The disclosed processes, methods, and devices are directed to a repairing defects in mammalian tissue, for example a cardiac patch device for repairing heart defects.
This application claims benefit of priority pursuant to 35 U.S.C. § 119(e) of U.S. provisional patent application No. 62/886,795 entitled “POLYURETHANE-REINFORCED HYDROGEL CARDIAC PATCH,” filed on 14 Aug. 2020, which is hereby incorporated by reference in its entirety.
Congenital heart defects (CHD) affect 1 of every 111 to 125 births in the United States. An estimated 40,000 infants are affected by CHD each year; of these, about 25% require invasive treatment in the first year of life. Surgical repair of CHD often requires the use of a polymer or fixed tissue patch to close septal defects or enlarge stenosed structures. Approximately 50% of Tetralogy of Fallot repairs include a patch in the right ventricular (RV) outflow tract. Currently, surgeons use synthetic or biological materials, including knitted polyethylene terephthalate (most commonly Dacron®), expanded polytetrafluoroethylene (such as Gore-Tex®), and glutaraldehyde-fixed bovine pericardium (such as SJ Medical and CardioCel®). These materials do not grow with the pediatric patients, are not electromechanically integrated, have mismatched mechanical properties compared with the surrounding tissue, and often become fibrotic, leading to an increased risk of malicious arrhythmia, sudden cardiac death, and heart failure. With the conventional patches discussed above, about 25% of patch-implanted patients require a second surgery.
The information included in this Background section of the specification, including any references cited herein and any description or discussion thereof, is included for technical reference purposes only and is not to be regarded subject matter by which the scope of the invention as defined in the claims is to be bound.
Disclosed herein are heart/cardiac patch devices useful in correcting various heart defects, as well as methods of manufacturing and methods of using same. The heart patch device of the present disclosure may be used to replace missing or damaged mammalian myocardium, wherein the devices provide for sufficient mechanical integrity, strength, electro- and bio-compatibility to support tissue repair, remodeling, and regeneration. In one embodiment, the disclosed cardiac patch device comprises a flexible but strong biodegradable polymeric scaffold capable of maintaining integrity, and a gel that enhances vascularization and cell infiltration. In many embodiments, the device may be seeded with a plurality of differentiated or non-differentiated cells, the cells may be autologous or non-autologous cells.
This disclosed device is capable of controlled degradation to allow sufficient time for the patient's cells and extracellular matrix to replace the degrading patch. The disclosed device may also be sufficiently porous to help initiate greater vascularization and tissue repair/remodeling while limiting the body's rejection, which may be in the form of fibrosis and/or a fibrotic response. The tendency of the disclosed devices to reduce fibrosis may result in fewer incidences of arrhythmia during repair/regeneration, and greater heart function compared with commercially available patches, e.g. those made of glutaraldehyde-crosslinked bovine pericardium, polyethylene terephthalate, and polytetrafluoroethylene.
Disclosed herein is a heart patch device comprising, a biodegradable polymeric mesh scaffold, and a biodegradable gel. In some embodiments, the the polymeric mesh scaffold may be made of a material comprising one or more of gelatin, polyurethane, and polycaprolactone, and the gel may comprises one or more of fibrin and polyethylene glycol. In many embodiments the device may further comprise one or more mammalian cells, that may be autologous cells and/or non-autologous cells. In most embodiments, where the device comprises mammalian cells, the cells may express one or more of CD31 (PECAM1), CD144 (VE-Cadherin), and CD309 (VEGFR2).
Also disclosed are methods for making a heart patch device, comprising forming a mesh scaffold by electrospinning, forming a gel layer, and combining the mesh scaffold with the gel layer. In some embodiments, the device may comprise an additional gel that may be added to the mesh scaffold after it is positioned adjacent the gel layer. In many embodiments the method may further comprise adding one or more mammalian cells to the device, wherein the cells may be autologous cells and/or non-autologous cells. In most embodiments, where mammalian cells are added to the device, the cells may express one or more of CD31 (PECAM1), CD144 (VE-Cadherin), and CD309 (VEGFR2). In embodiments wherein cells are added to the device, the device may be incubated for 1 or more days.
Also disclosed is a method of treating a subject having a heart defect, the method comprising, obtaining a heart patch device comprising a mesh scaffold and a biodegradable gel, placing the heart patch device at or near the heart defect, connecting the heart patch device to the heart with one or more sutures, and thereby treating the subject having the heart defect. In many embodiments the treatment may further comprise adding one or more mammalian cells to the device prior to implantation, wherein the cells may be autologous cells and/or non-autologous cells. In most embodiments, where mammalian cells are added to the device, the cells may express one or more of CD31 (PECAM1), CD144 (VE-Cadherin), and CD309 (VEGFR2). In embodiments wherein cells are added to the device, the device may be incubated for 1 or more days prior to implantation.
This Summary is provided to introduce a selection of concepts in a simplified form that are further described below in the Detailed Description. This Summary is not intended to identify key features or essential features of the claimed subject matter, nor is it intended to be used to limit the scope of the claimed subject matter. A more extensive presentation of features, details, utilities, and advantages of the present invention as defined in the claims is provided in the following written description of various embodiments and implementations and illustrated in the accompanying drawings.
The patent or application file contains at least one drawing executed in color. Copies of this patent or patent application publication with color drawing(s) will be provided by the Office upon request and payment of the necessary fee.
This disclosure is related to cardiac implants for correcting heart defects and methods of manufacturing such implants. An implant of the present disclosure may be used to replace heart tissue such as myocardium. The disclosed implants are sufficiently flexible, porous, and strong to provide a stable environment for new cell growth/invasion during repair/regeneration. Thus, the disclosed implants are suitable for use in non-heart tissues and organs, for example muscle tissue, and organs such as diaphragm, bladder, and uterus. In some cases the implants may be used to aid in wound repair. The disclosed devices provide mechanical integrity, biocompatibility, and support for repair and regeneration of missing or damaged tissue. In many embodiments, the disclosed biodegradable polymer-reinforced hydrogel heart/cardiac patch device is designed to maintain its structural integrity it degrades to allow the body's cells and secreted extracellular matrix to repopulate the defect or injury. The device is also designed to initiate and support great vascularization while limiting the body's rejection, in the form of fibrosis. The disclosed heart patch device may result in fewer incidences of arrhythmia, during repair/remodeling, as well as provide for greater heart function compared with commercially available natural and synthetic patches. In most embodiments, the disclosed cardiac patch device possesses mechanical properties approximating native tissue, allowing it to more closely resemble the performance of native tissue.
The presently implant device may be used in various tissues for repair and/or remodeling of the tissue. In many embodiments, the device may be useful in repairing congenital defects. In other embodiments, the disclosed device may be useful in repair of injuries such as hernias, for example diaphragmatic hernias. In other embodiments, the disclosed device may be useful in repair of injuries such as wounds, especially large wounds.
The presently disclosed heart patch device may include a gel portion and a porous mesh scaffold. In many embodiments, the gel may comprise one or more bioactive, biodegradeable materials, and the porous mesh framework may be manufactured of one or more biodegradable polymeric materials, in one example biodegradable polyurethane (BPUR). In many embodiments, the gel portion may be one or more discrete layers and/or may be contained within the porous mesh scaffold. In some embodiments, heart patch device may include a gel layer and a porous mesh scaffold layer adjacent the gel layer. In some embodiments, additional gel material may be contained within the mesh scaffold layer.
The disclosed cardiac patch device may be of various sizes and shapes. In some embodiments, the cardiac patch device may be sized while it is fabricated, and in other embodiments it may be reduced in size after it is fabricated, for example by cutting or trimming. Adjusting the size of the heart patch device may aid in it being readily degradable and may facilitate its integration into mammalian tissue.
The disclosed implant patch may have various shapes. In many embodiments, the shape of the implant device may be modified by a medical professional prior to implantation. In some embodiments, the device may be, for example, generally round, oval, square, rectangular or irregularly shaped. In many embodiments, the implant device may define an average diameter of about 0.5-5 cm, and an average thickness of about 0.05-2 cm. Average diameter is not meant to limit the shape or size of patches to round objects, rather all shapes (e.g. squares, ovals, triangles, etc.) may define an average diameter.
The disclosed mesh scaffold may be constructed using various methods. In some embodiments, the mesh scaffold is constructed by one or more methods that produce a porous material that is greater than 60% porous, as measured gavimetrically. For example, the porosity of the mesh scaffold may be greater than about 65%, 70%, 72%, 74%, 76%, 78%, 80%, 82%, 84%, 86%, 88%, 90%, 91%, 92%, 93%, 94%, 95%, 96%, 97%, 98%, or 99%, and less than about 100%, 99%, 98%, 97%, 96%, 95%, 94%, 93%, 92%, 91%, 90%, 89%, 88%, 87%, 86%, 85%, 80%, 75%, 70%, or 65%. In some embodiments the mesh framework may be constructed from a method that creates an irregular or regular pattern. In one embodiment, the method of manufacture may be electrospinning. In other embodiments, the method of manufacture may be selected from 3D printing, additive manufacturing, molding, spin coating, electrospinning, and combinations thereof. The disclosed mesh scaffold may have an average thickness between about 20 μm and 400 μm, for example greater than about 20 μm, 30 μm, 40 μm, 50 μm, 60 μm, 70 μm, 80 μm, 90 μm, 100 μm, 110 μm, 120 μm, 130 μm, 140 μm, 150 μm, 200 μm, 250 μm, 300 μm, or 350 μm, and less than about 4000 μm, 350 μm, 300 μm, 250 μm, 200 μm, 150 μm, 140 μm, 130 μm, 120 μm, 110 μm, 100 μm, 90 μm, 80 μm, 70 μm, 60 μm, 50 μm, 40 μm, and 30 μm.
One or more nanoparticles may be incorporated into the mesh scaffold during fabrication to help regulate the size and number of pores. In most embodiments, the nanoparticle is removable, for example dissolvable or degradable. In one embodiment, the nanoparticle is composed of polyethylene oxide (PEO).
The mesh scaffold may be trimmed and/or shaped after fabrication. In many embodiments, for example wherein the mesh is electrospun, the edges of the mesh may be manually trimmed to remove edges and/or to attain a desired shape. In some embodiments, channels through the mesh scaffold may be created after fabrication, for example by inserting a needle or other object through the scaffold. In some embodiments, the channels are formed by extracting material from the fabricated scaffold.
The polymeric material of the mesh scaffold may be biodegradeable, elastic, and with a tensile strength suitable for resisting stresses in a target tissue. In many embodiments, that target tissue is cardiac tissue, and the suitable tensile strength is greater than about 100 kPa. In various embodiments, the disclosed mesh scaffold may have an average Young's Modulous of less than about 50 Mpa, 45 MPa, 40 MPa, 35 MPa, 30 MPa, 25 MPa, 20 MPa, 15 MPa, 10 MPa, 5 MPa, 4 MPa, 3 MPa, 2 MPa, 1 MPa, 0.9 MPa, 0.8 MPa, 0.7 MPa, 0.6 MPa, 0.5 MPa, 0.4 MPa, 0.3 MPa, 0.2 MPa, 0.1 MPa, 0.09 MPa, 0.05 MPa, or 0.02 MPa, and greater than about 0.01 MPa, 0.02 MPa, 0.03 MPa, 0.04 MPa, 0.05 MPa, 0.06 MPa, 0.07 MPa, 0.08 MPa, 0.09 MPa, 0.1 MPa, 0.2 MPa, 0.3 MPa, 0.4 MPa, 0.5 MPa, 1.0 MPa, 1.5 MPa, 2 MPa, 3 MPa, 4 MPa, 5 MPa, 10 MPa, 15 MPa, 20 MPa, 25 MPa, 30 MPa, 35 MPa, 40 MPa, or 45 MPa. In many embodiments, the mesh scaffold material may comprise one or more natural and/or synthetic compounds, for example, proteins, polymeric proteins, monomers, polymers, diols, diisocyanates, esters, ethers, urethanes, gelatins, lactones, etc. In some embodiments, the material may be selected from one or more of poly lactone, polycaprolactone, poly urethane, poly urethane urea, poly (ether-ether) urethane, poly (ether-ester) urethane, and poly (ether-ester) urethane urea. In many embodiments, the material may comprise a polymer having a ‘hard segment,’ rigid group,' or ‘hard group,’ and a non-hard segment or linker. In some embodiments, the hard group may comprise about 5-50% of the polymer, for example 10% of the polymer. The amount of hard segment may be varied to alter the performance characteristics (such as stiffness, elasticity, degradability, etc.) of the polymer.
The gel may be comprised of various compounds, for example natural and/or synthetic compounds that degrade in the mammalian body, support vascularization, and cellular reconstruction. In one embodiment, the gel may comprise a mammalian clotting agent, such as fibrin. The compounds may be combined with other compounds to aid in stability, vascularization, etc. In one embodiment the compounds may be covalently decorated with polyethylene glycol (PEG), for example PEGylated human fibrinogen. The disclosed gel may be in the form of a layer, having a thickness between about 800 μm to 2 mm, for example greater than about 810 μm, 820 μm, 850 μm, 900 μm, 950 μm, 1000 μm, 1100 μm, 1200 μm, 1300 μm, 1400 μm, 1500 μm, 1600 μm, 1700 μm, 1800 μm, or 1900 μm, and less than about 2000 μm, 1900 μm, 1800 μm, 1700 μm, 1600 μm, 1500 μm, 1400 μm, 1300 μm, 1200 μm, 1100 μm, 1000 μm, 950 μm, 900 μm, and 850 μm.
The gel may be formed by allowing the gel to solidify in a mold of various shapes and sized. In many embodiments the gel may be formed between plates of a given separation distance. The gel may compressed after formation by about 5% to 40%, for example more than about 5%, 6%, 7%, 8%, 9%, 10%, 11%, 12%, 13%, 14%, 15%, 16%, 17%, 18%, 19%, 20%, 21%, 22%, 23%, 24%, 25%, 30%, or 35%, and less than about 40%, 35%, 30%, 25%, 24%, 23%, 22%, 21%, 20%, 19%, 18%, 17%, 16%, 15%, 14%, 13%, 12%, 11%, 10%, 9%, 8%, 7%, or 6%.
The presently disclosed patches may be fabricated by various methods. In one embodiment, the patches may be reinforced fibrin-containing patches. In some embodiments, the patches are fabricated by mixing PEGylated human fibrinogen with human thrombin plus. In some embodiments, one or more cells may be mixed with the fibrinogen and thrombin, for example stem cells and or endothelial cells, such as human umbilical vein endothelial cells (HUVECs) or amniotic fluid stem cells (AFSCs). In various embodiments the patch may be reinforced with one or more BPUR layers. In many embodiments, the cell-impregnated patches may be cultured in a tissue culture environment in various appropriate media. In one embodiment, the media may be GEM-2 media. The patch may be cultured for various periods, for example from about 0-24 hours or 1-14 days, for example 3 days, before the patch is implanted.
The disclosed devices could be seeded with a combination of endothelial cells and/or support cells. In some embodiments, seeding may be useful to enhance vascularization and other aspects of the device's integration. The seeded cells could be autologous or provided from matched donors and/or cells. Seeded cells could be differentiated or undifferentiated. In many embodiments, seeded cells may be positive for CD31 (PECAM1), CD144 (VE-Cadherin), and CD309 (VEGFR2) and may be obtained from amniotic fluid, umbilical cord, adult stem cells, or induced pluripotent stem cells. The seeded cells may be endothelial cells that may be sorted or verified positive for one or more surface markers, for example one or more of. In some embodiments, the seed cells may be one or more of fractionated or unfractionated mesenchymal stem cells, general fibroblasts, smooth muscle cells, etc. The seeded cells may be combined and seeded into the patch along with the PEG-fibrin. These patches could be implanted immediately, or cultured for up to 2 weeks to allow for microvessel formation, and then implanted. In some embodiments the seeded cells may be c-Kit+ cells, for example stem cells and/or endothelial cells, for example umbilical epithelial cells. In various embodiments, the cells may be c-Kit+ cells from amniotic fluid.
Disclosed herein is an improved implant device, with beneficial characteristics. In some embodiments the implant device is a cardiac/heart patch device. The disclosed heart patch device may comprise an angiogenic polymer, for example a poly(ethylene glycol) fibrin-biodegradable hydrogel, and may be reinforced with polymeric scaffold, for example an electrospun biodegradable poly(ether ester urethane) urea (BPUR) mesh layer. The disclosed implant device may enhance cell invasion, angiogenesis, and regenerative remodeling. The disclosed heart patch device may aid in improving defect closure, cardiac output, and heart function as measured by echocardiogram, electrocardiogram, and pressure loop measurements. In many embodiments, effectiveness of the implant device may be measured by analyzing extent of fibrosis, macrophage infiltration, vascularization, etc. post-implantation. Compared with traditional fixed pericardium patches, the disclosed reinforced hydrogel patches result in fewer arrhythmias and greater ventricular ejection fraction, fractional shorting, stroke work, and cardiac output. Further, implanted patches of the present disclosure may degrade at a higher rate. Less of the disclosed implant device was shown to remain in the body at 4- and 8-weeks post-implantation, as compared to conventional patches.
Experiments were performed to create a cardiac patch with reduced induction of fibrosis and improved vascularization. Previous research indicated that right ventricular (RV) wall replacement with a multi-layered patch composed of a chitosan-gelatin-heart matrix hydrogel reinforced with a polycaprolactone (PCL) membrane resulted in higher RV ejection fractions compared with fixed bovine pericardium at 8 weeks post surgery. However, the multi-layered patch induced significant fibrosis in the RV wall and relatively poor vascularization. In order to increase vascularization, a gel composed of fibrin covalently decorated with poly(ethylene glycol) (PEG) was developed and tested in a subcutaneous mouse model. The disclosed gel induced rapid gel vascularization with increased scaffold stability compared with fibrin alone. In addition, biodegradable polyurethanes with hydrolytically or enzymatically cleavable moieties were selected as an alternative to PCL, which takes 2-3 years to resorb in vivo, to reduce fibrosis.
A myocardial replacement patch of PEG-fibrin reinforced with an electrospun poly(ether ester urethane) urea mesh layer was fabricated. This engineered cardiac patch was tested in an RV wall replacement model in adult rats and compared with a sham surgery control and a clinical control of glutaraldehyde-fixed pericardium. Heart function was measured at 4- and 8-weeks post surgery, and histologic sections were evaluated for fibrosis, macrophage infiltration, vascularization, and defect size.
A biodegradable poly(ether ester urethane) urea (BPUR) with 10% hard segment was synthesized, for example, as described in A. P. Kishan, T. Wilems, S. Mohiuddin, E. M. Cosgriff-Hernandez, Synthesis and Characterization of Plug-and-Play Polyurethane Urea Elastomers as Biodegradable Matrixes for Tissue Engineering Applications, ACS Biomaterials Science & Engineering 3(12) (2017) 3493-502, which is hereby incorporated by reference herein in its entirety. For example, a poly(ether ester) triblock was synthesized by reacting poly(ethylene glycol) diisocyanate (PEG-DI) and polycaprolactone (PCL, MW=530 Da). PCL with stannous octoate (0.1 wt % with respect to the polymer) was added dropwise into a flask containing PEG-DI to a final PEG-DI:PCL molar ratio of 1:2 under nitrogen with stirring at 80° C. for 7 hours. BPUR was then synthesized in a two-step process from this triblock diol and hexane diisocyanate (HDI) using ethylene diamine (ED) as a chain extender at a molar ratio of 1:2:1 triblock diol:HDI:ED. A 10 wt % solution of the triblock diol in N,N-dimethylformamide (DMF) containing 0.1 wt % stannous octoate was first added dropwise to a flask containing a 10 wt % solution of HDI in DMF under nitrogen. The reaction proceeded at 80° C. under a nitrogen blanket with constant stirring until no change in the hydroxyl stretch was observed via transmission Fourier transform infrared (FTIR) spectroscopy (about 5 hours). The reaction was then cooled to room temperature. Chain extension was then performed by adding a 10 wt % solution of ED in DMF dropwise to the prepolymer solution under vigorous stirring. The BPUR chemical structure was confirmed using transmission FTIR spectroscopy. Neat BPUR films were cast onto KBr pellets from 1,1,1,3,3,3-hexafluoro-2-propanol (HFIP, Sigma, St. Louis, Mo.) solutions (5 wt %) and placed under vacuum for 1 hour under ambient conditions to remove the solvent. Spectra were recorded using a Nicolet iS10 (Thermo Scientific) FTIR spectrometer at a resolution of 2 cm-1 for 64 scans. Full reaction was confirmed by absence of the isocyanate peak at 2267 cm-1 (
A 10 wt % solution of BPUR in HFIP was used to electrospin fibrous meshes. BPUR solutions were dispensed using a syringe pump at a constant rate of 0.3 ml/hour. A positive voltage of 7.5 kV was applied at the needle tip, which was placed 17 cm from a −5-kV charged copper plate. Electrospinning was performed at ambient conditions (25° C., 45-55% relative humidity). The resulting fabricated electrospun meshes were vacuum dried for a minimum of 12 hours prior to characterization. The average thickness of BPUR meshes was 80±10 μm (n=5) (
BPUR support layers (7 mm in diameter) were cut from the electrospun BPUR mesh. To enhance gel integration with the mesh, 5 evenly spaced holes were made with a 22-gauge needles at about 2-mm away from the edge of the mesh, and BPUR meshes were sterilized by UV exposure in a culture hood for 1 h before use. PEGylated human fibrinogen (Sigma, St. Louis, Mo.) was prepared, for example, as described in O. M. Benavides, J. P. Quinn, S. Pok, J. Petsche Connell, R. Ruano, J. G. Jacot, Capillary-like network formation by human amniotic fluid-derived stem cells within fibrin/poly(ethylene glycol) hydrogels, Tissue Eng. Part A 21(7-8) (2015) 1185-94, which is hereby incorporated by reference herein in its entirety. For example, Human fibrinogen (F3879; Sigma-Aldrich, St Louis, Mo.) was solubilized in phosphate-buffered saline (PBS) at a concentration of 40 mg/ml. After 1 h of incubation at 37° C. and brief vortexing, the solution was sterilized using a 0.20-μm filter. Succinimidyl glutarate-modified bifunctional PEG (3.4 kDa SG-PEG-SG; NOF America Corporation, White Plains, N.Y.) was dissolved in PBS at 4 mg/ml and syringe filtered. Fibrinogen and PEG solutions were combined in a 1:1 volume ratio, mixed thoroughly, and incubated at 37° C. for 1 h.
Gel patches were fabricated in a sterile Teflon mold 7-mm in diameter and 2-mm deep. The PEG-fibrin gel was made with 30 μl sterile PBS containing 20 mg/ml of PEGylated human fibrinogen, adding 30 μl sterile ice-cold 40 mM CaCl2 containing 20 U/ml human thrombin (Sigma, St. Louis, Mo.) and mixing well to promote gel formation. The storage and loss moduli of the PEGylated fibrin gels were measured using a parallel-plate rheometer (Discovery Hybrid 2; TA Instruments). PEG-Fibrin gels (80 μl) were formed directly between the plates at a gap of 1400 μm. Samples were allowed to gel for 5 minutes at 37° C. before being compressed 20% (gap of 1120 μm). Four replicates were subjected to shear at 1% strain through a dynamic angular frequency range of 0.1 to 100 rad/s. The elastic modulus was calculated from the linear region of the storage modulus using Hooke's law. Poisson's ratio was assumed to be 0.5, corresponding to an incompressible material (
For a control, bovine pericardium patches (7 mm in diameter, 270 μm in thickness) were cut from commercially available glutaraldehyde-fixed bovine pericardium (St. Jude Medical, Saint Paul, Minn.).
Experiments were performed to implant the cardiac patch, as described above in Example 1, over a mammalian heart defect. Sixty-two male Sprague-Dawley rats weighing 400 ±10 g (Envigo, Cambridgeshire, UK) were randomly assigned to a 4-week sham group (Sham, n=10), a 4-week glutaraldehyde-fixed bovine pericardium group (Pericardium, n=10), a 4-week BPUR-reinforced hydrogel patch group (BPUR PEG-fibrin, n=10), an 8-week sham group (n=11), an 8-week pericardium group (n=10), or an 8-week BPUR PEG-fibrin group (n=11). Rats were anesthetized using 5% isoflurane inhalation with 100% oxygen followed by intubation and respiratory support with a rodent mechanical ventilator (Harvard Apparatus, Holliston, Mass.) at a peak inspiratory pressure of 11 cmH2O and 75 beats/minute. Surgery procedures were performed in a sterile environment on a controlled heating pad. An Animal Bio Amp (FE 136, ADInstruments) and an Animal Oximeter Pod (ML325, ADInstruments) that attached to a PowerLab 4/30 system (ADInstruments, Spring, Colo.) were used to monitor electrocardiogram and SpO2. Anesthesia was maintained with 2% isoflurane inhalation with 100% oxygen. The rat heart was exposed via a 3-cm incision through a 4th left thoracotomy, for example, as described in Z. W. Tao, L. G. Li, Z. H. Geng, T. Dang, S. J. Zhu, Growth factors induce the improved cardiac remodeling in autologous mesenchymal stem cell-implanted failing rat hearts, J Zhejiang Univ Sci B 11(4) (2010) 238-48 (herein “Tao”), which is hereby incorporated by reference herein in its entirety. FIGS. 2A1-4 show an exemplary surgical procedure on a rat heart to create a defect and implant a patch of the present disclosure. For example, as shown in FIG. 2A1, a purse string suture, 4 mm in diameter, was created on the RV free wall with a 6-0 polypropylene suture (Ethicon, US). Both ends of the stitch suture were passed through a 22-gauge plastic vascular cannula (VWR International, Radnor, Pa.) that served as a tourniquet to secure the purse string. As shown in FIG. 2A2, the distend portion was excised to create a full-thickness defect. For example, as shown in FIGS. 2B1 and 2B2, three quarters of the bulging part of the purse string was excised to create approximately a 2-3 mm full-thickness defect contacting blood. As shown in FIG. 2A3, the patch was stitched over the defect with 7-0 polypropylene suture (Ethicon, US), first fixed by 4 stitches at positions of 90°, 180°, 270° and 360°, then continuously sutured fully around the patch. As shown in FIG. 2A4, the purse string suture was released. Animals in the sham group experienced the same chest opening and pericardium tearing, but no defect was created and no patch was sutured onto the RV free wall. The muscle layers of the chest and the skin were closed with a 4-0 polyglactin absorbable suture (AD Surgical, Sunnyvale, Calif., USA).
Isoflurane supply was stopped immediately after the skin layer was closed. Before animals regained consciousness, Meloxican (5 mg/ml, MWI Animal Health, Grand Prairie, Tex., USA) 0.5 mg/kg was administered subcutaneously once to reduce post surgery pain. One animal in the 4-week pericardium group and 1 animal in the 8-week BPUR PEG-fibrin group died from massive bleeding during surgery, and 1 animal in the 4-week BPUR PEG-fibrin group died from acute RV myocardial infarction during surgery.
Experiments were performed to assess cardiac function resulting from implantation of a cardiac patch of the present disclosure over a mammalian heart defect. As one example, to assess left ventricular (LV) function, echocardiography was performed at the end of the 8-week time point. Animals were anesthetized and placed on a controlled heating pad, and anesthesia was maintained with 2% isoflurane inhalation with 100% oxygen. Standard transthoracic echocardiography was performed using the GE Vivid 7 system (GE Vingmed Ultrasound AS, N-3190 Horten, Norway) fitted with an GE S10 transducer. LV parameters were obtained from two dimensional images and M-mode interrogation in parasternal short-axis and long-axis view as described in Tao, and then LV end-diastolic dimensions (LVEDd), ejection fraction (LVEF) and fractional shortening (LVFS), and end-diastolic area (LVEDA) were calculated. Echocardiographic measurements were averaged from at least five cardiac cycles.
Electrocardiogram (ECG) may also be used to assess cardiac function. For example, ECG signals were recorded at 4- and 8-week post-implantation endpoints. Animals were anesthetized using 5% isoflurane inhalation with 100% oxygen followed by intubation and respiratory support with a rodent mechanical ventilator, placed on a controlled heating pad, and then anesthesia was maintained with 2% isoflurane inhalation with 100% oxygen. An ECG signal was recorded for 30 min with an Animal Bio Amp that attached to a PowerLab 4/30 system, by inserting a needle anode (MLA1213, ADInstruments) into the left front leg of the animal, a needle cathode into the right front leg, and using the testis skin as ground. LabChart (ADInstruments) was used for analysis of malicious arrhythmia, categorized as frequent atrial premature beats (APBs), atrial tachycardia (AT), atrial fibrillation, frequent ventricular premature beats (VPBs), ventricular tachycardia, and ventricular fibrillation.
Cardiac catheterization may be performed to assess hemodynamics in hearts implanted with the disclosed device. For example, after recording ECG signals, the heart was exposed through a 5th left thoracotomy and a 2F micromanometer tipped catheter (SPR-869 Millar Instruments, Houston, Tex.) was inserted into the LV apex, and advanced into the LV to obtain LV pressure and conductance. After stabilization for 15 minutes, the signals were digitized at a sampling rate of 1 kHz/s using MPVS-300 (Millar Instruments) and were acquired to a PowerLab 4/30 system at steady state. LabChart Pro v.8.10 software with the pressure-volume (PV) loop module (ADInstruments) was utilized for subsequent assessment of LV hemodynamic parameters. Heart rate (HR), LV systolic pressure (LVSP), LV end-diastolic pressure, maximal slope of systolic pressure increment (LV dP/dtmax) and diastolic pressure decrement (LV dP/dtmin), ejection fraction (LVEF), stroke volume (SV), end-diastolic volume (EDV), cardiac output (CO), and stroke work (SW) were computed using the cardiac PV-loop module. After completion of the hemodynamic assessment from of the LV, the catheter was inserted into the RV apex and advanced into the RV to acquire RV hemodynamics including systolic pressure (RVSP), RV end-diastolic pressure (RVEDP), maximal slope of systolic pressure increment (RV dP/dtmax), and diastolic pressure decrement (RV dP/dtmin). A 20-gauge IV catheter (VWR International, Radnor, Pa.) was inserted into the right jugular vein. After a stable signal was recorded from either LV or RV, 20 μl hypertonic saline (30%) bolus injection at least 2-3 times were performed for both ventricles to obtain a value for Vp for the saline calibration. After hemodynamic measurements were made under anesthesia, animals were euthanized with cardiac arrest by apical injection of 1 ml of 10% KCl. Hearts were excised, weighed, placed in a peel-away disposable embedding mold (VWR International, Radnor, Pa.), frozen in liquid N2, and then immediately immersed in Tissue Tek OCT compound (VWR International, Radnor, Pa.) and placed in a −80° C. freezer.
In some embodiments, immunohistochemistry may be conducted and histology may be studied. As one example, heart samples were sliced using a cryostat (Cryotome E, Thermo Shandon). Whole heart longitudinal sections, directly through the middle of the defect, from the base to apex of the heart were cut at a thickness of 10 μm. The sections were placed on VWR Microslides for preparation of morphological and immunofluorescence examinations. For measurements of patch implantation-induced fibrosis and defect thickness, whole heart sections were stained with Masson's trichrome reagents (Sigma) according to the manufacturer's protocol. Section images (200× magnification) were taken under Zeiss 2.1 microscope (Germany), and the images of whole heart sections were stitched together using the Series feature within the Zeiss microscopy software. The whole scar area (mm2) and the patch material remaining area (mm2) were measured by tracing the edge of the scar and the edge of the remaining patch materials in each patch area. The patch implantation-induced fibrosis area was calculated as the whole defect area minus the patch material remaining area. The size of the external scar was measured and expressed as the external curve length (mm) in each sample and averaged; the size of the internal defect was measured between the internal muscle breaks and expressed as the internal curve length (mm) in each sample and averaged.
For immunofluorescence staining, whole heart sections of 10 μm thickness directly through the middle of the defect were fixed in 4% paraformaldehyde at 4° C. for 20 min; nonspecific epitope antigens were blocked with 10% goat serum (Sigma) at room temperature for 45 minutes. Sections were incubated with specific mouse anti-α-actinin antibody (1:200, Sigma, A7811), rabbit anti-cardiac troponin T (1:200, Invitrogen, MA5-12960), mouse anti-vimentin (1:200, Sigma, C9080), rabbit anti-von Willebrand factor (vWF; 1:750, Abcam, ab6994), mouse anti-α-smooth muscle actin (α-SMA; 1:200, Sigma, C6198), rabbit anti-CD45 (1:200, Abcam, ab10558), mouse anti-CD68 (1:200, Invitrogen, MA5-16654), rabbit anti-CD206 (1:200, Abcam, ab64693), Alexa Fluor 488 (1:40, Invitrogen, A12379) and 546 (1:40, Invitrogen, A22283) phalloidin at room temperature for 1 hour. Subsequently, sections were treated with goat anti-mouse or goat anti-rabbit secondary antibodies (1:400, Invitrogen, Alexa Fluor 488, Alexa Fluor 546, and Alexa Fluor 647) at room temperature for 1 hour. Nuclei were counterstained with 4,6-diamidino-2-phenylindole (DAPI; 2.5 μg/ml) for 5 minutes at room temperature. Fluorescent images were obtained with a Zeiss 2.1 microscope. For determination of granular-like tissue, the volume of vimentin positive signals (staining fibroblasts and endothelial cells) were measured by whole defect area×intensity mean value in the section. For evaluation of blood vessels, the total number of vWF positive signals was counted from 5 random 400×magnification patch material-centered ocular fields in the section. For evaluation of acute inflammation, the volume of CD45, CD68 and CD206 positive signals (staining leukocytes, pan-macrophages and M2 macrophages respectively) were measured with the average of 3 random 200×magnification patch material-centered ocular fields calculated by area×intensity mean value in each section.
Experimental results were analyzed to determine outcomes resulting from the implantation of a cardiac patch of the present disclosure over a mammalian heart defect. Results are presented as mean±standard deviation. ECG arrhythmia was analyzed by Chi-square (and Fisher's exact) test. Comparisons between two groups were made using the independent-samples t-test, and comparisons among three groups were made using a one-way analysis of variance followed by a Tukey post hoc comparison test. Differences were considered statistically significant at a value of p<0.05.
In some embodiments, survival rates may be determined. For example, in the study discussed above, one animal in the 8-week pericardium group died from cardiac arrest the second day post surgery, and 1 animal in the 8-week BPUR PEG-fibrin group became paraplegic and was euthanatized the second day post surgery. All other animals surviving the surgery survived to the endpoint (57/62 total rats).
ECG was also used to assess arrhythmia in the test subjects. These results are depicted at
As shown, implantation of a fixed pericardium patch at 8 weeks post surgery resulted in a significant decrease in LVEDA (67.0±9.7 mm2, n=6;
Hemodynamics were also assessed. For example, Table 1 depicted below shows resulting body weight (BW), heart weight (HW), HW/BW, and heart rate (HR), as well as other ventricular hemodynamics measurements at 4- and 8-weeks post surgery. For example, the table below further shows right ventricular systolic pressure (RVSP); right ventricular end-diastolic pressure (RVEDP); right ventricular maximal slope of pressure increase (RV dP/dtmax); right ventricular maximal slope of pressure decrease (RV dP/dtmin); right ventricular end-diastolic volume (RVEDV); right ventricular stroke volume (RV SV); right ventricular ejection fraction (RVEF); right ventricular stroke work (RV SW); right ventricular cardiac output (RV CO); left ventricular systolic pressure (LVSP); left ventricular end-diastolic pressure (LVEDP); left ventricular maximal slope of pressure increase (LV dP/dtmax); left ventricular maximal slope of pressure decrease (LV dP/dtmin); left ventricular end-diastolic volume (LVEDV); left ventricular stroke volume (LV SV); left ventricular ejection fraction (LVEF); left ventricular stroke work (LV SW); and left ventricular cardiac output (LV CO). Values are mean±standard deviation. The single asterisk (*) represents p<0.05, and the double asterisk (**) represents p<0.01 vs Sham; while the single hashtag (#) represents p<0.05, and the double hashtag (##) represents p<0.01 vs Pericardium.
As shown in Table 1, four weeks after surgery, BW, HW, HW/BW, and HR were not significantly different between any experimental groups. Eight weeks after surgery, BW was significantly decreased (p<0.05) and HW/BW ratio was significantly increased (p<0.05) in both the Pericardium and the BPUR PEG-fibrin groups compared with the Sham group. As shown, LVSP, LVEDV, LVEF, LV SW and LV CO at both 4- and 8-weeks were dramatically decreased in the pericardium group, and LVSP and LVEF at 4 weeks, and LVEF at 8 weeks were dramatically decreased in the BPUR PEG-fibrin group compared with the sham control (p<0.05; p<0.01). LVEF was higher at 8 weeks, and LV SW and LV CO were higher at both 4-and 8-weeks post surgery in the BPUR PEG-fibrin group compared with the pericardium group (p<0.05).
RV pressure-volume (PV) may also be measured to assess heart function. For example,
As shown in
Histology may be performed to assess levels of fibrosis, infiltration, degredation, etc. in the inplanted heart patches. For example,
At 4- and 8-weeks post surgery, one third of patch-implanted hearts exhibited minimal thoracic adhesions. Neither group showed any dehiscence or aneurysm formation at the site of the implanted patch. In the BPUR PEG-fibrin group, the BPUR support layer degraded to an apparent loss of structure integrity and was replaced with native-like tissue at 4 weeks (FIG. 6A4) and further at 8 weeks (FIG. 6A5); however, the pericardium group showed no degradation or native-like tissue replacement at both 4 weeks (FIG. 6A2) and 8 weeks (FIG. 6A3) post surgery.
Masson's trichrome staining was used to evaluate fibrosis and wall thickness. As shown in FIGS. 6B1-B5, the pericardium group had significantly higher fibrotic area at both 4- and 8-weeks (10.93±1.97 mm2, n=7 and 8.43±1.67 mm2, n=7) compared with the BPUR PEG-fibrin group (5.52±1.06 mm2, n=7 and 6.23±1.39 mm2, n=7) (p<0.01; p<0.05), but the fibrotic area was significantly smaller at 8 weeks compared with at 4 weeks post surgery in the pericardium group (p<0.05) (
Areas of remaining patch material were measured to quantify degradation. s shown in FIGS. 6C1-C5, the patch material area in the pericardium group at 4 weeks (2.68 ±0.80 mm2, n=7) and 8 weeks (2.65±0.73 mm2, n=7) was greater than in the BPUR PEG-fibrin group (1.09 ±0.35 mm2, n=7 and 1.12±0.23 mm2, n=7) (p<0.01) (
Vimentin expression and vascularization was determined to assess tissue repair. For example, immunofluorescence staining of vimentin in the center of the patched area at 4- and 8-weeks post surgery was qualified and used as a measure of granulation tissue (FIG. 6D1-5). There were fewer vimentin-positive cells in the patch area in the pericardium group (FIG. 6D2 and D3) compared with the BPUR PEG-fibrin group (FIG. 6D4 and D5), indicating less granulation tissue (e.g., tissue repair) with the pericardium group than with the BPUR PEG-fibrin group. As shown in
Vascularization and macrophage infiltration was measured, as shown in
To measure vascularization, the number of blood vessels may be counted using immunofluorescence staining. For example, the number of blood vessels was counted using the immunofluorescence staining of vWF in the patch material-centered area at 4- and 8-weeks post surgery (see, e.g., FIGS. 7B1-B5). As shown, there was little ingrowth of blood vessels in the pericardium group at 4 weeks (FIG. 7B2) and 8 weeks (FIG. 7B3) compared with the BPUR PEG-fibrin group at 4 weeks (FIG. 7B4) and 8 weeks (FIG. 7B5). The number of blood vessels in the BPUR PEG-fibrin group at 4 weeks (142.4±17.6, n=7) and 8 weeks (162.1 ±14.9, n=7) was greater than in the pericardium group at 4 weeks (59.4±19.9, n=7) and 8 weeks (59.0±8.2, n=7) post surgery (p<0.01); the number of blood vessels in the BPUR PEG-fibrin group at 8 weeks was greater than at 4 weeks (p<0.05). However, the number of blood vessels in the sham group at 4 weeks (365.0±36.1, n=7) and 8 weeks (360.5±38.2, n=7) was much greater than in the two patch implantation groups (p<0.01) (
Macrophage infiltration may also be measured with immunofluorescence staining. For example, immunofluorescence staining of CD45, CD68 and CD206 was used to evaluate the infiltration of leukocytes, especially neutrophils and monocytes, pan-macrophages, and M2 macrophages (FIGS. 7C1-C5, D1-5 and E1-5) respectively at 4- and 8-weeks post surgery. There was a significant difference in CD45 expression between the pericardium group (1.40±0.54, n=7) and the BPUR PEG-fibrin group (2.34±0.47, n=7) at 8 weeks (p<0.01), and the CD45 expression significantly decreased between 4 weeks (2.77±0.75, n=7) and 8 weeks in the pericardium group (p<0.01) (
In several embodiments, the external scar size and the internal defect size may be measured. For example,
The above examples demonstrate that a cardiac patch comprised of PEG-fibrin reinforced by a BPUR mesh induced greater muscular and vascular ingrowth with a limited foreign body response compared to a commercial glutaraldehyde-crosslinked pericardium patch, resulting in improved heart function in an adult rat RV wall replacement model. At 8 weeks post surgery, rat hearts patched with BPUR PEG-fibrin had less fibrosis, a decreased patch material size, and increased infiltration of endothelial cells, leukocytes, pan-macrophages, and M2 macrophages compared to hearts patched with fixed pericardium. These regeneratively remodeled patches at 8 weeks resulted in fewer incidences of arrhythmia, greater RV function as shown by RVEF and RV CO, and greater LV function as shown by LVEF, LVFS, LV SW and LV CO compared to fixed pericardium. However, all patched hearts exhibited arrhythmias, decreased RV and LV function, and enlarged defect sizes compared with sham controls.
This study additionally found that the multilayer cardiac patch provided mechanical support of the full thickness defect and continued to support the ventricular wall as the material degraded and invading cells and secreted extracellular matrix replaced the implanted materials, as evidenced by the lack of dehiscence or aneurysm formation at the site of the implanted patch at both 4- and 8-weeks post surgery. The mechanical strength of the patch was provided by the electrospun BPUR mesh that had an average tensile modulus of 2.9±0.4 MPa, similar to the elastic modulus of a native muscle (approximately 10 MPa). Furthermore, this study showed that both the formation of granular tissue, indicated by vimentin-positive cell staining, and the infiltration of neutrophils, monocytes, and M2 macrophages were higher in rat hearts with BPUR-reinforced hydrogel patches than with pericardium patches at 8 weeks post surgery. Tissue regeneration and repair proceed in a cascade fashion beginning with a coagulation and inflammatory phase, followed by granulation tissue formation, which is characterized by proliferation of fibroblasts and new thin-walled, delicate capillaries, as well as infiltrated inflammatory cells in a loose extracellular matrix. Within the first few days after scaffold implantation, disruption of the tissue structure and subsequent cell damage initiates an acute inflammatory response with a rapid influx of innate immune cells, predominantly neutrophils, mast cells, and monocytes. Neutrophils and monocytes are of hematopoietic origin and are involved in phagocytosis and pathogen clearance. Upon activation, resident tissue macrophages are supplemented by an active recruitment of blood monocytes, which then differentiate into macrophages and dendritic cells in the scaffold. Depending on the scaffold properties, this is followed by an M1/TH1 cell dominated pro-inflammatory response or an M2/TH2 cell dominated pro-regenerative response. The former is characterized by the prolonged presence of M1 macrophages, and recruited fibroblasts typically acquire an activated phenotype, producing fibrous scar tissue. In contrast, the pro-regenerative process is dominated by M2 macrophages under influence of TH2 cell secreted cytokines.
This study also found that the BPUR-reinforced patch was more rapidly resorbed than the glutaraldehyde-fixed pericardium patch. This could be because M2 macrophages, which mediate regenerative remodeling, were more populous in the BPUR PEG-fibrin group than the pericardium group at 8 weeks. Additionally, faster degradation and more M2 macrophages coincided with a smaller defect size at 8 weeks post surgery in the BPUR PEG-fibrin group compared with the pericardium group. A regenerative remodeling response to the BPUR-reinforced hydrogel patch likely paved the way for better action potential conduction, resulting in the absence of arrythmia at 8 weeks post surgery, and improved mechanical performance in the patched area.
The defect sizes in this study, confirmed from images at postmortem, are about 2-3 mm in diameter; however, defects will be bigger in a heart under pressure and beating. In both patch materials, the defects grew larger both between 4- and 8-weeks post surgery, and wall thickness were thinner than the normal RV wall in the sham group, especially at 8 weeks post surgery.
Experiments were performed to create a cardiac patch with increased porosity, pore stability, cell infiltration. For example, sacrificial polyethylene oxide (PEO) particles were embedded within the electrospun scaffolds to increase both their pore size and porosity. As discussed above, polyurethane (PU) was first selected for electrospinning for its elastic behavior and strength, which simulate native cardiac tissue mechanics. However, when PU was electrospun with PEO, the PU was stretched and removal of the PEO particles resulted in pore collapse. To increase retention of pore size and shape, polycaprolactone (PCL) was incorporated into the PU material forming the scaffolds. PCL acts a stiffer material, and has been shown to be compatible with PEO,
Five different scaffold materials were made and tested. The materials comprised 25% gelatin mixed with 0% (magenta bars), 10% (orange bars), 20% (lime bars), 30% (green bars), or 75% (blue bars) biodegradable PU (e.g., synthesized according to Guan et al). The remaining portion of the material was PCL (80 kDa). These five composites were electrospun in the presence and absence of PEO (MW 8000) using a custom rotating co-electrospinning apparatus. Fiber size and pore size were measured with DiameterJ analysis of Scanning Electron Microscopy (SEM) images. A dynamic mechanical analyzer was used to measure tensile strength of the resulting scaffolds. Porosity was measured with the gravimetric method. Cell infiltration depth was assessed via H&E-stained slides of scaffolds seeded with human dermal fibroblasts (hDF) for two weeks.
As demonstrated by the graph of Fif. 9A, the addition of PEO significantly lowered the Young's Modulus for scaffolds fabricated from each PU formulations, except the 10% PU group (orange bar; p<0.05). The modulus of the 75% PU+PEO material was 2.19 MPa, making it the group closest to the control modulus of porcine ventricular tissue (0.05 MPa), but the Young's Modulous of scaffolds from this group was not significantly different those of the other groups (i.e. 0%, 20%, or 30% PU+PEO).
As depicted in
In this study, we fabricated a prevascularized BPUR reinforced PEG-fibrin patch as presently disclosed, for example by seeding human umbilical endothelial cells (HUVECs) and human c-Kit+amniotic fluid stem cells (AFSCs). The disclosed patch was tested in an athymic nude rat right ventricle wall defect replacement model. Heart function was measured, and histologic sections were evaluated for fibrosis, macrophage infiltration, vascularization and muscularization at 2 months postsurgery. As shown below, the prevascularized cardiac patch would induce better muscular and vascular ingrowth, resulting in improved heart function over a non-cell seeded patch.
As shown in
The Young's modulus of the PEG-fibrin was 713±226 Pa (n=5). BPUR fiber meshes had an average fiber diameter of 1.5±0.8 μm (n=5) and an average tensile modulus of 2.9±0.4 MPa (n=5). All animals surviving the surgery survived to the 2-month endpoint. One animal in the shame group had sporadic atrial premature beats (APBs) (1/5), one animal in the patch group had frequent ventricular beats (1/5) and one animal in the patch+cells group had sporadic APBs (1/6).
In this study, we found that a prevascularized fibrin gel patch induced less fibrosis and material remaining, better vascularization and muscularization, larger amount of M2 macrophage infiltration, and improved heart function compared witha non-cell seeded patch at 2 months post-surgery. Surgical correction of CHDs relies on biomaterials that are feasible, biocompatible, prone to endothelialization, disposed to remodeling and integration, and functionally long-lasting. In this study, HUVECs and c-Kit+AFSCs were cultured inside PEG-fibrin gel and at 3 days they successfully formed vascular branches before implantation (
While multiple embodiments are disclosed, still other embodiments of the present invention will become apparent to those skilled in the art from the detailed description. The invention is capable of modifications in various obvious aspects, all without departing from the spirit and scope of the present invention. Accordingly, the detailed description is to be regarded as illustrative in nature and not restrictive.
All references disclosed herein, whether patent or non-patent, are hereby incorporated by reference as if each was included at its citation, in its entirety. In case of conflict between reference and specification, the present specification, including definitions, will control.
Although the present disclosure has been described with a certain degree of particularity, it is understood the disclosure has been made by way of example, and changes in detail or structure may be made without departing from the spirit of the disclosure as defined in the appended claims.
Filing Document | Filing Date | Country | Kind |
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PCT/US2020/046231 | 8/13/2020 | WO |
Number | Date | Country | |
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62886795 | Aug 2019 | US |