具体实施方式：
[0031]Exemplary embodiments are illustrated in the referenced figures of the drawings. It is intended that the embodiments and figures disclosed herein are to be considered illustrative rather than restrictive. No limitation on the scope of the technology and of the claims that follow is to be imputed to the examples shown in the drawings and discussed herein.
[0032]The embodiments are mainly described in terms of particular processes and systems provided in particular implementations. However, the processes and systems will operate effectively in other implementations. Phrases such as ‘an embodiment’, ‘one embodiment’ and ‘another embodiment’ may refer to the same or different embodiments. The embodiments will be described with respect to methods and compositions having certain components. However, the methods and compositions may include more or less components than those shown, and variations in the arrangement and type of the components may be made without departing from the scope of the present disclosure.
[0033]The exemplary embodiments are described in the context of methods having certain steps. However, the methods and compositions operate effectively with additional steps and steps in different orders that are not inconsistent with the exemplary embodiments. Thus, the present disclosure is not intended to be limited to the embodiments shown, but is to be accorded the widest scope consistent with the principles and features described herein and as limited only by the appended claims.
[0034]Furthermore, where a range of values is provided, it is to be understood that each intervening value between an upper and lower limit of the range and any other stated or intervening value in that stated range is encompassed within the disclosure. Where the stated range includes upper and lower limits, ranges excluding either of those limits are also included. Unless expressly stated, the terms used herein are intended to have the plain and ordinary meaning as understood by those of ordinary skill in the art. The following definitions are intended to aid the reader in understanding the present disclosure, but are not intended to vary or otherwise limit the meaning of such terms unless specifically indicated.
[0035]Embodiments of the present disclosure provide for the generation of a pre-operative patient-specific electro spun fiber vascular graft. The electro spun fiber vascular graft is shaped to match a specific patient's anatomy and is optimized for fluid flow that is computed based on flow dynamic simulations. A electro spun fiber vascular graft as described herein requires minimal modifications during implantation surgery, and thereby makes surgery easier and simpler. Furthermore, the electro spun fiber graft promotes cell adhesion and proliferation, and thereby prevents conduit stenosis (i.e., abnormal narrowing) and maintains good long-term cardiac functioning.
[0036]In what follows, is provided a non-limiting example illustrating the requirement and importance of having a patient specific graft design in the treatment of congenital heart disease (CHD).
[0037]FIG. 1 illustrates exemplary staged surgical palliations for single ventricle anomalies of congenital heart disease. As stated previously, single-ventricle-anomalies (SVAs) make up one of the largest and most severe groups of CHDs. In untreated SVAs, only one of two ventricles is of functional size and is associated with a 70% mortality rate during the first year of life.
[0038]In order to treat SVAs, a three-staged surgery is performed as shown in FIG. 1. The first stage includes establishing stable sources of aortic and pulmonary blood flow using a shunt. Specifically, a modified BT-shunt is created to channel aortic flow to the pulmonary arteries (PAs). This procedure is typically performed in the first week of life. In the second stage (referred to as the bidirectional Glenn procedure), the superior vena cava (SVC) is disconnected from the heart and re-implanted into the PAs at about 4-6 months of age. In the third and final stage (also referred to as the Fontan procedure), the inferior vena cava (IVC) is connected to the PAs via a lateral tunnel (LT) or an extra-cardiac (EC) Gore-Tex tube, approximately 2-4 years after the first stage.
[0039]The Fontan pathway, as shown in Stage 3 of FIG. 1 is highly variable in each patient. FIG. 2 depicts an exemplary illustration of the anatomical variations of Fontan pathway. Specifically, FIG. 2 depicts three unique anatomical variations of Fontan pathway (201, 203, and 205, respectively), for three different patients via angiography after a Fontan operation. The Fontan pathway is seldomly straight and has many variances. For such different anatomies, surgeons routinely need to construct optimal shapes during surgery to fulfill their needs. Such variations in architecture demonstrate the requirement for customizing a graft prior to implantation, in order to accommodate proper blood flow and connection between the IVC and pulmonary artery.
[0040]Turning to FIG. 3, there is depicted an exemplary illustration showing variances in anatomies of right-ventricular-pulmonary artery (RV-PA) routes. As shown in FIG. 3, three-dimensional (“3D”) models of Fontan conduits from different patients (top row, 310), and the hepatic flow distribution between the right pulmonary artery (RPA) and left pulmonary artery (LPA) (bottom row, 320), demonstrate the clinically significant variability in outcome performance for the patient, and that commercially available conduits do not fit well for all patients. Since the blood in the Fontan circulation is passively pumped to the lung by the pressure difference between the vena cava and the left atrium, a small increase of venous resistance in Fontan pathway results in a significant drop in the cardiac output for a modified circulation bypassing the right ventricle, which could cause long-term complications.
[0041]There is a strong link between the specific characteristics of blood flow through the Fontan route and the cause or exacerbation of the complication. The most prominent example of such a relationship is the development of pulmonary arteriovenous malformations (PAVM) (causing progressive hypoxia) because of the maldistribution of hepatic factors produced by the liver to the pulmonary vasculature. Another example is the case of impaired exercise capacity, due in part to the non-linear increase in the energy dissipated through the Fontan pathway with increased cardiac output. Accordingly, ensuring an efficient Fontan pathway design with a balanced hepatic flow distribution (HFD) yields long-term benefits for patient health and quality of life.
[0042]Furthermore, conduit selection for right ventricular outflow tract (RVOT) and pulmonary artery reconstruction presents a major challenge in the treatment of many CHDs. This is due to the fact that the angle, length, and diameter of RV-PA route varies widely even amongst those individuals with normal cardiac anatomy. In patients with congenital heart disease, such variety is much more pronounced. Moreover, there is no ideal commercially available material for reconstruction of the route from the RV to PA without compromising laminar flow in the graft.
[0043]Despite the above stated complexities in surgery for the diverse anatomies, the surgeons have no information of flow-dynamics and hemodynamics data of the reconstructed route during procedure, as the surgical field is required to be bloodless. Thus, surgeons usually design the pathway, based on prior experiences in a limited amount of operation time. Surgeons can obtain the hemodynamics data of the constructed pathway only after the surgery is complete. If there are significant hemodynamic issues, surgeons could perform reconstruction again. However, minor energy loss or unbalanced flow cannot be identified immediately in operating room. Therefore, ensuring a patient-specific graft design for ideal reconstructed route before surgery with a balanced flow distribution and minimum energy loss may yield a long-term benefit for the patient's health and quality of life.
[0044]By one embodiment of the present disclosure, is provided a technique of creating patient-specific tissue engineered vascular grafts (TEVGs) for reconstruction in congenital heart disease. The TEVG include biodegradable scaffolds on which autologous cells proliferate and provide physiologic functionality. Specifically, the patient-specific TVEGs are created based on 3D-printing and electrospinning technology.
[0045]Imaging technologies such as computed tomography (CT) and magnetic resonance imaging (MRI) provide surgeons detailed, three-dimensional (3D) views of complex cardiovascular anatomies before surgery. Such technologies offer significantly more utility since the advent of 3D-printing technology, and greatly enhance available treatment options for CHD patients with complex anatomical or physiological requirements that are not easily addressed using currently available prostheses. Moreover 3D models of Fontan conduits provision for the analysis of hepatic flow distribution between the right and left pulmonary arteries. As described next with reference to FIG. 4, a method of creating a patient-specific TEVG utilizes pre-operative 3D imaging data to 3D-print a customized mandrel upon which a polyglycolic acid (PGA)/poly (L-lactide-co-ϵ-caprolactone) i.e., PLCL electro spun fiber blend is electrospun.
[0046]FIG. 4 depicts a flowchart 400 outlining the steps performed in designing and manufacturing a patient-specific vascular graft.
[0047]The process commences in step 401 wherein geometry of the patient's anatomy is determined by acquiring 3D images of target tissue using plenoptic, stereo, time of flight, or structured light cameras, or clinical imaging techniques. Specifically, imaging techniques such as CT, MM (described later with reference to FIG. 14), 3D-ultrasound, PET-CT, laser and the like may be utilized to obtain images of the patient's anatomy.
[0048]In step 403, signal processing is performed on the acquired images in order to refine the images. Various signal processing operations such as low pass filtering, high pass filtering, Gaussian filtering, de-noising, smoothing, sharpness, and contrast operations and the like may be performed to enhance the tissue edges in the acquired images.
[0049]In step 405, the process implements edge detection algorithms in order to detect the edges of target tissue and distinguish them from surrounding tissues.
[0050]The process further proceeds to step 407, wherein a patient-specific implant is designed. Specifically, using dimension, vector, or freeform software (e.g., computer-aided-drafting (CAD)), a patient specific implant (model) is designed from measurements acquired from target anatomy. By one embodiment, the implant may be designed to match existing geometry, or optimized per procedure i.e., increasing flow rates or maximizing functional output requirements. For instance, one or more flow simulations could be applied to the generated patient specific implant model. The model could then be adjusted based on the applied one or more flow simulations to enhance flow in the generated patient specific implant model. Furthermore, it must be appreciated that superficial or surface features may be included while determining the design of the patient-specific implant.
[0051]In step 409, a patient-specific mold (or model) is designed. By one embodiment, a patient specific mold may be designed by computing the negative of the patient specific implant model, where the negative represents a fill within the void areas within the patient specific implant. In other words, the patient specific implant includes one or more solid areas and one or more void areas encompassed within the one or more solid areas. The void areas are the negative of the solid areas. By one embodiment, the mold may include one or more pieces or halves, or may be a mandrel.
[0052]Further, in step 411, the designed patient-specific mold is manufactured to include an electrical conductive layer. By one embodiment, the mold may be realized using additive manufacturing process or by additive manufacturing process with post processing to make the mold electrically conductive.
[0053]Upon manufacturing the mold, the process deposits nano-fibers on the mold (i.e., a mandrel) via electrospinning process (step 413). The electrospinning includes depositing nano-fibers on the mold to create the vascular implant, whereafter the mold is removed prior to surgery. It must be noted that electrospinning provides the advantageous ability of being a highly tunable process by which a wide variety of polymer types and fiber sizes can be spun into various shape of mandrels, thus allowing for the rational design of custom made scaffolds for tissue engineering.
[0054]In step 415, the implant can be sterilized and surgically fixated in the patient. For instance, the implant may be delivered using catheter based tools, or fixated through surgical intervention to the target tissue in a patient. It must be appreciated that the implant may support, enhance, or replace target tissue in the patient.
[0055]Accordingly, the ‘patient-specific’ vascular graft, which is manufactured by the process as described above, is sized and shaped for surgery for the specific patient, thus expediting the corrective surgical procedure. Moreover, the vascular graft improves the quality and safety of vascular graft implantations, and is able to maintain optimal flow through vasculature reconstruction, and recapitulates the native mechanical properties.
[0056]By one embodiment, as stated above, based on pre-operative angiography images, the diameter and length of the patient's IVC can be measured and matching graft models can be designed using CAD software. The final mandrel design can be converted to a stereo-lithography (STL) format and exported for 3D fabrication out of material such as stainless steel-420.
[0057]According to one embodiment, the mandrel may be manufactured using additive manufacturing techniques. The mandrel may be made from solid metals (including titanium, steel, stainless, aluminum, brass, bronze, gold); of hollow metal structures with a thin section to break away the mandrel before surgery; of liquefiable metals e.g. through heating that allow easy removal of mandrel; of polymers with conducting coating, plating or paint (silver, chrome, gold, etc); of liquefiable polymers with conductive coating, plating or paint; of hollow polymer structures with thin section to break away mandrel with conducting coating, plating or paint; and of conductive polymers (e.g. resin contains conductive elements). Additionally, metal and coated polymer mandrels may also be manufactured by subtractive methods such as cutting, turning, etc. The manufactured mandrel can be matched with the patient's anatomy or can be shaped to create an implant that optimizes functional outcomes such as burst pressure, compliance, and the like (described below) for the patient.
[0058]By one embodiment, the mandrel may be manufactured to have a smooth finish for easy removal of the electrospun implant from the mandrel. Additionally, the mandrel may undergo post processing in order to polish, buff, smooth, apply a non-stick coating, and reduce surface friction, such that the implant can be easily removed from the mandrel after electrospinning. The mandrel can be coated with an electrically conductive substance such as a coating of electroplating nickel, copper, or other metal. Alternatively, the coating can also be a conductive paint, conductive glue or epoxy.
[0059]Furthermore, by one embodiment, the mandrel can be manufactured by an additive manufacturing process using a non-electrically conductive material. The non-conductive material can be mixed with a conductive material at the time of manufacturing in order to generate an electrically conductive mandrel. For instance, the mandrel may be a mixture of non-conductive material and aluminum shavings, wherein the aluminum shavings compose a substantially large concentration in order to make the mandrel electrically conductive.
[0060]Additionally, by one embodiment and as stated previously, a subtractive manufacturing process can be used to create a mold that is the negative of the mandrel. A conductive material (e.g., stainless steel) can be cast in the mold and used as the mandrel. Furthermore, a liquefiable metal or polymer can be cast in the mold to take the shape of the mandrel. After electrospinning, the mandrel can be liquefied by application of heat or other medium such as warm water to liquefy the mandrel for removal from the electrospun implant.
[0061]According to one embodiment, the mandrel may be manufactured from a plurality of pieces that can be snapped together or taken apart to create more complex geometry. The use of several pieces also enables easy removal of the mandrel after electrospinning. The mandrels can also include a lock and key feature that allows portions (i.e., pieces) of the mandrel to be aligned and be snapped together. Furthermore, the mandrels may be thin walled, or include thin walled portions that facilitate the breaking of the mandrel into smaller mandrels in order to facilitate easy removal of the mandrel from the electrospun implant. Additionally, by one embodiment, the mandrels may be hollow with thin walls and perforations, so that bifurcated mandrels can be manufactured. The perforations also enable easy removal of the individual portions of the mandrel upon electrospinning.
[0062]By one embodiment, the manufactured mandrels may include one or more conductive and non-conductive portions. The conductive portions of the mandrel may include more than one conductive material. In this manner, by varying the location of the conductive portions of the mandrel, one can generate a conductivity gradient across the mandrel. Such a feature provides the advantageous ability of depositing electro spun fiberin a concentrated fashion in different areas of the graft via electrospinning.
[0063]For the treatment of CHD, in addition to the complexity of surgery for the diverse anatomies, another significant source of morbidity and mortality arises from the use of synthetic biomaterials for various reconstructive cardiac operations. These materials do not grow and cause calcification, which require the patients multiple surgeries in the long term. Accordingly, by one embodiment of the present disclosure, FDA approved biodegradable materials are used in the manufacturing the vascular scaffold. The biodegradable materials offer a potential strategy for overcoming calcification complications by providing a biodegradable scaffold for the patient's own cells to proliferate and provide physiologic functionality over time.
[0064]Despite the strong potential of 3D printing to improve regenerative strategies, there are many challenges that relate to the biomaterials that are available for printing. Materials used for 3D bio-printing must adhere to three key characteristics: the scaffold materials should be a) biocompatible, b) support cell growth and differentiation, and c) be able to sufficiently retain their shape in order to preserve scaffold integrity until solidification locks in the scaffold geometry. Thus, by one embodiment, the scaffolds can be made from materials such as polyglycolide (PGA) and polylactide-co-caprolactone (PLCL)) in a electro spun fiber form. Other materials may include PDO (polydioxanone), PCL (polycaprolactone), PLGA (poly lactic-co-glycolic acid), polyurethane (PU), polyethylene terephthalate (PET) or any combination of these materials.
[0065]To reiterate, the above described patient-specific vascular grafts are created by utilizing pre-operative 3D imaging, followed by design of a vascular graft using computer aided design (CAD) model, 3D printing of a mandrel, and electrospinning of the electro spun fiber vascular grafts. The patient is imaged using a 3D medical imaging technique such as MRI, whereafter the images are processed and segmented to reveal the relevant anatomy. A graft is further designed using for instance, CAD software to optimally (both with respect to anatomical fit and optimal flow characteristics) fix the defect. Further, an electrically conductive mandrel is created from the CAD model of the graft, e.g. using 3D metal printing techniques. A polymer solution is electrospun to deposit electro spun fiberonto the conductive mandrel. According to one embodiment, by changing electrospinning parameters such as fiber material, fiber size, fiber thickness, voltage, spin speed, spin duration, the created grafts are manufactured to match the material properties (e.g. compliance) of the native vessels and promote neo-tissue formation (e.g. through porosity). The graft is then removed from the mandrel, sterilized, and implanted into the patient.
[0066]In what follows, is provided a detailed description of an experimental evaluation of the above method of manufacturing the patient-specific graft. Specifically, an electrospinning mandrel is 3D-printed after computer-aided design based on preoperative imaging of the ovine thoracic inferior vena cava. TEVG scaffolds are then electrospun around the 3D-printed mandrel. Six patient-specific TEVGs are implanted as cell-free inferior vena cava (IVC) interposition conduits in a sheep model, and were explanted after 6 months for histologic, biochemical, and biomechanical evaluation. The experiments as described below confirm the feasibility of utilizing a patient-specific electro spun fiber TEVG by evaluating neo-tissue formation, biocompatibility, and mechanical properties after the 6-month implantation period in the sheep model.
[0067]The experiment utilizes pre-operative 3D imaging data to 3D-print a customized mandrel upon which a PGA/PLCL electro spun fiber blend is electrospun, yielding a patient specific TEVG as shown in FIG. 5.
[0068]FIG. 5 depicts an exemplary flow chart illustrating the process of manufacturing patient-specific tissue engineered vascular grafts. As shown in FIG. 5, the 3D image of vasculature is segmented (step 501) to create a patient specific graft design from the preoperative CT image using computer aided design system (step 502). Further, a finalized mandrel (step 503) is 3D printed from stainless steel (step 504). The patient-specific electro spun fiber graft (as shown in step 505) is obtained via electrospinning and removal of the mandrel.
[0069]According to one embodiment, based on preoperative angiography images, the diameter and length of the sheep IVC are measured and matching graft models are designed using CAD software. The final mandrel design is converted to STL format and exported for 3D fabrication which is performed using materials such as stainless steel. Furthermore, the mandrel manufactured by the additive manufacturing process provides the advantageous ability of having a quick turn-around time. Specifically, the mandrel manufactured by the additive manufacturing process can be used to create patient-specific grafts within a week or less of the surgery. Additionally by one embodiment of the present disclosure, the mandrel may be made of a liquefiable material, thereby allowing the release of the mandrel from the electro spun graft in an easy fashion. Furthermore, the use of liquefiable mandrels also provisions for forming complex shapes of the graft.
[0070]FIG. 6 depicts an exemplary schematic workflow illustrating the manufacturing and surgical implantation of a nano-fiber graft for sheep. As shown in FIG. 6, the dimension and shape of the thoracic inferior vena cava (IVC) is measured from angiography prior to surgery in sheep model (step A, and step B). In step C, an electrospinning mandrel is modeled by computer aided design and subsequently 3D-printed. Thereafter, in step D, a nano-fiber scaffold is electrospun onto the 3D printed mandrel. A patient-specific cell-free electro spun fiber TEVG is implanted as IVC interposition conduit in the sheep model (steps E and F). Steps G and H in FIG. 6, depict a scanning electron microscope (SEM) image of the scaffold at a magnification scale of 500× and 4000×, respectively. For sake of completeness, step I depicts an intraoperative picture of the implanted graft in the sheep.
[0071]According to an embodiment, in order to create the co-electrospun polyglycolic acid (PGA) and polylactide-co-caprolactone (PLCL) scaffolds, 10 wt % PGA is dissolved in hexafluoroisopropanol (HFIP) and 5 wt % PLCL is dissolved in HFIP. Each solution is stirred via a magnetic stir bar for at least 3 hours at room temperature. In separate syringes, the PGA solution is dispensed at a flow rate of 2.5 mL/hr and the PLCL solution is dispensed at a flow rate of 5.0 mL/hr to create a graft with a 1:1 PGA: PLCL ratio. Both solutions are simultaneously electrospun onto the custom 3D-printed mandrel that was positioned 20 cm from the needle tip and rotated at 30 RPM. A +25 kV charge is applied to each syringe tip and electrospun electro spun fiberare deposited onto the grounded mandrel until the desired wall thickness was achieved. Eventually, the electrospun scaffold is removed from the mandrel, and terminally sterilized.
[0072]Upon implantation of the scaffold within the test subject (sheep), a plurality of analysis and testing is performed to qualitatively determine the performance of the scaffold. By one embodiment, the plurality of conducted tests include at least a mechanical testing process that determines compliance and burst pressure of the scaffold, a histological and quantitative analysis, and a biochemical analysis.
[0073]By one embodiment, six custom made cell-free electro spun fiber TEVGs are implanted as IVC interposition grafts in sheep whose body weights are in the range of 23.9±5.0 kg. The IVC is exposed and heparin (100 IU/kg) is administered intravenously. The TEVG is implanted as an IVC interposition graft using standard running 6-0 prolene suture. Further, angiography is performed to assess any potential graft complications at the 3- and 6-month time points. A catheter is inserted into the jugular vein to the IVC, and intravenous contrast manually-injected into the IVC and mid-graft. Additionally, the IVC blood pressure is measured at the proximal and distal anastomoses to evaluate the pressure gradients (PG) across the graft.
[0074]According to one embodiment, compliance and burst pressure data is acquired using a universal mechanical testing machine Specifically, data is acquired using a load frame fitted with a 50 lb. load cell with a force resolution of 10−4 pounds, and a linear displacement resolution of 10−8 inches. Compliance testing is performed using a displacement velocity of 1.5 mm per minute, and acquisition rate of 4 data points per second utilizing Laplace's Law to correlate linear force and displacement to compliance.
[0075]In a similar manner, burst pressure testing is performed using a displacement velocity of 50 mm per minute and acquisition rate of 4 data points/sec utilizing Laplace's Law to correlate linear force and displacement to burst pressure. By one embodiment, ring samples are placed around two parallel L-shaped steel rods, wherein one rod is attached to the base of the testing machine and the other to the load cell. The samples are strained perpendicular to the length of the sample. Compliance is calculated using systolic and diastolic pressures of 120 mmHg and 80 mmHg, respectively, and burst pressure is calculated as the maximum pressure immediately preceding failure.
[0076]According to one embodiment, elastin content is determined using a Fastin colorimetric assay. 100 mg dry weight of each sample is measured and transferred to 1.5 ml micro-centrifuge tubes containing 750 μl 0.25 M oxalic acid. The tubes are further placed on a heat block for 60 min at 100° C. to convert insoluble elastin to water-soluble a-elastin. The elastin content in each sample is determined by detection at 513 nm and interpolation to a standard curve after precipitation and dye binding following the manufacturer's protocol.
[0077]Additionally, by one embodiment, collagen content is determined by a Sircol colorimetric assay. 100 mg dry weight of each sample is measured and transferred to low protein binding 1.5 ml conical micro-centrifuge tubes containing 1.0 ml of pepsin (Sigma-Aldrich), with a concentration of 0.1 mg/ml of 0.5 M acetic acid to solubilize the collagen by means of overnight incubation. The collagen content in each sample is determined by detection at 555 nm and interpolation to a standard curve after precipitation and dye binding following the manufacturer's protocol.
[0078]For all experiments, data is represented as mean±standard error of mean (SEM). The parameter SEM quantifies how precisely one knows the true mean of the population. SEM takes into account both the value of the standard deviation and the sample size. Further, statistically significant differences between groups are determined using Student's t-test and Pearson correlation test. A paired t-test is performed comparing native tissue to tissue-engineered samples, wherein a value of p<0.05 is considered as the two samples being statistically significant.
[0079]In what follows is described with reference to FIGS. 7A-7B to FIG. 13, the results of evaluating the tissue engineered samples as compared to native tissue.
[0080]By one embodiment, a tubular scaffold created via the previously described electrospinning technique has a uniform wall thickness of 657 μm±36 μm, which is significantly less than the native IVC wall thickness of 1365±476 μm. The scanning electron microscope (SEM) images of the scaffold are shown in FIG. 6.
[0081]FIG. 7A depicts a graph illustrating burst pressure of the graft before implantation. The burst pressure of the graft before implantation is significantly less than that of the native IVC case. However, it is noted that there is no significant difference in burst pressure between the native IVC and TEVG at 6 months (i.e., a value of 1.74±0.91 MPa, as compared to a value of 1.56±1.53 MPa, for p=0.4110).
[0082]The pre-operative graft compliance is significantly higher when compared to the native IVC case as shown in FIG. 7B. However, there is no significant difference in compliance between a native IVC and TEVG after 6 months (i.e., a value of 2.37±0.85% as compared to a value of 2.29±0.46%, for p=0.4221).
[0083]FIGS. 8A-8E illustrate exemplary results pertaining to quantitative biochemical analysis of the implanted grafts. Graft biocompatibility including patency (i.e., a condition of showing detectable parasite infection) and tissue remodeling is observed to be excellent within the 6-month study duration. Furthermore, all sheep survived until the study end point without any graft-related complications, such as stenosis, dilation, or rupture as shown in FIG. 8A. In FIG. 8A, the ‘*’ indicates the right atrium, and the arrows corresponds to the anastomosis site of graft.
[0084]Further, native IVC and the patient-specific TEVG are evaluated at 3 months and 6 months to determine evidence of stenosis or dilation with contrast enhanced angiography. As shown in FIG. 8B, both the IVC and patient-specific TEVG displayed no significant diameter changes between the 3- and 6-month time points.
[0085]As shown in FIG. 8C, the pressure gradient (PG) across the TEVG at 6 months is observed to be significantly less than that at 3 months, suggesting advantageous remodeling and scaffold degradation (i.e., a value at 3 months of: 6.29±1.97 vs. a value at 6 months of: 2.08±2.15 mmHg, for p=0.0055).
[0086]According to one embodiment, elastin and collagen are important parameters for venous function and are well-studied markers of vascular graft remodeling. Referring to FIGS. 8D and 8E, the biochemical quantification revealed that the TEVG's elastin content of 6.74±3.13 vs. 7.40±0.88 μg/mg, for p=0.7416, and collagen content of 1.78±0.40 vs. 1.81±0.34 μg/mg, for p=0.9297 were equivalent to that of the native ovine IVC.
[0087]By one embodiment of the present disclosure, the TEVG samples were analyzed to determine histology and immunochemistry. The explanted TEVG samples are fixed in 10% formalin for 24 hours at 4° C., and thereafter embedded in paraffin. For standard histology, tissue sections are stained with hematoxylin and eosin (H&E), Masson's trichrome, Picrosirius Red, Hart's, and von Kossa stains. For immunohistochemistry, tissue sections were deparaffinized, rehydrated, and blocked for endogen