IMPLANTABLE VASCULAR GRAFTS

Abstract

Described are methods of making vascular grafts from man-made tubular scaffolds, tubular scaffolds, and methods implanting vascular grafts comprising tubular scaffolds into subjects. The tubular scaffolds of the present invention are made of hydrogel nanofibers that have internally aligned polymer chains and may be cellularized.

Claims

1. A non-cellularized vascular graft comprising: a tubular scaffold comprising a hollow core surrounded by one or more sheets comprising dehydrated or hydrated hydrogel nanofibers having internally aligned polymer chains.

2. The non-cellularized vascular graft of claim 1 wherein at least one of the sheets comprises longitudinally aligned dehydrated or hydrated hydrogel nanofibers.

3. The non-cellularized vascular graft of claim 1 wherein at least one of the sheets comprises circumferentially or otherwise angled relative to the tubular scaffold longitudinal axis aligned dehydrated or hydrated hydrogel nanofibers.

4. The non-cellularized vascular graft of claim 1 wherein the hollow core has an inner diameter in the range of 0.1 mm to 6 mm;

5. The non-cellularized vascular graft of claim 1 wherein the one or more sheets have a combined thickness in the range of 5 nm to 2000 m.

6. A cellularized vascular graft comprising: a tubular scaffold comprising a hollow core surrounded by one or more sheets comprising hydrated hydrogel nanofibers with internal polymer alignment; and one or more cell layers attached to the tubular scaffold.

7. The cellularized vascular graft of claim 6 wherein at least one of the sheets comprises longitudinally aligned hydrated hydrogel nanofibers.

8. The cellularized vascular graft of claim 6 wherein at least one of the sheets comprises circumferentially or otherwise angled relative to the tubular scaffold longitudinal axis aligned hydrated hydrogel nanofibers.

9. The cellularized vascular graft of claim 6 wherein the hollow core has an inner diameter in the range of 0.1 mm to 6 mm;

10. The cellularized vascular graft of claim 6 wherein the one or more sheets have a combined thickness in the range of 5 nm to 2000 m.

11. The cellularized vascular graft of claim 6 wherein the one or more cell layers have a combined thickness in the range of 10 m to 300 m.

12. The cellularized vascular graft of claim 6 wherein the cells comprise endothelial colony forming cells that align in the direction of flow on the tubular scaffold.

13. A method of using a vascular graft to treat vascular damage comprising the steps of: administering a vascular graft of claim 1 to a subject with vascular damage; treating the vascular damage of the subject.

14. The method of claim 13 wherein the vascular graft is administered by vascular bypass surgery.

15. The method of claim 14 wherein the vascular damage is to an artery or vein.

16. The method of claim 13 wherein the vascular damage is caused by a disease or trauma.

17. The method of claim 16, wherein the disease is selected from the group consisting of congenital cardiovascular defect (CCD), coronary artery disease (CAD), or peripheral artery disease (PAD).

18. A mesh comprising sheets comprising dehydrated or hydrated hydrogel nanofibers having internally aligned polymer chains wherein each sheet has a controlled nanofiber orientation that is longitudinal, perpendicular, or otherwise angled.

Description

BRIEF DESCRIPTION OF THE DRAWINGS

[0039] FIG. 1 illustrates a schematic of microvascular development process in three-chamber bioreactor using solid microfibers. Bioreactor is divided into three compartments by two Polydimethylsiloxane (PDMS) walls. Longitudinally aligned fibrin microfibers described in US20150118747 are tethered in the central compartment in between the two walls. ECFCs are seeded in this compartment and cultured for 5 days, after which SMCs are seeded on top and cultured for 10-15 more days. The core is then degraded using a plasmin solution. Not drawn to scale.

[0040] FIG. 2A-2E illustrates a three-chamber bioreactor design. (A) Design schematic and specifications of top and bottom plates of the bioreactor. Top view of (B) complete assembled bioreactor and (C) bottom plate only showing imaging window. (D) Assembled bioreactor with glass slide and PDMS walls dividing the chamber into three separate compartments filled with PBS (left and center) or DMEM (right). Seeding set-up on (E) 5 fibrin microfibers with a cell suspension in the center compartment and PBS in the outer compartments.

[0041] FIG. 3A-3F illustrates microvascular structures. Structures were developed in three-chamber bioreactor on fibrin microfibers with ECFCs for 5 days followed by culture with SMCs for 10-15 days before treatment with fibrin degradation media. (A) Epifluorescence images of whole microvascular structure. Scale=500 m. Confocal projections of structures at (B) low (scale=200 m) (C) mid (scale=100 m) and (D) high magnification (scale=20 m). (E) Orthogonal projections of structures at (I) luminal space, (II) inner cell layer, (III) mid cell layer, and (IV) outer cell layer. Scale=50 (F) High magnification orthogonal projections of structures at (I) mid cell layer and (II) outer cell layer. Scale=20 m. (G) Epifluorescence images of cross-sectional cryo-sections of structures with incomplete degradation. Scale=50 m. Elastin (Eln, red), Col IV (magenta), F-actin (green), and nuclei (blue). Double-headed arrow shows microfiber longitudinal direction.

[0042] FIG. 4A-4G illustrates fabrication and properties of fibrin microfibers and tubes, or cylinders. Fibrin hydrogel microfibers are spun into a sheet (A) by rastering the landing position of the biopolymer jet on a collection solution. (B) Sheets are collected by placing a PTFE coated mandrel on the resultant hydrogel sheet either perpendicular or parallel to fiber orientation. (C) Sheets are then (I) wrapped and (II) the PTFE coated mandrel is removed after the wrapped sheet is dehydrated via lyophilization or ethanol treatment, (III) yielding hollow fibrin tubes with longitudinal or circumferential alignment. (D) SEM micrographs of fiber alignment in (I) circumferentially aligned hollow tube, (II) a stretch dried fibrin fiber bundle lacking open lumen longitudinally aligned hollow tube, cross sectional view showing walls of lyophilized fibrin tube at (III) low and (IV) high magnifications. (E) SEM of (I, III) longitudinally aligned hollow tube alignment and (II, IV) cross sectional views of the open lumen for fibers dried using a (I, II) lyophilization treatment and (III, IV) a graded ethanol treatment, showing improved alignment definition and decreased wall thickness with the later treatment. (F) Mechanical properties, including (I) elastic modulus, (II) strain to failure, (III) axial ultimate tensile strength (UTS,), and (IV) outer diameter, of fibrin tubes dried using the lyophilization (Lyo) and graded ethanol (EtOH) treatments (n=3-5, * p<0.05). (G) Rehydrated EtOH dried fibers show increased (I) wall porosity and (II, III) surface texture as a result of tube swelling upon rehydration.

[0043] FIG. 5A-5D illustrates a single chamber bioreactor design and microvascular development process on fibrin tubes. (A) Longitudinally aligned fibrin tubes are tethered on the 25 Ga needles between the two PDMS walls within the glass chamber. ECFCs are seeded in this compartment and cultured for 5-7 days, after which SMCs are seeded on top and cultured for 5-7 more days. Resulting structures are then perfused using a peristaltic pump. (B) Side view design schematic and specifications of single chamber bioreactor. Not drawn to scale. (C) Top view of completely assembled bioreactor filled with PBS (top) and a cell suspension (bottom). (D) Two parallel perfusion set-ups powered by a peristaltic pump, featuring a media reservoir, air filter, and two single chamber bioreactor chambers.

[0044] FIG. 6A-6H illustrates microvascular structures developed in single chamber bioreactor. Structures at day 7 of ECFC culture (n=3) presented as (A) maximum intensity and (B) 3D cross-sectional reconstruction at low magnification (scale=200 m); (C) maximum intensity projections at mid magnification (scale=100 m); and (D) 3D cross-sectional reconstruction at high magnification. Structures after 7 days of ECFC culture followed by 7 days of SMC co-culture (n=2) presented as (E) stitched maximum intensity projection at low magnification (scale=200 m); (F) maximum intensity projection at mid magnification (scale=100 m); (G) Orthogonal projections at (I) inner and (II) outer cell layers showing ECFC and SMC layers, respectively (scale=50 m); and (H) 3D cross-sectional reconstruction at mid magnification. All 3D reconstructions depict a luminal cross-sectional view of the microvascular tube structures at a slight angle to enable visualization of the lumen and outer surface. Double-headed arrow shows microfiber longitudinal direction. Yellow arrowheads indicate cells underneath outer cell layer.

[0045] FIG. 7A-7C illustrates perfusion through engineered microvasculature. Hollow fibrin microfibers cultured with ECFCs for 5 days followed by co-culture with vSMCs for 5 more days and further cultured under static conditions or with luminal perfusion for 3 days (A) Manual perfusion. Arrow points to leading edge of medium during flow. (B) Confocal projections of ECM in (I) static control (II) 3 day perfusion. Scale bars=200 m. (C) Measurement of lumen diameter. Error bars represent SEM. Significance levels in the distribution represented by ***p<0.001.

[0046] FIG. 8 illustrates tubular scaffolds of the present invention used to create in-vitro vascular structures such as grafts having a hollow core. These tubular scaffolds have been dried to extend shelf life and for shipping.

[0047] FIG. 9. Fiber Dehydration Process for Storage of Hollow Fibrin Tubular scaffolds. The number of layers of fibrin used to wrap the mandrel determines the wall thickness of the hollow tubular scaffolds, which can be easily and precisely altered. Pre-dehydration hollow fibrin tubular scaffolds that are left on the mandrels and in DI water (left). During the graded ethanol dehydration, the hollow tubular scaffolds are left on the mandrel and placed in increasing concentrations of ethanol (center). Finally, once the hollow tubular scaffolds are completely dehydrated, the mandrels are removed and the self-standing tubular scaffolds can be placed in storage or transported (right).

[0048] FIG. 10. Internal cellularization of hollow fibrin tubular scaffolds. Samples at day 3-4 of ECFC culture (n=3) presented as maximum intensity confocal projections at, (A) low (scale=200 m), (B) mid (scale=200 m), peak magnifications (Ci, scale=20 m) (Cii, scale=50 m), and (D) orthogonal projection (scale=50 m). F-actin (green), nuclei (blue), lumen (L), and sTEVG wall (W). Double-headed arrows show microfiber longitudinal direction.

[0049] FIG. 11. Implantation of Acellular Hollow Fibrin Tubular scaffolds in Murine Model. The fibrin tubular scaffold (double headed arrow) was implanted as an abdominal aorta (Ao) interposition graft in a murine model (n=1). The inferior vena cava (IVC) is indicated. The graft was anastomosed to the native abdominal aorta (single headed arrows) and blood flow was observed after the clamps were removed. Pulsation was visible in the graft and artery, which was indicative of arterial flow.

[0050] FIG. 12. sTEVGs at 8 weeks post-implantation. Staining for the various markers indicated (arrowheads indicate elastin fibrils or fibers). Lumen (L), Sleeve (S), and fibrin sTEVG (F) (scale=30 m, n=4).

DETAILED DESCRIPTION OF THE INVENTION

Biodegradable Tubular Microfibers

[0051] The inventors used biodegradable microfibers (described in Zhang Biomaterials article and US20150118747, both incorporated by reference into this patent application) to create tubular scaffolds and sTEVGs. Briefly, solid fibrin microfibers are fabricated by electrospinning a fibrin biopolymer jet onto a thrombin collection solution. Microfiber bundles of varying diameters are collected and dehydrated to create the solid microfiber with controlled outer diameters. Dehydration can include freezing and lyophilization or graded ethanol treatments. These solid microfibers may have alignment of their surface topography, which can be used to culture cells. Once the fibers have been cellularized with the desired cell populations, the solid microfiber fibrin core can be degraded with plasmin to create a hollow structure with cell walls. The sTEVGS of the present invention have been successfully implanted into mammals and may be used to treat vascular disease such as pediatric CCD, CAD, or PAD.

Microvascular Development Protocol in Three-Chamber Bioreactor

[0052] Following the inventor's recent success in creating self-standing luminal multicellular microvascular structures in vitro utilizing fibrin hydrogel microfibers as molds for structure development (Barreto-Ortiz, et al., 2015), the inventors designed a bioreactor device that would allow complete control over culture conditions while having built-in capabilities to perfuse the microvessel. The inventors designed a bioreactor comprised of two polycarbonate plates with a single rectangular chamber inside to achieve this goal (FIG. 1). This space was divided into three separate compartments by two PDMS walls, which also served as anchors, to which a fibrin microfiber was tethered. Each of these three compartments featured two luer lock ports to allow controlled media changes and perfusion. Using this bioreactor and following our recently published protocol (Barreto-Ortiz, et al., 2015) the inventors devised a step-wise process for the generation of an in vitro microvessel model (FIG. 1).

[0053] The bioreactor preparation consisted of placing two PDMS walls in the bioreactor compartment and tethering the fibrin hydrogel microfibers in between the two walls before sealing and sterilizing the complete system. ECFCs were then seeded on the hydrogel microfibers by adding a cell suspension in the middle chamber and tumbling overnight to optimize cell attachment. ECFCs were cultured for 5 days to allow a confluent endothelial layer to form, after which vSMCs were seeded on top in a similar manner. Culture was continued for 10 to 15 more days to obtain a robust tunica media. After this, the system was treated with a plasmin solution to degrade the fibrin hydrogel microfiber core of developing microvessel grafts.

Three-Chamber Bioreactor Design and Implementation

[0054] Two different bioreactor prototypes were designed and tested, each composed of two symmetrical plates enclosing three compartments: a middle compartment to hold the microfibers and serve as the seeding and culture chamber and two outer compartments to serve as inlet and outlet media ports for perfusion. The final bioreactor was designed to have the same dimensions as a standard cell culture well plate (12585 mm) and contain an imaging window in the bottom plate (4215 mm) within an inner compartment fitting a standard microscopy glass slide (7525 mm), as shown in FIG. 2A-C. This enabled live imaging of developing structures within the device, providing the opportunity for detailed monitoring. The bioreactor plates were designed to be held in place by two standard jackscrews and nuts, and vacuum grease was used to seal the microscopy slide to the bottom plate as well as to seal the two plates together. Two indentations were fabricated on each side of the bottom plate to allow leverage between the plates when dismantling the device.

[0055] The inner chamber was designed to have a total height of 11 mm, 8 mm within the bottom plate and 3 mm from the top plate (FIG. 2A). A total of 6 Luer lock ports were built in to allow individual media changes within each of the three separate compartments created after the PDMS walls are placed within the chamber (FIG. 2A-D). As shown in FIG. 2D, this can be done without leaking between the 3 chambers, allowing for detailed control of seeding and culture conditions. Up to five fibrin microfibers were tethered in between the PDMS walls in the middle compartment of the inner chamber by feeding them through plastic tubing traversing each PDMS wall (FIG. 2E). This design would allow the developing cell wall to grow both on the outer surface of the microfiber and, near the PDMS walls, also on the outer surface of the tubing, creating a self-cannulating system for the developing microvasculature (data not shown).

Development of Microvascular Structures with a Tunica Media

[0056] Previously the inventors showed that fibrin microfibers seeded with ECFCs for 5 days and co-cultured for 5 more days with vSMCs created a multicellular structure with a fully invested mural cell layer (Barreto-Ortiz, et al., 2015). However, native microvasculature larger than 250 m in diameter has a tunica media of at least 3 layers (Mulvany, et al., 1978). In the current invention, the inventors cultured vSMCs on ECFC-seeded microfibers for 10-15 days to achieve a multi-layer tunica media before treating the fibrin microfiber core with plasmin for degradation. As shown in FIG. 3A-D, the resulting structures were uniform throughout their full length and express necessary ECM proteins for vascular stability, including Col IV and elastin.

[0057] Orthogonal projections of confocal images showed structures with a circular cross-section, and a vessel wall of 57.83.9 m and inner diameter of 1858.2 m (FIG. 3E-F). Furthermore, these projections revealed the presence of multiple cell layers, starting with a confluent aligned cell layer at the inner wall (FIG. 3E II), a layer with cells exhibiting diagonal projections in the middle (FIG. 3E III, FIG. 3F I), and a confluent aligned cell layer in the outer part of the cell wall (FIG. 3 E IV, FIG. 3F II). Cross-sectional cryo-sections of these structures showcased the circular cross-section of the structures, robust cell wall, and marked deposition of ECM proteins (FIG. 3G). However, the development of a multilayered robust tunica media created a new challenge for complete microfiber degradation with some fibrinolysis products remaining inside the structure's lumen even after plasmin treatment, leading to the development of hollow, fibrin microfibers. The development of microvascular structures occurred step-wise and could be precisely controlled at each step.

Fabrication and Manufacturing of Tubular Scaffolds Comprising Hydrogel Nanofibers Having a Hollow Core

[0058] In microvasculature, it has been shown that vessels collapse in the absence of flow (Berger and Jou, 2000). Indeed, the inventors observed structures would begin losing patency after 10 or more days following coculture with vSMCs, even without complete fibrin microfiber degradation. Longer culture time points resulted in complete diameter reduction and structure failure, and perfusion through developing structures was obstructed by remaining fibrinolysis products. In order to have better control over perfusion and prevent lumen occlusion, the inventors developed hollow fibrin hydrogel tubes with the same aligned nanotopography and bioactive substrate as our fibrin microfibers to give rise to tubular scaffolds that are hollow and used to create sTEVGs. WO 2013/165975, US20150118747, (incorporated by reference) discloses a manufacture of solid hydrogel microfibers. Here the inventors hypothesized that the hollow space of these tubes could be used to apply flow through developing structures at earlier stages of microvascular development, independent of fibrin degradation, thus preventing structure collapse.

[0059] Tubular scaffolds of the present invention including hydrogel nanofibers are made by a different process than hydrogel microfibers described in US20150118747. Instead of creating a bundle of hydrogel nanofibers having polymer chains with an internal alignment, the fibrin hydrogel nanofibers are spun into a sheet by rastering the landing position of the fibrin biopolymer jet back-and-forth across a thrombin collection solution. To create hollow microfibers, the inventors fabricated tubular structures by wrapping the aligned fibrin sheets on PTFE coated mandrels (FIG. 4A-B) (Zhang, et al., 2014). Wrapping the fibrin sheets with their alignment perpendicular to the mandrel's axis resulted in a circumferentially oriented nanotopography on the hollow microfibers, while keeping the fibrin sheet's alignment parallel to the mandrel's axis resulted in longitudinally aligned nanotopography (FIG. 4C). Other angles of fiber orientation may be achieved by altering the angle of the PTFE mandrel relative to the sheet alignment. The resulting alignment was confirmed with SEM for both circumferentially (FIG. 4D I) and longitudinally aligned (FIG. 4D II) microfibers. Detail of cross-sectional morphology are shown in FIG. 4D After wrapping, microfiber sheets were frozen and lyophilized to ease removal of the mandrel from resulting hollow fibrin microfibers. Lyophilized tubular scaffolds had a highly internal porous structure following removal of water but tube walls were very thick (FIG. 4E I-II). Dehydration with ethanol followed by air-drying reduced wall thickness from several hundred m down to 100 m (FIGS. 4E III-IV, F IV). The revised dehydration process using ethanol gradient resulted in reduced tube diameters in both the dry and rehydrated state (gradient drying: 592.417.1 m and 898.741.4 m; lyophilization: 1329.282.4 m and 1186.9120.6 m, respectively), improved tube elastic moduli and ultimate tensile strength (gradient drying: 20.94.5 kPa and 185.051.1 kPa; Lyo: 8.31.3 kPa and 51.127.5 kPa, respectively), and maintained similar longitudinal strain to failure relative to the lyophilized tubes (gradient drying: 128.88.5%; Lyo: 122.215.5%) (FIG. 4F). Rehydrated EtOH-dried fibers show increased wall porosity (G I) and surface texture (G as a result of tube swelling upon rehydration.

[0060] Endothelial colony forming cells (ECFCs) adhere to microfibers used in the present invention (described in Biomaterials journal article (Creating Polymer Hydrogel Microfibers with Internal Alignment via Electrical and Mechanical Stretching, Zhang et al., Biomaterials. 2014 March; 35(10): 3243-3251), and US20150118747, hereby incorporated by reference) that have structural characteristics that align the ECFCs in the direction of flow to form a microvascular structure. Other cells such as mural cells may be random, longitudinal, or circumferentially aligned. The microfiber may be removed from the microvascular structure by enzyme treatment creating a hollow microvascular core. A biodegradable tubular microfiber used in the present invention has a longitudinally aligned nanotopography comprising biodegradable, electrostretched hydrogel polymer fibers with internal polymer alignment. The term longitudinally aligned nanotopography means the nanofibers are aligned longitudinally with each other within a microfiber. The term internal alignment means the polymer chains in a nanofiber are aligned by mechanical and electrical methods. The longitudinally aligned nanotopography means the structure of a microfiber resulting from the longitudinally aligned nanofibers having polymer chains with internal alignment.

[0061] Tubular scaffolds of the present invention are made of the hydrogel nanofibers used to form the microfibers discussed US20150118747. However, the tubular scaffolds of the present invention have a different structure then the US20150118747 microfibers. Specifically, the microfibers described in US20150118747 bundle hydrogel nanofibers to form a solid, without a hollow core, microfiber. The tubular scaffolds of the present invention form a sheet of hydrogel nanofiber, in some embodiments the nanofibers are longitudinally aligned and in other embodiments the nanofibers are random, circumferential, or otherwise angled in a controlled manner, to form a sheet. One or more sheets of these nanofibers is then formed into a cylinder shape having a hollow core creating a tubular scaffold of the present invention. Each sheet may have a separate and distinct alignment of nanofibers in any angle. Some sheets may be random and others aligned. Consequently, the tubular scaffolds of the present invention are designed to novel structures with specific axial and radial strength, and circumferential or longitudinal topography, by layering sheets having the same or different angle alignments or no angle alignment. The tubular scaffolds of the present invention are then used in vascular grafts such as sTEVGs having structural characteristics not seen before.

[0062] Other characteristics of tubular scaffolds include in some embodiments that they clearly exhibit longitudinally aligned nanotopography resulting from the bundling of aggregated polymeric nanofibers with internal polymer chain alignment. In addition, these electrostretched hydrogel tubular scaffolds used in the present invention are mechanically stronger and easier to handle than typical hydrogels of the same composition and dimensions and the electrostretched hydrogel tubular scaffolds exhibited preferential alignment along the nanofiber axis. When the crosslinking mechanisms are compatible, multi-component hydrogel tubular scaffolds can be produced with similar degree of alignment. In addition, the nanofibers making up the tubular scaffolds have a highly porous and aligned surface texture (polymer chains are internally aligned) that is also very different from recently developed fibrin microthreads, which are dense and smooth on the surface.

[0063] The inner diameter of the tubular scaffold with a hollow core can be altered and controlled by changing the diameter of the mandrel used. Multiple layers of sheets can be wrapped around the mandrel to modulate tubular scaffold wall thickness and resulting outer diameter. Additionally, the direction of wrapping can be altered so that the inner layers of the graft have a longitudinal or circumferential alignment, while the outer layers have a circumferential or longitudinal alignment. This can aid in internal and external cellularization of the fibers, enabling control of cellular orientation in both areas. The tubular scaffolds are then dehydrated via lyophilization or graded ethanol treatments and removed from the mandrel yielding hollow fibrin tubular scaffolds with longitudinal, circumferential, or mixed alignment.

[0064] The dehydration of these tubular scaffolds with solid and hollow cores allows them to be stored for extended periods at room temperature or in a fridge (4 C.). While dehydrated or after rehydration, these tubular scaffolds can be shipped to laboratories, hospitals, or other facilities for use to create microvascular structures using the cells of a subject who is suffering from vascular injury or cardiovascular disease. These tubular scaffolds can also be shipped for immediate implantation in a subject who has a severe, emergency condition requiring immediate surgical intervention. Transport in a dehydrated state or in a rehydrated state with sterile solution is possible. In either state, the hollow microfibers can be cannulated on a mandrel or needle. The solid and hollow tubular scaffolds can also be shipped independent of cannulation or other supporting structures in a vial, test tube, plate, well, dish, or other closed container. To facilitate culture, tubular scaffolds can also be shipped in a sterile bioreactor, which would enable the shipment of cellularized or acellular vascular grafts for culture or implantation purposes. Before implanting acellular structures or creating cellularized microvascular structures for implantation or study, the tubular scaffolds should be rehydrated in a reverse graded ethanol treatment. The ethanol treatment further sterilizes the fiber and slowly begins the rehydration process. The fiber is slowly moved to PBS solutions with decreasing ethanol concentrations and rinsed several times to ensure removal of all ethanol before use. After rehydration, the tubular scaffolds with solid and hollow cores are ready for use.

Implementation of Hydrogel Microfibers in Single-Chamber Bioreactor

[0065] sTEVGs were prepared by growing cells around tubular scaffold of the present invention. This process required a custom bioreactor to encase the developing microvascular structures and support perfusion. For this, we designed a simple yet effective single chamber bioreactor composed of rectangular borosilicate glass tubing capped on both ends with custom fitted PDMS walls. These walls could be traversed with large diameter luer lock needles to create media change ports and with small gauge needles to cannulate and perfuse the tubular scaffolds from day zero, without the need to degrade the fibrin core (FIG. 5A-B).

[0066] To assemble the bioreactor, PDMS walls were first custom cut to fit the 1123 mm inner rectangular cross-section of the chamber. Then, 14 and 25 gauge needles were punctured though the top and bottom corners of the PDMS, as shown in FIG. 5B. After all needles were in place, one PDMS wall was placed on one end of the glass chamber and the microfiber was cannulated between the two 25 gauge needles and secured with sutures before closing the second PDMS wall. The chamber was then flushed with ethanol and washed with water or PBS before seeding. Final assembled single chamber bioreactors with microfibers cannulated in between two needles are pictured in FIG. 5C. As shown here, the needles were capped with standard luer lock caps until use, efficiently creating a sealed space within the chamber.

[0067] This new system enabled the step-wise development of vasculature (FIG. 5A), starting with the seeding of ECFCs on a tubular scaffold of the present invention and tumbling the device overnight to optimize cell attachment. After 5-7 days, vSMCs were seeded on top in the same manner and further cultured before perfusion. With this system, there is no need to apply a plasmin degradation treatment since a solid hydrogel microfiber described in US20150118747 is not used, and tubular scaffolds of the present invention can be perfused at any time point by attaching the needles cannulating a tubular scaffold to a closed loop flow system powered by a peristaltic pump (FIG. 5D). Vascular grafts of varying lengths and diameter could be cultured in this bioreactor system by altering the length and gauge of the needle, respectively, used to cannulate the fibrin tubular scaffold.

Fabrication of Perfusable Microvascular Structures or Grafts Using Tubular Scaffolds

[0068] The feasibility of immediately beginning to perfuse or implant the sTEVG by perfusing the acellular tubular scaffolds with the pulsatile perfusion system at various physiologic shear stresses was also assessed. To definitively reveal any perfusion issues, PBS with blue dye (dye concentration: 6.78 ppm) was flowed through the fiber, while clear PBS (dye concentration: 0.00 ppm) was used to fill the chamber, which would drastically change color from leaks. Perfusion for 24 hours at 7 dyn/cm.sup.2 or 14 dyn/cm.sup.2 was successful, with no leaks from the fiber. After 24 hours of high shear stress exposure, the PBS filled chamber was only tinted blue. There was no significant difference in the absolute dye concentration in the chamber between the low and high shear stress conditions until the 24 hour timepoint (0.1600.020 ppm and 0.2000.035 ppm, respectively). The uniform distribution of blue tint in the chamber and the small overall increase in absolute dye concentration after 24 hours showed that only diffusion through the fibrin graft wall occurred and flow could be maintained. With this new fibrin hollow tubular scaffold (or tubular scaffold), there was no need to apply a plasmin degradation treatment and structures could be perfused at any time by attaching the needles cannulating the microfibers to a closed loop flow system powered by a peristaltic pump. Thus, immediate, continuous, and controlled perfusion could be applied without damage to the acellular sTEVG. Additionally, there was no significant difference in the absolute dye concentration in the chamber between the acellular and cellularized conditions after 24 hours of perfusion at 7 dyn/cm.sup.2 (0.1600.020 ppm and 0.1940.069 ppm, respectively). Therefore, no loss of graft functionality occurred with adding cells to the natural scaffold of the present invention and lumen patency of the sTEVGs was maintained

[0069] The inventors were uncertain whether solid hydrogel microfibers described in US20150118747 would provide similar results to the tubular scaffolds of the present invention, given the previous dependence ECM orientation on curvature and the significant increase in curvature for the tubular scaffolds, so they performed the following test. Since the tubular scaffolds had a larger surface area due to an increased outer diameter, the inventors increased ECFC seeding concentration and culture time to maximize cell coverage. Due to this increased surface area, performing two rounds of cell seeding of ECFCs on day 0 and 4 enhanced the formation of a confluent endothelial layer. Given the increased surface area of the tubular scaffolds made out of the microfibers, altering the length of the microfiber used had marked effects on the attainment of a confluent endothelium (data not shown). After 5-7 days of ECFC culture on tubular scaffolds, the formation of an aligned, confluent endothelium expressing mature endothelial markers vWF and CD31 was observed (FIG. 6 A-D). In congruity with our previous results, a diameter of over 450 m resulted in random deposition of ECM proteins Col IV and Fn by ECFCs (FIG. 6 A, C).

[0070] vSMCs seeded on top of ECFC-seeded tubular scaffolds and cultured for 5-7 more days resulted in fully-invested microvascular structures expressing vSMC marker SM22, EC marker CD31, and evidencing both Col IV and elastin deposition (FIG. 6E-H). Similarly to the ECFCs, performing multiple rounds of vSMC seedings every two days resulted in improved multi-layered cellular constructs. At this time point the structures contained two distinct cell layers, with only the outer cell layer expressing SM22 and Elastin (FIG. 6G).

[0071] The inventors then verified the perfusion capability of this new system by culturing ECFCs for 5 days followed by co-culture with vSMCs for 5 more days, after which samples were either maintained in static culture conditions for 3 more days or perfused at 5 mL/hr (about 5 dyne/cm.sup.2) for 3 days. As observed in FIG. 5D and FIG. 7A the structures could be either attached to a peristaltic pump for continuous long-term perfusion or manually perfused, and perfusion can be visualized following the medium flow from the inlet to the outlet port (FIG. 7A, Supp. Video 1). Interestingly, the inventors observed higher elastin deposition after 3 days of perfusion compared to static control, as well as a more aligned Col IV and F-actin organization (FIG. 7B). More importantly, the lumen of microvascular structures under static conditions started collapsing while perfused developing microvasculature maintained a cylindrical cross-section and a diameter larger than its static counterpart (FIG. 7B-C). As in the three-chamber bioreactor, the development of cellularized tubular scaffolds occurred step-wise and could be precisely controlled at each step.

[0072] To enhance the hollow tubular scaffolds of the present invention for implantation, we began to internally seed endothelial colony forming cells (ECFCs) onto the grafts via injection into the luminal space. After 3-4 days of static ECFC culture in tubular scaffolds (Internal, Static), the formation of an aligned, confluent endothelium expressing the mature endothelial markers: vWF, VECad, and CD31 was observed (FIG. 10). The fibrin wall of the tubular scaffold was overexposed in FIG. 10B to clearly demonstrate the cellular attachment to the luminal surface. By seeding the cells via injection into the luminal space or by dropwise seeding on the external graft surface, the seeding became more efficient, requiring both fewer culture days to attain a confluent endothelium and fewer cells to be seeded on day 0.

[0073] Recently the inventors achieved for the first time the generation of self-standing luminal multicellular microvascular structures or grafts such as sTEVG in vitro. However, the final step in the blood vessel maturation process requires the presence of luminal flow. Most groups that have attempted to create microvascular structures or grafts with perfusion have done this by creating void-space channels within a hydrogel or microfluidic device (Khan, et al., 2011, King, et al., 2001, Zheng, et al., 2012). However, none of these are self-standing or have demonstrated the development of both a tunica intima with a continuous endothelium, a robust tunica media, and their corresponding organized ECM proteins. Studies that have presented perfusion of tissue engineered blood vessels have done so in structures above 3 mm in diameter (Niklason, et al., 2001, Williams and Wick, 2004), with perfusable microvasculature below 3 mm remaining a challenge. Similarly, significant advancement has recently been made in the realm of acellular vascular grafts, but most of these vascular grafts have been larger than 4 mm in diameter (Fukunishi, et al., 2016, Syedain, et al., 2016).

[0074] Previous work by Neumann et al. described the use of a perfusion chamber wherein nylon strands were enclosed and used to support the growth of vSMC sleeves in this size range for 21-28 days, after which the nylon strands were mechanically removed, creating perfusable vSMC structures (Neumann, et al., 2003). While this approach allowed for the creation and perfusion of vSMC sleeves in vitro, it did not permit the formation of an aligned endothelium, possibly due to inner cell layer damage upon nylon strand removal (Neumann, et al., 2003). The formation of a confluent endothelium is critical for the reduction of thrombosis risk in vivo and has been a significant focus for small diameter vascular graft development, especially as decreasing diameter is correlated with increased thrombus formation (Brisbois, et al., 2015, Fleser, et al., 2004, Pashneh-Tala, et al., 2015, van Hinsbergh, 2012). For this reason, our bioreactor designs focused on the formation and maintenance of a robust cell wall with a confluent endothelium.

[0075] The inventors' established hydrogel fibrin microfiber system allowed the inventors to study growing microvasculature, and was key in understanding the biochemical and biomechanical cues that guide endothelial cell alignment, vascular smooth muscle cell investment, and organized ECM deposition by each cell type. However, longer culture time points were required to develop a multilayered tunica media, and constructing a perfusion system to support nascent vasculature was critical. To be able to extend the culture time of developing structures, we engineered a custom three-chamber bioreactor with specific design features enabling the controlled, progressive fabrication of microvasculature starting with the inventors' established fibrin microfibers.

[0076] The newly developed bioreactor allowed up to five structures to be grown at the same time, and in situ monitoring through its imaging window permitted the inventors to observe the formation of the microvasculature and optimize culture conditions and time points in order to achieve a robust microvessel wall. Developing structures could also be fixed, stained, and imaged either inside or outside of the device, allowing the analysis of both cell and ECM markers in the developing structures. As such, structures cultured with ECFCs for 5 days followed by vSMC seeding and culture for 10 to 15 more days evidenced a multilayer, multicellular microvascular structure with a robust expression of ECM proteins Col IV and Eln.

[0077] Following the inventors previous microvessel model protocol, the inventors used plasmin to degrade the fibrin core in order to obtain a luminal vessel. Previous results demonstrated this degradation was feasible in structures developed by culturing ECFCs for 5 days followed by 5 days more of vSMC co-culture, a time point where the structures had one layer of each cell type. In the new bioreactor of the present invention, cryo-sections performed on structures cultured with vSMCs for 10-15 days and treated with plasmin for 24 hours revealed a circular lumen and thick vessel wall, but also evidenced remaining fibrinolysis products within the lumen that were not able to fully degrade. As such, longer culture time points after vSMC seeding resulted in more robust microvascular structures, but presented an issue for full fibrin microfiber degradation as the plasmin had to diffuse through a thicker vessel wall to have an effect on the fibrin core. Similarly, fibrinolysis products now had a barrier to diffuse through, and accumulation inside of the vessel wall would prevent flow. Further attempts to strengthen the degradation protocol, including performing several rounds of degradation treatments, were ultimately unsuccessful in completely dissolving the fibrin core (data not shown).

[0078] In developing microvasculature, vessels will collapse in the absence of flow. In the inventors' system, the inventors observed partial lumen occlusion, likely caused by increased cellular weight and vSMC contractility, in structures that had been in culture with vSMCs for more than 15 days and that had partial microfiber degradation. Without full degradation, flow was unachievable and these structures collapsed. In contrast, degrading the microfiber core at earlier time-points resulted in structures that had a distinct lumen but were not robust enough to withstand flow. Therefore, controlling the time-point for perfusion became of critical importance to maintaining lumen stability.

[0079] The hollow tubular scaffolds of the present invention are made of nanofibers described in US20150118747 and have been demonstrated to surprisingly maintain their longitudinal nanotopography even after being wrapped around a mandrel. However the tubular scaffolds of the present invention are structurally different from the microfibers described in US20150118747. The tubular scaffolds of the present invention are made of sheets of hydrogel nanofibers having internal polymer alignment. The nanofibers of a sheet may be circumferentially aligned or aligned in other angles and one or more sheets may be used to make tubular scaffolds having specific suturability and/or strength depending upon the alignment of nanofibers in a sheet, Furthermore, the wall thickness of a tubular scaffold can be easily controlled by varying the number of sheets made of hydrogel nanofibers, including fibrin, that is wrapped around the mandrel and lumen size can be controlled by changing the mandrel diameter.

[0080] In order to create vascular constructs using the tubular scaffolds made of hydrogel nanofibers, the inventors designed a new single chamber bioreactor composed of a glass rectangular chamber sealed on each side by PDMS blocks. The PDMS blocks were punctured with two 14 gauge needles to act as media changing ports and two small diameter needles to cannulate the developing microvascular structures. This allows different size vessels to be generated by simply changing the needle gauge, which also act as the connecting ports for either manual perfusion with a syringe or constant perfusion with a peristaltic pump. This simple design allowed the fabrication of multicellular microvascular constructs with a preformed lumen that when perfused for three days evidenced distinct circular lumen stability and patency, compared to static controls that experienced significant lumen occlusion. This new system is easy to set-up and allows for in situ monitoring of the developing structure. The current prototype allows for the development of one structure per bioreactor at a time. However, several devices can be run concurrently.

[0081] Using this new setup, the inventors achieved for the first time the generation of perfusable, self-standing multicellular vascular structures, or grafts such as sTEVGs, in vitro with an aligned endothelium and full vSMC investment. Perfusion can either be done manually for short-term or with a peristaltic pump for prolonged experiments. Furthermore, perfusion can be conducted at any time during the blood vessel development time-line, opening the door for a wide array of 3D flow experiments in a setting recapitulating the cellular and extracellular organization of native vasculature for the investigation of arteriogenesis. As a proof of concept, the inventors demonstrated that perfusion for three days supported lumen stability in developing structures, compared to static culture conditions that resulted in partial lumen occlusion in a fully cellularized construct (having ECFCs and vSMCs).

[0082] The inventors established a three-compartment bioreactor system and culture protocol, which can be used to generate multicellular microvessels with a robust endothelial vessel layer supported by a fibrin hydrogel microfiber scaffold. The developed construct showed enhanced deposition of ECM proteins Col IV and Eln, as well as a vessel wall composed of three different cell layers. This system allowed structures to be developed in a controlled environment while enabling in situ monitoring, revealing real-time information about microvessel development, including lumen occlusion caused by increased cellular weight and vSMC contractility in the absence of flow. The inventors also developed a novel hollow fibrin microfiber tubular scaffold platform to make vascular grafts including sTEVGs that encompasses the strengths of hydrogel microfibers, while also allowing early stage perfusion through a developing microvessel in order to prevent lumen occlusion. Using this tubular scaffold platform, the inventors successfully generated perfusable multicellular sTEVGs in vitro recapitulating the cell and ECM organization of native vasculature for the first time.

Implantation of Tubular Scaffold as Vascular Graft Including sTEVGs

[0083] Small-diameter tissue engineered vascular grafts (sTEVGs) were made from the tubular scaffolds of the present invention having structural characteristics that mimic topographical and biomechanical features of the ECM. Our tubular scaffolds were prepared from hydrogel nanofibers, such as fibrin and exhibits a microscale, longitudinally aligned surface topography and tunable stiffness. These longitudinally aligned fibrin nanofibers were prepared to form a hollow, tubular scaffold, serving as a vascular graft such as a sTEVG matrix template, as discussed previously. Fibrin was chosen to develop sTEVGs as it has been shown to improve elastin deposition, a critical ECM component for sTEVGs. The unique surface topography induces endothelial alignment with increased ECM deposition, as discussed above.

[0084] 1 mm diameter microvascular grafts were used as a testing ground to allow efficacy studies in an infrarenal abdominal aorta mouse model, which faithfully recapitulates the process of neovessel integration that occurs in large animals and humans, but over a shorter time course. By starting with the smallest, most challenging size, future scale-up to 1-6 mm diameter vascular grafts will be minimally challenging as increased diameter is correlated with decreased thrombus formation and increased patency.

[0085] Importantly, our sTEVG design affords the flexibility to create both cellularized and acellular vascular grafts, depending on the application. Acellular tubular scaffolds, can be used as an off-the-shelf product for emergency vascular operations; while cellularized sTEVGs can be manufactured for CCD populations that do not require emergency procedures. In order to improve patency and reduce thrombus formation, endothelial cells will be uses as a bioactive component to encourage remodeling of the graft by host cell infiltration. By using acellular and cellularized, bioactive sTEVGs using a natural, fibrin scaffold the negative outcomes associated with conventional synthetic sTEVGs is minimized.

Perfusion and Cellularization of Fibrin Tubular Scaffolds in a Single-Chamber Bioreactor

[0086] For cellularized sTEVGs, human endothelial colony forming cells (ECFCs) were internally seeded on the sTEVGs and cultured in a bioreactor for 3-4 days before implantation in a SCID mouse as an abdominal aorta interposition graft (FIG. 11). We focused on 1 mm diameter microvascular grafts for testing in an infrarenal abdominal aorta mouse model for up to 24 weeks (n=4 for each timepoint).

[0087] Dehydrated, acellular sTEVGs were stored for several months before use. Grafts were implanted and ruptured on post-operative day (POD) 21 due to a tear in the fibrin wall along the longitudinal topography emanating from the sutures (n=5). An improved graft composed of longitudinally and circumferentially aligned microfibers (Multidirectional) resulted in a significantly increased mouse lifespan to POD 92 (n=6), but rupture with the same tear propagation pattern from the suture points still occurred. The suture retention strength (SRS) for the Longitudinal and Multidirectional grafts was 436 mN and 49114 mN, respectively, far below the native mouse aorta SRS of 12833 mN (n=2). In all cases, the grafts were able to withstand aortic flow for several days; yet, over time, these natural grafts began to fail due to mechanical inadequacies, as the field would suggest. Using a surgical sleeve as an intervention to prevent tearing at the suture sites resulted in a significantly larger SRS of 20787 mN and survival for up to 24 weeks (n=4).

[0088] Acellular grafts occasionally developed clots on the luminal walls by week 1, which did not appear in the cellularized grafts. At later time points, no evidence of these clots was visible in either acellular or cellularized sTEVGs, suggesting an antithrombotic benefit of ECFCs was acute. By week 8, significant host cell infiltration could be seen throughout the fibrin sTEVG with delamination and fragmentation of the fibrin (FIG. 12 H&E). The regenerating tissue was densely populated with circumferentially oriented SMA-positive smooth muscle cells (SMCs), which had a confluent, luminal lining of CD31-positive endothelial cells (FIG. 12 SMA, CD31). Von Kossa staining indicated that the acellular graft had significantly increased calcification relative to the cellular sTEVG (FIG. 12). Collagen and elastin staining indicated significant, organized ECM deposition similar to, but not as defined as, the mature native aorta (FIG. 12 MT and VVG). To longitudinally assess graft function, color doppler echocardiography was used to visualize blood flow through the graft and measure blood velocity in the distal aorta prior to the iliac bifurcation. Over 20 weeks, both the acellular and cellularized graft groups had no significant difference in peak systolic velocity, end diastolic velocity, or pulsatility and resistivity indices relative to the baseline native aorta, indicating no change in vascular function due to sTEVG implantation. We fabricated a man-made, hollow fibrin tubular scaffold with controlled surface topography, enabling controlled endothelialization or immediate implantation as a potential candidate for arterial bypass surgery.

Kits of the Disclosure

[0089] Any of the compositions described herein may be comprised in a kit. In a non-limiting example, a tubular scaffold of the present invention, that is dehydrated or rehydrated, may be comprised in a kit. The kits may comprise a tubular scaffold of the present invention, preferably a dehydrated tubular scaffold or tube, and a suitable aliquot of one or more reagents to rehydrate the tubular scaffold. Such reagents would include those used in a reverse graded ethanol treatment, for example. The component(s) of the kits may be packaged either in aqueous media or in lyophilized form. The container means of the kits will generally include at least one vial, test tube, flask, bottle, syringe or other container means, into which a component may be placed, and preferably, suitably aliquoted. Where there is more than one component in the kit, the kit also will generally contain a second, third or other additional container into which the additional components may be separately placed. However, various combinations of components may be comprised in a vial. The kits of the present invention also will typically include a means for containing the tubular scaffolds of the present invention and any other reagent containers in close confinement for commercial sale. Such containers may include injection or blow-molded plastic containers into which the desired vials are retained.

[0090] When the components of the kit are provided in one and/or more liquid solutions, the liquid solution is an aqueous solution, with a sterile aqueous solution being particularly preferred. However, the components of the kit may be provided as dried powder(s). When reagents and/or components are provided as a dry powder, the powder can be reconstituted by the addition of a suitable solvent. It is envisioned that the solvent may also be provided in another container means.

Methods/Examples

Cell Culture

[0091] Cells were cultured in a humidified incubator at 37 C. and 5% CO.sub.2 as previously described (Barreto-Ortiz, et al., 2015). Human ECFCs (Lonza, Walkersville, Md.) were expanded on collagen I (BD Biosciences, Franklin Lakes, N.J.) coated flasks in Endothelial Growth Medium-2 Bulletkit (Lonza) supplemented with 10% fetal bovine serum (FBS; Hyclone, Logan, Utah) and used for experiments between passages 7 and 10. Medium was changed every other day and cells were passaged every 5 to 7 days with 0.05% trypsin (Invitrogen, Carlsbad, Calif.).

[0092] Human vSMCs (ATCC, Manassas, Va.) were used between passages 7 and 10 and cultured in F-12K medium (ATCC) supplemented with 0.01 mg/ml insulin (Akron Biotech, Boca Raton, Fla.), 10% FBS (Hyclone), 0.05 mg/ml ascorbic acid, 0.01 mg/ml transferrin, 10 ng/ml sodium selenite, 0.03 mg/ml endothelial cell growth supplement, 10 mM HEPES, and 10 mM TES (all from Sigma-Aldrich, St. Louis, Mo.). Medium was changed every third day and cells were passaged every 5 to 7 days with 0.25% trypsin (Invitrogen).

Preparation of Fibrin Hydrogel Microfibers

[0093] Fibrin hydrogel microfibers were fabricated as previously described (Barreto-Ortiz, et al., 2015, Barreto-Ortiz, et al., 2013, Zhang, et al., 2013). Briefly, 1.5 wt % alginate (Sigma-Aldrich) was mixed in-line with 2.0 wt % fibrinogen (Sigma-Aldrich) at flow rates of 2 ml/h and 1 ml/h, respectively. Both solutions were dissolved in 0.2 wt % PEO (Sigma-Aldrich). A 4 kV electric potential was applied to a 25-gauge needle through which the solution was extruded. The resulting jet was collected in a grounded, rotating bath containing a crosslinking solution of 50 mM CaCl.sub.2 with 10 units/ml thrombin (Sigma-Aldrich) for 35 min. Fibers were left in the crosslinking solution for an additional 10 min and then soaked overnight in 0.25 M sodium citrate to remove alginate from the fibrin fibers. Fibers were then washed in DI water for 60 min, bundled and stretched to 150% of their initial length, and air-dried for 60 min.

Preparation of Longitudinally Aligned Fibrin Hydrogel Tubes

[0094] Fibrin hydrogel microfiber sheets were prepared similarly to the hydrogel microfibers by electrospinning 2.0 wt % fibrinogen solution co-dissolved in 0.2 wt % polyethylene oxide (PEO) in water under the effects of an applied electric field (4.5 kV) to propel the resultant fiber jet across an air gap of 2 cm and onto a rotating collection bath (45 rpm) containing 50 mM calcium chloride and 20 U/mL thrombin. The landing position of the spinning jet was rastered back and forth via use of a linear stage during the spinning step to yield a uniform aligned fibrin sheet. Hollow fibrin tubes with longitudinal alignment were formed by rolling sheets arranged parallel to the fiber orientation onto polytetrafluoroethylene (PTFE)-coated stainless steel mandrels to generate tubes. Tube wall thickness was controlled by altering the number of wraps around the mandrel. Following wrapping, fibrin tubes were further crosslinked in 100 U/ml thrombin for 2 h before lyophilization. Alternatively, fibrin tubes were dehydrated in a series of 25, 50, 60, 70, 80, 90, 95, 100, 100, and 100% EtOH solutions for a minimum of 15 minutes per step and then allowed to air dry. Dried fibrin tubes were removed from the PTFE mandrels following either drying method.

SEM

[0095] Dried fibrin tubes (lyophilized or EtOH treated) were attached to conductive carbon tape on metal stubs and then sputter coated with a 15-nm layer of Au/Pd (Hummer 6.2 Sputter System, Anatech UDA, Hayward, Calif.). Samples were imaged using a JEOL 6700F field emission electron microscope at an accelerating voltage of 5 kV.

Mechanical Testing

[0096] Fibrin tube mechanical testing was done using a Q800 DMA (TA Instruments, New Castle, Del.) under tensile loading conditions in controlled ramp force mode. Hydrated samples were quickly removed from solution and loaded onto the instrument clamps with a preload force of 0.001 N. Tubes were then subjected to increasing force load (ramp rate of 0.05 N/min) until tube failure.

[0097] To determine fibrin tube ultimate suture retention strength (SRS), the sample was mounted onto an electromechanical puller (DMT560; Danish Myo Technology A/S, Aarhus, Denmark). After calibration and alignment, the pins were slowly moved apart using an electromotor at a rate of 100 m/sec to pull 10-0 nylon sutures looped through the vascular graft wall and over the pins, similar to ultimate radial tensile testing methods used previously..sup.4,5

Three-Chamber Bioreactor Design

[0098] Initial 3D designs were created using SketchUp (Trimble Navigation, Sunnyvale, Calif.), while final computer-aided designs were made using SolidWorks (Concord, Mass.). Johns Hopkins Whiting School of Engineering Manufacturing assisted with the final design and manufacturing of the devices.

Three Chamber Bioreactor Preparation and Cell Seeding

[0099] Prior to each experiment, the bioreactor components were autoclaved including top and bottom plates, 6 Luer lock caps (Qosina, Edgewood, N.Y., USA), two jackscrews and nuts, and a 75 mm25 mm glass slide (Fisher Scientific). Two PDMS blocks (1:7 ratio, Sylgard 184, Dow Corning, Midland, Mich.) were cut and fitted to the bioreactor's inner chamber. Vacuum grease (Dow Corning, Midland, Mich., USA) was then applied generously to the bottom of the glass slide and the interface of the two plates, leaving a 1.5 cm barrier around the inner chamber of the device on both plates. All surfaces coated with vacuum grease were left under UV light for 15 minutes. The glass slide was then placed in the inner compartment of the bottom plate and sealed by pressing down firmly across the surface. The PDMS blocks were punctured with a 25 gauge needle and hydrogel microfibers were tethered in between the blocks using polyester shrink tubing (Advanced Polymers, Salem, N.H., USA) through these puncture openings. Fibers were secured in place using 6-0 sutures (Henry Schein, Melville, N.Y.) before sealing the bioreactor. The entire system was sterilized with 75% ethanol and rinsed with sterile water 3 times. After the sterilization process, the exterior compartments were filled with PBS or media and the inner chamber with a cell suspension containing 210.sup.6 ECFCs or vSMCs in ECFC media supplemented with 1% penicillin/streptomycin (Life Technologies). Bioreactors were tumbled for 24 hours to optimize cell seeding and medium was changed every other day thereafter.

Plasmin Treatment

[0100] Fibrin hydrogel microfibers with cells were treated with 9 g/mL plasmin from human plasma (Athens Research & Technology, Athens, Ga., USA) and 2 u/mL alginate lyase (Sigma-Aldrich) in Dulbecco's Modified Eagle Medium (DMEM; Life Technologies) for the time periods specified.

Live/Dead Assay

[0101] Samples were incubated with 2 M calcein AM and 4 M ethidium homodimer (Invitrogen) in PBS for 30 min at 37 C. and 5% CO.sub.2. Samples were imaged immediately after.

Single Chamber Bioreactor Preparation

[0102] Borosilicate tubing (Friedrich and Dimmock Glass, Millvile, N.J.) with dimensions 13 mm26 mm cut in 38 mm long pieces were cleaned and autoclaved to ensure sterility. Two PDMS blocks were custom cut to each end of the bioreactor. A one inch 25 gauge blunt tip luer lock needle was then punctured 3 mm from the bottom of both PDMS blocks. Similarly, a size 14 gauge blunt tip luer lock needle was punctured through both blocks at the top left corner of each block and capped with luer locks for media changes. The conduit fibers were then cannulated and sutured between the two 25 gauge needles. The bioreactor was sealed before sterilizing with 75% ethanol and washing three times with sterile distilled water or PBS. The needles cannulating the microfiber are capped until perfusion.

[0103] ECFCs were seeded on day 0 and 4 in the bioreactor at 510.sup.6 cells in ECFC media supplemented with 1% penicillin/streptomycin (Life Technologies) and 50 ng/mL VEGF (Pierce, Rockford, Ill., USA). Bioreactors were tumbled for 24 hours to optimize cell seeding and medium was changed every other day. vSMCs were seeded on top of the ECFCS 5 to 10 days after ECFC seeding at 1-310.sup.6 cells using a single seeding or repeated seedings every 2 days in ECFC media supplemented with 1% penicillin/streptomycin. Structures were cultured for 5-13 more days before perfusion.

Perfusion

[0104] The 25 gauge needles cannulating the microfibers were connected to either a luer lock syringe for manual perfusion or to silicone tubing for perfusion with a peristaltic pump. A media reservoir was used with an air filter to allow gas exchange, and the whole set-up was placed in an incubator. Flow rate was set to 5 mL/hr (equivalent to 5 dyne/cm.sup.2) or above. Spectrophotometry was used to determine the absolute dye concentration in the chamber after perfusion of fluid with a blue dye (792.8 g/mol) was flowed through the tubular scaffold.

In Vivo Implantation

[0105] To internally seed cells on a vascular graft comprising a tubular scaffold for implantation, a cell suspension of 1.410.sup.3 cells/L was injected through the bioreactor cannulation needles and into the vascular graft comprising a tubular scaffold comprising fibrin. Subsequently, bioreactors were tumbled for 24 hours to optimize cell seeding. All cellularized grafts were cultured for 3-4 days.

[0106] Vascular grafts were implanted as abdominal aorta interposition grafts, as previously. Briefly, anesthetized SCID Beige mice (n=54) (Charles River, Frederick, Md.) was placed in the supine position and opened with an abdominal midline incision. A portion of the infrarenal abdominal aorta was exposed, cross-clamped, and excised. A 3-5 mm length of fibrin nanofiber scaffold was then inserted as an interposition graft using a 10-0 nylon suture for the end-to-end proximal and distal anastomoses. Finally, the intestines are returned to the abdominal cavity and the abdominal musculature and skin are closed in 2 layers with 5-0 absorbable suture. Aspirin (30 mg/L, Bayer) in their drinking water was administered to prevent excess graft thrombosis.

Blood Flow Measurements

[0107] Graft patency and blood flow was monitored with sonography at 2, 3, 8, 12, 16, 20 and 24 weeks post-implantation, which was then compared to blood flow from measurements taken prior to implantation of the grafts (baseline). Using color doppler, the abdominal aorta and grafts were visualized. In the distal abdominal aorta, before the iliac bifurcation, a pulse wave doppler spectrum was collected and analyzed for peak systolic velocity (PSV), end diastolic velocity (EDV), and mean velocity (MV) for the 3 largest waveforms for each mouse and timepoint. The resistivity index (RI) and pulsatility index (PI) were calculated as follows: RI=(PSVEDV)/PSV and PI=(PSVEDV)/MV.

Cryo-Sectioning

[0108] Samples were embedded in OCT compound (Electron Microscopy Sciences, Hatfield, Pa.) and frozen down on a dry ice/ethanol bath. Sections were cut at 50 m using a Cryostat Microtom HM550 (Thermo Fisher Scientific) and collected on positive charged microscopy slides (Thermo Fisher Scientific).

Immunofluorescence Staining and Imaging

[0109] Samples were processed as previously described (Barreto-Ortiz, et al., 2015, Barreto-Ortiz, et al., 2013). Briefly, samples were fixed with 3.7% formaldehyde (Fisher Chemical, Fairlawn, N.J.) for 30 min, permeabilized with 0.1% Triton X-100 (Sigma-Aldrich) in PBS for 20 min, washed three times with PBS, and blocked overnight with 1% BSA. Samples were then incubated overnight at 4 C. with the indicated primary antibodies. Samples were rinsed three times with PBS before being incubated with the appropriate secondary antibodies or conjugated phalloidin at room temperature for 2 h. Samples were then rinsed with PBS three times, and counterstained with DAPI for 15 min. Images were obtained and processed using confocal microscopy (LSM 780, Carl Zeiss Inc., Thornwood, N.Y.). 2D epifluorescence images were obtained using an Olympus BX60 microscope.

[0110] The Examples/Methods have been included to provide guidance to one of ordinary skill in the art for practicing representative embodiments of the presently disclosed subject matter. In light of the present disclosure and the general level of skill in the art, those of skill can appreciate that the following Examples are intended to be exemplary only and that numerous changes, modifications, and alterations can be employed without departing from the scope of the presently disclosed subject matter. The following Examples are offered by way of illustration and not by way of limitation.

[0111] All references, including publications, patent applications, and patents, cited herein are hereby incorporated by reference to the same extent as if each reference were individually and specifically indicated to be incorporated by reference and were set forth in its entirety herein.

[0112] The use of the terms a and an and the and similar referents in the context of describing the invention (especially in the context of the following claims) are to be construed to cover both the singular and the plural, unless otherwise indicated herein or clearly contradicted by context. The terms comprising, having, including, and containing are to be construed as open-ended terms (i.e., meaning including, but not limited to,) unless otherwise noted. Recitation of ranges of values herein are merely intended to serve as a shorthand method of referring individually to each separate value falling within the range, unless otherwise indicated herein, and each separate value is incorporated into the specification as if it were individually recited herein. All methods described herein can be performed in any suitable order unless otherwise indicated herein or otherwise clearly contradicted by context. The use of any and all examples, or exemplary language (e.g., such as) provided herein, is intended merely to better illuminate the invention and does not pose a limitation on the scope of the invention unless otherwise claimed. No language in the specification should be construed as indicating any non-claimed element as essential to the practice of the invention.

[0113] Preferred embodiments of this invention are described herein, including the best mode known to the inventors for carrying out the invention. Variations of those preferred embodiments may become apparent to those of ordinary skill in the art upon reading the foregoing description. The inventors expect skilled artisans to employ such variations as appropriate, and the inventors intend for the invention to be practiced otherwise than as specifically described herein. Accordingly, this invention includes all modifications and equivalents of the subject matter recited in the claims appended hereto as permitted by applicable law. Moreover, any combination of the above-described elements in all possible variations thereof is encompassed by the invention unless otherwise indicated herein or otherwise clearly contradicted by context.

REFERENCES

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