Abstract
Objectives
The customized vascular graft offers potential to simplify the surgical procedure, optimize physiological function, and reduce morbidity and mortality. This experiment evaluated the feasibility of a flow dynamic optimized branched tissue engineered vascular graft (TEVG) customized based on medical imaging and manufactured by 3D printing for a porcine model.
Methods
We acquired magnetic resonance angiography (MRA) and 4D flow data for the native anatomy of the pigs (N=2) to design a custom-made, branched vascular graft of the pulmonary bifurcation. An optimal shape of the branched vascular graft was designed using computer aided design (CAD) system informed by computational flow dynamics (CFD) analysis. We manufactured and implanted the graft for pulmonary artery (PA) reconstruction in the porcine model. The graft was explanted four weeks after implantation for further evaluation.
Results
The custom-made branched PA graft had a wall shear stress and pressure drop (PD) from the main PA to the branch PA comparable to the native vessel. At the end point, , the MRI revealed comparable left/right pulmonary blood flow balance. PD from main PA to branch between before and after the graft implantation was unchanged. Immunohistochemistry showed evidence of endothelization and smooth muscle layer formation without calcification of the graft.
Conclusion
Our animal model demonstrates feasibility of design and implantation of image-guided, 3D printed, customized grafts. These grafts can be designed to optimize both anatomically fit and hemodynamic properties. This study demonstrates the tremendous potential structural and physiological advantages of a customized tissue engineered vascular graft in cardiac surgery.
Graphical Abstract
A, The native pulmonary artery (PA) vasculature model was created via the MRI segmentation technology. B, Computational Fluid Dynamic (CFD) simulation was done to assess the flow dynamics in each PA models. C, Optimization of the graft design by manual selection of the graft design with the best performance with the reference of the native flow dynamics. D, 3D printing of the graft design for the in vivo implantation surgery.
INTRODUCTION
One of the leading causes of death in newborns is congenital cardiac anomalies. [1]Due to the anatomical complexity, the uniqueness of each vascular defect and the individual physiology of the circulatory system, congenital vascular reconstructive surgery can be very challenging. There is an unmet clinical need for individual-customized grafts in congenital vascular repair. In pediatric vascular reconstruction, there is an increased risk of graft complication such as rapid graft dysfunction, anastomotic stricture formation, size mismatch related geometric disruption, and pulmonary artery obstruction due to the rapid growth rate of the patient. [2–5] The unmatched sizing of the vasculature caliber, lack of growth potential, suboptimal biocompatibility and risk of thromboembolic events makes the use of synthetic implant for congenital vascular reconstructive an unsettling choice. [6–8]
In the last decade, 3D printing has been used in the manufacture of tissue engineered vascular grafts (TEVG). This technology offers the potential to develop a customized biodegradable scaffold to promote cellular proliferation and maturation, leading to the formation of a physiologically functional blood vessel. For congenital vascular repair, the customized TEVG provides an opportunity for high fidelity anatomical reconstruction of vasculature with growth potential. This can significantly improve the clinical outcomes. Our previous study in an ovine model validated the concept of 3D-printed customized TEVGs using a simple straight vascular graft. The results showed that the TEVGs were biocompatible with adequate neotissue formation and had mechanical properties comparable to the native tissue with 6 months follow up after INFERIOR VENA CAVA (IVC) transposition surgery. [9]
The unique aspects to consider for cardiovascular surgery include a dynamic fluid profile of the circulatory system and the physiological compliance of the vasculature. Thus, the long-term surgical outcomes of the grafts are not only based on anatomic specificity, but also on the growth potential of the TEVG to match the size of the growing patient. The TEVG must maintain the fluid dynamic profile of the circulation system in order to achieve a successful long-term clinical outcome. Our group previously introduced a novel preoperative virtual surgical planning strategy in silico for the Fontan procedure. [10] This study established that preoperative virtual surgical planning can be used to optimize a conduit design that improves the hemodynamics profile after the surgery.
Though a feasibility analysis was performed for a simple straight TEVG in our previous study, a curvilinear or branched TEVG is needed to address cases involving complex anatomy for clinical use. Thus, in this study we build on the previous work by creating a branched and curved TEVG to evaluate the potential to design and implant complex shape TEVG in vivo. (Video 1) As a preclinical experiment, the goal of this study is to investigate the feasibility to fabricate patient specific TEVGs with the Computational Fluid Dynamic (CFD) optimization for reconstruction of central PA. (Figure 1) The central PA reconstructive surgery plays an important role in many surgical procedures repairing congenital cardiac defects in neonates and young infant. [11] We then evaluated the short-term outcomes of the conduit optimized by CFD hemodynamic analysis in a porcine model for a one month trial period. (Graphical abstract: The graft manufacture begins from the Magnetic Resonance Imaging (MRI) acquisition. It follows by CFD analysis and optimization of the flow dynamics of the graft design. Finally, the graft is fabricated by 3D printing technology. )
Figure 1. Flow chart descripting the manufacture process of image guided customized tissue-engineered graft.
A, Creation of the 3-Dimensional (3D) vasculature image of the native model from the Magnetic Resonance Imaging (MRI) segmentation technique, in which the 3 dimensional image of the vasculature is created from the summation of individual 2 dimensional image. B, Optimization of the conduit design with the iteron using iterative strategy. C, Computational Fluid Dynamics (CFD) optimization of 3D printed customized graft, D Manufacture of the nanofiber graft via electrospinning technology. (Left: stainless-steel mandrel, Right: graft)
METERIALS AND METHODS
Preoperative Imaging, 3D model design, and Scaffold Fabrication
Cardiac MRI was performed 4–5 weeks prior to implantation, including an MR angiogram and phase contrast flow analysis using both 2D and 4D flow acquisitions. Following the MRI, the MR angiogram was used as a roadmap to build a 3D digital model of the central PA reconstruction, using a commercially available image segmentation software (Mimics, Materialise, Belgium). Both automatic thresholding and manual methods were used to identify the blood pool of the right ventricular outflow tract and branch pulmonary arteries in each slice of the MR angiogram. Based on species-specific growth curves, a small increase in dimension (both length of the outflow tract and diameter of the pulmonary arteries) was added to the blood pool segmentation to account for growth of the animal during the 30 days between imaging and surgery. This segmentation was converted into a 3D digital model, which was then exported using the stereolithography (STL) file format. This STL file was converted into a computer aided design (CAD) file, smoothed, and hollowed using computer aided design software (3-matic, Materialize, Belgium and SolidWorks Dassault Systemes, USA). After validation of the design, a biodegradable version of each conduit was manufactured using a 3D electrospinning technique. For the electrospinning, a stainless-steel mandrel in the shape of the optimized graft was 3D printed by exporting the STL file of the graft to an external printing house (Proto Labs INC, MN). The mandrel was designed in a way that the electrospun polymer graft could be removed. The mandrel was a five-part piece consisting of main pulmonary artery (MPA), distal right pulmonary artery (RPA), proximal RPA, distal left pulmonary artery (LPA), and proximal LPA pinned together that could be taken apart for the graft removal step. An additional pin was added to the mandrel to clamp it into the electrospinning setup.
Next, the biodegradable nanofiber material composed of a 1:1 ratio of polycaprolactone (PCL) and poly-L-lactide- co-ε-caprolactone (PLCL) were electrospun to coat the 3D printed mandrel (Nanofiber Solutions, OH). By applying a high voltage to the polymer solution, polymer fibers were deposited on the mandrel to create the graft. After depositing the fibers, the mandrel was disassembled and the electrospun graft removed. The graft was placed in a standard Tyvek pouch and sterilized using low temperature sterilization with vaporized hydrogen peroxide and ozone (STERIZONE VP4, Getinge, NY).
Graft implantation in vivo
The Animal Care and Use Committee at Johns Hopkins Hospital (Baltimore, Maryland) approved the care, use, and monitoring of animals for these experiments. All animals (N= 2) received humane care in compliance with the Guide for the Care and Use of Laboratory Animals. The graft was implanted as central PA reconstruction in the porcine model (mean body weight, 35 kg). The animal size increased by roughly one third over the 1-month study interval (mean body weight, 45 kg). The pigs were anesthetized with 1.5% isoflurane during surgery. The surgery was performed at the left thoracotomy position, the left and the right PA were exposed and isolated, and heparin (100 IU/kg) was administered intravenously. During the procedure, cardiopulmonary bypass (CPB) was performed. The graft was implanted to replace main and branch PAs as an interposition graft using 6–0 Prolene suture. Antibiotic treatment (cefazolin 22mg/kg IV) was administered intraoperatively and 7 days postoperatively. All pigs were maintained on a daily oral dose of aspirin (81 mg/day) until the 1-month endpoint. Animals were imaged with an identical cardiac MRI protocol 1month post-surgery, prior to sacrifice.
Histology and Immunohistochemistry
Explanted TEVG samples were fixed in 10% formalin for 24 hours at 4°C, and then embedded in paraffin. For the standard histology, staining of tissue sections with hematoxylin and eosin, Masson trichrome, Verhoeff-Van Gieson, and von Kossa stains were done. For the immunohistochemistry, we used the primary antibodies including von Willebrand factor (vWF) (1:2000; Dako, Cat:A0082), α-smooth muscle actin (SMA) (1:500; Abcam, Cat:ab5694), and CD68 (1:200, Abcam,ab31630). Detection of the antibody binding was done by using biotinylated secondary antibodies (Vector Laboratories, Burlingame, Calif), followed by incubation with streptavidinated horseradish peroxidase (Vector Laboratories). The chromogenic reaction with 3,3-diaminobenzidine (Vector Laboratories) was performed for the development of the immunohistochemistry. Counterstaining of the Nuclei was done with Gill’s hematoxylin (Vector Laboratories).
Histological and Quantitative Analysis
The remaining scaffold area was measured from hematoxylin an eosin staining and plain (polarized light) histology. The thickness of smooth muscle layer from the graft was measured with the α-smooth muscle actin staining using ImageJ software (National Institutes of Health, Bethesda, MD). The thickness of the muscle layer was quantified by analyzing the muscular thickness to full thickness ratio of the graft to that of the native tissue from a representative section of each sample and averaging the value for all samples (n=5). The remaining scaffold area was quantified by analyzing the light intensity under polarized light microscopy the same exposure time and the pixel calculation from a representative section of each sample and averaging the value for all samples (n=5).
Biochemical Analysis
The Sircol colorimetric assay (Biocolor, Carrickfergus, United Kingdom) was used to evaluate the collagen content. According to the assay’s protocol, we measured 100 mg dry weight of each sample and transferred to low-protein-binding 1.5-mL conical microcentrifuge tubes with 1.0 mL of pepsin (Sigma-Aldrich, St Louis, MO), with a concentration of 0.1 mg/mL of 0.5 M acetic acid to solubilize the collagen by overnight incubation. The collagen content in each sample was evaluated by assessing the absorbance intensity at 555 nm after dye binding according to the manufacturer’s protocol.
Statistics Analysis
In this experiment, data from CFD simulation and histological analysis are represented graphically as bar or line charts with error bars which represents the mean with standard error of the mean. The unpaired 2-tailed t-test was done for the analysis of collagen. The Pearson correlation coefficient was used to determine the significance of any correlation of SMA / high-power fields. A P value less than 0.05 was considered statistically significant. Statistical analysis was performed using the Prism version 8 software (GraphPad Software, La Jolla, CA).
Results
Optimization of the custom made graft with CFD simulation
Optimization of the graft was done using an iterative strategy as previously described. [10] Firstly, with all the designed CAD models, the model of the surrounding heart and vasculature anatomy was overlaid with the new graft designs, to minimize any overlap between the optimized graft and other thoracic structures. (Figure 2A) The grafts with the different feasible anatomical designs were chosen. The CFD simulation optimized the graft based on three parameters: i) the reduction of power loss across the conduit inlet and outlet, ii) the adjustment of the even fluid flow between the right and left branches, and iii) the wall shear stress distribution optimization. The design with the best performance was determined for the next iteration of the simulation. (Figure 2B)
Figure 2. A & B, Diagram illustrating the iteration optimization strategy via surgical virtual planning.
A, 3D model of the surrounding heart and vasculature anatomy was overlaid with the new graft designs, to minimize any overlap between the optimized graft and other thoracic structures. B, The design with the best performance was determined for the next iteration of simulation using CFD. C & D, The wall shear stress distribution of the porcine models is represented with different colors, in which blue is the minimal and red is the maximum as the reference on the left sidebar. C&D. The power loss between the inlet and the outlet of the graft compared to the native model was shown. The wall shear stress distribution in the native tissue and the graft. C. In the porcine model #1, Power loss: Native vs Designed, 1.5×10−2 W vs 1.4×10−2 W. Wall shear stress distribution: Native vs Designed, 2Pa vs 2.16Pa. D. In the porcine model #2, Power loss: Native vs Designed, 2.35×10−3 W vs 1.65×10−3 W. Wall shear stress distribution: Native vs Designed, 1.14Pa vs 1.02Pa. Overall, the wall shear stress distribution of two porcine models were similar in native tissue and the graft.
The computational fluid dynamics analysis was performed with the hemodynamic parameters considered in this study including power loss, flow distribution (LPA: RPA) and wall shear stress. The power loss was calculated based on the changes in pressure and flow rates on inlets and outlet of each model as previously described. [10] The wall shear stress distribution of the porcine model was shown in Figure 2C & D. In the porcine model #1 & #2, in comparison to the preoperative native model, the simulation model has a similar power loss between the inlet and the outlet of the graft (Pig #1: Native vs Designed, 1.5×10−2 W vs 1.4×10−2 W, Pig #2: Native vs Designed, 2.35×10−3 W vs 1.65×10−2 W). The flow distribution was maintained with the similar flow in the both the native and the designed conduit. (Pig #1: LPA: RPA, 47:53, Pig #2: LPA: RPA, 42:58). The wall shear stress was comparable in native tissue and the graft as shown in Figure 2 C & D, respectively (Pig #1: Figure 2C, Native versus (vs) Designed, 2Pa vs 2.16Pa, Pig #2: Figure 2D, Native vs Designed, 1.14Pa vs 1.02Pa). Overall, the designed models demonstrated comparable hemodynamic and Wall Shear Stress to the native models.
In vivo implantation of the Customized conduit
The preoperative vasculature images were created with MRI for CAD image creation. (Figure 3A) The preoperative left/right pulmonary blood flow balance was measured. (Figure 3F, LPA: RPA, pre-op:45:55 ) The customized branched TEVG was fabricated via 3D electrospinning technology. (Figure 3B) The graft implantation surgery was performed as described above. (Figure 3C) One month after graft implantation, MRI post operation imaging was performed. (Figure 3D) The left/right pulmonary blood flow balance in the optimized graft design was measured. (Figure 3G, LPA: RPA, post-op:52:48). At the endpoint of the study, the graft was explanted for macroscopic inspection. (Figure 3E) The direct blood pressure measurement of RPA and LPA was measured. (Figure 3H) The magnitudes of RPA and RPA pressure were about 11–12 mmHg. The absence of clinically significant stenosis was demonstrated with an almost even flow distribution and the similar pressure within the PA branches.
Figure 3. Study design of the in vivo implantation of CFD optimized TEVG.
A, Preoperative Magnetic Resonance Imaging (MRI) images of the porcine model’s vasculature. B, 3D-printed branched Tissue Engineered Vascular Graft (TEVG) with the optimized design. C, Intraoperative picture of the implanted conduit. The blue lines outlined the implanted graft intraoperatively. D, Post-operative Magnetic Resonance Imaging (MRI) MRI images of the porcine model’s vasculature. E, The explanted graft at the end point of the experiment. F &G, Flow distribution measurement via MRI across left pulmonary artery (LPA) and right pulmonary artery (RPA) preoperatively and postoperatively F. Flow distribution measurement via MRI across left pulmonary artery (LPA) and right pulmonary artery (RPA) preoperatively. G. Post operation imaging analysis demonstrated comparable left/right pulmonary blood flow balance with improvement of the flow balance in the optimized graft design ( LPA:RPA, 3F, pre op: 45:55, 3G, post op:48:52). H. The intraoperative direct RPA and LPA blood pressure was measured. Both LPA and RPA pressure were of the same magnitudes. ( Mean pressure of LPA: RPA, 11.93: 11.80mmHg). The middle horizontal line represents the median. The upper and lower whiskers represent the maximum and minimum values of non-outliers.
Biocompatibility of the Graft
The biocompatibility of the graft including the patency and the tissue degradation were evaluated after 1-month. There were no graft-related stenosis, dilation or rupture within the 1-month endpoint of the study. There was no clot formation macroscopically and microscopically. (Figures 4A–D) Verification of the presence of collagen in the graft was done by collagen assay, suggesting ongoing collagen tissue formation. (Figure 4E) The graft degradation was detected by the presence of the remaining scaffold area from the polarized light microscopy. (Figure 4F) The remaining percent scaffold area in the individual part of the graft was measured. (Figure 4G) The average of the remaining scaffold of MPA, RPA, and LPA in the models were 11.01 ± 2.64 %, 9.76 ± 3.18% and 8.93 ± 2.14%, respectively.
Figure 4.
A-D, Graft evaluation explanted at one month after surgery. There are no clot formation macroscopically. A&B: Pig#1. A. The explanted graft as a whole. B. The anastomosis interface of the graft. C&D: pig#2. C. The explanted graft as a whole. D. The anastomosis interface of the graft. E, The collagen content determined by Sircol colorimetric assay. ( Native vs MPA vs LPA vs RPA, 1.89 ± 0.18 vs 0.57 ± 0.23 vs 0.48 ± 0.19 vs 1.13 ± 0.16 μg/mg ) F, Representative of the visualization of the remaining scaffold from the polarized light microscopy. The blue color indicates the remaining scaffold.* = lumen side Scale bar = 1000μm G, The remaining scaffold percentage in individual part of the graft were measured. The average of remaining scaffold of main PA, right PA, left PA in porcine model were 11.0 ± 2.64, 9.76 ± 3.18 and 8.93 ± 2.14 respectively. Scale bar = 1000 μm
Formation of Neotissue and the Vasculature
The extracellular matrix formation was evaluated with Mason Trichrome staining. Compared to the native tissue, (Figure 5A,F&K, 5A: Native PA, 5F&K: Graft#1) the vascular graft showed ongoing extracellular matrix formation. (Figure 5B,G&L, 5B: Native PA, 5G&L: Graft#1 ) Von Kossa staining demonstrated no ectopic calcification. (Figure 5C,H&M, 5C: Native PA, 5H&M: Graft#1 ) Smooth muscle cells (SMCs) layer contribute to vascular function and were evaluated by α-smooth muscle actin immunohistochemistry staining (SMA). The presence of multilayered SMA-positive cells in the graft, which compared to the native pulmonary tissue, suggests an ongoing active vascular muscle remodeling take place. (Figure 5D,I&N, 5D: Native PA, 5I&N: Graft#1 ).
Figure 5. Neotissue formation of the vascular graft.
A-E: Native pulmonary artery (PA), F-J: Pig#1 Graft at 1 month endpoint. K-O: Pig#2 Graft at 1 month endpoint * = lumen side. A,F&K: Hematoxylin & Eosin (H&E) staining of the native PA tissue (A) and the vascular graft (F&K) (Scale bar = 1000 μm). B,G&L: Mason Trichrome (MT) staining of the native PA (B) and the vascular graft (G&L) (Scale bar = 1000 μm). The extracellular matrix formation was evaluated with Mason Trichrome staining. Compared to the native tissue, the vascular graft showed ongoing extracellular matrix formation. C,H&M: Von Kossa (VK) staining of native PA (C) and vascular graft (H&M) (Scale bar = 1000 μm). The staining demonstrated no ectopic calcification. D,I&N: Smooth Muscle Actin (SMA) staining of the native PA (D) and the vascular graft (I&N) (Scale bar = 500 μm) The presence of multilayered SMA-positive cells in the graft, which compared to the native pulmonary tissue, suggests an ongoing active vascular muscle remodeling take place. E,J&O: E. Von Willebrand factor (vWF) staining of the native PA (E) and the vascular graft (J&O) (Scale bar = 100 μm). The endothelial cell formation in the graft was shown with the vWF-positive cell in the graft.
The endothelial cell formation in the graft was shown with the vWF-positive cell in the graft. (Figure 5E,J&O, 5E: Native PA, 5J&O: Graft#1) The presence of CD68 positive macrophages with graft degradation is indicative of the inflammatory process of tissue remodeling in the graft. (Figure 6A: the center of the graft, Figure 6B: the edge of the graft) The thickness percentage of the smooth muscle layer (percentage = thickness of smooth muscle / total thickness of the vessel x 100%) demonstrated some neotissue formation. (Figure 6C)
Figure 6. Macrophages recruitment inside the graft & Mechanical properties of the graft with comparison to the native tissue.
A&B, The CD 68 positive macrophages were noted throughout the graft including the inside (A) and the edge (B) of the graft (Scale bar = 100 μm). The presence of CD68 positive macrophages with graft degradation is indicative of the inflammatory process of tissue remodeling in the graft. Figure 6C, The ratio of smooth muscle layer / wall thickness ( Mean value of Native vs MPA vs LPA vs RPA, 33.89 ± 1.18 vs 15.26 ± 2.15 vs 16.87 ± 1.70 vs 15.54 ± 1.75 %) Plots: The middle horizontal line represents the median. The upper and lower whiskers represent the maximum and minimum values of non-outliers. Figure 6D-F, Mechanical Properties of the Customized conduit. Native Tissue: the pulmonary artery from the same porcine model. Graft (Day 0): Graft before implantation, Graft (1 month): Graft 1 month after implantation. F, The circumferential tensile strength of the native tissue and graft after implantation have comparable values. (native vs graft (Day 0) vs graft (1 month), 0.65 vs 1.31 vs 0.45 N/mm) B, The compliance of the graft after implantation was noted to be slightly higher than that of the native tissue. (native vs graft (day 0) vs graft (1 month), 9.89 vs 2.18 vs 21.10 % mmHg ) C, The inner diameter of the graft was comparable to the native PA. (native vs graft (day 0) vs graft (1 month), 10 vs 15 vs10mm) Due to the limitation of the mechanical testing, the tissue tested was destroyed. There was only one sample tested in each parameter.
Mechanical Properties of the Customized TEVG Conduit
The mechanical properties of the graft were measured at different time points including before implantation and 1 month after implantation. The circumferential tensile strength, an analog of the ultimate tensile strength in the tubular structure, of the native tissue and graft after implantation have comparable values. (Figure 6D, native vs graft (1 month), 0.65 vs 0.45 N/mm) The compliance of the graft after implantation was slightly higher than that of the native tissue. (Figure 6E: Compliance, native vs graft (1 month), 9.89 vs 21.10 % mmHg ) At the endpoint, the inner diameter of the graft was similar to the inner diameter of the native PA. (Figure 6F, native vs graft (1 month), 10 vs10mm) Due to the short follow up, the mechanical property measurements of the graft indicate that the customized graft degradation process was in progress with a higher compliance from the remaining scaffold though a similar tensile strength was noted.
DISCUSSION
The results of this study validate the feasibility of the fabrication of the complex shape 3D printed cell-free nanofiber TEVG in an in vivo experiment. Previously, we demonstrated the capability of creation of customized TEVGs using 3D printing and electrospinning technology for straight IVC conduit. [9] Electrospinning provides a unique advantage of the use of various of polymer and fiber sizes in manufacture of the various shape of the vascular graft. [12] This study demonstrated that a complex curved and branched “Y” shape conduit can be fabricated by the electrospinning process. By combining computational fluid dynamic technology, the development of more physiologically compatible patient specific hemodynamics can be achieved. [13]
The materials used in this experiment show adequate physical properties in a low pressure venous system in this in vivo experiment with 1 month time frame. By optimizing the ratios of the combination of different polymers, the materials can display the advantageous mechanical properties of each individual polymer. The biocompatibility profile of some of the polymers, for example polyglycolic acid (PGA), has been investigated in vivo. [12] Previous human trial used scaffolds with PGA in a ratio of 1:1 with the PLCL. [14] In this study, the materials used to make the TEVG are biodegradable and have been previously deployed as scaffolds in the low-pressure circulatory system in patients with congenital heart disease. [15] Because the pressure in the pulmonary artery is higher than that in the IVC, we used PCL instead of PGA due to the slower degradation profile which will allow more extracellular matrix to be deposited. [16] We tested the graft with a similar diameter and demonstrated adequate mechanical properties and 1-month survival without graft-related complications. In order to extend the potential clinical use of this combination of polymers, the in vivo degradation profile of the TEVG made from the combined polymers will be studied in a long term animal survival study.
One of the innovations of this study was the bifurcated shape of the TEVG. Sugiura et al. showed a promising result of no graft-related fatal complications in a clinical trial of TEVG use in children with congenital heart disease. [17] The group followed up 25 patient for 11years who had the implantation of a linear TEVG as an extracardiac total cavopulmonary conduit. The beneficial potential would be extended to a larger population of patients who may need a curved or even complex shaped TEVG. Best et al. has intended to use a CMR guided patient specific TEVG for a 10-year-old patient for total anomalous pulmonary venous return (TAPVR) repair. [18] Unfortunately, the excessively complex anatomy made the patient ineligible for the linear TEVG clinical use. This study established that the production of a complex shaped, customized 3D printed and electrospun TEVG is a feasible technology to manufacture a customized vascular graft with predictable flow dynamics. It would provide the possibility for future development of complex shaped grafts to cater to the anatomy of the patient.
The goal for the TEVG is to support autologous tissue growth and to remain patent without affecting the blood flow inside the conduit and causing diameter mismatch. [19] The cell seeded graft has been investigated to alleviate the intimal hyperplasia in other animal models. [20, 21] The dose-related response in the cell seeded graft was also advocated for the optimization of cell concentration seeded in the TEVG to prevent the progress of stenosis. [22] In this study, we used a cell-free electrospun nanofiber vascular graft which demonstrated satisfactory patency and tissue remodeling in 6 months in our previous study of straight TEVG in the sheep model. [9] In this study, we demonstrated the complex graft in vivo implantation without stenosis formation in 1-month long study. The promising data supports the need for a study with a longer follow up period.
The tissue remodeling potential was seen in the graft by the presence of a single layer of endothelial cells, an organized SMC layer, and collagen deposition. As a pilot study, we aimed to demonstrate the feasibility of neotissue formation in a custom-made graft utilizing a porcine model with short term survival of one month. Specifically, we demonstrated no graft-related mortality and identified no cases of aneurysm formation, graft rupture, or ectopic calcification using routine imaging modalities for TEVG observation and histological assessment.
In this study, we proved the feasibility of the in vivo implantation of a complex shaped, patient specific TEVG. We developed several key components including specific parameters for optimal graft design, unique electrospinning technology, and the technique for surgical experiments using cardiopulmonary bypass. The novelty of this experiment may provide a potential clinical application in congenital heart surgery which demands complex shaped grafts with optimal hemodynamic profiles. Our future work will include a study with higher number of animals and a longer follow up period. Our CFD optimization strategy will be refined to evaluate the low flow regions and areas of low wall shear stress for the susceptibility of thrombosis formation. We will also include different simulation strategies such as pulsatile flow circulation and hyperdynamic conditions in the future. In long term, we aim to create an animal model like Fontan physiology which would be amenable to the real clinical encounters such as vessels stenosis or hypoplasia to provide a possible application close to clinical needs.
In conclusion, we have demonstrated the manufacturability and biocompatibility of the virtual surgical optimization strategy of the 3D printed patient specific TEVG conduit in a large animal model. Although this study was limited by its short survival duration and small sample number, the results of this study are promising and distinctly warrant further investigation.
Supplementary Material
Video 1. Video summarizes the objectives and methodology of optimization and in vivo implantation of the customized branched Tissue Engineered Vascular graft (TEVG).
Central Message.
Our integrated virtual surgical planning (left) and the flow dynamic profile optimization (middle) prove the feasibility of providing an optimized anatomical and hemodynamic parameter of the 3D printed, customized vascular graft (right) in vivo. The three pictures show the concept of the combined innovation (left and middle) and the product (right) for implantation.
In vivo evaluation of our flow optimized customized vascular graft via virtual planning.
Perspective Statement.
Each congenital heart patient has a unique vascular defect and hemodynamic profile. Therefore, there is a need for a ‘best fit’ conduit to provide the best long-term outcome. This study demonstrates that our integrated approach combining virtual surgical planning, 3D printing and electrospinning to create a customized vascular graft can provide an optimized anatomical and hemodynamic result in vivo.
ACKNOWLEDGEMENT
We thank Dr. Henry Halperin, Dr. Cecillia Lui, Dr. Sara Abdollahi, Mr.Tom Loke, Mrs. Melissa Jones, Mr. Sean Kearney, for their surgical support, Dr. Ehud Schmidt, Mr. Michael Guttman, Mr. Rick Tunin, and Ms. Sarah Fink for the perioperative surgical support and the veterinary imaging technical expertise.
The authors acknowledge the University of Maryland supercomputing resources (http://hpcc.umd.edu) made available for conducting the research reported in this paper.
Funding of the work: NIH R21/33 R21HD090671, 4R33HD090671-03
Disclosures: Justin Opfermann: Patent WO2017035500A1. Jed Johnson: NIH NHLBI, Equity holder in Nanofiber Solutions, Patent WO2017035500A1. Kevin Nelson: NIH NHLBI. Laura Olivieri: NIH R01 HL141612, R21HD090671, 4R33HD090671. Axel Krieger: NIH R01 HL141612, R21HD090671, 4R33HD090671, Patent WO2017035500A1. Narutoshi Hibino: NIH R01 HL141612, R21HD090671, 4R33HD090671, Nanofiber Solution: Graft creation, Patent WO2017035500A1. All other authors have no disclosures.
Glossary of abbreviation
- 3D
3-Dimensional
- CAD
Computer Aided Design
- CFD
Computational Flow Dynamics
- CPB
Cardiopulmonary Bypass
- IVC
Inferior Vena Cava
- LPA
Left Pulmonary Artery
- MPA
Main Pulmonary Artery
- MRA
Magnetic Resonance Angiography
- MRA
Magnetic Resonance Imaging
- PA
Pulmonary Artery
- PCL
Polycaprolactone
- PD
Pressure drop
- PLCL
Poly-L-lactide- co-ε-caprolactone
- RPA
Right Pulmonary Artery
- SMA
α-smooth muscle actin
- SMC
Smooth muscle cells
- STL
Stereolithography
- TEVG
Tissue Engineered Vascular Graft
- vWF
von Willebrand Factor
How the authors with patent managed to avoid bias in this paper:
Drs. Hibino and Krieger are inventors on the International Patent WO 2017/035500 Al (PATIENT-SPECIFIC TISSUE ENGINEERED VASCULAR GRAFT UTILIZING ELECTROSPINNING). The patent filing has been disclosed for grant applications and to institutions.
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Supplementary Materials
Video 1. Video summarizes the objectives and methodology of optimization and in vivo implantation of the customized branched Tissue Engineered Vascular graft (TEVG).