DECELLULARIZED TISSUE/POLYMER MULTI-COMPONENT BIOMATERIALS

Abstract

The technology concerns a construct comprising at least one tissular region and at least one polymeric region for use as an implant.

Claims

1. A construct comprising at least one decellularized tissue, at least one element, and at least one polymeric component; said at least one polymeric component is in physical contact with at least a portion of at least one surface region of the at least one decellularized tissue; wherein at least one polymeric component at least partially coating at least one surface region of the element; further wherein said at least one surface region of the decellularized tissue being in physical contact with said polymer component comes into contact with the at least one surface region of the element coated by said polymer component.

2. A decellularized tissue physically associated to at least one element by means of at least one polymer component, said association comprising or consisting at least partial physical contact of the polymer component with at least a portion of the surface region of the at least one decellularized tissue; and coating by said at least one polymer component of at least a portion of said at least one element; wherein said at least one surface region of the decellularized tissue being associated with said polymer component comes into contact with the at least one surface region of the element coated by said polymer component.

3. A construct comprising at least one decellularized tissue, at least one element, and at least one polymeric component; said the polymeric component having at least one surface feature associated at least partially with at least one face of the decellularized tissue and at least partially at least one surface region of the element; further wherein said at least one surface region of the decellularized tissue being in physical contact with said polymer component comes into contact with the at least one surface region of the element coated by said polymer component.

4. The construct according to claim 1, wherein said at least one element is selected from a group consisting of a metal element, polymeric material, any material having mechanical properties substantially different than said at least one decellularized tissue and any combination thereof.

5. The construct according to claim 4, wherein said at least one decellularized tissue being at least partially in contact with said at least one polymeric component are thermally bonded to provide said association.

6. The construct according to claim 5, wherein said thermal bonding is achieved by raising the temperature of the polymeric component above either (a) the transition temperature thereof, for amorphous thermoplastic polymers; or, (b) the melting temperature thereof for semi-crystalline polymers to result, at least partially, in a polymer chain entanglement along at least one region of the polymeric component.

7. The construct according to claim 4, being in a form of a multisheet construct.

8. The construct according to claim 4, wherein said coating is performed by at least one method selected from a group consisting of spraying, brushing, dip coating and any combination thereof.

9. The construct according to claim 4, wherein said at least one decellularized tissue is dried before being in physical contact with said polymer.

10. The construct according to claim 4, wherein at least one or any of the at least one sheets of a decellularized tissue is/are confined between any two sheets of the polymeric component.

11. The construct according to claim 4, wherein the metal element is a metal wire.

12. The construct according to claim 4, additionally comprising at least one solvent; wherein said solvent is being added to the area at which said at least one surface region of the decellularized tissue being in physical contact with said polymer component comes into contact with the at least one surface region of the element coated by said polymer component.

13. The construct according to claim 12, wherein said addition of said solvent results in at least partial dissolvement of (a) at least a portion of the polymer component in physical contact with said the decellularized tissue; and, (b) at least a portion of said polymer component coating at least one surface region of the element.

14. The construct according to claim 13, wherein said construct is obtained by a reflow effect in which chain entanglement of the along at least one region of the polymeric component of (a) at least a portion of the polymer component in physical contact with said the decellularized tissue; and, (b) at least a portion of said polymer component coating at least one surface region of the element.

15. The construct according to claim 4, additionally comprising at least one solvent, wherein said polymer component is dissolved in said at least one solvent.

16. The construct according to claim 12, wherein said solvent is selected from a group consisting of THF, acetone, dioxane, DMSO, halogenated solvents such as chloroform and dichloromethane, alcohols and polyols, polyethers, acetonitrile, ethyl acetate, dimethylformamide (DMF), Dimethylacetamide, DMAC, and any combination thereof.

17. The construct according to claim 4, wherein the polymeric component is in a form of a layer or a coat of particles, a polymeric sheet, a polymeric film, a polymeric fiber or a polymeric mesh, a gel, a hydrogel or as a liquid or fluidic film and any combination thereof.

18. The construct according to claim 4, wherein the decellularized tissue is obtained from a tissue selected from oral mucosa, small intestinal submucosa and bladder-decellularized matrixes, pericardium, omentum or small intestine mucosa, bovine pericardium, swine pericardium and any combination thereof.

19. The construct according to claim 4, wherein the polymer component is or comprises a polymer is selected from a group consisting of hydrophobic, hydrophilic, amphiphilic polymers, blend, an IPN, or a semi-IPN, an acrylic or a methacrylic polymer, polyolefin, silicone polymer, polycarbonate, a polyurethane, a polyurea or a polyamide, polyurethane, polymethyl methacrylate (PMMA), poly(n-butyl methacrylate) (PBMA), poly(hexyl methacrylate) (PHMA), polystyrene (PST), poly(2-hydroxyethyl methacrylate) (PHEMA), poly(N-(2-hydroxypropyl)methacrylamide) (PHPMA), polycyanoacrylate (PCA), a polyethylene/polypropylene copolymer, a polyethylene/polybutylene copolymer, a polypropylene/polybutylene copolymer, poly-isobulyene, polydimethylsiloxane (PDMS), phenyl-containing PDMS, polyester urethanes, polyether urethanes, polycarbonate, silicone-containing polyurethanes, polyglycolic acid, polylactic acid, polycaprolactone, polylactide-caprolactone copolymer, polyglycolic acid-lactic acid copolymer, polyethylene oxide-polylactic acid copolymer, polyethylene oxide-polycaprolactone copolymer, polytetramethylene oxide-caprolactone copolymer, polyhydroxy butyrate, polyhydroxy valerate, polyethylene adipate, polybutylene adipate, polyethylene succinate polybutylene succinate and polybutylene terephthalate and polyethylene/butylene terephthalate copolymers and combinations and copolymers thereof, shape memory materials, urethane, Pellethane, Elastane, Elastolan, Tecoflex, Biomer, Chronoflex, Biospan, Bionate, silicone-containing polyurethane selected from CarboSil, PurSil, Avcothane and Cardiothane, Chronoflex, Tecoflex and any combination thereof.

20. The construct according to claim 4, wherein one or more of the sheets is designed as a material reservoir for releasing active or non-active materials.

21. The construct according to claim 20, wherein the active material is selected from analgesics; antianxiety drugs; antiarrhythmics; antibacterial agents; antibiotics; anticoagulants and thrombolytics; anticonvulsants; antidepressants; antidiarrheals; antiemetics; antifungals; antihistamines; antihypertensives; anti-inflammatories; antineoplastics; antipsychotics; antipyretics; antivirals; beta-blockers; corticosteroids; cytotoxics; hormones and sex hormones; enzymes; and vitamins.

22. The construct according to claim 4, wherein said metal element is selected from stents, metallic stents, vascular grafts, heart valves, membrane, sealing devices, suture or staple lines, hernia meshes or hernia repair devices, pelvic floor reconstruction devices, wound or burn dressings, dural closures and cardiac patches.

23. The construct according to claim 4, wherein at least two polymeric components are used; further wherein at least one polymer component has a substantially different viscosity than at least one second polymer component.

24. The construct according to claim 4, wherein at least one second decellularized tissue, at least a portion of at least one surface region thereof being at least partially in physical contact with (a) said at least one polymeric component, and (b) in physical contact with said at least one polymeric component being in physical contact with said at least one decellularized tissue, such that application of thermal bonding connects said at least one decellularized tissue and said at least one second decellularized tissue.

25. The construct according to claim 4, wherein the polymeric component being in physical contact with at least a portion of at least one surface region of the decellularized tissue is either (a) at least partially penetrates at least one surface region of a second decellularized tissue; and/or (b) having at least one surface feature protruding one face of a second decellularized tissue, crossing it to the other face through at least one hole formed in the tissue.

26. The construct according to claim 24, being in a form of a multisheet construct.

27. The construct according to claim 26, wherein the multi-sheet construct comprising at least one sheet of a second decellularized tissue and at least one sheet or segment of a polymeric component, wherein any sheet of the second decellularized tissue is adjacent to or in contact with at least one sheet or segment of the polymeric component.

28. The construct according to claim 24, wherein said at least one second decellularized tissue is dried before being coated by said polymer.

29. The construct according to claim 4, wherein the polymeric component is in a form of a layer or a coat of particles, a polymeric sheet, a polymeric film, a polymeric fiber or a polymeric mesh, a gel, a hydrogel or as a liquid or fluidic film and any combination thereof.

30. The construct according to claim 25, wherein the at least one hole is pre-formed or is present in the second decellularized tissue.

31. A device comprising a construct according to claim 4.

32. The device according to claim 31, configured as an implant.

33. The device according to claim 31, the device is selected from stents, metallic stents, vascular grafts, heart valves, membrane, sealing devices, suture or staple lines, hernia meshes or hernia repair devices, pelvic floor reconstruction devices, wound or burn dressings, dural closures and cardiac patches.

34. A process for manufacturing a construct according to claim 1, the process comprising contacting at least one surface region of at least one decellularized tissue with at least one polymer, and contacting at least one surface region of at least one metal element with at least one polymer; contacting said at least one surface region of at least one metal element with said at least one polymer with said at least one surface region of at least one decellularized tissue with said polymer; thereby forming said construct.

35. The process according to claim 34, wherein the polymer is coating said at least one surface region of at least one metal element; and/or said at least one surface region of at least one decellularized tissue.

36. The process according to claim 34, wherein the polymer is cured.

37. The process according to claim 34, additionally comprising step of adding at least one solvent; wherein said solvent is being added to the area at which said at least one surface region of the decellularized tissue being in contact with said polymer component comes into contact with the at least one surface region of the metal element coated by said polymer component.

38. The process according to claim 34, wherein said step of adding said solvent results in at least partial dissolvement of (a) at least a portion of the polymer component in physical contact with said the decellularized tissue; and, (b) at least a portion of said polymer component coating at least one surface region of the element.

39. The process according to claim 38, wherein said construct is obtained by a reflow effect in which chain entanglement of the along at least one region of the polymeric component of (a) at least a portion of the polymer component in physical contact with said the decellularized tissue; and, (b) at least a portion of said polymer component coating at least one surface region of the element.

40. The process according to claim 34, additionally comprising at least one solvent, wherein said polymer component is dissolved in said at least one solvent.

41. The process according to claim 37, wherein said solvent is selected from a group consisting of THF, acetone, dioxane, DMSO, halogenated solvents such as chloroform and dichloromethane, alcohols and polyols, polyethers, acetonitrile, ethyl acetate, dimethylformamide (DMF), Dimethylacetamide, DMAC, and any combination thereof.

42. The process according to claim 34, additionally comprising a. contacting a pierced surface region of at least one decellularized tissue with at least one polymer, and b. permitting said at least one polymer to penetrate into the piercings (holes).

43. The process according to claim 34, additionally comprising a. contacting a surface region of at least one decellularized tissue with at least one polymer, and b. permitting said at least one polymer to at least partially coat said at least one decellularized tissue.

44. The process according to claim 42, further comprising at least one step selected from (a) piercing or forming holes in a surface region of at least one decellularized tissue; (b) inserting the polymer into the piercings (holes); (c) curing the polymer; (d) permitting said at least one polymer to penetrate into the piercings (holes) and form a polymeric sheet on the surface region; (e) any combination thereof.

45. The process according to claim 34, further comprising at least one step selected from (a) thermal bonding said at least one decellularized tissue and said at least one polymer by raising the temperature of the polymeric component above either (i) the transition temperature thereof, for amorphous thermoplastic polymers; or, (ii) the melting temperature thereof for semi-crystalline polymers; to result, at least partially, in a polymer chain entanglement along at least one region of the polymeric component to manufacture the construct; (b) selecting the polymeric to be in a form of a layer or a coat of particles, a polymeric sheet, a polymeric film, a polymeric fiber or a polymeric mesh, a gel, a hydrogel or as a liquid or fluidic film and any combination thereof; (c) any combination thereof.

46. The process according to claim 34, wherein at least one of the following is being held true (a) the polymer fully penetrates through the one or more holes; (b) the polymer partially penetrates the one or more holes; (c) the polymer fully penetrates through the one or more holes to form a polymer sheet on both faces of the surface region, to thereby form a construct assembly; and, (d) any combination thereof.

Description

BRIEF DESCRIPTION OF THE DRAWINGS

[0260] In order to better understand the subject matter that is disclosed herein and to exemplify how it may be carried out in practice, embodiments will now be described, by way of non-limiting example only, with reference to the accompanying drawings, in which:

[0261] FIGS. 1A-1D are SEM cross-section micrographs of a lyophilized decellularized bovine pericardium (DBP) component.

[0262] FIG. 2 shows a high magnification (200,000) SEM micrograph of the DBP.

[0263] FIGS. 3A-3C present structures of tissues after applying increasingly high pressures: 100 MPa, 200 MPa and 500 MPa, respectively, far beyond the pressure required to engineer the DBP/polymer construct, which are typically below 1 MPa, or below 0.3 MPa. The fact that the collagen fibrils fully retain their structure after the exceedingly high pressures, demonstrates the robustness of DBP and its remarkable resistance to extremely high pressures.

[0264] FIGS. 4A-4B show holes made with 25G and 27G needles, respectively, filled with a polymer phase, Tecoflex.

[0265] FIG. 5 presents coated struts of a metallic stent, when compressed in the left-hand side, and once expanded, in the right-hand side. The weldable polymer used consists of PCL and HDI.

[0266] FIG. 6 shows partial welding of a patch to only a small part of coated struts of a metallic stent (around 15% of the area of the stent).

[0267] FIG. 7 shows a stent stretched three times its initial length, with no detachment of the patch from the stent occurring.

[0268] FIG. 8 shows a structure comprising DBP and the bi-component construct comprising DBP and the polymer phase, in this case Tecoflex. The six holes made in the DBP to enhance the flow of Tecoflex aimed to achieve a strong connection between the two DBP phases and the polymeric phase are readily seen. By doing so, in addition to the diffusion of Tecoflex molecules in the DBP phase, a laminate comprising two external thin Tecoflex films that are continuously connected through the holes made in two DBP phases, to a central Tecoflex film, is formed.

[0269] FIGS. 9A-9C show DBP leaflets of a heart valve comprising also a metallic frame connected together via polymer connections, e.g., a polyurethane, Tecoflex. In this case, the DBP leaflets are welded together via the polymer phase of the DBP/polymer construct and also to the metallic frame of the heart valve via its coated struts.

[0270] FIG. 10 shows a stent-graft used to treat Abdominal Aortic Aneurysms (AAA) among other indications, comprising a metallic stent and a fabric where the latter is sewn together to the former via multiple suturing points.

[0271] FIG. 11 demonstrates a polymeric phase consisting of a PCL/HDI poly(ester-urethane) and a low molecular weight polymerizable molecule being hydroxyethyl methacrylate (HEMA). It further reports that CLUR2k has a tensile modulus of around 180 MPa when not plasticized (100:0), and with the addition of increasing amounts of the smart HEMA component, it significantly decreases, showing a value of about 60 MPa in the presence of 20% monomeric HEMA, down to 6 MPa in the 50:50 composition. As apparent from the data presented, polymerized PHEMA results in a significant increase in CLUR's modulus, reaching a value above 250 MPa, for the CLUR2K:PHEMA 50:50 composition.

[0272] FIG. 12 presents CLUR2K's behavior when the low molecular component is not only polymerizable but crosslinkable, as in the case of triethyleneglycol dimethacrylate, comprising two carbon double bonds.

[0273] FIG. 13 demonstrates a further crosslinking methodology.

[0274] FIGS. 14A-14B schematically show A) the generation of the tissue/polymer construct, with the first step being the formation of holes in the tissue components, the size, number and array of which is optimized. An embodiment illustrating a sandwich of two tissues welded together through and mediator polymer sheet is shown in B).

[0275] FIGS. 15A-15B and 16A-16C are illustrations of the tissue-polymer-tissue multisheet construct.

[0276] FIGS. 17-33 are examples are given to evaluate the different parameters and the effect thereof on the strength of the welding.

[0277] FIGS. 34A-34B are examples of different tissue piercing shapes (slots and zig-zag).

[0278] FIG. 34C demonstrates a tissue to metal connection.

[0279] FIGS. 35A-B-36A-C demonstrate a heart valve, according to one aspect of the present invention.

[0280] FIGS. 37A and 37B illustrate the hemodynamic behavior of the heart valve, according to one aspect of the present invention.

DETAILED DESCRIPTION OF EMBODIMENTS

[0281] Thus, for the purpose of providing a superior class of biomedical constructs and medical devices, that would lack the deficiencies associated with currently available equivalents, the inventors of the technology disclosed herein have developed a novel and unique class of constructs comprising [i] a decellularized tissue and [ii] a polymeric component, in which each of its components forms a different phase in space, wherein said decellularized tissue and polymeric components are associated with each other via one or more physical anchoring means. Unexpectedly, the constructs and medical devices disclosed hereby display properties previously unattainable.

[0282] The decellularized tissue/polymer constructs taught by this invention constitute the whole medical device or are part of it and said constructs include, among others, a vascular graft, a cardiac patch, a stent, a heart valve, a wound or burn dressing, a membrane, a sealing device, or devices that reinforce a suture or staple line, are used in hernia repair, in pelvic floor reconstruction, or in dural closure.

[0283] The teachings of the present invention are readily applicable to a diversity of decellularized tissues and polymers. For the sake of clarity, conciseness and simplicity, though, and without detracting from the generality of the scope of the present invention in any form or fashion, the inventors have chosen to illustrate the invention hereby disclosed, by focusing on constructs comprising decellularized pericardium and, more precisely, on decellularized bovine pericardium. For the sake of clarity, conciseness and simplicity, and without detracting from the generality of the scope of the present invention in any form or fashion, the inventors have chosen to illustrate the invention hereby disclosed, by focusing on a medical device where the decellularized pericardium/polymer construct is part of a larger medical device, and more precisely, a heart valve. For the sake of clarity, conciseness and simplicity, though, and without detracting from the generality of the scope of the present invention in any form or fashion, the inventors have chosen to illustrate the invention hereby disclosed, by focusing on polyurethanes as the polymeric phase of said decellularized pericardium/polymer construct and, more precisely, where said polyurethanes is Tecoflex.

[0284] Among them and without any limitation, the decellularized bovine pericardium/polymer constructs disclosed by the current invention were developed in the inventor's laboratory. The mechanical properties of the decellularized pericardium somewhat vary with the batch and as a function of the decellularization technique used. Typically, the stress at break of the decellularization technique falls in 10-80 MPa range, its typical Young's modulus values span between 80 and 300 MPa, exhibiting 20-50% strain at break values. The stress at break, Young's modulus and strain at break values measured at the inventors' lab for lyophilized decellularized bovine pericardium were 75 MPa, 280 MPa and 43%, respectively.

[0285] Tecoflex polyurethane is an aliphatic polyether urethane comprising a polytetramethylene oxide soft segment and a methylene dicyclohexane diisocyanate (MDI Richards J M, McClennen W H, Meuzelaar H L C, Shockcor J P, Lattimer R P: Determination of the structure and composition of clinically important polyurethanes by mass spectrometric techniques. Journal of Applied Polymer Science 1987, 34:1967-1975).

##STR00001##

[0286] Tecoflex's mechanical properties were measured at the inventors' lab and are shown in the Table 1.

TABLE-US-00001 TABLE 1 mechanical properties of Tecoflex Young's Stress at Modulus Strain at Strain at Stress at yield (MPa) (MPa) peak (%) break (%) break (MPa) 175 50 500 520 169

[0287] It is apparent from the mechanical data presented that Tecoflex is a very strong material and will not represent the weak component of the constructs. Other polyurethanes such as Elastolan can be used, as well as other polymers displaying similar rheological and mechanical properties. In cases that the construct requires that the polymeric component is biodegradable, polymers such as polylactic acid (PLA), poly(lactic acid/glycolic acid) (PLGA), polycaprolactone (PCL) and their copolymers, as well as more flexible biodegradable polymers such as biodegradable block copolymers, plastic or elastomeric, comprising hydrophilic polyethers such as polyethylene oxide (PEO) or their hydrophobic counterparts, e.g., polypropylene oxide (PPO) or polytetramethylene oxide (PTMO), or flexible aliphatic polyesters such as amorphous polycaprolactones, or silicone-based segments comprising polydimethyl siloxane (PDMS), among numerous others. The molecular weight of the soft segments typically varies from 600 to 20,000 dalton.

[0288] For the sake of clarity, conciseness and simplicity, though, and without detracting from the generality of the scope of the present invention in any form or fashion, the inventors have chosen to illustrate the invention hereby disclosed, by focusing on constructs wherein the polymeric phase of the construct is able to connect two or more of the tissue phases or other phases comprising materials of all types selected from a group including polymers, metals, ceramics, carbonaceous materials and combinations thereof. More precisely, the inventors have chosen to illustrate the invention hereby disclosed, by focusing on constructs wherein the polymeric Tecoflex phase of the construct connects two decellularized bovine pericardium tissue samples. Even more specifically, the inventors have chosen to illustrate the invention hereby disclosed, by focusing on constructs that are part of a heart valve comprising also a metallic frame, and connect the leaflets between them and also to the metallic frame of the valve. In some embodiments of the present invention, the tissue/polymer construct is part of decellularized pericardium leaflets. In yet other embodiments of the invention, the polymeric Tecoflex phase of the construct connects two tissue leaflets, whereby a multilayered, integrative tissue/polymer construct is formed comprising three layers of the polymer, two of them external and one in the middle of said construct and two layers of tissue internal to said two external layers of the polymer layers, forming a laminate. In other embodiments, said multilayered, integrative tissue/polymer construct formed consists of polymer layers connected by polymeric connections spanning in dimensions from the nanometric to the centimetric scale.

Materials and Methods

Lyophilization of Pericardium

[0289] Since in most cases the decellularized tissue component of the construct has to be lyophilized prior to the generation of the construct, the decellularized tissue component was studied following lyophilization. The structure of the lyophilized decellularized bovine pericardium (DBP) component is shown in the SEM cross-section micrographs exhibited in the FIGS. 1A-1D below.

[0290] The magnification increases from a relatively low 500 value, with the SEM micrograph showing the whole thickness of the DBP, up to a high 120,000 value, where the well-known structure of collagen fibrils is readily observed.

[0291] FIG. 2 shows the high magnification (200,000) SEM micrograph of the as-received DBP.

Preparation of the Polymer

[0292] A selected polymer such as Tecoflex, was added to the welding area in two different states: solution in THF and preprepared film of Tecoflex. The Tecoflex solution was in the range of 5-25% of Tecoflex in THF (refers hereinafter as the polymer solution). For the film preparation 3.78 gr of Tecolfex was added to 100 ml of THF until it fully dissolved. This solution was poured into a glass petri dish, and the THF was left to evaporate slowly. Films had an average thickness of several micron.

Welding of the Pericardium

[0293] Two dry pericardium stripes were prepared for the welding process. On each one of them, area of 2 mm from the edge was marked and 6 holes were generated. The holes were introduced with the polymer solution by syringe with a 25G needle at its tip. The needle was inserted inside the fabric from the rough side to the smooth side. Once it pulled back outside, the solution was poured, simultaneously. The holes were aligned in a straight, parallel to the edge line, in the middle of the welding area.

[0294] The pierced fabrics were left for several minutes for further solvent evaporation and then welded at the sealer, with the rough sides pressed towards each other. A sheet of Tecoflex film was inserted in between the pressed fabrics right before the welding process took place.

[0295] For the welding process three rounds of sealing were applied in delicate regime which is a combination of temp, pressure, and time of the press with relaxation between each cycle out of three cycles. After the welding was performed, the welded materials were dried at room temperature.

Welding Characterization Methods

[0296] The strength of the welded stripes was analyzed by the Instron instrument by a 10 mm/min tensile testing, whereas the leaflets were gripped 10 mm from the welded area at both its sides.

[0297] Also, the structure of the welded stripes was tested by SEM to ensure that the structure had no major changes because of the press and the heat.

[0298] The results are provided in Tables 2-4 below:

TABLE-US-00002 TABLE 2 Tensile test results for a Tissue to Tissue welded model with no holes Stress @ Force @ Width Thickness Peak Test No Peak (N) (mm) (mm) (Mpa) 1 0.447 6.0 0.4 0.186 2 0.914 6.0 0.4 0.381 3 0.426 6.0 0.4 0.178 Min 0.426 6.0 0.4 0.178 Mean 0.596 6.0 0.4 0.248 Max 0.914 6.0 0.4 0.381 S.D. 0.276 0.0 0.115

TABLE-US-00003 TABLE 3 Tensile test results for a Tissue to Tissue welded model with 4 holes Stress @ Force @ Width Thickness Peak Test No Peak (N) (mm) (mm) (Mpa) 1 9.084 6.0 0.4 3.785 2 7.38 6.0 0.4 3.075 3 9.017 6.0 0.4 3.757 Min 7.38 6.0 0.4 3.075 Mean 8.494 6.0 0.4 3.539 Max 9.084 6.0 0.4 3.785 S.D. 0.965 0.0 0.402

TABLE-US-00004 TABLE 4 Tensile test results for a Tissue to Tissue welded model with 6 holes Stress @ Force @ Width Thickness Peak Test No Peak (N) (mm) (mm) (Mpa) 1 9.057 6.0 0.4 3.6228 2 16.156 6.0 0.4 6.4624 3 8.586 6.0 0.4 3.4344 Min 8.586 6.0 0.4 5.248 Mean 11.266 6.0 0.4 4.507 Max 16.156 6.0 0.4 13.463 S.D. 4.241 0.0 1.696

[0299] The data presented in Tables 2-4 support the understanding that the more holes filled with the polymer are formed on the tissues, the stronger the association or welding is.

DISCUSSION

[0300] Applying mild pressure may be necessary in some embodiments of the invention, especially when the polymeric phase connects more than one DBP and/or polymeric phases. The objective of applying pressure in these embodiments of the invention aims at causing the polymer phase or phases to flow, in some embodiment in contact with the DBP and in others when in contact with other polymeric phases, and combinations thereof.

[0301] In some embodiments of this invention, the polymeric phase can flow through pores or holes made in the tissue phase, and polymer molecules can also diffuse and intermingle with molecules of other phases. In some embodiments, the flow of the polymeric molecule causes the interface between the phases to vanish, welding them together.

[0302] FIGS. 3A-3C present the structure of the tissue after applying increasingly high pressures, 100 MPa, 200 MPa and 500 MPa, respectively, far beyond the pressure actually required to engineer the DBP/polymer construct, which are typically below 1 MPa, preferably below 0.5 MPa and even more preferably below 0.3 MPa. The fact that the collagen fibrils fully retained their structure after these exceedingly high pressures, demonstrates the robustness of DBP and its remarkable resistance to extremely high pressures.

[0303] Several polymers that are able to form the polymer phase of the DBP/polymer construct have demonstrated their ability to flow under physiologically acceptable pressure and temperature conditions. Thermograms of two biodegradable poly(ester-urethane)s consisting of polycaprolactone (PCL) of different molecular weights (2,000 and 14,000 dalton) and hexamethylene diisocyanate (HDI), and of their welded phase, obtained by Differential Scanning Calorimetry (DSC), demonstrated new, broader endotherms, shifted to lower temperatures. This provides an indication that the molecules of both polymers have inter-diffused, one hampering the crystallizability of the other.

[0304] This phenomenon and the strength of the welding connection were further explored where two films of two different polyurethanes, Tecoflex and PEU, were welded together. The results demonstrate the ability of Tecoflex and PEU chains to flow across the interface between the two films, providing a strong connection of long term clinical importance. Additionally, both polymers are similar since they are polyurethanes, even though in the first case they are poly(ester-urethane)s and in the latter case they are poly(ether-urethane)s. To illustrate the broad scope of this phenomenon, the DSC thermograms of two polymers were explored.

[0305] DSC demonstrated that the two polymers were very different, one being flexible poly(ester-urethane) and the other a stiff polymethacrylate. The flexible polyurethane consists of PCL 2000 segments and HDI as their coupling agent, while the rigid polymethacrylate is poly(ethylmethacrylate) (PEMA). Also in this case, even though the polymers substantially differ in their composition and mechanical properties, it is apparent from the thermogram of their welded phase, that the phenomenon entails intermingling and entanglement of the chains of both components at the molecular level. In several embodiments of this invention, the mobility of the polymer chains and their ability to disentangle, cross an interface with another polymeric or decellularized tissue phase, diffuse into the second phase and then re-entangle, plays a key role in generating the constructs or medical devices taught by this invention.

[0306] The strength of the connection between the DBP and polymer phases, e.g., Tecoflex, was quantitatively determined using the Instron machine. It was found that the construct failed cohesively, within the tissue, and no de-welding failures were observed.

[0307] In some embodiments, where an especially strong connection between the DBO phase and the polymer phase was aimed at, holes were made in the DBP phase, to maximize the flow of the polymer phase through them. In some embodiments the holes were made using needles in the 16G to 27G range. FIGS. 4A-4B show the holes made with 25G and 27G needles, respectively, filled with a polymer phase, Tecoflex in these cases.

[0308] The DBP phase and the polymer phase can be of any size, spanning from nanometric to centimetric and can adopt any shape, including, without limitation, spherical, fibrous, strips, ribbons, a film, porous or not, and combinations thereof. The polymeric phase/s can also be present in phases the size of which range from being nanosized up to being in the centimeters scale, and they can be on the surface of a DBP phase and/or in its bulk, and each of these cases being of a size ranging from nanometers to centimeters and adopting any geometry.

[0309] It is an objective of this invention to connect the DBP/polymer construct or the device the construct is part of to an additional component of the device. In the case of heart valve, a much-preferred embodiment of this invention, said additional component is the metallic frame of the heart valve.

[0310] In some embodiments of the present invention, when welding the DBP/polymer construct or the device the construct is part of to a metallic stent, the struts of the stent are coated with a weldable polymer that can be the same or different from the polymer constituting the polymer phase of the DBP/polymer construct. FIG. 5 presents the coated struts of a metallic stent, when compressed in the left hand side, and once expanded, in the right hand side. In this embodiment, due to the large expansion of the stent, the weldable polymer coating its struts was chosen to be especially flexible and displaying a high strain at break value. In the case shown below, the weldable polymer used consisted of PCL and HDI. Coatings of different thicknesses were prepared, starting with coating as thin as 5 micrometers and increasing as required.

[0311] Other highly flexible polymers were used as well. Among others, various polyurethanes were used, their soft segment consisting of polyethers or polyesters varying in their molecular weight, their hydrophilicity and, in the case of biodegradable polymers, also in their rate of degradation. One of the polymers used consists of poly(tetramethylene oxide) (PTMO) (MW=650) segments chain extended via HDI. As reported in Table 5, the polymer was welded within 20 seconds at a 47-48 C. temperature.

TABLE-US-00005 TABLE 5 Polymer properties P(PTMG650) Patch material P(PTMG650) T at balloon surface 47-48 C. Welding time ~20 s

[0312] The same polymer was used to coat the struts of the stent s also to create the polymer phase of the DBP/polymer construct and of the additional polymer phase of the medical device, the polymer is part of. To illustrate not only the speed of the welding process but also the strength of the polymer/polymer connection formed, a patch of the PTMO650/HDI polymer was welded in 20 seconds to the struts of the stent coated with the same polymer. Furthermore, the patch was only allowed to weld to only a small part of the coated struts of the metallic stent, as shown in FIG. 6 (around 15% of the area of the stent).

[0313] Initially, forceful manual efforts were made to de-weld the patch from the coated stent but, as shown in FIG. 7, the stent was stretched three times its initial length, with no detachment of the patch from the stent occurring.

[0314] When finally the Instron instrument was used to de-weld the patch from the stent, unexpectedly and surprisingly, it was the metallic stent that failed, while the welded connection remained unaffected. This compelling demonstrates the advantageous features that stem from the ability of the polymer phase of the DBP/polymer construct or of the polymer phase of the medical device the construct is part of to, rapidly and strongly, form long-term, stable connections between the different phases.

[0315] FIG. 8 shows a structure comprising DBP and the bi-coponent construct comprising DBP and the polymer phase, in this case Tecoflex. The six holes made in the DBP to enhance the flow of Tecoflex aimed to achieve a strong connection between the two DBP phases and the polymeric phase are readily seen. By doing so, in addition to the diffusion of Tecoflex molecules in the DBP phase, a laminate comprising two external thin Tecoflex films, that are continuously connected through the holes made in two DBP phases, to a central Tecoflex film, is formed.

[0316] In a much-preferred embodiment, DCP leaflets of a heart valve comprising also a metallic frame are connected together via polymer connections, preferably a polyurethane, more preferably Tecoflex, as shown in FIGS. 9A-9C. In this case, the DBP leaflets are welded together via the polymer phase of the DBP/polymer construct and also to the metallic frame of the heart valve via its coated struts.

[0317] As usually done and routinely reported in the literature, the long-term dynamic stability of heart valves is determined in vitro by the Accelerated Wear Tester (AWT), under accelerated conditions.

[0318] References to three articles, representative of the many studies that describe the use of AWT follow in heart valves follow: {1} A correlation between long-term in vitro dynamic calcification and abnormal flow patterns past bioprosthetic heart valves, Oleksandr Barannyk, Robert Fraser and Peter Oshkai, J Biol Phys (2017) 43:279-296); {2} Pitfalls and outcomes from accelerated wear testing of mechanical heart valves, A Campbell, T Baldwin, G Peterson, J Bryant and K Ryder, J Heart Valve Dis, 1996 June; 5 Suppl 1:S124-32; discussion 144-8; {3}. A study in the design of an Accelerated Wear Tester that is compatible with a particle image velocimetry and high-speed camera setup, Edward A. Brown, Master of Science thesis, The Pennsylvania State University, The Graduate School, College of Engineering, 2015).

[0319] During these tests, a maximum cyclic stress of 100 mmHg is applied, which equals to 0.0133 MPa (12.3 kPa). The failure values measured using the Instron instrument typically fell in the 5 to 10 MPa range, more than 300 higher than the maximum cyclic stress applied during the AWT determinations.

[0320] The advantageous features of the DBP/polymer constructs taught by this invention, become even more striking and surprising, when compared to the routinely used alternative suturing technique.

[0321] Connecting the phases of the construct and the construct to other phases of the medical device the construct is part of have important advantages over suturing: [a] It is very fast as opposed to the extremely tedious and time consuming suturing process, via numerous suturing points, [b] generates strong connections, [c] is reproducible and not technician dependent, [d] achieves much better compliance match between the phases of the construct and the device, in striking contrast to the much stiffer sutures used, [e] achieves a much more homogeneous distribution of stresses, as opposed to the extremely detrimental and sometimes life threatening stress concentrating effect of sutures, and [f] is inexpensive.

[0322] In some embodiments, any of the phases of the construct and/or the medical device the construct is part of or any element of the invention may comprise at least one additional material to improve any aspect of the clinical performance of any of the embodiments of the invention and combinations thereof, including its biocompatibility, its hemocompatibility, the cellular response they trigger, among others. The at least one additional material may be selected amongst active and non-active materials. In some embodiments, the active materials are selected from a variety of bioactive agents. Exemplary bioactive agents include, for example, anticoagulants, such as heparin and chondroitin sulphate; fibrinolytics such as tPA, plasmin, streptokinase, urokinase and elastase; steroidal and non-steroidal anti-inflammatory agents such as hydrocortisone, dexamethasone, prednisolone, methylprednisolone, promethazine, aspirin, ibuprofen, indomethacin, ketoralac, meclofenamate, tolmetin; calcium channel blockers such as diltiazem, nifedipine, verapamil; antioxidants such as ascorbic acid, carotenes and alpha-tocopherol, allopurinol, trimetazidine; antibiotics, such as noxythiolin and other antibiotics to prevent infection; prokinetic agents to promote bowel motility; agents to prevent collagen crosslinking such as cis-hydroxyproline and D-penicillamine; and agents which prevent mast cell degranulation such as disodium chromolglycate, among numerous others.

[0323] In addition to the above agents, which generally exhibit favorable pharmacological activity related to promoting wound healing or reducing infection or having hemostatic properties or enhancing hemocompatibility, other bioactive agents may be delivered by the constructs or the medical devices of the present invention that include, for example, amino acids, peptides, proteins, including enzymes, carbohydrates, growth factors, antibiotics (treat a specific microbial infection), anti-cancer agents, neurotransmitters, hormones, immunological agents including antibodies, nucleic acids including antisense agents, fertility drugs, psychoactive drugs and local anesthetics, among numerous additional agents. The delivery of these agents and others will depend upon the pharmacological activity of the agent, the site of activity within the body and the physicochemical characteristics of the agent to be delivered, the therapeutic index of the agent, among other factors. One of ordinary skill in the art will be able to readily adjust the physicochemical characteristics of the present polymers and the hydrophobicity/hydrophilicity of the agent to be delivered in order to produce the intended effect. In this aspect of the invention, bioactive agents are administered in concentrations or amounts which are effective to produce an intended result. It is noted that the chemistry of polymeric phases according to the present invention can be modified to accommodate a broad range of hydrophilic and hydrophobic bioactive agents and their delivery to sites in the patient.

[0324] In some embodiments, the non-active materials are selected amongst dyes, polymeric materials, thickening agents, plastizicers, agents affecting hydrophilicity, agents affecting lubricity and others.

[0325] The constructs and medical devices taught by the invention may be manufactured by any of the existing manufacturing techniques, such as extrusion, compression molding, injection molding, dip coating, solvent casting, welding, any of the numerous 3D printing techniques, and in each case the specific manufacturing technique being used will be tailored so it is compatible with the constructs and medical devices taught by the invention.

[0326] In some embodiments where at least one of the polymer phases is biodegradable, said biodegradable polymer is selected from a group comprising lactic acid, lactide, glycolic acid, glycolide, or a related aliphatic hydroxycarboxylic acid or ester (lactone) selected from the group consisting of. -propiolactone, -caprolactone, -glutarolactone, -valerolactone, -butyrolactone, pivalolactone, ,-diethylpropiolactone, ethylene carbonate, trimethylene carbonate, -butyrolactone, p-dioxanone, 1,4-dioxepan-2-one, 3-methyl-1,4-dioxane-2,5-dione, 3,3,-dimethyl-1-4-dioxane-2,5-dione, cyclic esters of -hydroxybutyric acid, -hydroxyvaleric acid, -hydroxyisovaleric acid, -hydroxycaproic acid, -hydroxy--ethylbutyric acid, -hydroxyisocaproic acid, -hydroxy--methyl valeric acid, -hydroxyheptanoic acid, -hydroxystearic acid, -hydroxylignoceric acid, salicylic acid and mixtures, thereof.

[0327] The polymeric phases according to the present invention comprise optionally low molecular weight molecules able to enhance the flowability of said polymeric and/or allow causing the polymer phase or part of it to flow under milder temperature and pressure conditions. It is a further object of the invention to provide low molecular weight molecules that are polymerizable or crosslinkable, so they soften the polymer phase or parts of it before they polymerize or crosslink and strengthen or stiffen said polymer phase once polymerized or crosslinked, where said low molecular weight molecules can polymerize or crosslink following any mechanism including, without limitation, addition and condensation polymerization reactions as well as additional reactions, including all types of click chemistry and combinations thereof. Among others, said polymerizable or crosslinkable low molecular weight molecules include precursors comprising one or more double bonds. A few examples are given in the figures below.

[0328] In FIG. 11, the polymeric phase consists of a PCL/HDI poly(ester-urethane) and the low molecular weight polymerizable molecule is hydroxyethyl methacrylate (HEMA).

[0329] FIG. 11 reports that CLUR2k has a tensile modulus of around 180 MPa when not plasticized (100:0), and with the addition of increasing amounts of the smart HEMA component, it significantly decreases, showing a value of about 60 MPa in the presence of 20% monomeric HEMA, down to 6 MPa in the 50:50 composition. As apparent from the data presented, polymerized PHEMA results in a significant increase in CLUR's modulus, reaching a value above 250 MPa, for the CLUR2K:PHEMA 50:50 composition.

[0330] FIG. 12 presents CLUR2K's behavior when the low molecular component is not only polymerizable but crosslinkable, as in the case of triethyleneglycol dimethacrylate, comprising two carbon double bonds.

[0331] Several additional chemistries can be employed to polymerize or cross-link the low molecular weight component, such as the epoxy-amine reaction. This chemistry is illustrated by using polyglycidyl methacrylate (PGMA) (see FIG. 13), is very stiff and somewhat brittle, blended with polyethylene imine molecules, which contain several amine groups. When initially blended with PGMA, PEI molecules lower PGMA's rigidity but, once it reacts with the epoxide ring (see FIG. 13), it crosslinks PGMA and stiffens the polymer.

[0332] FIGS. 14A-14B schematically show the generation of the tissue/polymer construct, with the first step being the formation of holes in the tissue components, the size, number and array of which is optimized. In this case a liquid polymer phase is added, which may be for example a polymer solution in an appropriate solvent, or when the polymer is above its glass transition or melting point, when these transitions occur at a suitably low temperature. The polymer penetrates the tissue, primarily through the holes made in it, the length, size, number and array of which is controlled and varies over a significant range. In some instances, the polymer phase contains holes of different depth of penetration. In some instances, holes penetrate the tissue phase only partially and in other embodiments, the depth of the holes cross face-to-face the tissue phase. In the case depicted below, the holes cross the whole tissue thickness, generating on both side of the tissue component, two layers of polymer. Since they are connected via the polymer connections filling the holes, the two polymer films become one integrative polymer phase. In this case, the mechanical properties of the construct created by the physical association of the tissue and polymer phases are especially high, since they derive from the cohesive strength of the polymer itself.

[0333] In other embodiments of the invention, a pre-formed polymer phase is first produced, the composition and morphology of which are such that it can flow under the right temperature and pressure conditions, as described before.

[0334] In some embodiments, said pre-formed polymer phase has different properties in its surface layers, as opposed to its bulk, which allow the polymer phase in contact with the tissue to flow into the holes made in the tissue. In some embodiments, the difference between the surface layers, the thickness of which is optimized, and the bulk of the polymer phase, is compositional, so that the surface layers exhibit the required flowability under the temperature and pressure conditions applied. In some embodiments, the difference between the surface layers, the thickness of which is optimized, and the bulk of the polymer phase, is morphological. In this embodiment, the surface layer is less crystalline, and in some embodiments, amorphous, while the polymer in the bulk of the polymer phase displays enhanced crystallinity and, therefore, higher rigidity at the relevant temperature. The morphological differential encoded in the polymer phase can be achieved following various strategies. Among other, this can be achieved conducting spatially confined thermal treatments that render the surface layer less crystalline or amorphous, as opposed to the bulk of the polymer phase. In some embodiments, the difference between the surface layers and the bulk of the polymer phase is achieved by the addition of a mobile component that tend to migrate and concentrate on the surface layers of the polymer phase. In yet other embodiments, the difference between the surface layers and the bulk of the polymer phase is achieved by judiciously choosing the surfaces in contact with which the polymer phase is produced. These and other similarly effective techniques can be used separately or can be combined.

[0335] In yet additional embodiments of the invention, an initially liquid polymer phase and a pre-formed polymer phase are combined in different ways. In some embodiments, for example, they are added simultaneously, or they can also be deployed sequentially, or in any other manner that will produce a tissue/polymer optimal association, as derived from its clinical use.

[0336] The procedure whereby a tissue/polymer consisting of at least two tissue phases is produced using a polymer solution, is exemplified hereby and depicted in FIG. 14B. In some instances, the term welding is used interchangeably with connecting and similar terms. Two lyophilized pericardium strips were used and the tissue/polymer association resulted in the connection between two tissue phases via the polymer phase. On each one of the pericardium strips, 6 holes were generated within an area of 2 mm from the edge, spaced throughout the width of the tissue strip. In this embodiment, the holes were formed and filled with a polymer solution simultaneously. In this case, an 8% Tecoflex/THF solution was prepared, and a syringe with a 25G needle at its tip was filled with the solution (for other purposes on in other experiments, a concentration of 1% or 4% Tecoflex was used). The needle was inserted in the pericardial tissue strip from its rough side puncturing it and creating the holes, and the polymer solution was poured initially distally and then while pulling back, filling the holes, and finally generating a polymer film on the tissue proximal face. By doing so, a laminate consisting of two external polymer phases connected via polymer connections (through the holes made in the tissue), was produced. The holes were aligned in a straight, parallel to the edge line, in the middle of the welding area. The tissue/polymer constructs consisting of the laminate described above were left for 2 minutes to allow optimal solvent (THF in this specific case) evaporation and then the two treated tissue/polymer constructs of each tissue strips were overlapped as required and welded together under the right temperature, pressure and time conditions, using a sealer, with the rough sides pressed towards each other. Three rounds at the sealer were applied at 85 C. Each round at the sealer consisted of a 2 bar, 12 seconds step, followed by 1 second of relaxation between the cycles. After the connection was performed, the welded tissue/polymer constructs were dried at room temperature for 5 hours, followed by an immersion in saline solution for 24 hours. The strength of the connection between the two connected tissue strips was studied at the Instron instrument using a 10 mm/min cross-head speed, with the strips being gripped 10 mm away from the connected area.

[0337] The strength of the connection formed was also determined under peel conditions and to this end, samples were connected accordingly. All other details of the procedure remained the same. The mechanical studies of these samples were conducted using the same conditions as previously indicated, with a sole differencethe grasp took place 20 mm from the welded area.

[0338] Some of the different conditions and connection parameters studied are listed below: [0339] 1) For 27G needle2/4/6/8 holes. [0340] For 25G needle2/4/6 holes. [0341] For 23G needle2/4/6 holes. [0342] For 21G needle2/4 holes. [0343] For 19G needle2/4 holes. [0344] For 16G needle2 holes. [0345] 2) Cycles performed at the sealerfrom 1 to 6. [0346] 3) Connection pressurefrom 2 bar to 10 bar. [0347] 4) Connection timefrom 8 seconds to 12 seconds.

[0348] The procedure whereby a tissue/polymer consisting of at least two tissue phases is produced by combining a polymer solution and a pre-formed polymer phase, is exemplified hereby. The steps involving the Tecoflex/THF solution were the same as described above.

[0349] The tissue/polymer constructs consisting of the laminate formed, as described above, were left for 2 minutes to allow optimal solvent (THF in this specific case) evaporation. Before welding the two tissue/polymer constructs prepared, a 2*6 mm Tecoflex film (thickness 160 m) was deployed between the two tissue/polymer constructs before the welding process using the sealer, took place. The Tecoflex film was pre-prepared using the solvent casting method using 3.78 g of Tecoflex in 100 ml of chloroform, even though other solvents were also used. The same steps described already were conducted at the sealer and the connected samples were studied at the Instron machine, as described above.

[0350] Some of the different conditions and connection parameters studied are listed below: [0351] 1) For 27G needle2/4/6/8 holes. [0352] For 25G needle2/4/6 holes. [0353] For 23G needle2/4/6 holes. [0354] For 21G needle2/4 holes. [0355] For 19G needle2/4 holes. [0356] For 16G needle2 holes. [0357] 2) The pre-generated film was prepared either from THF or Chloroform solvents using 1.5-4 g of Tecoflex in 100 ml of the mentioned solvents. [0358] 3) Cycles performed at the sealerfrom 1 to 6. [0359] 4) Welding pressurefrom 2 bar to 10 bar. [0360] 5) Welding timefrom 8 seconds to 12 seconds.

[0361] Thus, according to one embodiment of the present invention, a tissue-to-tissue connection process is enabled.

[0362] According to a preferred embodiment, the process includes the following main 2 steps: [0363] (a) two tissue sheets 10 are positioned from both sides of a polymer film 20 (see FIGS. 15A and 15B). [0364] (b) rivets 30 are created, which functions as polymer tubes, by injecting polymer from side to side of the tissue-polymer-tissue sandwich (see FIGS. 16A-16C). Such rivets function similarly to classic wall flange or rivets. According to preferred embodiments, a sandwich of polymer-tissue-polymer-tissue-polymer is created (as illustrated in FIG. 14B).

[0365] In other words, the connection/welding of the 2 tissues is a superposition of two mechanisms: [0366] 1. surface attachment (like glue); and, [0367] 2. physical connectionrivets-like.

[0368] It should be noted that it is possible for each of the tissue sheet to be coated on both sides thereof by the polymer. Thus, the construct sandwich would be polymer-tissue-polymer-tissue-polymer.

[0369] According to another embodiment of the present invention, laser cutting is used to cut the tissue at the connection (weld) area (namely, at the areas of connection with the polymer).

[0370] According to another embodiment of the present invention, the tissue is freeze and dried (i.e., lyophilization process) for a period of 5-20 hours, with or without vacuum, at a temperature of 70-20 degrees Celsius, before being welded with the polymer.

EXAMPLES

[0371] The following examples are given to evaluate the different parameters and the effect thereof on the strength of the welding.

[0372] The tests were mainly divided into three groups: [0373] Welding-related parameters: heat, pressure, and time were examined to see the effect thereof on the welding's strength. It should be noted that the sensitivity of pericardium to high temperature is limited to 85 C. Therefore, the main applied experiments were 2, 4, and 6 bars pressure, and heat time of 36 sec (3 pulses) 50, and 120 seconds (one pulse). [0374] Holes-related parameters: number of the holes, the distance between the holes, holes' size and depth. [0375] Polymer-related parameters: two kinds of polymers and different thicknesses of used polymer films.

[0376] It is noted that all holes were prepared by laser cutting machine and were in one row. Generally, the used protocol is as follows: [0377] a pericardium sheet was divided into 6 pairs of strips, each pair is referred to as one welding sample.

[0378] At 1 mm of the upper side of each strip, X holes were conducted with a distance of Y therebetween (X, Y are disclosed in each of the experiments detailed below).

[0379] The laser cutting process was done using the following parameters: 100% power, 50% speed, 1000 points per inch, PPI.

[0380] Next tensile tests were conducted. In each test, where tissue was secured to metallic handles covered with plastic to increase the resistance of the surface (in Testometric X250-2.5 device with a cell load up to 250N) and stretched at a rate of 10 mm/min till the failure point.

[0381] In the following examples, the polymer used is Chronoflex.

Welding-Related Parameters:

[0382] 1) Pressure Effects on the Welding Strength:

[0383] Pressures of 2, 4, and 6 bars and three pulses with 12 seconds of heating were applied. In the test, the strength needed to break the welding was measured. The measurements were made with a Testometric X250-2.5 device and stretched at a rate of 10 mm/min till the failure point.

[0384] Reference is now made to FIG. 17 illustrating the effect the pressure on the stress of the welding. It is noted that in light of the differences in the thicknesses of the pericardium tissues stress was measured and not strength.

[0385] As can be seen, applying higher pressure results in stronger welding (higher force needs to be applied to break the welding). Where, in the FIG., W represents failure in welding site; PW represents failure in pericardium at the welding site; and P represents failure in the pericardium.

[0386] Using high pressure enforces the polymeric chains to move at the welding site leading to higher entanglement. Therefore, as shown, the welding strength is greater.

[0387] 2) Time of Heat and Pulse Number Effects on the Welding Strength:

[0388] Higher temperatures could not be used due to the sensitivity of pericardium to high temperatures. However, heating at 85 C. for a longer time and without relaxation could be tested. Thus, in this example, the heating time was tested as follows: 36 sec in three pulses, 50 sec in one pulse, and 120 sec in one pulse.

[0389] Reference is now made to FIG. 18, illustrating the effect the temperature on the strength of the welding. As disclosed above, due to the differences in the thicknesses of the pericardium tissues stress was measured and not strength.

[0390] As can be seen in the figure, the higher heating time has a max at 50 sec then the welding strength becomes weaker.

[0391] Where, in the FIG., W represents failure in welding site; PW represents failure in pericardium at the welding site; and P represents failure in the pericardium.

[0392] The results can be explained by some damages caused in the pericardium leading to weakness in its mechanical properties. In other words, a higher heating time has the same effect as using a higher temperature.

Holes-Related Parameters:

[0393] 1) Holes Size Effects on the Welding Strength

[0394] Reference is now made to FIG. 19 illustrating the effect of the size of the hole on the welding strength. As disclosed above, due to the differences in the thicknesses of the pericardium tissues stress was measured and not strength.

[0395] As seen in the figure, the greater holes' diameter results in increase in the welding's strength till the 0.4 mm holes; thereafter, the greater holes' diameter results in a decrease in the welding's strength. Where, in the FIG., W represents failure in welding site; PW represents failure in pericardium at the welding site; and P represents failure in the pericardium.

[0396] 2) Holes Number Effects on the Welding Strength

[0397] Reference is now made to FIG. 20 illustrating the effect of the numbers of the hole on the welding strength. Where, in the FIG., W represents failure in welding site; PW represents failure in pericardium at the welding site; and P represents failure in the pericardium. As disclosed above, due to the differences in the thicknesses of the pericardium tissues stress was measured and not strength.

[0398] For the purpose of this example, 0.2 holes were tested. As seen in this figure, a maximum strength was achieved at four holes with a distance of 1.5 mm between each other; then, the strength decreases as the no. of holes increases.

[0399] On the one hand, fewer holes have less polymer; on the other hand, more holes have less tissue. In other words, the welding became weaker for additional holes due to less tissue in the welding site.

Polymer-Related Parameters:

[0400] Film Thickness Effects on the Welding Strength:

[0401] Reference is now made to FIG. 21 illustrating the effect of the film thickness on the welding strength. Where, in the FIG., W represents failure in welding site; PW represents failure in pericardium at the welding site; and P represents failure in the pericardium. As disclosed above, due to the differences in the thicknesses of the pericardium tissues stress was measured and not strength.

[0402] As seen, the results indicate that thicker polymers lead to a weaker welding area. Thicker polymer means that it requires higher energy (temperature in our case) to de-entangled the chains of the polymer, and due to the limitation of temperature thicker polymer is not preferred to be used, resulting in an increased welding's strength.

[0403] In the following examples, the polymer used is polycaprolactone 43K, PCL 43K. It is noted that PCT is a biodegradable polyester with a low melting point of around 60 C.

[0404] 1) Holes Size Effects on the Welding Strength:

[0405] The melting point of PCL is about 60 C., therefore the welding process of the PCL was performed at 55 C., pressure of 4 bars and 50 sec in one pulse.

[0406] As described above, the insertion of holes strengthens the welding site since the holes include polymer, increasing the welding occurrence.

[0407] As PCL, 8% solution is not as viscous as Chronoflex, the holes can accommodate decrease amounts of the polymer. In addition, the used temperature of the process essentially melts PCL (as the same has a low melting temperature); thus, the welding becomes weaker, as illustrated FIG. 22 (Where, in the FIG., W represents failure in welding site; PW represents failure in pericardium at the welding site; and P represents failure in the pericardium). As disclosed above, due to the differences in the thicknesses of the pericardium tissues stress was measured and not strength.

[0408] 2) Holes Number Effects on the Welding Strength:

[0409] In this section, the number of holes was tested, and the used size of holes used is 0.2 mm.

[0410] Reference is now made to FIG. 23 illustrating the effect of the film thickness on the welding strength. Where, in the FIG., W represents failure in welding site; PW represents failure in pericardium at the welding site; and P represents failure in the pericardium. As disclosed above, due to the differences in the thicknesses of the pericardium tissues stress was measured and not strength.

[0411] As seen in FIG. 23, an increase no. of holes increased the welding strength.

CONCLUSION

[0412] All the conducted work indicated the importance of holes for strengthening the welding site and a better understanding of the welding process and all the parameters to achieve welding with enhanced mechanical properties.

[0413] To sum up, both PCL 43K and Chronoflex insertion of holes improved the welding strength and increased its mechanical properties. However, the best results were not achieved using the same set of parameters. PCL got the best results by creating six holes of 0.2 mm, and the stress at the peak was almost doubled. While in the Chronoflex case, the best result was achieved by inserting four holes of 0.4 mm, where the strength was nearly tripled of the no holes result.

[0414] Reference is now made to FIGS. 24-33 illustrating the following:

[0415] FIG. 24 illustrates the dry pericardium tissue with no holes.

[0416] FIG. 25 illustrates the wet pericardium tissue with no holes after tensile test.

[0417] FIG. 26 illustrates the dry pericardium tissue with 0.6 mm, 4 holes.

[0418] FIG. 27 illustrates the wet pericardium tissue 0.6 mm, 4 holes wet after tensile.

[0419] FIG. 28 illustrates the dry pericardium tissue 0.2 mm, 3 holes.

[0420] FIG. 29 illustrates the wet pericardium tissue 0.2 mm, 3 holes wet after tensile

[0421] FIG. 30 illustrates the dry pericardium tissue 0.2 mm, 4 holes dry.

[0422] FIG. 31 illustrates the wet pericardium tissue 0.2 mm, 4 holes wet after tensile.

[0423] FIG. 32 illustrates the dry pericardium tissue 0.2 mm, 6 holes dry.

[0424] FIG. 33 illustrates the wet pericardium tissue 0.2 mm, 6 holes wet after tensile.

[0425] The following examples are given to evaluate the different steps in the production of the multi-sheet construct and the effect thereof on the strength of the welding. [0426] 1. Freeze dry (lyophilization) impact on the tissue's strength (before the welding):

[0427] In the following testing, the impact of the shock freeze on the mechanical properties of the tissue was evaluated. The dried tissue's mechanical properties were compared with/without liquid nitrogen, N2.

[0428] The following testing protocol was employed and performed: [0429] 1. laser cutting a tissue area having dimensions of 6 mm40 mm and thickness of 250 microns+50 microns; [0430] 2a. measurements of tensile strength at the middle of the 6 mm width and along the 40 mm at the ends and middle; [0431] 2b. for the liquid nitrogen, N2, testing the tissue was submerged in liquid Nitrogen; [0432] 3. Next, the tissue samples were placed in a freeze-dry machine for a period of 7-24 hours. in vacuum at a temperature range of 70-0 degrees Celsius; [0433] 4. Next, the dry tissue's thickness was measure to compare the difference between the treatment with and without liquid Nitrogen. [0434] 5. Then, the tissue samples were placed in saline for 12 hr. for recovery [0435] 6. Next, the dry tissue's thickness was measure again [0436] 7. Lastly, the tissues throw a tensile test to define at what force it failed (rapture), and the stress results were are calculated.

[0437] The mechanical properties as measured were as follows:

TABLE-US-00006 Number of tested samples Force(N) Stress (MPa) Strain (%) Wet tissue 14 13.57 5.69 13.93 5.44 572.3 303.5 Wet tissue after freeze drying 9 18.14 4.53 18.62 6.90 782.5 371.4 Wet tissue after freezing dry with 10 13.24 4.44 13.45 5.46 266 79.4 N.sub.2

[0438] As seen from the table, although freeze drying the tissue with N2 gave fairly good results, the max. stress was obtained treating the tissue after freeze drying thereof without the use of N.sub.2.

[0439] When the water is frozen at a very fast cooling rate (un-controlled cooling rate), the water molecules within the tissue's cells expand and damage the collagen matrix of the tissue. Thus, when Nitrogen is used, the mechanical properties of the tested tissued are inferior to those where Nitrogen is not used. On the contrary, controlled cooling rate of the water molecules will cause less damage to the tissue and hence will results in superior mechanical properties.

[0440] 2. Welding Strength: [0441] a) In this section the drying time in the freeze-dry machine was tested and the influence thereof on the welding strength was measured:

TABLE-US-00007 Number of tested Stress samples Force(N) (MPa) Baseline (4 holes of 0.3 mm each) 2 2.9 0.9 4.2 1.9 1 Evaporation Day Baseline 4 Evaporation days 3 3.3 0.4 4.8 0.9 Baseline 7 Evaporation days 3 2.8 0.3 3.8 0.4

[0442] Although it is not a substantial difference, as seen from the table, the max. stress was obtained after 4 evaporation days. [0443] b) Testing the sticking strength of the tissue-to-tissue connection (without piercing of the tissue and allowing the polymer to cross the same; namely, the use of polymer-based rivets):

[0444] As pericardium tissue has different properties on each surface, one side is smooth, the other side is rough, the mechanical properties of both sides were tested.

TABLE-US-00008 Number of tested samples Force(N) Stress (MPa) Baseline (no holes, smooth to smooth) 3 0.5 0.3 0.7 0.5 Rough side to rough side 3 2.26 0.47 2.29 0.46 Thermo-bond for each side (namely, 3 0.40 0.15 0.5 0.2 application of thermal, heat, energy) engraving the tissue (i.e., poking the tissue 3 1.3 0.2 1.87 0.45 with the needle along all the connection area without fully pricing in from side to side) engraving the tissue (i.e., poking the tissue 3 2.85 0.49 3.58 1.34 with the needle along all the connection area without fully pricing in from side to side) with viscous polymer

[0445] As seen, the best results were received with the rough side faced with the rough side or tissue engraving. All the failure points were at the welding connection. [0446] c) Testing the Rivets

[0447] In the following tests, the mechanical properties of the rivets were measured according to the following protocol: [0448] measurements of tissue thickness area that are 250 microns+50 microns [0449] 1. laser cutting tissue thickness of 250 microns with dimensions of 6 mm40 mm and an adequate amount of holes at 1 mm from the edge of the tensile model [0450] 2. measurements of the tensile model at the middle of the 6 mm width and along the 40 mm at the ends and middle; [0451] 3. Next, the tissue samples were placed in a freeze-dry machine for a period of 7-24 hours in vacuum at a temperature range of 70-0 degrees Celsius; [0452] 4. Next, the dry tissue polymer films were prepared by attaching the dried tissue, together and injecting the polymer rivet into the prepared holes (or injecting the polymer into the prepared hole) [0453] 5. the Construct (tissue-polymer-tissue) is heat pressed at the Thermo-bond at 60 Celsius degrees at 2 bar. Namely, a solvent (e.g., THF) was steamed for 4 days. Its main purpose is to melt the polymer components and connect them together to provide the welding. [0454] 6. Then, the tissue samples were placed in saline for 12 hr. for recovery [0455] 7. Next, the dry tissue's thickness was measure again [0456] 8. Lastly, the tissues throw a tensile test to define at what force it failed (rapture), and the stress results were are calculated.

TABLE-US-00009 Number of tested samples Force(N) Stress (MPa) Failure point Baseline 4 Evaporation 3 3.3 0.4 4.8 0.9 The connection days 4 holes of 0.8 millimeter 3 5.9 0.1 7.1 0.1 The tissue 3 holes of 0.8 millimeter 3 4.2 1.4 5.3 0.8 The tissue 4 holes of 0.6 millimeter 3 9.2 0.7 11.4 1.4 The tissue 3 holes of 0.6 millimeter 3 7.5 1.4 10.2 1.2 The tissue 4 holes of 0.8 millimeter 3 6.9 1.4 10.5 1.9 The tissue (Far from the tissue's edge) 5 holes of 0.8 millimeter 3 7.4 1.5 9.9 2.9 The tissue (Far from the tissue's edge) 4 holes of 0.6 millimeter 3 6.92 1.22 10.1 2.9 The Connection (Chronoflex C75D)

[0457] As seen in the above, inferior results were obtained when evaporation was used. Also, the more holes, each being with smaller diameter, in the tissuethe better results obtained. [0458] d) Holes size and shape and different arrangements of the holes

TABLE-US-00010 Number of tested Stress Failure samples Force(N) (MPa) point Slot holes (see 3 3.13 0.84 3.5 1.2 The tissue FIG. 34A, where numerical reference no. 10 denotes the tissue and numerical reference 11 denotes the holes), where each slot was successive filled with the polymer Slot holes (see 3 6.9 1.0 8.64 1.67 The tissue FIG. 34A), where each slot was fully filled with the polymer (in one shot) Slot holes 0.4 mm 3 7.3 0.8 8.82 1.13 The tissue Zig-zag holes 3 5.61 0.47 6.57 1.02 The tissue (see FIG. 34B) [0459] e) Different polymers:

[0460] In the below table, a summary of the stress results as measured are given for a different polymer used.

[0461] As seen in the table above, the Chronoflex C 75D provides the best results.

TABLE-US-00011 Number of tested samples Force(N) Stress (MPa) Failure point Chronoflex C75A 3 3.3 0.4 4.8 0.9 The connection Chronoflex C93A 3 2.74 0.68 4.47 1.55 The connection Chronoflex C75D 3 5.05 0.23 8.96 1.16 The connection A combination of 2 2 5.64 0.07 9.3 0.3 The connection Chronoflex C: the films are from shore 75 A the rivets are from shore 75 D A combination of 2 3 3.7 0.3 4.9 0.4 The connection Chronoflex C: the films are from shore 75 D the rivets are from shore 75 A [0462] f) Alignment of the holes, offset testing between the holes pre-created on the two tissues:

[0463] In this testing the holes in the tissue are not alighted. They have some overlap but they not concentric.

TABLE-US-00012 Number of tested Stress samples Force(N) (MPa) Failure point Non-aligned 3 3.5 0.3 3.4 0.2 The connection holes

[0464] It should be noted that according to another embodiment of the present invention, the tissue-polymer-tissue connection can be achieved by at least partially coating/bringing into physical contact the 2 tissues with the polymer (without the need to create the holes). According to this embodiment, each of the tissues is at least partially coated (or comes into contact with the polymer) and then, thermal bonding takes place to weld the two tissues together.

[0465] In a mechanical property testing 2 6 mm*6 mm tissue samples were coated with the polymer (coating thickness was 0.25-0.26 mm)

TABLE-US-00013 Stress @ Strain @ Stress @ Strain @ Sample Peak Peak Break Break # Force(N) (MPa) (%) (MPa) (%) 1 3.455 2.215 135.721 0.078 342.093 1.426 0.985 65.55 0.076 113.084 2 5.906 3.937 133.144 0.057 461.226 1.426 0.985 65.55 0.076 113.084 3 4.447 2.965 102.055 0.162 185.395 1.426 0.985 65.55 0.076 113.084 4 6.623 4.415 251.12 0.023 340.376 1.426 0.985 65.55 0.076 113.084

[0466] As seen in the table above, such tissue-tissue connection (welding) is achieved to withstand a force of 4.5 MPa.

[0467] Thus, the construct, according to one embodiment of the present invention comprising: at least two decellularized tissue and at least one polymeric component, wherein the polymeric component at least partially coating at least one surface region of each of the two decellularized tissue. The welding between the two decellularized tissue is achieved by thermal bonding the layers together.

[0468] As noted above, thermal energy is applied to raise the temperature of the polymer above the appropriate transition temperature, i.e., the glass transition temperature, Tg, for amorphous thermoplastic polymers, or the melting temperature, Tm, for semi-crystalline polymers. When two sheets or segments of the polymeric component are brought into intimate contact under these conditions, polymer chain entanglement will occur resulting in a weld. According to aspects of the invention, welding need not be achieved over the full surface of the sheet or segment. Point welding at one or more regions of the polymeric component may suffice to provide a robust association of the plurality of assemblies or any two polymeric sheets or segments.

[0469] According to another embodiment, a tissue-to-metal connection/welding is provided. According to this embodiment, a single attachment mechanism is provided. According to which a tissue (being is in physical contact with the polymer) is in contact with a polymer coated metal to facilitate the welding. In this embodiment, the construct is an assembly of (a) n numbers of decellularized tissues, at least one of which is at least partially associated with at least one polymer; and m numbers of the metal element, at least one of which is at least partially coated with the polymer. N and m are integer numbers higher than 0. According to one embodiment, m equals n; according to another embodiment, m is substantially different than n.

[0470] According to one embodiment, the association of said at least one decellularized tissue with said polymer refers to physical contact/interaction, coating, adhesion, welding, robust anchoring and any combination thereof. Alternatively, or additionally, the association is referred to: [0471] (i) at least partial embedding of the polymeric sheet or segment in a surface region of a layer of the tissue, whereby the embedding may be through a single point on the surface region of the layer, or through two or more points, with or without penetrating the tissue, or [0472] (ii) anchoring of the polymeric sheet or segment into the tissue to a tissue depth that permits secured association, whereby the anchoring may be through a single anchoring point, or through two or more anchoring points, whereby the depth of anchoring or penetration of the polymer sheet or segment may vary, not including puncturing of the tissue, or [0473] (iii) anchoring the polymeric sheet or segment fully into the tissue to completely penetrate the tissue, from one face of the tissue to the other, whereby the anchoring may be through a single anchoring point, or through two or more anchoring points. Typically, such penetration would involve a surface feature that is configured to protrude one face of the decellularized tissue to the other through at least one hole formed in the tissue. As will be detailed below, such a feature may be formed in situ after holes are formed in the tissue or may be provided on the polymer sheet or segment in a form selected and configured to puncture or pierce the tissue.

[0474] The polymer sheets or segments are said to be associated to the tissue in a way that secures association with the decellularized tissue. One type of association or interaction present in constructs of the invention is via anchoring or piercing of the decellularized tissue surface, as detailed herein, or by forming holes in the tissue through which two polymer sheets or segments may be associates. In some embodiments of the invention, where the construct is made of a plurality of construct assemblies each assembly comprising e.g., a decellularized tissue confined between two sheets or segments of a polymer component, association of the plurality of assemblies may be achieved by polymer-to-polymer welding.

[0475] According to this embodiment, the welding between the tissue (associated with the polymer) and the polymer-coated metal is enabled by associating (e.g., physically contacting) the tissue (already associated with the polymer) and the at least partially polymer-coated metal element together to facilitate the welding.

[0476] According to another embodiment, at least one solvent (e.g., THF, DMF) is dripped into the contact area between tissue (associated with the polymer) and the at least partially coated metal element. According to another embodiment the solvent partially dissolves the coating polymer (of both the polymer that is associated with the tissue and the polymer-coated metal, at the welding area) and fuses them together to provide the connection/welding of the tissue and the metal.

[0477] According to another embodiment, the use of the solvent at least partially to dissolves the uppermost surface of the coating polymer of the metal and the uppermost surface of the polymer associated with the tissue; thus, the polymer chains at said uppermost surfaces are mobile enough to entangle with chains in the other sheet or segment (and thereby fuses the tissue and metal together). Such an effect is a reflow effect. It is within the scope of the present invention where the term reflow are referred to as a process in which at least partial dissolvement of the polymers occurs, and the rearrangement of the polymer's chains creates the welding.

[0478] According to another embodiment, the polymer is first dissolved in the at least one solvent and only then applied to the coat the tissue/metal elements.

[0479] According to one embodiment, the solvent can be THF, acetone, dioxane, DMSO, halogenated solvents such as chloroform and dichloromethane, alcohols and polyols, polyethers, acetonitrile, ethyl acetate, dimethylformamide (DMF), Dimethylacetamide, DMAC, and any combination thereof.

[0480] According to another embodiment, the use of the solvent, when applied to thermal heating, at least partially to melt/dissolves uppermost surface of the coating polymer of the metal and the uppermost surface of the polymer associated with the tissue. It is within the scope of the present invention where the terms thermal reflow or thermal bonding are referred to as a process in which dissolving occurs, after which the entire assembly is subjected to controlled heat (which results in a molten state of at least part of the polymer) to creating permanent casting (and welding) thereof. The coating of the metal element could be provided by dip coating and/or brushing and/or spraying.

[0481] Reference is now made to FIG. 34C illustrating the encapsulated frame 40 with the polymer (performed by dip coating or spraying) and the tissue 10 associated with the polymer on the specific connection area with the metal frame.

[0482] Thus, according to this embodiment, the metal frame is coated with the polymer and the tissue is coated by the same polymer at the connection area/welding areas.

[0483] The metal frame could be any element selected from stents, metallic stents, vascular grafts, heart valves, membrane, sealing devices, suture or staple lines, hernia meshes or hernia repair devices, pelvic floor reconstruction devices, wound or burn dressings, dural closures and cardiac patches.

[0484] According to another embodiment, the tissue is freeze dried before the polymer is applied thereto.

[0485] It should be pointed out that although the above and below discloses tissue to metal connection, the same can be applied to any other material (e.g., polymeric material) having substantially the same mechanical properties.

[0486] The following measurements provides testing of the mechanical strength of the polymers and their welding to the tissue/metal frame: [0487] 1. The strength of the polymers' thickness on the welding to each other on top of metal stent:

TABLE-US-00014 Frame coated Frame coated Frame coated Chronoflex C75A Chronoflex C93A Chronoflex C75D Chronoflex C75A, Force(N): 13.3 Force(N): 5.63 Force(N): Thickness of 60 Stress (MPa): 37.0 Stress (MPa): 15.63 Stress(MPa): micron Chronoflex C75A, Force(N): 13.1 Force(N): 14.1 Force(N): 11.15 Thickness of 100 Stress (MPa): 21.78 Stress (MPa): 23.5 Stress (MPa): 18.58 micron Chronoflex C75A, Force(N): 16.3 Force(N): 9.5 Force(N): 10.33 Thickness of Stress (MPa): 20.94 Stress (MPa): 12.1 Stress (MPa): 13.25 130 micron Chronoflex C93A, Force(N): 8.77 Force(N): Force(N): Thickness of 60 Stress (MPa): 24.4 Stress (MPa): Stress (MPa): micron Chronoflex C93A, Force(N): 11.89 Force(N): 16.2 Force(N): 15.28 Thickness of 100 Stress (MPa): 19.81 Stress (MPa): 26.96 Stress (MPa): 25.5 micron Chronoflex C93A, Force(N): 16.24 Force(N): 18.22 Force(N): 20.55 Thickness of 130 Stress (MPa): 20.82 Stress (MPa): 23.36 Stress (MPa): 26.35 micron Chronoflex C75D, Force(N): 14.51 Force(N): 6.36 Force(N): 5.69 Thickness of 60 Stress (MPa): 40.3 Stress (MPa): 17.67 Stress (MPa): 15.8 micron Chronoflex C75D, Force(N): Force(N): 13.25 Force(N): 18.234 Thickness of 100 Stress(MPa): Stress (MPa): 22.08 Stress (MPa): 30.39 micron Chronoflex C75D, Force(N): 26.84 Force(N): 22.3 Force(N): Thickness of 130 Stress (MPa): 34.42 Stress (MPa): 28.6 Stress (MPa): micron

[0488] As seen in the table, the best mechanical properties were obtained by the combination of Chronoflex 75A and Chronoflex 75A; and, Chronoflex 75A with Chronoflex 75D.

[0489] A medical device examplea stichless heart valve using either the tissue-to-tissue welding and/or the tissue to metal welding, as described in any of the above, is provided as follows:

[0490] According to this example, the heart valves include 4 components:

[0491] Metal stent (or frame) made of nitinol, cobalt chrome or stainless steel 11; usually, three leaflets 12, inner skirt 13 and outer skirt 14 (see FIG. 35A), in addition, as disclosed above and below, the heart valve includes polymeric-based components. As disclosed hereinafter, the polymer component is used both in its liquid form or as a sheet or thin film.

[0492] According to one embodiment of the present invention, first the all-tissue parts are laser cut and holes are created on the leaflets, inner skirt and the outer skirt, made in 2D/3D laser cutting (see FIG. 36A).

[0493] Next, the polymer is either dissolved in at least one solvent (e.g., THF) or taken as is, casted as a thin film (thickness of 10-100 micron) and then laser cut to create polymer sheets and rivets (for future connection of the different parts of the tissue).

[0494] According to one embodiment a polymer sheet is casted and laser cut with the shape of the 2D projection of the assembled heart valve (see FIG. 35B).

[0495] According to one embodiment, the polymer sheet is pierced and according to another embodiment, the polymer sheet is maintained pierce-less.

[0496] Next, freeze-drying (lyophilization) of the tissues takes place at a temp. 54 C., vacuum pressure decreases from 110K Pa to 2-3 Pa overnight.

[0497] Then, coating of the stent (frame) begins. Chronoflex in Tetrahydrofuran, THF, is used to provide at least one layer of coating on the stent, preferably more than 10 layers of coating is provided on the frame. Preferably, the metal stent (frame) is coated by means selected from brushing, spraying or dip coating the same. According to one embodiment of the present invention, in between at least one layer of coating to another, heating is applied to dry the coating layer.

[0498] Next, welding of the different tissue components to each other is achieved by welding the polymer sheets to the tissues on both sides. The welding is achieved by injecting through the holes in the tissues a solution of Chronoflex dissolved in THF. The solution of Chronoflex dissolved in THF results in adhesion of the tissues to each other. For further strengthening the adhesion thermo-bond heating is applied. The thermal bonding is performed in at least 6 bars and heating temperature of 80 C. for all the welded parts to enhance and strengthen the welding.

[0499] It is noted that at all times, all holes, of e.g., the leaflets and of the skirts, must be substantially aligned to result in improved welding.

[0500] Thermo-bond can be used to enhance and strengthen the welding.

[0501] Next, after the welding, the assembly can be placed in saline for the recovery of the tissues for at least 14 hours.

[0502] Next the tissue to metal process starts to provide the finalized heart valve. It is noted that while the tissue to tissue welding was performed while the tissues were dried, the tissue to metal welding is obtained when the tissue is wet (post recovery in saline.

[0503] In tissue to metal welding process the tissue-polymer-tissue assembly is placed inside the stent/frame and welding of the tissue-polymer-tissue to the polymer-coated stent is obtained by, again, the use the polymer. Namely, the polymer is used to coating the tissue-polymer-tissue assembly and the relevant area of the polymer-coated stent. Next, thermal bonding is applied. As noted above, welding of the polymers sheets occurs when the polymer chains at the surface of one sheet or segment are mobile enough to entangle with chains in the other sheet or segment. To achieve welding, thermal energy may be applied to raise the temperature. When two sheets or segments of the polymeric component are brought into intimate contact under these conditions, polymer chain entanglement will occur resulting in a weld. According to aspects of the invention, welding need not be achieved over the full surface of the sheet or segment. Point welding at one or more regions of the polymeric component may suffice to provide a robust association of the plurality of assemblies or any two polymeric sheets or segments.

[0504] It should be pointed out that the welding can be achieved in numerous placed along the circumference of the stent.

[0505] According to another embodiment the coated metal (i.e., polymer-metal-polymer) is encapsulated within an assembly of tissue-polymer-tissue; namely, a construct of tissue-polymer-tissue-polymer-metal-polymer-tissue-polymer-tissue is provided (see FIGS. 36B-36C). As seen in FIG. 36C part of metal frame is encapsulated within the tissue-polymer-tissue assembly.

[0506] It should be pointed out that according to one embodiment, at least one type of polymer component can be used. alternatively, two different polymers, each having a different viscosity, are used.

[0507] The all-welded-no-stiches-valve assembly was hemodynamic tested & durability tested. Reference is now made to FIG. 37, illustrating hemodynamic behavior of the valve according to the present invention.

[0508] In the test 2 major parameters presented:

[0509] EOAeffective orifice areawhich provides number to how well the valve is opened and allows blood to flow throw it (the higher the number the better it functions); and,

[0510] R.Fregurgitation fractionwhich provides % for how much blood leaks throw the valve backflow is in every pulse (namely, how much the valve is leaking).

[0511] As seen in the figures, the valve according to the present invention reached to 130M cycles out of 200M cycles required.

[0512] The terms comprises, comprising, includes, including, having and their conjugates mean including but not limited to

[0513] The term consisting of means including and limited to.

[0514] The term consisting essentially of means that the composition, method or structure may include additional ingredients, steps and/or parts, but only if the additional ingredients, steps and/or parts do not materially alter the basic and novel characteristics of the claimed composition, method or structure.

[0515] As used herein, the singular form a, an and the include plural references unless the context clearly dictates otherwise. For example, the term a compound or at least one compound may include a plurality of compounds, including mixtures thereof.

[0516] Throughout this application, various embodiments of this invention may be presented in a range format. It should be understood that the description in range format is merely for convenience and brevity and should not be construed as an inflexible limitation on the scope of the invention. Accordingly, the description of a range should be considered to have specifically disclosed all the possible subranges as well as individual numerical values within that range. For example, description of a range such as from 1 to 6 should be considered to have specifically disclosed subranges such as from 1 to 3, from 1 to 4, from 1 to 5, from 2 to 4, from 2 to 6, from 3 to 6 etc., as well as individual numbers within that range, for example, 1, 2, 3, 4, 5, and 6. This applies regardless of the breadth of the range.

[0517] Whenever a numerical range is indicated herein, it is meant to include any cited numeral (fractional or integral) within the indicated range. The phrases ranging/ranges between a first indicate number and a second indicate number and ranging/ranges from a first indicate number to a second indicate number are used herein interchangeably and are meant to include the first and second indicated numbers and all the fractional and integral numerals therebetween. As used herein the term method refers to manners, means, techniques and procedures for accomplishing a given task including, but not limited to, those manners, means, techniques and procedures either known to, or readily developed from known manners, means, techniques and procedures by practitioners of the chemical, pharmacological, biological, biochemical and medical ails. As used herein, the term treating includes abrogating, substantially inhibiting, slowing or reversing the progression of a condition, substantially ameliorating clinical or aesthetical symptoms of a condition or substantially preventing the appearance of clinical or aesthetical symptoms of a condition.

[0518] It is appreciated that certain features of the invention, which are, for clarity, described in the context of separate embodiments, may also be provided in combination in a single embodiment. Conversely, various features of the invention, which are, for brevity, described in the context of a single embodiment, may also be provided separately or in any suitable subcombination or as suitable in any other described embodiment of the invention. Certain features described in the context of various embodiments are not to be considered essential features of those embodiments, unless the embodiment is inoperative without those elements. Various embodiments and aspects of the present invention as delineated hereinabove and as claimed in the claims section below find experimental support in the following examples.