pores in the construct of the invention form a pattern having pore density of about 1 to about 60 pores/cm . In other embodiments, the pore density is about 40 to about 240 pores/cm .

[020] In some aspects, the pores in the construct of the invention have a diameter of about 50 μιη to about 150 μιη. In other aspects, the diameter is about 100 μιη to about 250 μιη, about 250 μιη to about 500 μιη, smaller than about 30 μιη, about 30 μιη to about 60 μιη, greater than about 60 μιη. In further aspects, the diameter is about 500 μιη, about 250 μιη or about 130 μιη.

[021] In some embodiments, the plurality of pores in the construct of the invention form a pattern having pore density of about 60 pores/cm . In other embodiments, the pore density is about 240 pores/cm . In a further aspect, the pores have a diameter of about 250 μιη. In another further aspect, the plurality of pores comprise about 10% or more of the total surface area of the construct.

[022] In certain embodiments, the plurality of pores in the construct comprise about 1% to about 5% of the total surface area of the construct. In other embodiments, the pores comprise about 5% to about 15% of the total surface area of the construct.

[023] In some aspects, collagen concentration in the plurality of collagen sheets in the construct of the invention determines the rate of construct degradation.

[024] In some embodiments, the construct of the invention is adapted to promote new tissue formation within the construct.

[025] In some aspects, the alginate present in the construct of the invention comprises a plurality of cells. In one further aspect, the plurality of cells are present at a concentration of about 4 x 10 ⁶ cells/mL. In other further aspects, the plurality of cells are present at a concentration of at least about lxlO ⁶ cells/mL, at least about 2xl0 ⁶ cells/mL, at least about 3xl0 ⁶ cells/mL, at least about 4xl0 ⁶ cells/mL, at least about 5xl0 ⁶ cells/mL, at least about 6xl0 ⁶ cells/mL, at least about 7xl0 ⁶ cells/mL, at least about 8xl0 ⁶ cells/mL, at least about 9xl0 ⁶ cells/mL or at least about lxlO ⁷ cells/mL.

[026] In some aspects, at least one mechanical characteristic of the construct of the invention controls secretion of cytokines and growth factors from the cells. In other aspects, the construct is adapted to maintain viability of the cells for at least about seven days after the construct is implanted in a subject. In a further aspect, the cells are selected from the group consisting of mesenchymal stem cell (MSCs), hematopoietic stem cells (HSCs) or progenitors thereof, and endothelial progenitor cells (EPCs).

[027] In some embodiments, the alginate present in the construct of the invention is cross- linked with calcium chloride. In another embodiment, the alginate is cross-linked covalently or non-covalently. In yet another embodiment, the alginate is cross-lined with a divalent cation, e.g., a divalent cation selected from the group consisting of Ca ²⁺ , Mg ²⁺ , Sr ²⁺ , Ba ²⁺ , or Be ²⁺ . In a further embodiment, the divalent cation is Ca ²⁺ .

[028] In some aspects, the alginate is conjugated to a cell-adhesive peptide.

[029] In some aspects, the construct of the invention is adapted for treating hernia. In other aspects, the construct is adapted to promote formation of a blood vessel network within the construct after implantation into a subject. In yet other aspects, the construct is adapted for use in at least one of the following: a graft for abdominal wall reconstruction, a cardiac patch, a skin replacement patch, a matrix for muscle regeneration, an engineered cartilage and bone tissue, a tendon replacement patch, an open wound mesh, a graft for oral tissue repair, a graft for plastic surgery, a matrix for wound healing in the extremities, a matrix for venous ulcer repair, or a graft for combat related blast injuries.

[030] In some embodiments, the present invention provides a method of preparing an artificial construct for tissue repair. The method comprises fabricating a plurality of collagen sheets; assembling the plurality of collagen sheets into a laminate having the plurality of collagen sheets; forming a plurality of pores in the laminate; and incorporating an alginate into the laminate, such that the alginate is at least partially deposited on both sides of the laminate, thereby preparing an artificial construct for tissue repair. In some embodiments, the method further comprises mixing the alginate with at least one agent selected from the group consisting of a cell, a growth factor and a cytokine prior to incorporating the alginate into the laminate.

[031] In some aspects, the plurality of collagen sheets are fabricated from a collagen gel. In other aspects, the collagen gel comprises a layer of collagen gel having a thickness of about 8 mm.

[032] In further embodiments, fabricating the plurality of collagen sheets comprises drying the collagen gel to provide a dried collagen gel; cutting the dried collagen gel to form the plurality of collagen sheets; and hydrating the plurality of collagen sheets prior to assembling the plurality of collagen sheets into the laminate. The method may further comprise, prior to forming the plurality of pores in the laminate, incubating the laminate in a fibril assembly buffer for about 48 hours; and drying the laminate after the incubation.

[033] The present invention also provides a method of repairing tissue in a subject in need thereof. The method comprises implanting into the subject a construct having a top side and a bottom side and comprising a plurality of collagen sheets and a plurality of pores, the construct further comprising alginate at least partially incorporated into both the top side and the bottom side of the construct. In some embodiments, the alginate present in the construct comprises at least one agent selected from the group consisting of: a cell, a growth factor and a cytokine. In some embodiments, the construct is adapted for use in at least one of the following: a graft for abdominal wall reconstruction, a cardiac patch, a skin replacement patch, a matrix for muscle regeneration, an engineered cartilage and bone tissue, a tendon replacement patch, an open wound mesh, a graft for oral tissue repair, a graft for plastic surgery, a matrix for wound healing in the extremities, a matrix for venous ulcer repair, or a graft for combat-related blast injuries.

[034] Unless otherwise defined, all technical and scientific terms used herein have the same meaning as commonly understood by one of ordinary skill in the art to which this invention belongs. Although methods and materials similar or equivalent to those described herein can be used in the practice or testing of the described embodiments, suitable methods and materials are described below. All publications, patent applications, patents, and other references mentioned herein are incorporated by reference in their entirety. In the case of conflict, the present specification, including definitions, will control. In addition, the materials, methods, and examples are illustrative only and are not intended to be limiting.

[035] Other features and advantages of the invention will be apparent from the following detailed description and claims.

BRIEF DESCRIPTION OF THE DRAWINGS

[036] Figs. 1A-1F are schematic illustrations of a process of fabrication of a collagen- alginate construct in accordance with some embodiments.

[037] Fig. 1A is an illustration of a collagen gel.

[038] Fig. IB is an illustration of a collagen sheet.

[039] Fig. 1C is an illustration of layering of multiple collagen sheets.

[040] Fig. ID is an illustration of lamination in a fibril assembly buffer under compression.

[041] Fig. IE is an illustration of a construct with pores formed therein.

[042] Fig. IF is an illustration of exemplary dimensions of pores in the construct.

[043] Fig 1G is a bright field microscopy image of a top view of the construct.

[044] Fig. 1H is a bright field microscopy image of a cross-sectional view of the construct.

[045] Fig II is a scanning electron microscopy (SEM) image of a prospective cross- sectional view of the construct.

[046] Fig. 1J is another SEM image taken under higher magnification and showing a pore in the construct.

[047] Figs. 2A-2D are images illustrating different pore patterns in a construct in accordance with some embodiments.

[048] Fig. 2A is an image of an area of a patterned construct.

[049] Fig. 2B is an image of an area of a patterned construct having pores that are larger than the pores in the construct of Fig. 2A occupying the same area.

[050] Fig. 2C is an image of an area of a patterned construct having the same number of pores as the construct of Fig. 2A which are smaller in size than the pores in the construct of Fig. 2A.

[051] Fig. 2D is a construct without pores. [052] Figs. 2E-2H are bar graphs illustrating different mechanical properties of the constructs of Figs. 2A-2D, with error bars representing standard error of the mean.

[053] Fig. 2E is a bar graph showing ultimate tensile strength (MPa) for constructs of Figs.

2A-2D.

[054] Fig. 2F is a bar graph showing Young's modulus (MPa) for constructs of Figs. 2A- 2D.

[055] Fig. 2G is a bar graph showing strain at failure (%) for constructs of Figs. 2A-2D.

[056] Fig 2H is a bar graph showing suture retention strength (g-f) for constructs of Figs. 2A-2D.

[057] Figs. 3A, 3B and 3C are schematic illustrations of a process of incorporating alginate into a construct in accordance with some embodiments.

[058] Figs. 4A and 4B are fluorescent microscopy images illustrating cell distribution on both sides ("Side 1" and "Side 2") of a construct in accordance with some embodiments. Cells were stained with calcein AM. Scale bars are 200 μιη.

[059] Figs. 5 A-5F are confocal microscopy images of human mesenchymal stem cells

(hMSCs) in a pore of a construct after incorporation of an alginate with hMSCs into the construct in accordance with some embodiments. Scale bars are 200 μιη.

[060] Figs. 5 A and 5B are images of hMSCs 1 day after incorporation.

[061] Figs. 5C and 5D are images of hMSCs 3 days after incorporation.

[062] Figs. Figs. 5E and 5F are images of hMSCs 7 days after incorporation.

[063] Fig. 6 A is a graph showing the number of hMSCs per square centimeter in a construct in accordance with some embodiments 1, 3, and 7 days after incorporation of an alginate with hMSCs into the construct. Error bars represent standard error of the mean.

[064] Fig. 6B is a graph showing viability of hMSCs in the construct 1, 3, and 7 days after incorporation of an alginate with hMSCs into the construct. Error bars represent standard error of the mean.

[065] Figs. 6C-6E are graphs illustrating the release of growth factors by hMSCs in a construct in accordance with some embodiments ("Patch") and a release of growth factors by hMSCs cultivated on a plate ("Plate") at days 2 and 8 (normalized to 10 ⁵ cells).

[066] Fig. 6C is a bar graph showing the release of monocyte chemotactic protein- 1 (MCP-

1) from hMSCs in a construct in accordance with some embodiments.

[067] Fig. 6D is a bar graph showing the release of vascular endothelial growth factor

(VEGF) from hMSCs in a construct in accordance with some embodiments. [068] Fig. 6E is a bar graph illustrating the release of platelet-derived growth factor subunit B (PDGF-B) release from hMSCs in a construct in accordance with some embodiments. Error bars represent standard error of the mean.

[069] Figs. 7A and 7B are images illustrating the repair of an abdominal wall defect (hernia) induced in Wistar rats over an 8-week period using a construct in accordance with some embodiments.

[070] Figs. 7C, 7D, and 7E are images illustrating a subcutaneous view of the abdominal wall defect shown in Figs. 7A and 7B treated with a construct seeded with hMSCs ("hMSCs- seeded").

[071] Fig. 7C is an image of the abdominal wall defect at 2 weeks after implantation.

[072] Fig. 7D is an image of the abdominal wall defect at 4 weeks after implantation.

[073] Fig. 7E is an image of the abdominal wall defect 8 weeks after implantation.

[074] Figs. 7F, 7G, and 7H are images illustrating a subcutaneous view of the abdominal wall defect shown in Figs. 7A and 7B treated with a construct without cells ("Acellular").

[075] Fig. 7F is an image of the abdominal wall at 2 weeks after implantation.

[076] Fig. 7G is an image of the abdominal wall at 4 weeks after implantation.

[077] Fig. 7H is an image of the abdominal wall at 8 weeks after implantation.

[078] Fig. 8A is an image of a Wistar rat demonstrating that no hernia is observed in animals repaired with acellular and with cell-seeded constructs in accordance with some embodiments.

[079] Figs. 8B, 8C, and 8D are images illustrating a peritoneal view of the abdominal wall defect shown in Figs. 7A and 7B treated with a construct seeded with hMSCs ("hMSCs- seeded").

[080] Fig. 8B is an image of the abdominal wall at 2 weeks after implantation.

[081] Fig. 8C is an image of the abdominal wall at 4 weeks after implantation.

[082] Fig. 8D is an image of the abdominal wall at 8 weeks after implantation.

[083] Figs. 8E, 8F, and 8G are images illustrating a peritoneal view of the abdominal wall defect shown in Figs. 7A and 7B treated using a construct without cells ("Acellular").

[084] Fig. 8E is an image of the abdominal wall at 2 weeks after implantation.

[085] Fig. 8F is an image of the abdominal wall at 4 weeks after implantation.

[086] Fig. 8G is an image of the abdominal wall at 8 weeks after implantation.

[087] Fig. 9A is a series of histological images showing Masson's Trichrome staining of extracellular matrix and blood vessels formed in a construct in accordance with some embodiments after implantation. Subcutaneous, middle and peritoneal views of constructs without cells ("Acellular") and constructs seeded with hMSCs ("hMSCs-seeded") are shown at 2 weeks, 4 weeks, and 8 weeks after implantation. Scale bars represent 200 μιη.

[088] Fig. 9B is a series of histological images showing vWF staining of extracellular matrix and blood vessels formed on a construct in accordance with some embodiments after implantation. Middle section areas of constructs seeded with hMSCs ("hMSCs-seeded") and constructs without cells ("Acellular") are shown at 2 weeks, 4 weeks, and 8 weeks after implantation. Scale bars represent 200 μιη.

[089] Fig. 9C is a graph illustrating a number of blood vessels per 20x field at 2, 4 and 8 weeks after implantation formed in the middle section of a construct in accordance with some embodiments. Error bars represent a standard error of the mean.

[090] Figs. 10A and 10B illustrate macrophage infiltration of acellular and hMSC-seeded patches after in vivo implantation.

[091] Fig. 10A is a series of images of CD68 staining of acellular and hMSC-seeded samples at 2, 4, and 8 weeks. Scale bar 200 μιη.

[092] Fig. 10B is a graph showing the amount of CD68 staining per 20x field at 2, 4 and 8 weeks (*p < 0.05, n = 3 - 7). Bars represent standard error of the mean.

[093] Fig. 11, Panels A-D are images of iNOS+ to CD68+ stained cells in serial sections samples at 2, 4 and 8 weeks after in vivo implantation. Scale bar 200 μιη.

[094] Fig. 11, Panel D is a graph showing the ratios of iNOS+ to CD68+ stained cells in serial sections normalized to acellular samples at 2 weeks, 4 weeks, and 8 weeks. Bars represent standard error of the mean.

[095] Fig. 12 illustrates the persistence of hMSCs after implantation of hMSC-seeded patches. Panel A of Fig. 12 is a series of confocal microscopy cross-sectional images through hMSC-seeded patches at 2 hours, 3 days, 1, 2, 4, and 6 weeks, with dashed lines representing a pore in the patch. Scale bar 40 μιη. Panel B of Fig. 12B is a graph showing the number of labeled hMSCs over a 42 day period in vivo (n = 1).

[096] Fig. 13 is a graph illustrating strength of integration of a construct in accordance with some embodiments at 2, 4, and 8 weeks after implantation. Error bars represent standard error of the mean.

DETAILED DESCRIPTION OF THE INVENTION

[097] Mechanically robust and compliant composite matrices in accordance with some embodiments, also referred to herein as constructs, that comprise collagen and alginate, are useful for tissue repair and regeneration. Collagen is a biodegradable, biocompatible, and non-immunogenic structural protein, and alginate is a natural biocompatible and biodegradable polysaccharide having many different applications in biomedicine.

[098] Commonly used prosthetic meshes include synthetic textiles of polypropylene, polytetrafluoroethylene, or polyethylene-terepthalate. Although these polymeric materials generally prevent certain conditions, for example, recurrent hernias, morbidity rates remain undesirably high. Moreover, more than 50% of patients may experience discomfort, decreased abdominal wall mobility, and/or seroma formation. Clinical studies have demonstrated that, when conventional implants are used, infection rates may range from approximately 4% to approximately 14%, and hernia recurrence rates may range from approximately 8% to approximately 35%. Use of synthetic hernia meshes may also be limited due to their predisposition to harbor bacteria which can lead to infections of the meshes with associated severe consequences.

[099] Biologic materials, which may be derived from human cadaver, porcine, or bovine donors, or that have been obtained from tissues such as the dermis, pericardium, or intestinal submucosa have also been problematic. Although attempts have been made to remove pathogens and living cells from the donor to create a network of structural proteins (e.g., collagen), the use of biologic materials, for example, for hernia repair has been compromised by infections that can accelerate degradation and lead to hernia recurrence. Porcine biologies have been shown to be associated with less recurrence of hernia, but patients exhibited elevated rates of seroma formation, significant discomfort, and pain. (Klosterhalfen, B., Junge, K. & Klinge, U. The lightweight and large porous mesh concept for hernia repair. Expert Rev Med Devices, 2, 103-117 (2005); Patel, K.M., & Bhanot, P. Complications of Acellular Dermal Matrices in Abdominal Wall Reconstruction. Plastic Reconstruction Surgery, Volume 130, Number 5S-2, 216-224 (2012).)

[0100] The embodiments described herein are based on a surprising and unexpected discovery that a construct including a plurality of collagen sheets or lamellae compressed into a stack having a porous structure and embedded with alginate provides improved integration strength and tissue healing response. The collagen in the construct provides strength, and its porous pattern leads to a construct with a desirable flexibility without compromising its mechanical stability. The alginate, integrated into both sides of the construct including inner surface of the pores, promotes integration strength and host tissue incorporation.

[0101] In some embodiments, the alginate is seeded with cells and/or bioactive molecules (e.g., growth factors and cytokines) and serves as a vehicle for cell delivery such that the construct retains cells in defective tissue areas and maintains viability of the cells for several weeks after implantation.

[0102] A plurality of collagen sheets prepared from a collagen gel are stacked on top of each other (laminated) to form a collagen construct. Multiple orifices or pores are then formed in the construct in a suitable pattern to form a porous structure. The pores can be formed using laser, stamping, or any other suitable machining technique.

[0103] Characteristics of the pores are tuned to modulate mechanical properties of the construct and provide tissue structure. The range of sizes of the pores can vary. For example, the pores can have a size of less than about 30 μιη (micrometers) (e.g. , for drug and growth factor delivery), from about 30 μιη to about 60 μιη (e.g. , for vascularization), and/or greater than about 60 μιη (e.g. , to promote faster tissue ingrowth). Pores smaller than 30 μιη may facilitate cytokine and growth factor transfer while substantially increasing tensile strength of the construct. This may be beneficial, for example, in tendon repair or other applications where constructs having increased tensile strength are required. For applications where strong and at the same time flexible constructs are required (e.g. , in hernia repair), a construct having larger pores can be used, for example, pores of greater than 60 μιη. Such larger size pores facilitate host tissue integration with enough space for vascular network formation, while maintaining the strength and flexibility of the construct.

[0104] After the construct having pores therein has been formed, an alginate is incorporated therein. The alginate in a liquid form is optionally mixed with cells, cytokines and/or growth factors prior to being incorporated into the collagen construct. The alginate acts as a temporary bioactive interface between the construct and the host environment and as a carrier of cells and bioactive molecules to promote a beneficial host response.

[0105] Many clinical applications of tissue repair and regeneration involve repairing contaminated tissue (e.g. , due to a trauma). For example, abdominal wall defects such as hernia can present a challenging task because of a potentially contaminated surgical field. The described constructs may limit the adverse effects produced by implantation of a material that cannot induce host responses that promote tissue healing in a potentially contaminated surgical field.

[0106] Constructs in accordance with some embodiments can be used in any clinical context to treat various tissue injuries in a subject. Tissue having different boundaries and

geometries, and any degrees of damage can be repaired using the described techniques.

[0107] As used herein, a subject means humans, nonhuman primates, dogs, cats, sheep, goats, horses, cows, pigs and rodents. [0108] The described constructs may be used in a variety of tissue repair applications. For example, the constructs may be used in a hernia repair mesh, grafts for abdominal wall reconstruction, cardiac patch, skin replacement patch, matrices for muscle regeneration, engineered cartilage and bone tissues, tendon replacement patch, open wound meshes, grafts for oral tissue repair, grafts for plastic surgery, matrices for wound healing in the extremities, matrices for venous ulcer repair, grafts for combat related blast injuries and any other applications.

[0109] Properties of the constructs can be adjusted to have desirable strength, mechanical integrity and flexibility while providing enhanced tissue repair. An amount of collagen removed by etching pores therein can be selected such that a resulting construct has appropriate strength and at the same time sufficiently flexible for a particular application. Thus, a pattern, density, number, size, shape and any other characteristics of pores can be selected to achieve desirable mechanical properties of the construct formed from multiple collagen sheets compressed together. In experiments conducted herein, the pores formed in tested constructs were through and through pores. A term through, or "through and through," pores refers to pores that start on one side and penetrate through the entire thickness of the construct. However, in some cases, blind pores, which terminate within the construct material and do not penetrate through the entire thickness of the construct, can be formed additionally or alternatively. Furthermore, the pores formed in constructs in accordance with some embodiments can be any holes, cavities, or openings, as the embodiments are not limited to any particular type of pores.

[0110] It was discovered that pores formed in constructs prepared using the described techniques provide access to the construct for cells of a host, due to internal accessibility created by the pores. The pores additionally or alternatively retain viable cells and/or bioactive molecules within a construct seeded with cells (e.g., human mesenchymal stem cells) and thus promote vascularization and tissue regeneration. As discussed below, in the experiments conducted herein, pores in constructs provided voids for blood vessels formation and tissue ingrowth. The pores also serve as areas of enhanced and controlled drug delivery.

[0111] The construct is fabricated such that its degradation rate matches a rate of new tissue formation. For example, various properties of the collagen can be modulated for this purpose. The main degradation mode of the collagen component is by enzymatic degradation (i.e., by collagenolytic enzymes). The dynamics of collagen turnover influence the structural or mechanical properties of the construct in vivo. The factors that can be modulated to influence the rate of degradation and rate of new tissue formation include collagen concentration, packing density of collagen fibers, thickness of the construct, number of collagen sheets, the number of pores, pore size distribution, addition of growth factors, and host dependent responses. Furthermore, properties of an alginate that is incorporated into a stack of collagen sheets prepared using the described techniques is altered to adjust a degradation rate of the alginate and its mechanical properties to control cell and bioactive molecule delivery from the construct and formation of new tissue within and around the construct.

[0112] Figs. 1 A to IF are schematic illustrations of an exemplary process of fabricating a construct in accordance with one embodiment. As shown in Fig. 1A, a layer of a collagen gel 102 can be placed on a support 104 and can be used to fabricate collagen sheets. In one embodiment, the collagen gel 102 can have a thickness of about 8 mm (2.5 mg mL ^"1 ).

However, it should be appreciated that the collagen gel 102 can have any suitable thickness, as embodiments are not limited in this respect. For example, in some embodiments, the thickness of the collagen gel may be from about 5 mm to about 20 mm, from about 20 mm to about 100 mm, or any other suitable thickness.

[0113] The collagen gel 102 may be dried in a suitable manner. For example, in one embodiment, the collagen gel 102 may be dried for about 24 hours at a temperature of about 140 °C and then dried for another about 24 hours at a temperature of about 37 °C. However, the collagen gel 102 can be dried at other temperatures for different periods of time, as embodiments described herein are not limited in this respect.

[0114] After the collagen gel 102 is dried, it may be rehydrated (for example, with water or a physiologically compatible solution such as phosphate buffered saline. Solutions used to rehydrate the collagen gel may include Lactated Ringer's solution (contains sodium chloride, potassium chloride, calcium chloride, and sodium lactate in sterile water), normal saline solution (0.9% sodium chloride salt in sterile water), and any other clinically suitable solutions that may include, for example, growth factors with or without heparin,

glycosaminoglycans, antibiotics, and/or other bioactive compounds or drugs.

[0115] In one embodiment, after the rehydration of the collagen gel 102, a collagen layer having a thickness of, for example, about 65 μιη can be created. It should be appreciated, however, that a collagen layer having any other thickness can be formed. The layer of the collagen gel 102 may then be cut into multiple collagen sheets. An example of a collagen sheet 106 thus formed is shown schematically in Fig. IB. It should be appreciated that multiple collagen sheets 106 can be formed from the collagen gel layer 102. The collagen sheet 106 can have any suitable dimensions. For example, in one embodiment, the collagen sheet 106 can be shaped as a rectangle having sides of about 25 mm and 15 mm in length. However, it should be appreciated that the collagen sheet 106 can have any other dimensions. Furthermore, the collagen sheet can have a variety of shapes which may be different from a rectangle (e.g. , round, oval, etc.). In addition, the collagen sheet can have irregular shapes, for example, when the resulting construct is intended to repair a wound or other tissue damage having an irregular shape.

[0116] The collagen gel 102 can be cut into separate sheets using, for example, a carbon dioxide (C0 ₂ ) laser. However, any other cutting technology may be used additionally or alternatively.

[0117] After multiple collagen sheets are formed by drying and then cutting the layer of the collagen gel, the hydrated sheets can be placed on a support 108 (e.g. , an acrylic plate or any other support) and can be allowed to dry in a suitable manner. For example, the collagen sheets may be dried at a room temperature for about 48 hours.

[0118] After multiple collagen sheets are prepared for assembly into a construct, the sheets can be stacked on top of each other to thus create a construct 110, as shown in Fig. 1C. Fig. 1C schematically shows that nine collagen sheets may be used to form the construct 110. However, it should be appreciated that any number of collagen sheets may be used to form the construct 110. For example, in some cases, the number of collagen sheets can range from 1 to 10, from 5 to 12, from 10 to 20, from 5 to 100, and/or from 50 to 150 sheets. The number of collagen sheets may be selected based on a desired thickness to allow for tissue repair which can depend on a type of the tissue, type of damage, patient's characteristics, and any other factors.

[0119] Preparing a construct by laminating multiple separate collagen sheets that were previously dried can provide improved cell incorporation into the construct upon

implantation. Small spaces can be formed within stacked sheets forming the construct and cells can penetrate within the construct through such spaces.

[0120] As shown in Fig. ID, the construct 110 may next be placed into a container 112 including a fibril assembly buffer 114 (7.89 mg mL ¹ sodium chloride, 4.26 mg mL ¹ dibasic sodium phosphate, 10 mM Tris, pH 7.4). The construct 110 can be placed in a compression component 116 configured such that the layered collagen sheets are compressed at both sides (e.g. , using plates or other elements made from acrylic or any other suitable material).

[0121] In one embodiment, the construct 110 may be incubated under compression in the fibril assembly buffer 114 for about 48 hours at 37 °C. Different other incubation time periods may be used as well. [0122] After incubation in the fibril assembly buffer, the laminated sheets may be rinsed and allowed to dry (e.g. , on a glass or any other surface). In one embodiment, nine collagen sheets may form a construct having a thickness of about 500 μιη after rehydration. However, it should be appreciated that the construct may include any number of collagen sheets. For example, in some embodiments, the construct can include two or more sheets.

[0123] Fig. IE illustrates that multiple pores can be formed in a construct in accordance with some embodiments, for example, the construct 110, to create a patterned construct 118. The pores may be formed using any suitable technique (e.g. , laser etching, stamping, etc.) and may form any suitable pattern(s) in the construct 110, as discussed in more detail below. FIG. IF illustrates exemplary dimensions of a pore pattern in the construct. As shown in this example, a distance between each pore, either vertically or distally, can be about 0.61 mm, a (vertical) diameter of each pore from the top to the bottom can be about 0.25 mm, and a side length of each pore can be about 0.14 mm. It should be appreciated, however, that these dimensions are shown by way of example only, as pores having any other suitable dimensions may be formed in the construct described herein.

[0124] Fig. 1G is a bright field microscopy image of a top view of an exemplary construct prepared using the described techniques that illustrates that approximately two pores span a distance of 1 mm along a length of the construct. Fig. 1H is a bright field microscopy image of a cross sectional view of the construct of Fig. 1G. Fig. 1H illustrates that, in this example, a thickness of the construct is about 500 μιη. As shown in Figs. 1G and 1H, the pores are formed in the construct may be "through" pores which span the entire thickness of the construct. However, in some embodiments, other types of surface features may be formed in the construct that do not traverse the entire thickness thereof.

[0125] Fig. II is a scanning electron microscopy (SEM) image of a prospective cross- sectional view of a pore in the construct of Fig. 1G, illustrating that a distance between adjacent pores is about 200 μιη. Fig. 1J is a higher magnification SEM image showing a prospective cross- sectional view of the pore in the construct.

[0126] Mechanical properties of constructs having different pore patterns formed therein and constructs without pores were analyzed. Figs. 2A to 2D are images illustrating different pore patterns formed in constructs in accordance with some embodiments. In some embodiments, a pore size can vary from about 5 μιη to about 1000 μιη. The overall empty space due to the presence of pores can vary from about 0.05% to about 30% of a total area of the construct.

2 2

Pore density can vary from about 10 pores/cm to about 1000 pores/cm . [0127] Fig. 2A shows a construct having a pore pattern such that an area of the construct (e.g. , having a length of about 2 mm and a width of about 2 mm) is occupied by

approximately 16 pores. For example, in one embodiment, the pore pattern of the construct shown in Fig. 2A may comprise 240 pores/cm with the pores having a diameter of approximately 250 μιη such that the pores occupy about 10% of a total area of the construct.

[0128] Fig. 2B illustrates a construct having a pore pattern such that an area of the same size (e.g. , having a length of about 2 mm and a width of about 2 mm) is occupied by

approximately four pores which are larger than pores formed in the construct of Fig. 2A. For example, in one embodiment, the pore pattern of the construct shown in Fig. 2B may comprise 60 pores/cm with the pores having a diameter of approximately 500 μιη such that the pores occupy about 10% of a total area of the construct. Accordingly, the pores formed in the construct of Fig. 2B can occupy the same portion (e.g. , 10%) of the total area of the construct as the pores formed in the construct of Fig. 2A.

[0129] Fig. 2C illustrates a construct having a pore pattern such that an area of the same size (e.g. , having a length of about 2 mm and a width of about 2 mm) is occupied by

approximately 16 pores (which can be the same number of pores as those of pores formed in the construct of Fig. 2A) which are smaller than pores formed in the construct of Fig. 2A. For example, in one embodiment, the pore pattern of the construct shown in Fig. 2C may comprise 240 pores/cm with the pores having a diameter of approximately 130 μιη such that the pores occupy about 3% of a total area of the construct.

[0130] In the constructs shown in Figs. 2A, 2B and 2C, a distance between pores can be constant throughout a construct. The pores can be linearly distributed with a staggered position from line to line. For example, in one embodiment, one pore can be surrounded by 6 other pores, 2 pores can be in line and separated from center to center from a middle pore by a distance of about 610 μιη, and 4 pores can be staggered and separated center to center from the middle pore by a distance of about 680 μιη.

[0131] In constructs in accordance with some embodiments, for example, the embodiments shown in Figs. 2A, 2B and 2C, the pores can have a hexagonal or substantially hexagonal shape. As used herein, "substantially hexagonal" can be defined as having a six sided radial cross section, even though some or all of the sides may be curvilinear rather than rectilinear, as in a regular hexagon. The hexagonal pores can be shaped as regular or irregular hexagonals.

[0132] The hexagonal shape of the pores can be preferable in some embodiments, because it provides an efficient use of the area of the construct and allows fine-tuning mechanical properties of the construct in a more controlled manner, as compared to constructs having pores of other shapes formed therein. Furthermore, the use of hexagonal pores allows creating a flexible and durable construct that has an amount of open area (i.e. , the pore space) sufficient for the desirable flexibility and at the same time has the area not occupied by pores that defines the strength of the construct. Because of the nature of a hexagon, in some cases, a larger number of pores can be formed in an area of a construct than in cases where round or other pores are created. Also, hexagonal or substantially hexagonal pores can be formed in many different patterns, which can allow creating different constructs with pore patterns tailored to a specific application. In addition, the hexagonal pores allow efficient use of materials used to make a construct.

[0133] In some embodiments, the constructs can have other shapes, such as substantially circular, substantially oval, regular or irregular rectangles, and any other suitable shapes. As used herein, a term "substantially" can be defined as identifying a shape that does not exhibit a perfect shape but has some deviations from that ideal state. For example, imperfections in the collagen structure and/or a cutting technique used to form pores may result in pores that deviate from a hexagonal or other intended shape. In some embodiments, a shape of a pore can deviate in a range from about 0% to about 25% from a perfect shape (e.g. , a hexagonal shape).

[0134] A diameter of hexagonal pores, as used herein, can be an average diameter of each pore, for example, an average of a sum of minimal and maximal diameters of the hexagonal pore. For example, a hexagonal pore having a side length of about 140 μιη can have a diameter of about 250 μιη.

[0135] It should be appreciated that the constructs shown in Figs. 2A, 2B and 2C are shown by way of example only, to illustrate the constructs which were analyzed. Thus, it should be appreciated that constructs having any suitable patterns of pores of any suitable size can be fabricated using the described techniques. In some embodiments, different portions of a construct may have pores of different sizes which can form different patterns in a construct. Such arrangements may allow creating collagen- alginate implants or meshes for tissue areas having irregular shapes, different degrees of damage, and/or to provide for any other conditions. Also, it should be appreciated that a construct fabricated using the described techniques may be flexible such that it can be folded or otherwise manipulated to conform to localized variations in the tissue in need of repair. In this way, a construct with a uniform or approximately uniform pore pattern can be used to repair tissues having any boundaries and geometries. [0136] Fig. 2D shows a construct having no pores formed therein.

[0137] Figs. 2E-2H are graphs illustrating different mechanical properties of the constructs of Figs. 2A-2D, with error bars representing a standard error of the mean. Exemplary parameters indicating mechanical properties of the constructs measured herein are ultimate tensile strength, Young's modulus, strain at failure, and suture retention strength. However, it should be appreciated that any other suitable parameters can be used to evaluate and adjust mechanical properties of the constructs described herein.

[0138] Ultimate tensile strength (UTS) can be defined as the maximum stress that a material can withstand while being stretched or pulled before failing or breaking. The UTS may be determined by performing a tensile test and recording an engineering stress versus strain. The highest point of a stress-strain curve may be taken as the UTS. The UTS can be defined as a stress, which can be measured as force per unit area or just force. In the embodiments described herein, the UTS can be determined using any suitable techniques as known in the art, and it is measured in the Pascal (Pa).

[0139] The Young's modulus, or elastic modulus, can be defined as a measure of stiffness of a material, such as collagen. The Young's modulus can be defined as the slope of the initial straight portion, e.g., the first 5-10% of strain, of a stress-strain curve. The modulus is the ratio of the change in stress to the change in strain expressed as a fraction of the original length. The Young's modulus has units of Pa (or N/m 2 or m 1 ^» kg ^» s—2 ). For materials that exhibit stress relaxation, the initial elastic modulus can be defined as the elastic modulus for a stress-strain measurement that is performed over a timescale at which there is minimal stress relaxation. The initial elastic modulus can be determined using standard methods available in the art, e.g., by a compression test or rheology.

[0140] Strain at failure may be defined as a measure of how much a specimen is elongated to failure when a tensile test is performed on the specimen. For example, a strain at failure 10% measured in a specimen of length 100 mm, the material will fail when it is elongated 10 mm.

[0141] Suture retention strength may be measured by placing a suture 2 mm from the end of a construct and measuring the force required to dislodge the suture by pulling it at a constant rate in a physiologically relevant condition and rate: 1 mm/s, 37°C. The peak force when the suture ripped through the construct was measured and shown in gram-force (g-f), which is a gravitational metric unit of force that is equal to the magnitude of the force exerted by one gram of mass in a 9.80665 m/s gravitational field. One gram-force is 9.80665 milliNeuton (mN). [0142] Fig. 2E is a graph illustrating the ultimate tensile strength of the constructs shown in Figs. 2A to 2D, Fig. 2F is a graph illustrating a Young' s modulus of the constructs shown in Figs. 2A to 2D, Fig. 2G is a graph illustrating a strain at failure of the constructs shown in Figs. 2A to 2D, and Fig. 2H is a graph illustrating a suture retention strength of the constructs shown in Figs. 2A to 2D. In some embodiments, a pore size can vary from about 5 μιη to about 1000 μιη. The overall empty space due to pores formed in the construct can vary from about 0.05% to about 30% of the total area of the construct. Pore density can vary from

2 2

about 10 pores/cm to about 1000 pores/cm . The ultimate tensile strength of the constructs can vary from about 0.05 MPa to about 10 MPa, and the Young' s modulus can vary from about 1 MPa to about 50 MPa. The strain at failure can vary from about 5 % to about 50 % and suture retention strength can vary from about 40 g-f to about 1000 g-f. Specific pore parameters can be selected based on a clinical application. For example, for hernia repair, the pores can have a size of from about 60 μιη to about 250 μιη microns, with 5-30 % of the area of the construct attributed to pores at a density of from about 50 to about 1000 pores/cm , to facilitate host tissue integration while maintaining the strength and flexibility of the construct.

[0143] Figs. 2E, 2F, 2G, and 2H illustrate that mechanical testing experiments demonstrate a decrease in the ultimate tensile strength as the overall pore percent area is increased. All patterned samples (the constructs of Figs. 2A, 2B, and 2C ) show an increase in suture retention strength compared to the non-patterned construct of Fig. 2D. The construct shown in Fig. 2A (pores having a diameter of 250 μιη, 240 pores/cm , pores occupying 10% of a total area of the construct) exhibits an advantageous increase in compliance and flexibility (e.g. , a decrease in Young's modulus) compared to the non-patterned (Fig. 2B) and other patterned samples (Figs. 2B and 2C) tested. As shown in the graphs, the construct of Fig. 2A exhibits the ultimate tensile strength of above 1 MPa, Young' s modulus of below 8 MPa, strain at failure of below 20%, and suture retention strength of about 60 g-f.

[0144] The improved flexibility of the construct of Fig. 2A may be useful for some tissue repair applications, for example, hernia repair. Although the construct of Fig. 2A can have a decreased ultimate tensile strength, particularly compared to the construct without pores, as shown in the graph of Fig. 2E, this construct beneficially exhibits a retention of strain at failure comparable to that of the non-patterned construct of Fig. 2D. Thus, mechanical properties of a construct can be modulated to select a combination of properties that allow the construct to maintain a desirable strength while having flexibility suitable for implantation into a subject. For example, the construct may be formed having an ultimate tensile strength of 1 MPa and a Young's modulus of 10 MPa. [0145] After a construct has been incubated in the fibril assembly buffer, rinsed, and dried, an alginate can be incorporated into the construct.

[0146] Alginates are versatile polysaccharide-based polymers that may be formulated for specific applications by controlling the molecular weight, rate of degradation and other parameters. Coupling reactions can be used to covalently attach bioactive epitopes, such as the cell adhesion sequence RGD to the polymer backbone. Alginate polymers are formed into a variety of scaffold types.

[0147] Alginate molecules include (l-4)-linked β-D-mannuronic acid (M units) and a L- guluronic acid (G units) monomers, which can vary in proportion and sequential distribution along the polymer chain. In some examples, the alginate can be modified (e.g. , covalently or noncovalently) with a cell adhesive peptide. Exemplary cell adhesive peptides include arginine-glycine-aspartate (RGD), RGDS (SEQ ID NO: 1), LDV, REDV (SEQ ID NO: 2), RGDV (SEQ ID NO: 3), LRGDN (SEQ ID NO: 4), IKVAV (SEQ ID NO: 5), YIGSR (SEQ ID NO: 6), PDSGR (SEQ ID NO: 7), RNIAEIIKDA (SEQ ID NO: 8), RGDT (SEQ ID NO: 9), DGEA (SEQ ID NO: 10), and VTXG (SEQ ID NO: 11). In some examples, the cell adhesive peptide comprises the RGD amino acid sequence.

[0148] Figs. 3A, 3B, and 3C illustrate schematically a process of incorporating the alginate into a construct in accordance with some embodiments.

[0149] A patterned collagen construct schematically shown in Fig. 3 A (e.g. , the construct 118 in Fig. IE or any other construct), can be exposed to alginate. For example, in one embodiment, the alginate can be reconstituted at 2% w/v in media without serum. The collagen construct can be sterilized in 70% ethanol for 30 minutes, and then rinsed with IX PBS three times. The collagen construct can then placed in a well on a mold component (e.g. , Teflon mold (2.5 x 1.5 x 0.5 cm)) and 425 μL· of 1% oxidized- RGD peptide conjugated alginate gel solution (20 mg mL ^"1 ) can be added on the top of the collagen construct, as schematically shown in Fig. 3B. The oxidation of the alginate gel makes the alginate degradable over time by hydrolysis. In some embodiments, the alginate gel can be replaced in about four weeks in vivo. Conjugation with the RGD peptide can allow for cell adhesion, which may facilitate cell migration and cell-gel interaction.

[0150] In some embodiments, the alginate gel can be cross-linked or polymerized (e.g. , using divalent cation, such as Ca ²⁺ , Mg ²⁺ , Sr ²⁺ , Ba ²⁺ , or Be ² ). For example, in one embodiment, to polymerize the alginate gel, a suitable amount (e.g. , 45μί) of CaCl ₂ (100 mM) can be added on top of the alginate in a suitable manner (e.g. , dropwise). The top cover of the mold component can then placed to cover the collagen construct with the alginate. The collagen construct coated with the alginate can be incubated in the mold component for a suitable period of time (e.g., about 20 min) at room temperature to allow the alginate to polymerize. The collagen construct can then be turned over and placed in the well in the mold component and the process of incorporation of the alginate can be repeated for the other side of the construct. Fig. 3C illustrates schematically a resulting collagen construct having alginate embedded therein at both sides thereof. The alginate can distribute around the entire surface of the construct including pores formed therein.

[0151] In some embodiments, prior to incubating the collagen construct with the alginate, the alginate may be mixed with cells, growth factors and/or cytokines. Thus, the alginate component may act as an interphase layer that facilitates a transfer of cells, growth factors, and cytokines from and to the construct during healing of a damaged tissue. It was shown that the alginate can serve as a structural scaffold for the initial vascularization of the construct in vivo.

[0152] The cells that can be incorporated into the alginate gel may be, for example, mesenchymal stem cells (MSCs), which are multipotent stromal cells that can differentiate into a number of different cell types, including osteoblasts, adipocytes, and chondrocytes. In some embodiments, MSCs comprise multipotent cells derived from non-marrow tissues, e.g., umbilical cord blood, adipose tissue, adult muscle, corneal stroma, human peripheral blood, amniotic fluid, or dental pulp of deciduous baby teeth.

[0153] In some cases, MSCs release proteins such as, for example, vascular endothelial growth factor (VEGF), platelet-derived growth factor-B (PDGF-B), monocyte chemotactic protein 1 (MCP-1), stromal cell derived factor 1 (SDF-1), tumor necrosis factor-inducible gene 6 protein (TSG-6), interleukin-6 (IL-6), interleukin-8 (IL-8), basic fibroblast growth factor (bFGF or FGF-2), insulin-like growth factor 1 (IGF-1), hepatocyte growth factor (HGF), thrombospondin-4, secreted frizzled-related protein 2 (Sfrp2), matrix

metalloproteinase 9 (MMP-9), tissue inhibitor of metalloproteinases (TIMP) metallopeptidase inhibitor 2 (TEVIP-2), thrombospondin 1 (TSP-1), chemokine (C-X-C motif) ligand 6

(CXCL6), or interferon gamma- induced protein 10 (IP-10).

[0154] Additionally or alternatively to adding cells to the alginate, various growth factors regulating cell growth and division can be added to the alginate. For example, the alginate can be mixed with vascular endothelial growth factor (VEGF), platelet-derived growth factor- B (PDGF-B or PDGF-BB), monocyte chemotactic protein 1 (MCP-1) or any other growth factors including basic fibroblast growth factor (bFGF), insulin growth factor (IGF), Bone morphogenetic proteins (BMPs), and Nerve growth factor (NGF). (Zhao, S. et al. Biological augmentation of rotator cuff repair using bFGF-loaded electrospun poly(lactide-co-glycolide) fibrous membranes. International Journal of Nanomedicine. 2014:9 2373-2385; Song, Y.H. et al. The therapeutic potential of IGF-I in skeletal muscle repair. Trends in Endocrinology and Metabolism. 24(6):310-319; Carreira, A. C. et al. Bone morphogenetic proteins:

structure, biological function and therapeutic applications. Archives of Biochemistry and Biophysics. 2014 July 17; Aloe, L. et al. Nerve growth factor: from early discoveries to the potential clinical use. Journal of Translational Medicine 2012, 10:239-254). In some embodiments, the alginate can additionally or alternatively include cytokines, such as, for example, granulocyte-colony stimulating factor (G-CSF), and interleukin 10 (IL-10)

(Marmotti, A. et al. Human cartilage fragments in a composite scaffold for single-stage cartilage repair: an in vitro study of the chondrocyte migration and the influence of TGF-bl and G-CSF. Knee Surgery, Sports Traumatology, and Arthroscopy 2013, 21(8): 1819-1833; King, A. Regenerative wound healing: the Role of Interleukin- 10. Advances In Wound Care 2014, 3 (4):315-323).

[0155] In one embodiment, the cells can be suspended in an alginate gel and then the alginate gel including the cells at an exemplary concentration of 4 xlO ⁶ cells mL ^"1 can be poured over the collagen construct as described above, which can be done in the same manner as when the alginate without cells is poured over the collagen construct. The cells can also have a concentration of at least about 1 xlO ⁶ cells mL ^"1 , at least about 2 xlO ⁶ cells mL ^"1 , at least about 3 xlO ⁶ cells mL ^"1 , at least about 4 xlO ⁶ cells mL ^"1 , at least about 5 xlO ⁶ cells mL ^"1 , at least about 6 xlO ⁶ cells mL ^"1 , at least about 7 xlO ⁶ cells mL ^"1 , at least about 8 xlO ⁶ cells mL ^"1 , at least about 9 xlO ⁶ cells mL ^"1 , at least about 1 xlO ⁷ cells mL ^"1 , and any other suitable concentration.

[0156] In some embodiments, the alginate, either with or without cells, can be adapted to modulate a rate of release of cells, cytokines, and growth factors from the construct, modulate a rate of degradation of the alginate, and modulate the dynamics of other processes involved is tissue repair. A degradation rate of the alginate and its mechanical properties {e.g., rigidity and elasticity) relating to its ability to support cells in a bioactive condition in the construct can be modulated, for example, by altering a molecular weight of the alginate, a ratio of high molecular weight (HMW) and low molecular weight (LMW) alginates, a density of conjugated peptides, a degree of oxidation of the alginate, and any other properties of the alginate. For example, a higher molecular weight alginate that has a decreased number of reactive positions can contribute to a slower degradation rate of the alginate. A HMW alginate can provide better mechanical properties and can therefore be useful in applications where stronger constructs are desirable. A LMW alginate can degrade and clear from the host organism faster. Thus, different ratios of LMW and HMW alginate can be used to control the extent to which MSCs differentiate and accumulate matrix. Alginates can be partially oxidized to create acetal groups that are susceptible to hydrolysis.

[0157] In some embodiments, a molecular weight of the alginate can vary from about 1 x 10 ⁴ g mol ^"1 to about 1.0 x 10 ⁶ g mol ^"1 . In one example, a molecular weight of a HMW alginate may be about 3 x 10 ⁵ g mol ^"1 and a molecular weight of a LMW alginate may be about 5.0 x 10 ⁴ g mol .

[0158] In some examples, a molecular weight of HMW alginates can range from about 100,000 g mol ^"1 to about 1,000,000 g mol ^"1 , with a preferred range being, for example, from about 100,000 g mol ^"1 to about 300,000 g mol ^"1 . A molecular weight of LMW alginates can range from about 10,000 g mol ^"1 to about 100,000 g mol ^"1 , with a preferred range being, for example, from about 10,000 g mol ^"1 to about 75,000 g mol ^"1 .

[0159] The ratio of LMW to HMW alginates used to form gels can be varied while maintaining the ability of the alginate to form gels. The rate of degradation of the construct can be controlled by both the oxidation and the ratio of high to low molecular weight alginates. In some examples, the HMW:LMW alginate ratios can range from about 90: 1 to about 10:90. In other examples, the HMW:LMW alginate ratios can range from about 1:99 to about 99: 1. In some examples, a preferred range of the HMW:LMW alginate ratios can be from about 1:50 to about 1:2. In other examples, a preferred range of the HMW:LMW alginate ratios can be from about 1:3 to 1:4. In other examples, the HMW:LMW alginate ratios of from about 1:3 to about 1:4 and 1% oxidation may be preferred for in vivo degradation of about 4 weeks.

[0160] In some embodiments, ratios of HMW:LMW alginates may be about 90: 10, about 75:25, about 50:50, or about 25:75. In some examples, the HMW:LMW alginate ratio of about 50:50 can be preferred. The HMW:LMW alginate ratio of 50:50 and 1% oxidation may be preferred for in vivo degradation of about 4 weeks.

[0161] In embodiments where the alginate is mixed with cells, the cells can be distributed on both sides of the collagen construct. Figs. 4A and 4B are fluorescent microscopy images illustrating hMSCs (human mesenchymal stem cells) distribution on both sides of the construct (prior to implantation of the construct) in accordance with some embodiments. Cells were stained with calcein AM (acetomethoxy derivate of calcein, as known in the art). Scale bars visible on lower right corners of Figs. 4A and 4B are 200 μιη. Figs. 4A and 4B, where the cells are shown with a lighter or white color illustrate that the cells can be distributed substantially evenly around a surface of a construct. The collagen component provides the mechanical properties of the construct.

[0162] The incorporation of the pores in the collagen construct allows modifying mechanical properties of the construct and provides discrete sites for both therapeutic delivery and host tissue incorporation. In some embodiments, the cells may remain viable in a pore for a time period of longer than 7 days. Figs. 5A-5E are confocal microscopy images illustrating distribution and release of hMSCs from a pore of a collagen construct prepared using the described techniques at 1 day (Figs. 5A and 5B), 3 days (Figs. 5C and 5D), and 7 days (Figs. 5E and 5F) after an alginate mixed with hMSCs has been incorporated into the collagen construct. These experiments were conducted in vivo. Figs. 5 A, 5C, and 5E illustrate are top views of the pore, and Figs. 5B, 5D, and 5F are three-dimensional views of the pore. Scale bars are in the lower right corner of 5A-5F 200 μιη. Figs. 5A-5F illustrate that a pore of the construct can retain hMSCs for at least 7 days. These data indicate that the construct retains hMSCs for over 7 days, for example, for up to 90 days depending on the alginate degradation rate. Human mesenchymal stem can release growth factors and cytokines that promote wound healing.

[0163] Fig. 6A is a graph illustrating a number of hMSCs per square centimeter in a construct in accordance with some embodiments, and Fig. 6B is a graph illustrating viability of hMSCs in that construct. Error bars represent standard error of the mean. In these experiments, the construct was prepared as described above, and hMSCs were added to an alginate gel in a concentration of 4 xlO ⁶ cells mL ^"1 . Cell viability was assessed on days 1, 2, and 7 after incorporation into the alginate. In these experiments, hMSC-seeded constructs were placed in a digestion solution of collagenase and alginate lyase to isolate the cells from the constructs. Trypan blue solution was added to the cells to identify dead cells. Live and dead cell numbers were obtained by counting cells with a hemacytometer. Cell viability was calculated as the number of live cells divided by the total number of cells and presented as the viability percentage.

[0164] Figs. 6A and 6B illustrate that the construct can have an improved ability to retain cells (e.g., hMSCs) and the cells can remain viable in the construct for at least 7 days. These properties of the construct improve its efficacy when used in a subject to repair tissue, because the construct can provide improved support for new tissue formation for a time sufficient for the tissue to heal.

[0165] Figs. 6C, 6D, and 6E are graphs illustrating release of growth factors by hMSCs in a construct in accordance with some embodiments ("Patch") and release of growth factors by hMSCs grown on a plate ("Plate") in vivo, at days 2 and 8 after an alginate with hMSCs has been incorporated into the construct. Fig. 6C shows release of MCP-1 (monocyte chemotactic protein- 1) (normalized to 10 ⁵ cells), Fig. 6D shows release of VEGF (vascular endothelial growth factor) (normalized to 10 ⁵ cells), and Fig. 6E shows release of PDGF-B (platelet-derived growth factor B) (normalized to 10 ⁵ cells). Error bars represent a standard error of the mean.

[0166] Figs. 6C-6E illustrate that hMSCs in the construct prepared using the described techniques released or secreted consistently larger amounts of MCP-1, VEGF, and PDGF-B than hMSCs cultivated on a plate. As shown in Fig. 6E, the difference was particularly notable for PDGF-B where essentially no release thereof was observed for hMSCs cultured on a plate compared to a higher than 150 picogram (pg) release of PDGF-B from hMSCs in the construct at both days 2 and 8.

[0167] Also, as the graph of Fig. 6C illustrates, larger amounts of MCP-1 were secreted by hMSCs in the construct at both days 2 and 8 and the amount of the secreted MCP-1 at day 8 increased relative to that at day 2. However, the amount of MCP-1 secreted by hMSCs cultured on a plate decreased notably at day 8.

[0168] Although the release of VEGF was comparable for hMSCs cultured on a plate and hMSCs seeded in the construct, the release was more consistent by hMSCs seeded in the construct at days 2 and 8.

[0169] These experiments demonstrate that the construct in accordance with the described embodiments can maintain hMSCs in a condition that promotes secretion of growth factors by hMSCs for prolonged periods of time. These results demonstrate that the constructs facilitate the release of therapeutic growth factors and cytokines by cells during the time that the cells remain incorporated within the construct.

EXEMPLIFICATION OF THE INVENTION

[0170] The embodiments described herein will be further illustrated in the following non- limiting examples.

Materials and Methods

Isolation and purification of monomelic collagen

[0171] Rat tail tendon monomelic Type I collagen was isolated by acid extraction from Sprague-Dawley rats (Pel-Freez Biologicals, Rogers AR) following an adapted procedure from Silver and Treslad, J. Biol. Chem. 1980; 255:9427-33. Frozen rat tails were allowed to thaw at room temperature. Tendons were extracted with sterile pliers and placed in 10 mM HC1 for 4 hours at room temperature (pH 2.0, 6 tendons in 1 L). Soluble collagen was isolated by centrifugation at 30,000 g at 4 °C for 30 minutes, followed by sequential vacuum filtration through 20 μιη, 0.45 μιη, and 0.2 μιη filter membranes. Precipitation of sterile collagen was achieved by addition of concentrated NaCl to a final concentration of 0.7 M with stirring for 1 hour. The precipitated collagen was centrifuged at 30,000 g for 1 hour, and the pellet was re-dissolved in 10 mM HC1 (-150 mL) overnight. The solution was subsequently dialyzed using Spectra/Por Dialysis membrane, 50,000 MW cut-off ) against 20 mM phosphate buffer at room temperature for 8 hours and then at 4 °C for at least 8 hours. This was followed by 10 mM HC1 dialysis and by deionized water dialysis at 4 °C overnight. The solution was frozen and collagen was lyophilized.

Fabrication of collagen sheets and laminated collagen constructs

[0172] Monomeric Type I collagen was dissolved in 10 mM HC1 (2.5 mg/mL) and a gel was cast with a neutralizing buffer (4.14 mg/mL monobasic sodium phosphate, 12.1 mg/mL dibasic sodium phosphate, 6.86 mg/mL TES (n-Tris(hydroxymethyl)methyl-2-aminoethane sulfonic acid sodium salt, 7.89 mg/mL sodium chloride, pH 8.0) at 4 °C in a rectangular mold (10 x 7 x 0.8 cm) for 24 hours. Gels were subsequently removed and allowed to warm to room temperature for 2 h and then incubated at 37 °C for 24 h. Gels were rinsed in deionized water for 4 hours with three water changes and then allowed to dry on a glass substrate overnight. Once the collagen gel was dry, the collagen film was cut into rectangles (25 mm x 15 mm) with a C0 ₂ laser (Universal Laser Systems, Scottsdale, AZ). The films were placed in glass container filled with ddH ₂ 0 for 30 minutes. Nine films were carefully layered on an acrylic plate and were allowed to dry completely. After the layered construct had dried, fiber incubation buffer (FIB, 7.89 mg/mL sodium chloride, 4.26 mg/mL dibasic sodium phosphate, 10 mM Tris, pH 7.4) was added to re-hydrate the construct. Subsequently, another acrylic plate was placed on top of the partially hydrated patch and then secured with screws (Figs. 1A-1D). The compression device was then incubated in FIB for 48 hours at 37 °C to promote fibrillogenesis, with slight tightening after the first 24 hours. Laminated constructs were removed from the compression device and rinsed in deionized water for 4 hours with three water changes and were allowed to dry on a glass slide. The constructs were then

micropatterned with a C0 ₂ laser (Versa laser). Imaging of laminated collagen patches

[0173] Optical and scanning electron microscopy (SEM) was used to image the laminated collagen constructs. For optical imaging, samples were hydrated in PBS for 24 hours and then imaged with a stereoscope (Zeiss Axio Zoom VI 6). For SEM studies, constructs were hydrated in DI water for 24 hours and dehydrated by serial incubation in ethanol/water mixtures from 30% to 100%. Samples were then critical point dried (Auto Samdri 815 series A, tousimis, Rockville, MD), sputter coated with 8 nm of gold (208HR Cressington, Watford, England), and imaged at an accelerating voltage of 10,000 eV using a field emission scanning electron microscope (Ziess Supra 55 FE-SEM, Peabody, MA).

Mechanical testing of laminated collagen constructs

[0174] Collagen constructs were cut using a dog-bone press to yield samples with a gauge length of 13 mm and 4.5 mm width. Tensile testing of collagen patches was done using an Instron 5566 (Instron, Norwood, MA). Samples were preconditioned 15 times to 66% of the average maximum failure strain determined from pilot samples and then tested to failure at 5 mrn/min. Hydrated thickness was measured using optical microscopy for calculation of cross- sectional area. Young's modulus was determined from the slope of the last 4% of the stress-strain curve. Suture retention strength testing was done using sutures (Prolene 4-0) that were passed through 5 mm square nine layer collagen samples, 2.5 mm from the patch edge and the suture fastened to the actuating arm of the Instron and pulled at a rate of 1 mm/s. The maximum force measured before the suture tore out of the patch was recorded as the suture retention strength, reported in grams-force (g-f).

Generation of RGD-derivatized alginate

[0175] Ultrapure alginates (ProNova Biomedical) were chemically modified as previously described by Silva and Mooney, Biomaterials, 2010; 31: 1235-41. Briefly, MVG alginate (M/G:40/60) was used as the high molecular weight (HMW, 2.65 x 10 ⁵ g/mol) component to form gels. Low molecular weig ht (LMW, 8.5 x 10 ⁴ g/mol) alginate was obtained by γ- irradiating HMW alginate with a cobalt-60 source for 4 hours at a dose of 5.0 Mrad (Kong et ah, Biomacromolecules, 2004, 5: 1720-7). Alginate treatment with sodium periodate (Sigma) for 17 hours in the dark at room temperature oxidized 1% of the sugar residues in the polymer. An equimolar amount of ethylene glycol (Fisher) was added to stop the reaction, and the solution was then dialyzed (MWCO 1000, Spectra/Por) over three days. Following oxidation, the adhesion peptide sequence GGGGRGDSP (Peptides international) was coupled to both the HMW and LMW alginate by carbodiimide chemistry. Following peptide modification, alginate was dialyzed, treated with activated charcoal, filter sterilized (0.22 μιη), freeze-dried, and stored at -20 °C.

Human mesenchymal stem cells

[0176] Frozen vials of passage one human mesenchymal stem cells (hMSCs) from bone marrow aspirates from the iliac crest were obtained from the Center for the Preparation and Distribution of Adult Stem Cells (Texas A&M, 5701 Airport Rd, Temple, TX,

http://medicine.tamhsc.edu/irm/msc-distribution.html), which supplies standardized preparations of MSCs under the auspices of an NIH/NCRR grant (P40 RR 17447-06). MSCs were obtained from two different donors (nos. 8001R, 8004L). After a 24 hour recovery period, hMSCs were seeded at a low density (100 cells/cm), incubated in complete culture medium (CCM) and allowed to proliferate to 50 to 70% confluency over 6 to 7 days. hMSCs were cultured in a-MEM medium (Gibco, CA) containing deoxy- and ribonucleosides, supplemented with 16% fetal bovine serum (FBS) (Atlanta Biologicals, Lawrenceville, GA), 100 units/mL penicillin and 100 mg/mL streptomycin (Gibco). Cell cultures were incubated in a humidified 37 °C and 5% C0 ₂ environment. Cells were harvested using 0.25% trypsin in 1 mM EDTA at 37 °C for 2 minutes. The trypsin was inactivated by adding CCM and the cells were washed with phosphate buffered saline (PBS) by centrifugation at 1,200 rpm for 5 minutes. The cells were frozen in a-MEM with 30% FBS and 5% dimethyl sulfoxide at a concentration of 1 x 10 ⁶ cells/mL. Only passages two or three were used to initiate experiments.

Fabrication of alginate-collagen composites

[0177] Partially oxidized and RGD-derivatized, high and low molecular weight alginates were mixed at a 1: 1 ratio in media without serum (2% w/v) to provide a low viscosity solution that facilitated impregnation of the collagen construct with uniform cell distribution throughout the pores. Collagen constructs were first sterilized in 70% ethanol for 30 min and then rinsed three times with lx DPBS. The collagen construct was placed on a sterile gauze to remove excess lx DPBS and then placed in a well on a Teflon mold (2.5 x 1.5 x 0.5 cm). A total of 425 μΐ _^ of RGD-derivatized, 1% oxidized alginate (20 mg/mL) solution was added on top the collagen construct followed by 45 μΐ _^ of CaCl ₂ (100 mM) to crosslink the alginate. The Teflon mold was covered and the top of the mold secured with screws. The coated construct was molded for 20 minutes at room temperature. The coated construct was then turned over, placed in the well on the Teflon mold, and the process repeated. Cells were incorporated within the RGD-alginate coated collagen constructs by suspending hMSCs (passage 3) in the RGD-alginate solution at a concentration of 4 xlO ⁶ cells/mL prior to placing alginate solution on the collagen construct. Seeded constructs were removed from the mold and placed in a 12 well plate with 2 mL of complete culture media for in vitro studies or in HBSS buffer prior to surgical implantation.

In vitro assessment of hMSCs within alginate-collagen composites

[0178] Cell viability was initially analyzed by calcein AM and ethidium homodimer staining (Life Technologies, Carlsbad, CA) at 1, 3, and 7 days after impregnation of collagen constructs with cell containing alginate gels. Briefly, constructs were washed with HBSS buffer and then incubated for 30 min with HBSS containing calcein and ethidium

homodimer. Cells were imaged using both a stereoscope (Zeiss Axio Zoom VI 6) and by confocal microscopy (Leica SP5 X MP Inverted Confocal Microscope). As a complementary approach, cell viability within the collagen-alginate constructs was also assessed by direct cell isolation at 1, 2, and 7 days. Briefly, hMSC- seeded constructs were placed in a solution of collagenase (1 mg/mL) and alginate lyase (250 μg/mL) prepared in serum free a- MEM media (5 mL/cm ) for 90 minutes at 37 °C with shaking, which was typically associated with patch dissolution. Cells were centrifuged for 5 minutes at 400 g, supernatant removed, and fresh media added to the pellet to re-suspend the cells. The number of live and dead cell and percent viable cells was obtained using Trypan blue exclusion by counting cells with a hemocytometer.

[0179] Analysis of cytokine secretion was performed on days 2 and 8 after cell incorporation within the construct. Complete culture media was added for collecting conditioned media. hMSC conditioned media was collected from cells cultured on tissue culture plastic and within alginate-collagen constructs. A total of 1.2 x 10 ⁵ cells were either plated directly in a well or added to a well within a composite patch (5 mm x 5 mm) in a 6 well plate.

Conditioned media was collected on days 2 and 7, after washing the construct twice in HBSS, followed by a 24 hour incubation period in 3 mL of collection media at 37 °C. On the day of collection, the media was collected, sterile filtered, and stored at -20 °C. The media was concentrated 5 times and analyzed for three different growth factors and cytokines

(MilliplexR multiplex assay, EMD Millipore, Billerica, MA). Data was normalized to cell number.

Rat abdominal wall repair model

[0180] Alginate-collagen composite constructs were evaluated in a full thickness abdominal wall defect (2 cm x 1 cm). Female Wistar rats (250 g) were repaired with either an acellular or hMSC-seeded alginate-collagen composite patch, as approved by the Beth Israel

Deaconess Medical Center Institutional Animal Care and Use Committee. Anesthesia was induced and maintained with isoflurane inhalation (2.5% and 1.5%, respectively). A 3 to 4 cm vertical midline incision was used to expose muscular and fascial layers followed by creation of a full thickness, rectangular (2 x 1 cm), ventral abdominal wall defect. The defect was repaired with a planar construct using an onlay technique without relaxing fascial incisions. The skin was closed and animals closely monitored for 1 to 2 hours, and then daily. Samples were retrieved at 2, 4, and 8 weeks for histological analysis and measurement of integration strength at the host-implant interface. All studies were approved by the BIDMC Animal Care and Use Committee.

Immunohistochemistry

[0181] Specimens were fixed overnight in 10% neutral buffered formalin, processed for paraffin embedding, and 5 μιη sections obtained and stained for extracellular matrix

(Masson's trichrome), macrophages (CD68, iNOS), and endothelial cells (VWF) (Abeam, Cambridge, MA). Blood vessels were analyzed by counting vessels that stained for VWF in 18 random fields of view at 20x magnification for all samples in each group using Image J (2 weeks samples: n = 3-4 animals/group; 4 weeks samples: n = 34 animals/group; 8 weeks samples: n = 6-7 animals/group). Macrophage (CD68) infiltration was measured in 12 random fields of view of three to four sections per sample using Image J (2 weeks samples: n = 3-4 animals/group; 4 weeks samples: n = 3-4 animals per group; 8 weeks samples: n = 6-7 animals per group). Data were presented as area covered per field of view at 20x

magnification. The ratio of 1NOS/CD68 staining was obtained in serially collected sections stained with CD68 and iNOS, respectively, by comparing CD68 rich areas to the same iNOS stained area in three fields of view for at least three sections per sample (2 weeks samples: n = 3-4 animals/group; 4 weeks samples: n = 3-4 animals/group; 8 weeks samples: n = 6-7 animals/group). Strength of host tissue-construct integration

[0182] To measure the strength of integration, 4 x 20 mm strips of patch and adjacent tissue were excised and mounted on opposing platens of a uniaxial tensile tester (DMTA V, Rheometric Scientific, Piscataway, NJ) to determine tension at failure. hMSC tracking after implantation of cell containing constructs

[0183] hMSCs were incubated in a 12 μΜ solution of carboxyfluorescein diacetate succinimidyl ester (Vybrant CFDA SE Cell Tracer kit; Life Technologies, Carlsbad, CA) in PBS for 15 minutes at 37 °C, followed by incubation in media for 30 minutes at 37 °C. Labeled cells were incorporated in the composite constructs as detailed previously and cell- seeded constructs used to repair a full thickness abdominal wall defect in Wistar rats.

Animals were sacrificed at 2 h, 3 days, 1 week, 2 weeks, and 6 weeks. Tissue was collected, fixed in 10% formalin, placed in OTC, frozen in liquid nitrogen and stored at -80 °C.

Samples were cryo sectioned, placed on a glass slide, stained with DAPI (SlowFadeR Gold Antifade Mountant, Life Technologies, Carlsbad CA) and imaged using a confocal microscope (Leica SP5 X MP Inverted Confocal Microscope). hMSCs were analyzed by counting stained cells in 12 random fields of view at 20x magnification for each time point group using Image J (n = 1 animal/group).

Statistical analysis

[0184] Mean values and standard deviation was obtained for all measurements, image analysis, and mechanical data. Comparisons were performed using the Student's t-test for unpaired data, ANOVA for multiple comparisons, and Holm's post hoc analysis for parametric data. Values of p < 0.05 were considered statistically significant.

Results

Fabrication of collagen sheets and multilayer constructs

[0185] We developed a mechanically robust collagen construct for tissue repair that can be easily tuned and modified for different applications. In fabricating the construct, 8 mm thick collagen gels were initially cast for 24 hours (Fig. 1A), rinsed, dried overnight on a glass substrate (Fig. IB), and 25 mm x 15 mm rectangular sheets cut with a C0 ₂ laser. Rehydrated collagen sheets were 65 μιη thick and a total of nine hydrated sheets layered on an acrylic plate were allowed to dry to form a single multilamellar construct, which was then placed in a fibril incubation buffer within a compression set up that promoted physical bonding (Figs. 1C-1D). After rehydration, the multilamellar collagen construct had a measured thickness of 500 μπι.

Patterning of constructs and mechanical properties of collagen constructs

[0186] Through and through pores were patterned into the construct using a C0 ₂ laser to tune mechanical properties and facilitate integration of host tissue (Fig. IE). In one patterned format, 0.05 mm hexagonal pores were generated with a side length of 140 μιη (Fig. IF). Scanning electron microscopy demonstrated that the samples remained laminated after pore formation (Figs. II- 1 J). The mechanical responses of three different pore patterns were analyzed and compared to a non-patterned construct. Specifically, constructs with 0.05 mm

9 2 2

pores (240 pores/cm ) (Fig. 2A) or 0.20 mm pores (60 pores/cm ) (Fig. 2B), each covering

10% of the total surface area, and constructs with 0.013 mm 2 pores (240 pores/cm 2 ) (Fig. 2C), covering 3% of the total surface area were analyzed. The distance between pores remained constant throughout the construct with pores linearly distributed with a staggered position from line to line. Each pore was surrounded by six other pores.

Mechanical testing of collagen constructs

[0187] Ultimate tensile strength decreased as the percentage of the surface area occupied by pores increased, but ranged between 1.39 ± 0.15 MPa and 2.19 ± 0.18 MPa for the porous constructs (Fig. 2E). The Young's modulus showed little difference between the non-porous and the majority of patterned porous constructs (Fig. 2F). The exception was a construct patterned with 0.05 mm 2 pores (240 pores/cm 2 , 10% surface area), which had a significantly lower Young's modulus (7.72 ± 1.1 MPa vs. 13.02± 2.2 MPa for the non-patterned construct, p < 0.05), while maintaining a strain at failure that was comparable to the non-patterned sample (p = 0.2; Fig. 2G). Suture retention strength was greater for all patterned samples when compared to the non-patterned construct (patterned constructs: 60.5-55.1 g-f vs. non- patterned constructs: 38.3 g-f, p < 0.05; Fig. 2H). The enhanced flexibility along with acceptable tensile strength (1.39 ± 0.15 MPa), failure strain (26.0 ± 5%), and suture retention strength (60.5 ± 5.3 g-f) of the sample patterned with 0.05 mm 2 pores (240 pores/cm 2 ) were seen as advantageous for hernia repair and, therefore, this construct was selected for subsequent studies. Characterization of a cell-populated, alginate-collagen composite construct

[0188] The porous collagen construct was embedded within a partially oxidized and RGD- derivatized alginate gel containing hMSCs using an in-house fabricated Teflon mold (Figs. 3A-C). Uniform cell distribution was observed over both sides and within the pores of the construct (Figs. 4A-4B and 5A-5F). Cells were released following 1, 3, and 7 days in culture with dead cells identified by trypan blue staining. Overall loading efficiency was 72 ± 11%. Over the seven day period, little proliferation was noted and a high level of cell viability (91%) maintained (Figs. 6A-6B). Consistent with these findings, an analysis of the secretome revealed that VEGF, PDGF-β and MCP-1 were produced at a relatively uniform rate from hMSC-populated constructs over the seven day period. hMSCs cultured on tissue culture plates displayed decreased levels over time or very low levels throughout, particularly in the case of MCP-1 and PDGF-β (MCP-1 day 8: construct 2867 pg vs. plate 509 pg, p < 0.0002; PDGF-B day 8: construct 163 pg vs. plate 0.0 pg, p <0.001; Figs. 6C-E).

Evaluation in an abdominal wall full thickness defect rat model

[0189] Alginate-collagen composite patches, with or without hMSCs, were implanted in a full thickness abdominal wall defect (1 cm x 2 cm) in the rat using an on-lay technique. Fig. 7 A shows an image the defect induced in a rat, while Fig. 7B shows an image of the defect repaired using a construct of the invention.

[0190] All composite constructs prevented hernia recurrence over an 8-week period, as is illustrated by Fig. 8A showing an image of a Wistar rat showing that no hernia is observed in animals repaired with acellular and with cell-seeded constructs.

[0191] Figs. 7C, 7D, and 7E are images showing a subcutaneous view of the abdominal wall defect 2, 4, and 8 weeks after the surgery using the construct seeded with hMSCs. Figs. 7F, 7G, and 7H are images showing a subcutaneous view of the abdominal wall defect 2, 4, and 8 weeks after the surgery using the construct without hMSCs . Figs 7C-7H illustrate that de novo tissue forms on top of a preserved patch, particularly after 4 weeks and 8 weeks, when the side of the construct interfacing with the subcutaneous tissue is examined.

[0192] Figs. 8B, 8C, and 8D are images showing a peritoneal view of the abdominal wall defect 2, 4, and 8 weeks after the surgery using the construct seeded with hMSCs. Figs. 8E, 8F, and 8G are images showing peritoneal view of the abdominal wall defect 2, 4, and 8 weeks after the surgery using the construct without hMSCs. [0193] These images demonstrate formation of vessels penetrating the construct and fatty tissue on the side facing the peritoneal cavity. Adhesions to the viscera were not observed.

[0194] These images illustrate that tissue repair is improved when the constructs seeded with hMSCs are used, as compared to when the constructs without hMSCs are used. None of the samples exhibited re-herniation at any of the studied time points. The collagen construct and sutures are more visible at weeks 2 and 4 than at week 8, and enhanced integration with host tissue was observed for the hMSCs-seeded constructs relative to the constructs without hMSCs. At week 8, the collagen constructs exhibit integration with host tissue. The strength of integration at 8 weeks was 0.9+0.20 N mm-1 for cell-seeded samples and 0.6+0.25 N mm- 1 for acellular samples. An increase in strength of integration is observed as the host incorporates additional tissue into the construct. This newly developed cellularized tissue is maintained by a blood vessel network, which developed within the construct.

Histological Analysis

[0195] Histological analysis was performed at 2, 4, and 8 weeks after implantation. The results of the histological analysis are shown in Figs. 9A and 9B. Fig. 9A is a set of histological images showing Masson's Trichrome staining of extracellular matrix and blood vessels formed on a construct. Subcutaneous, middle and peritoneal views of a cross-section of constructs without cells and a cross-section of constructs seeded with hMSCs are shown at 2, 4, and 8 weeks after implantation. Scale bars represent 200 μιη.

[0196] Fig. 9B is a set of histological images showing vWF staining of extracellular matrix and blood vessels formed on a construct. Middle views of a cross-section of constructs seeded with hMSCs and constructs without cells are shown at 2, 4, and 8 weeks after implantation. Scale bars represent 200 μιη.

[0197] Fig. 9C is a graph illustrating a number of blood vessels per 20x field at 2, 4 and 8 weeks formed in the middle section of a construct. Error bars represent a standard error of the mean.

[0198] Fig. 9A demonstrates that at 2 weeks, constructs were present with alginate observed on both the subcutaneous and peritoneal sides of the implants, and host cell responses noted to acellular and cell-seeded samples. New blood vessels were also observed in the subcutaneous side, the peritoneal side, and in the pores of the construct. At 4 weeks, most of the alginate component has been replaced by new tissue and maintenance of a vascular network is observed. At 8 weeks, the collagen component of the construct remains, although there appears to be partial replacement with collagen produced by the host. Overall, acellular and cell- seeded constructs were vascularized by two weeks and remained vascularized over the 8 week study period. Fig. 9B demonstrates that there are no differences in the number of blood vessels observed in the subcutaneous side at all the time points analyzed. On the peritoneal side, no difference in the number of blood vessels was observed at 2 and 4 weeks, but an increase in the number of blood vessels in hMSC-seeded samples was noted at 8 weeks (3.5 ± 1.1 vs. 1.8 ± 0.5 VWF/hpf, p = 0.01). However, a significantly greater number of blood vessels were observed in the mid-section of cell-seeded as compared to acellular constructs at all time points (acellular: 1.7-2.1/hpf vs. cell seeded: 2.7-2.1/hpf, p < 0.03; Figs. 9B-9C).

[0199] Acellular and hMSC-seeded patches were analyzed for macrophage infiltration following in vivo implantation. Fig. 10A is an image of the CD68 staining of acellular and hMSC-seeded samples at 2, 4, and 8 weeks. Scale bar 200 μιη. Fig. 10B is a graph showing the amount of CD68 staining per 20x field at 2, 4 and 8 weeks (*p < 0.05, n = 3 - 7). Bars represent standard error of the mean.

[0200] Figs. 10A-10B demonstrate that there are significant differences in macrophage infiltration between acellular and hMSC-seeded patches at all time points. hMSC-seeded samples had increased macrophage infiltration as compared to acellular constructs (acellular: 1570-2530 μηι ² /1ιρΐ vs. cell seeded: 2021- 3630 μπί ² / ιρΐ, p < 0.05).

[0201] To determine the proportion of Ml macrophages in acelluar and hMSC-seeded patches after in vivo transplantation, serially collected sections were immunostained for iNOS (an Ml associated marker) or for CD68 (marker which is common to both Ml and M2 macrophages). Fig. 11, panels A-C are images of iNOS+ to CD68+ stained cells in serial sections samples at 2 weeks (A), 4 weeks (B), and 8 weeks (C), *p < 0.05, n = 3-7. Scale bar 200 μιη. Panel D of Fig. 11 is a graph showing the ratios of iNOS+ to CD68+ stained cells in serial sections normalized to acellular samples at 2 weeks, 4 weeks, and 8 weeks. Fig. 11 demonstrates that there is a statistically significant decrease in the ratio of 1NOS/CD68 for hMSC-seeded samples at all the time points analyzed (p < 0.05).

[0202] Persistence of hMSCs after implantation of hMSC-seeded patches was also investigated by labeling hMSCs with carboxyfluorescein diacetate succinimidyl ester and incorporating the cells within composite constructs. Panel A of Fig. 12A is a series of confocal microscopy cross- sectional images through hMSC-seeded patches at 2 hrs, 3 days, 1, 2, 4, and 6 weeks, with dashed lines representing a pore in the patch. Scale bar 40 μιη. Panel B of Fig. 12 is a graph showing the number of labeled hMSCs over a 42 day period in vivo (n = 1). Panel B of Fig. 12 indicates that hMSCslevels were greatest during the first 3 days after implantation and decreased substantially after 7 days.

Tensile strength of integration

[0203] An increase in the strength of integration was observed from 4 to 8 weeks as host tissue was incorporated into the construct and maintained by new blood vessel formation. Fig. 13 is a graph illustrating a strength of integration of the construct at 2, 4, and 8 weeks after implantation. Error bars represent a standard error of the mean. Cell-seeded constructs had a significantly greater strength of integration at 4 and 8 weeks when compared to the acellular constructs at those time points. At 8 weeks, the tensile strengths of integration were 0.59 ± 0.25 N/mm and 0.96 ± 0.19 N/mm for acellular engineered composite constructs and hMSC-seeded constructs, respectively (p = 0.01). A decellularized hexamethylene diisocyanate (HMDI)-crosslinked porcine dermis (PermaColTM) with 1 mm thickness was also analyzed after an 8-week implant period (n=7). The tensile strength of integration was somewhat greater in strength than the acellular sample, but lower in strength than the cell- seeded construct (0.72 ± 0.29 vs. 0.96 ± 0.19 N/mm, p = 0.1). These differences were not statistically significant. During tensile testing, samples failed either within the patch itself or the abdominal muscle rather than at the tissue material interface.

Conclusion

[0204] Alginate-collagen composites containing human mesenchymal stem cells (hMSCs) effectively bridged a full thickness abdominal wall defect and prevented hernia recurrence in Wistar rats over an 8 week period. De novo tissue and vascular network formation was observed without peritoneal adhesions. Cell-laden constructs displayed improved strength of integration, which correlated with increased neovascularization, increased macrophage infiltration, and a reduced proportion of Ml macrophages. Laser micromachining facilitated the fabrication of porous collagen sheets and provided a convenient means to tailor the mechanical properties of the multilamellar construct, as well as its capacity to harbor cells and locally integrate with host tissue.

Discussion

[0205] Fabrication of a composite construct for stem cell delivery and tissue repair is described herein. The engineered patch comprises a mechanically robust, multilamellar collagen sheet impregnated with an alginate gel containing human mesenchymal stem cells. Laser machining of pores within the collagen component provides a means to enhance construct flexibility and suture retention strength, while facilitating cell delivery and subsequent integration with host-derived tissue. The alginate provides a temporary interphase layer for localized release of hMSC secreted factors during the initial phase of wound healing.

[0206] In the study, the construct consisted of multiple dense collagen films that were physically laminated together to create a strong, but flexible, 500 μιη thick sheet. It was previously shown that collagen concentration does not affect ultimate tensile strength, since increasing collagen mass resulted in an increase in sheet thickness. However, strength does increase after lamination, presumably due to interpenetration of adjacent layers with an observed decrease in layer thickness. Significantly, load induced construct failure was not associated with delamination. Although three empirically selected pore patterns were analyzed in this study, multiple options exist that can be selected to a desired application. It was determined that a particular pattern (0.05 mm 2 pores, 240 pores/cm 2 , 10% surface area) was sufficiently strong while providing suitable flexibility and suture retention strength. Increased suture retention strength for laser patterned as compared non-patterned constructs may be related to the presence of thermally induced crosslinks between collagen fibers and individual lamellae that constitute the multilamellar collagen sheet. As a consequence, this may be associated with a more controlled distribution of locally induced stress patterns that limit crack propagation.

[0207] Bioactive constructs are designed to improve tissue repair by presenting adhesive ligands, as well as soluble signals, such as cytokines and growth factors that guide the host cell inflammatory and reparative response. The incorporation of an RGD-derivatized alginate gel within and surrounding the collagen construct facilitates the delivery of cells in a manner that enhances cell survival and preserves cell functionality. Human MSCs can be generated with well-defined and reproducible phenotypic properties and low passage and low density cultures of these hMSCs are enriched for early progenitor cells. As an isolated cell therapy, hMSCs have been shown to improve hind limb ischemia in a number of animal models due to their capacity to produce VEGF, basic fibroblast growth factor, and other angiogenic factors. The secretome of hMSCs incorporated within the construct demonstrated constant levels of VEGF, PDGF-s and MCP-1 for at least seven days in vitro, while cells in 2-D culture were unable to sustain this response. Of note, the stability of the neovascular response depends on the coordinated interaction of VEGF and PDGFs, which recruit endothelial cells and pericytes, respectively. While MCP-1 promotes a monocyte and macrophage response, these cells, in turn, release multiple angiogenic factors.

[0208] Following repair of a full thickness abdominal wall defect, de novo tissue formation with associated neovascularization was observed for both acellular and cell-seeded constructs. Significantly, more blood vessels were noted within hMSCs-seeded samples as compared to acellular samples, which persisted through the 8 week study period. Likewise, macrophage infiltration was greater among hMSC-seeded, but the ratio of Ml macrophages to total macrophages was lower in these constructs at all time points, which indicates an increase in alternatively activated M2 macrophages. These results are consistent with prior studies that have shown that hMSCs modulate macrophages from Ml to M2 phenotype (Kim et al, Exp Hematol. 2009, 37: 1445-53; Maggini et al, PLoS One, 2010, 5:e9252; Nemeth et al, Nat. Med., 2009, 15:42-9).

[0209] In the constructs studied herein, the collagen component can degrade primarily by enzymatic degradation in vivo. In animal studies conducted herein, it was observed that the host cells infiltrate the construct and begin to deposit additional collagen. Thus, a cycle of collagen degradation and deposition was observed. The dynamics of this collagen turnover may influence the structural and/or mechanical properties of the construct in vivo. An initial collagen amount and density can play a role in the overall structure and mechanical strength of the constructs, particularly at initial stages of host integration. As time progresses, a host response may also impart changes to the strength of the construct. In some embodiments, a concentration of collagen can be increased to delay loss of mechanical strength due to degradation. Furthermore, a number of collagen sheets in a construct can be increased to add strength and provide additional framework for the host cells to regenerate supplementary tissue.

[0210] hMSCs were present in substantial numbers for at least one week after in vivo implantation and then at lower levels thereafter, despite the absence of immunosuppression in what was essentially a xenotransplant model. The results of these studies demonstrate that hMSCs can produce a significant therapeutic benefit without cell engraftment or

differentiation at sites of host tissue injury. The mechanism of the therapeutic effect is likely paracrine signaling through the hMSC secretome during the acute inflammatory phase.

Overall, improved blood vessel formation and a lower ratio of Ml macrophage to total macrophages correlated with a significantly higher strength of integration at four weeks and eight weeks when compared to the acellular constructs. The collagen- alginate constructs provided outstanding mechanical support for the full thickness abdominal wall defect, even though they were 50% thinner than the commercially available patch. This new therapeutic platform can be customized for numerous applications and may be further modified to improve clinical outcomes.

OTHER EMBODIMENTS

[0211] While some embodiments have been described above in conjunction with the detailed description thereof, the foregoing description is intended to illustrate and not limit the scope of the invention, which is defined by the scope of the appended claims. Other aspects, advantages, and modifications are within the scope of the following claims.