TY - JOUR
AU - Ito,, Masaya
AB - Abstract OBJECTIVES Materials used in paediatric cardiac surgery have drawbacks of deterioration, calcification and pseudointimal proliferation resulting in haemodynamic disturbance. The aim of this study was to investigate whether these drawbacks can be overcome by in situ tissue regeneration using a newly developed synthetic hybrid fabric (SHF). METHODS The SHF is an expandable, warp-knitted fabric composed of a combination of biodegradable [poly-l-lactic acid (PLLA)] and non-biodegradable (polyethylene terephthalate) yarns. The fabric is coated with cross-linked gelatin. Mechanical properties of the SHF were compared with those of 2 commercial products: expanded polytetrafluoroethylene sheet and glutaraldehyde-treated bovine pericardium. An oval-shaped defect created in the canine descending aorta or inferior vena cava was filled with the SHF patch. After 2 weeks and 1, 3, 6 and 12 (or 24 in the inferior vena cava) months, the patch was removed for histological examination and evaluation of the remaining PLLA. RESULTS The SHF exhibited satisfactory tensile and suture retention strength for surgical implantation similar to or better than the 2 commercial products. Tissue regeneration was induced with multilayered smooth muscle cells and collagen fibres on both sides of the patch, along with a mature endothelial layer and tissue connections containing vasa vasorum across the patch in the aorta and inferior vena cava. Inflammatory reactions were minimal, and no calcium deposition occurred. The molecular weight of PLLA was reduced to half at 12 months after implantation. CONCLUSIONS The SHF may solve the drawbacks of the existing products. Further studies of the expandability of the SHF patch after degradation of PLLA are warranted. Congenital cardiac surgery, Surgical material, Tissue regeneration, Biodegradable polymer INTRODUCTION Synthetic materials and xeno-biomaterials [expanded poly tetrafluoroethylene (ePTFE) and glutaraldehyde-preserved bovine pericardium or jugular vein] have been used as a patch, conduit or valve leaflet to correct anatomical disorders in paediatric cardiac surgery. However, these products have innate drawbacks such as tissue degeneration due to calcification or chronic inflammatory reaction, excessive growth of a new intimal layer due to no vasa vasorum development and lack of growth potential leading to patients outgrowing the implanted materials [1, 2]. These problems often require correction of haemodynamic disturbance with reoperations to replace failed materials over a lifetime. Decellularization of heterologous tissue material to obtain skeletonized collagen and a new modification for cross-linking collagen to reduce calcification and inflammatory responses have been proposed to overcome these drawbacks. Decellularized porcine small intestinal submucosa extracellular matrix (CorMatrix®, CorMatrix Cardiovascular Inc., Alpharetta, GA, USA) and a porcine pulmonary valve conduit (No-React®, Shelhigh Inc., Millburn, NJ, USA) have shown favourable results in animal models [3, 4], but failed tissue integration, recellularization and replacement with native tissue and ingrowth of materials with a thickened and shrunk decellularized porcine bioprosthesis have occurred in paediatric cardiac surgery [5–7]. Decellularized bovine pericardium treated with a modified formula for cross-linking collagen with less use of cytotoxic glutaraldehyde (ADAPT® and CardioCel®, Celxcel Pty Ltd., Perth, WA, Australia) has been investigated clinically after the initial favourable results [8, 9]. However, calcification and pseudointimal proliferation causing luminal stenosis were observed in the new treatment with bovine pericardium 1 or 2 years after implantation in porcine pulmonary arteries and in human neonatal aortic walls [10, 11]. Vascular tissue regeneration by tissue engineering using biodegradable polymer scaffolds preincubated with various kinds of cells in vitro is an evolving technology developed to overcome the drawbacks that the existing products have [12–14]. This approach has been applied in a clinical setting as an external tube graft implanted in the human inferior vena cava (IVC) [13, 14]. However, several problems are still unsolved in classical in vitro tissue engineering, such as environment for incubation, need for cell seeding, cell source, required cell number, cost and practicality as a routine method [15]. Recent progress has been made in vascular tissue regeneration by in situ tissue engineering with fabrications using biodegradable polymers without cell seeding [15–18], with the goal of inducing growth of functional tissue by endogenous cells after simple implantation. This approach has the potential for use as an off-the-shelf medical implant. In this paradigm shift from in vitro to in situ tissue engineering, we investigated whether our simple synthetic hybrid fabric (SHF) could promote in situ tissue regeneration in large canine vessels. MATERIALS AND METHODS Design and fabrication of the synthetic hybrid fabric The SHF is an expandable, warp-knitted fabric composed of biodegradable [poly-l-lactic acid (PLLA)] and non-biodegradable [polyethylene terephthalate (PET)] yarns. The yarns generally run lengthwise in the fabric and are horizontally connected through a longitudinal series of loops of yarns (PET) interlocked with diagonally running hoops of yarns (PLLA) to maintain the initial structure before implantation (Fig. 1A and B). The PET portion is designed to be stretchable and quadruple in area after degradation of PLA (Fig. 1B and C). The warp structure was designed using the software connected to a specially modified industrial warp-knitting machine (HKS, Karl Meier Textilemachinenfabrik GmbH, Obertshausen, Germany). The fabric is coated with cross-linked gelatin (Nitta Gelatin Inc., Yao, Osaka, Japan) to seal its porous structure (Fig. 1C). Figure 1: Open in new tabDownload slide Design of the SHF. (A) Representative scanning electron microscope findings for the SHF. (B) Expandability of the SHF, a warp-knitted fabric composed of a combination of biodegradable (poly-l-lactic acid, purple) and non-biodegradable (polyethylene terephthalate, sky blue) yarns. The polyethylene terephthalate portion is designed to be stretchable and quadruple in area after complete degradation of the poly-l-lactic acid portion. (C) The fabric is coated with cross-linked gelatin. Figure 1: Open in new tabDownload slide Design of the SHF. (A) Representative scanning electron microscope findings for the SHF. (B) Expandability of the SHF, a warp-knitted fabric composed of a combination of biodegradable (poly-l-lactic acid, purple) and non-biodegradable (polyethylene terephthalate, sky blue) yarns. The polyethylene terephthalate portion is designed to be stretchable and quadruple in area after complete degradation of the poly-l-lactic acid portion. (C) The fabric is coated with cross-linked gelatin. Mechanical properties of the synthetic hybrid fabric Basic mechanical properties of the SHF were measured in uniaxial tensile testing and compared with those of a commercial ePTFE sheet of 0.4 mm in thickness (Gore-Tex® cardiovascular patch, W. L. Gore & Associates Inc., Flagstaff, AZ, USA) and a glutaraldehyde (GA)-preserved bovine pericardial patch (Model #700, Edwards Lifesciences Corp., Irvine, CA, USA). Each sample was cut as a 5 × 1 cm rectangle and hung in a tensile testing machine (EZ-test, Shimadzu Corp., Kyoto, Japan). The tensile test was performed by stretching the samples at a speed of 50 mm/min at room temperature, in accordance with ISO 13937-1: 2000 ‘Textiles-tear properties of fabrics Part 1’. The strain–stress relationship was obtained, and the ultimate tensile strength was determined in each sample. Suture retention strength Suture retention strength was tested to compare the force exerted by suture threads that cause samples to lacerate. The test was performed by pulling a suture through the sample to determine the force necessary to tear the sample, in accordance with the ISO 7198 protocol. Each sample was cut as a 5 × 1 cm rectangle, and a 5-0 monofilament polypropylene suture (Prolene®, Ethicon Inc., Somerville, NJ, USA) was passed through one end. Using the tensile testing machine described above, the suture was pulled at a speed of 50 mm/min, and the force for complete laceration was determined in each sample. Water leakage from a needle hole Water leakage from one needle hole created in the middle of each sample sheet using a 5-0 monofilament polypropylene suture (as above) was measured. The sample was fixed on the end cap of a well-sealed container of pure water, and the internal pressure was varied. The amount of water dropping from the needle hole was weighed using an electric balance at a pressure of 150 mmHg for 3 min. This method is a modification based on the ISO 7198 protocol ‘Cardiovascular implants and extracorporeal systems’. Implantation of the synthetic hybrid fabric in the aorta or inferior vena cava Sixteen adult male beagles (age 6–12 months, body weight 8–12 kg, Hamaguchi Laboratories LLC, Osaka, Japan) were used in the study. Under general anaesthesia, animals were placed on a ventilator. In a right decubitus position, the descending aorta was exposed via a left thoracotomy through the 4th intercostal space. After injecting heparin (400 IU), a temporary bypass between the aortic arch and lower descending aorta was established to avoid paraplegia during cross-clamping of the descending aorta. After the cross-clamp, an oval-shaped defect was made in the upper descending aorta. The defect was filled with a patch (20 mm × 10 mm) trimmed from the SHF using a 6-0 polypropylene continuous suture (as above). The animal was extubated after chest closure and received all necessary care. Implantation of the SHF in the IVC was performed in a similar fashion via a right thoracotomy through the 4th intercostal space, except for the temporal bypass used in the aortic implantation model. After euthanization with general anaesthesia followed by intravenous injection of a fatal dose of a barbiturate, the patch with the surrounding native vascular wall was removed en bloc for histological examination and measurement of the molecular weight (MW) of the remaining PLLA and PET at 2 weeks (in the aorta) and 1, 3, 6 and 12 (or 24 in the IVC) months after implantation. The study was approved by the Animal Research Committee of Osaka Medical College (approval ID #27019 for the IVC and #28086 for the aorta), and the animals received humane care according to the guidelines of this committee and in accordance with the National Institute of Health Guidelines for the Care and Use of Laboratory Animals. Histological examination The excised tissue block was fixed with 4% paraformaldehyde. The ring-shaped sample was cut open by a longitudinal incision made in the opposite side of the patch. A sample slice was cut out from the centre of the patch along the circumference and embedded in paraffin. Microtome-sliced samples of 5 μm thickness were stained with haematoxylin and eosin (HE) to examine the general status of the SHF patch and the inflammatory response. In addition, Masson’s trichrome, elastica van Gieson and Alizarin red staining were performed to evaluate collagen fibres, elastic fibres and calcium deposition in the regenerative tissue, respectively. Immunohistochemistry with von Willebrand factor and alpha-smooth muscle actin staining was performed to detect endothelial cells/neoarteries and myofibroblasts/smooth muscle cells, respectively. Tissue regeneration, infiltration of inflammatory cells, calcium deposition at the patch implant site and status of the implanted patch were evaluated semi-quantitatively by scoring as follows: 0: none, 1: very slight, 2: mild, 3: moderate, 4: marked (see Supplementary Material, Table S1). Evaluation of polymer remnants To evaluate the extent of PLLA degradation, the weight-averaged MW of each polymer was analysed by gel permeation chromatography via a modification of a previously reported method [19]. Briefly, the polymers were extracted from the excised SHF samples with chloroform, and the tissue residue was removed by filtration with deactivated glass wool. Measurements were performed with a TSK-gel GMHHR-M column (Tosoh Corporation, Tokyo, Japan) equipped with a refractive index detector for PLLA and an ultraviolet detector for PET. The MW of each polymer was calculated from the calibration curves obtained for polystyrene standards (MW range 0.5–2110 kDa). Statistical analysis One-way analysis of variance followed by a Dunnett’s multiple comparisons test was performed using GraphPad Prism version 7.03 for Windows (GraphPad Software, La Jolla CA, USA). A P-value <0.05 was considered significant. RESULTS Mechanical properties of the synthetic hybrid fabric Representative strain–stress curves for the SHF, ePTFE sheet and GA-preserved bovine pericardium are shown in Fig. 2A. The SHF patch exhibited satisfactory ultimate tensile strength for surgical use that was equivalent to those of ePTFE and the GA-preserved bovine pericardium (Fig. 2B). Figure 2: Open in new tabDownload slide Mechanical properties of the SHF. (A) Representative strain–stress curves of the SHF, ePTFE sheet and bovine pericardium. (B) Comparison of the ultimate tensile strength among the sheets. Results are shown as raw data and mean ± 1SD. ePTFE: expanded polytetrafluoroethylene; SHF: synthetic hybrid fabric. Figure 2: Open in new tabDownload slide Mechanical properties of the SHF. (A) Representative strain–stress curves of the SHF, ePTFE sheet and bovine pericardium. (B) Comparison of the ultimate tensile strength among the sheets. Results are shown as raw data and mean ± 1SD. ePTFE: expanded polytetrafluoroethylene; SHF: synthetic hybrid fabric. Suture retention strength and leakage from a needle hole The SHF exhibited significantly greater suture retention strength than those of ePTFE and the GA-preserved bovine pericardium (Fig. 3A). Water leakage from the needle hole in the SHF and the GA-preserved bovine pericardium were negligible and significantly less than that in the ePTFE sheet (Fig. 3B). Figure 3: Open in new tabDownload slide Suture retention strength (A) and needle hole water leakage (B). Results are shown as raw data and mean ± 1SD. ePTFE: expanded polytetrafluoroethylene; SHF: synthetic hybrid fabric. Figure 3: Open in new tabDownload slide Suture retention strength (A) and needle hole water leakage (B). Results are shown as raw data and mean ± 1SD. ePTFE: expanded polytetrafluoroethylene; SHF: synthetic hybrid fabric. Surgical handling The SHF was pliable without difficulty in suturing and fitted well to the native vessel wall. Bleeding from needle holes and the suture line was easily controlled in both the aorta and the IVC (Fig 4A and D). Figure 4: Open in new tabDownload slide Representative gross appearance of the SHF implanted into the aorta (A–C) and IVC (D–F). (A and D) At implantation in the aorta or IVC. (B and E) At explanation, after implantation for 6 months in the aorta and 24 months in the IVC. (C and F) Luminal surfaces of the explanted SHF patch after implantation for 6 months in the aorta and 24 months in the IVC. White arrowheads: the SHF. IVC: inferior vena cava; SHF: synthetic hybrid fabric. Figure 4: Open in new tabDownload slide Representative gross appearance of the SHF implanted into the aorta (A–C) and IVC (D–F). (A and D) At implantation in the aorta or IVC. (B and E) At explanation, after implantation for 6 months in the aorta and 24 months in the IVC. (C and F) Luminal surfaces of the explanted SHF patch after implantation for 6 months in the aorta and 24 months in the IVC. White arrowheads: the SHF. IVC: inferior vena cava; SHF: synthetic hybrid fabric. Macroscopic and histological findings Macroscopic findings for the explanted SHF are shown in Fig. 4. At the time of explantation, there were no aneurysmal or stenotic changes in patch-implanted regions of both great vessels (Fig 4B and E). The luminal surfaces of the explanted SHF from both great vessels were covered with the smooth intimal tissue (Fig. 4C and F). Histological findings of the specimen from the aorta at 2 weeks and 1, 3 and 6 months after implantation are shown in Fig. 5 and Table 1. At 2 weeks, myofibroblasts regenerated from the cut end of the aortic wall towards the luminal surface of the SHF patch (Fig. 5A–C). No inflammation was observed around the patch, while neutrophils and lymphocytes infiltrated on the pleural side of the surgical area (Fig. 5A, asterisk). At 1 month, luminal and pleural sides of the SHF were covered with the regenerative tissue (Fig. 5D). Macrophages aggregated around the patch yarns (Fig. 5E). Lining of endothelial cells on the luminal surface and neovascularization around the patch were visible (Fig. 5F). At 3 months, the gelatin layer (basophilic matrix around the patch yarns) decreased (Fig. 5G). The regenerative aortic wall consisted of multilayered smooth muscle cells and collagen fibres (Fig. 5H and I). At 6 months, connective tissue with vasa vasorum infiltrated into the space where gelatin had disappeared, namely the ‘bridging tissue’ (Fig. 5J and K). There was no calcium deposition on the SHF or the regenerative tissue (Fig. 5L). Development of elastin was not detected at this time (Fig. 5M). Findings at 12 months were similar to those at 6 months, except for gelatin remaining due to stronger cross-linking of the gelatin in the initial protocol (Table 1). Table 1: Semi-quantitative scoring of histological characteristics over time after implantation Findings Implantation site Aorta IVC Duration after implantation 2 weeks 1 month 3 months 6 months 12 months 1 month 3 months 6 months 12 months 24 months Number of animals 2 2 2 1 1 2 2 2 1 1 Vascular wall, implant site  Tissue regeneration   Connective tissue 1, 1 1, 1 2, 1 1 1 2, 1 2, 2 1, 1 1 2   Smooth muscle 0, 0 3, 2 1, 2 1 1 3, 3 3, 3 3, 3 3 3   Collagen fibres 0, 0 3, 3 3, 3 3 4 4, 4 3, 4 4, 4 4 3   Endothelial cells (luminal surface) 1, 1 3, 3 3, 3 3 3 3, 3 3, 3 3, 3 3 3  Infiltration, inflammatory cells 2, 4 0, 1 1, 1 0 0 3, 1 0, 0 0, 0 0 0  Calcium depositiona 0, 1 0, 0 2, 0 0 1 0, 0 0, 0 0, 0 0 0 SHF condition  Gelatin remaining 4, 4 2, 2 1, 1 0 2 3, 3 2, 1 1, 1 1 0  PLLA remaining 3, 3 3, 3 3, 3 3 3 3, 3 3, 3 3, 3 3 2b  Bridge formation 0, 0 1, 1 1, 3 4 1 1, 0 2, 4 4, 2 3 4 Cell infiltration into SHF  Neutrophils 0, 0 0, 0 0, 0 0 0 0, 0 0, 0 0, 0 0 0  Lymphocytes 0, 0 0, 0 1, 0 0 0 0, 0 0, 0 0, 0 0 0  Plasma cells 0, 0 0, 0 0, 0 0 0 0, 0 0, 0 0, 0 0 0  Macrophages/giant cells 0, 0 1, 1 2, 2 1 1 2, 1 2, 3 4, 3 3 3  Connective tissue 1, 1 1, 1 2, 1 1 1 1, 1 2, 4 4, 2 3 3 Findings Implantation site Aorta IVC Duration after implantation 2 weeks 1 month 3 months 6 months 12 months 1 month 3 months 6 months 12 months 24 months Number of animals 2 2 2 1 1 2 2 2 1 1 Vascular wall, implant site  Tissue regeneration   Connective tissue 1, 1 1, 1 2, 1 1 1 2, 1 2, 2 1, 1 1 2   Smooth muscle 0, 0 3, 2 1, 2 1 1 3, 3 3, 3 3, 3 3 3   Collagen fibres 0, 0 3, 3 3, 3 3 4 4, 4 3, 4 4, 4 4 3   Endothelial cells (luminal surface) 1, 1 3, 3 3, 3 3 3 3, 3 3, 3 3, 3 3 3  Infiltration, inflammatory cells 2, 4 0, 1 1, 1 0 0 3, 1 0, 0 0, 0 0 0  Calcium depositiona 0, 1 0, 0 2, 0 0 1 0, 0 0, 0 0, 0 0 0 SHF condition  Gelatin remaining 4, 4 2, 2 1, 1 0 2 3, 3 2, 1 1, 1 1 0  PLLA remaining 3, 3 3, 3 3, 3 3 3 3, 3 3, 3 3, 3 3 2b  Bridge formation 0, 0 1, 1 1, 3 4 1 1, 0 2, 4 4, 2 3 4 Cell infiltration into SHF  Neutrophils 0, 0 0, 0 0, 0 0 0 0, 0 0, 0 0, 0 0 0  Lymphocytes 0, 0 0, 0 1, 0 0 0 0, 0 0, 0 0, 0 0 0  Plasma cells 0, 0 0, 0 0, 0 0 0 0, 0 0, 0 0, 0 0 0  Macrophages/giant cells 0, 0 1, 1 2, 2 1 1 2, 1 2, 3 4, 3 3 3  Connective tissue 1, 1 1, 1 2, 1 1 1 1, 1 2, 4 4, 2 3 3 Grading: 0, none; 1, very slight; 2, mild; 3, moderate; 4, marked (Supplementary Material, Table S1). Raw data are shown, for example ‘2, 1’ means scores were 2 and 1 in the 2 animals. a Calcium deposition was observed on the necrotizing tunica media at the cut end of the aorta but not on the SHF or regenerated tissues. b The PLLA fibre showed weekly eosinophilic appearance in HE section, and there were some cracks in the cross-section of the fibre. In other specimens, the PLLA fibre was colourless and transparent with no cracks. HE: haematoxylin and eosin; IVC: inferior vena cava; PLLA: poly-l-lactic acid; SHF: synthetic hybrid fabric. Table 1: Semi-quantitative scoring of histological characteristics over time after implantation Findings Implantation site Aorta IVC Duration after implantation 2 weeks 1 month 3 months 6 months 12 months 1 month 3 months 6 months 12 months 24 months Number of animals 2 2 2 1 1 2 2 2 1 1 Vascular wall, implant site  Tissue regeneration   Connective tissue 1, 1 1, 1 2, 1 1 1 2, 1 2, 2 1, 1 1 2   Smooth muscle 0, 0 3, 2 1, 2 1 1 3, 3 3, 3 3, 3 3 3   Collagen fibres 0, 0 3, 3 3, 3 3 4 4, 4 3, 4 4, 4 4 3   Endothelial cells (luminal surface) 1, 1 3, 3 3, 3 3 3 3, 3 3, 3 3, 3 3 3  Infiltration, inflammatory cells 2, 4 0, 1 1, 1 0 0 3, 1 0, 0 0, 0 0 0  Calcium depositiona 0, 1 0, 0 2, 0 0 1 0, 0 0, 0 0, 0 0 0 SHF condition  Gelatin remaining 4, 4 2, 2 1, 1 0 2 3, 3 2, 1 1, 1 1 0  PLLA remaining 3, 3 3, 3 3, 3 3 3 3, 3 3, 3 3, 3 3 2b  Bridge formation 0, 0 1, 1 1, 3 4 1 1, 0 2, 4 4, 2 3 4 Cell infiltration into SHF  Neutrophils 0, 0 0, 0 0, 0 0 0 0, 0 0, 0 0, 0 0 0  Lymphocytes 0, 0 0, 0 1, 0 0 0 0, 0 0, 0 0, 0 0 0  Plasma cells 0, 0 0, 0 0, 0 0 0 0, 0 0, 0 0, 0 0 0  Macrophages/giant cells 0, 0 1, 1 2, 2 1 1 2, 1 2, 3 4, 3 3 3  Connective tissue 1, 1 1, 1 2, 1 1 1 1, 1 2, 4 4, 2 3 3 Findings Implantation site Aorta IVC Duration after implantation 2 weeks 1 month 3 months 6 months 12 months 1 month 3 months 6 months 12 months 24 months Number of animals 2 2 2 1 1 2 2 2 1 1 Vascular wall, implant site  Tissue regeneration   Connective tissue 1, 1 1, 1 2, 1 1 1 2, 1 2, 2 1, 1 1 2   Smooth muscle 0, 0 3, 2 1, 2 1 1 3, 3 3, 3 3, 3 3 3   Collagen fibres 0, 0 3, 3 3, 3 3 4 4, 4 3, 4 4, 4 4 3   Endothelial cells (luminal surface) 1, 1 3, 3 3, 3 3 3 3, 3 3, 3 3, 3 3 3  Infiltration, inflammatory cells 2, 4 0, 1 1, 1 0 0 3, 1 0, 0 0, 0 0 0  Calcium depositiona 0, 1 0, 0 2, 0 0 1 0, 0 0, 0 0, 0 0 0 SHF condition  Gelatin remaining 4, 4 2, 2 1, 1 0 2 3, 3 2, 1 1, 1 1 0  PLLA remaining 3, 3 3, 3 3, 3 3 3 3, 3 3, 3 3, 3 3 2b  Bridge formation 0, 0 1, 1 1, 3 4 1 1, 0 2, 4 4, 2 3 4 Cell infiltration into SHF  Neutrophils 0, 0 0, 0 0, 0 0 0 0, 0 0, 0 0, 0 0 0  Lymphocytes 0, 0 0, 0 1, 0 0 0 0, 0 0, 0 0, 0 0 0  Plasma cells 0, 0 0, 0 0, 0 0 0 0, 0 0, 0 0, 0 0 0  Macrophages/giant cells 0, 0 1, 1 2, 2 1 1 2, 1 2, 3 4, 3 3 3  Connective tissue 1, 1 1, 1 2, 1 1 1 1, 1 2, 4 4, 2 3 3 Grading: 0, none; 1, very slight; 2, mild; 3, moderate; 4, marked (Supplementary Material, Table S1). Raw data are shown, for example ‘2, 1’ means scores were 2 and 1 in the 2 animals. a Calcium deposition was observed on the necrotizing tunica media at the cut end of the aorta but not on the SHF or regenerated tissues. b The PLLA fibre showed weekly eosinophilic appearance in HE section, and there were some cracks in the cross-section of the fibre. In other specimens, the PLLA fibre was colourless and transparent with no cracks. HE: haematoxylin and eosin; IVC: inferior vena cava; PLLA: poly-l-lactic acid; SHF: synthetic hybrid fabric. Figure 5: Open in new tabDownload slide Histology and immunohistochemistry of the explanted SHF patch from the aorta after 2 weeks (A–C), 1 month (D–F), 3 months (G–I) and 6 months (J–M). (A, D, E, G, J) Haematoxylin and eosin staining. (B and M) Elastica van Gieson staining. (C and I) alpha smooth muscle actin immunohistochemistry. (F and L) von Willebrand factor immunohistochemistry. Arrows indicate endothelial cells. Asterisks indicate neovascularization or vasa vasorum. (H) Masson’s trichrome staining. (K) Alizarin red staining. Figure 5: Open in new tabDownload slide Histology and immunohistochemistry of the explanted SHF patch from the aorta after 2 weeks (A–C), 1 month (D–F), 3 months (G–I) and 6 months (J–M). (A, D, E, G, J) Haematoxylin and eosin staining. (B and M) Elastica van Gieson staining. (C and I) alpha smooth muscle actin immunohistochemistry. (F and L) von Willebrand factor immunohistochemistry. Arrows indicate endothelial cells. Asterisks indicate neovascularization or vasa vasorum. (H) Masson’s trichrome staining. (K) Alizarin red staining. Histological findings of the specimen from the IVC at 1, 3, 6, 12 and 24 months after implantation are shown in Fig. 6 and Table 1. At 1 month, both sides of the SHF were covered with the regenerative tissue (Fig. 6A and B). Minimal lymphocyte infiltration was observed around the gelatin layer (Fig. 6B). Regenerative tissue consisted of myofibroblasts with the luminal surface covered with endothelial cells (Fig. 6C and D). At 3 months, the gelatin layer was still present between the SHF yarns (Fig. 6E). There was no inflammation in this period; instead, macrophages and multinucleated giant cells aggregated around the patch yarns (Fig. 6F). Neovascularization was visible in the regenerative tissue (Fig. 6G). At 6 months, the regenerative venous wall consisted of abundant collagen fibres and modest smooth muscle cells that were arranged in layers (Fig. 6H–J). At 12 and 24 months, mature collagen fibres covered the SHF with almost the same thickness as the native IVC wall (Fig. 6K–O). At these times, gelatin had disappeared, connective tissue with vasa vasorum infiltrated between the patch yarns (Fig. 6L and M) and there was no calcium deposition (Fig. 6P). Figure 6: Open in new tabDownload slide Histology and immunohistochemistry of the explanted SHF from the inferior vena cava after 1 month (A–D), 3 months (E–G), 6 months (H–J), 12 months (K–M) and 24 months (N–P). (A, B, E, F, H, K, L, N) Haematoxylin and eosin staining. (C and J) alpha smooth muscle actin immunohistochemistry. (D, G, M) von Willebrand factor immunohistochemistry. Arrows indicate endothelial cells. (I and O) Masson’s trichrome staining. (P) Alizarin red staining. Figure 6: Open in new tabDownload slide Histology and immunohistochemistry of the explanted SHF from the inferior vena cava after 1 month (A–D), 3 months (E–G), 6 months (H–J), 12 months (K–M) and 24 months (N–P). (A, B, E, F, H, K, L, N) Haematoxylin and eosin staining. (C and J) alpha smooth muscle actin immunohistochemistry. (D, G, M) von Willebrand factor immunohistochemistry. Arrows indicate endothelial cells. (I and O) Masson’s trichrome staining. (P) Alizarin red staining. Degradation of poly-l-lactic acid yarn In vivo degradation of PLLA of the SHF expressed as the weight-averaged MW is shown in Fig. 7A. The MW of PLLA gradually decreased over time to about 75% and 50% of baseline at 6 and 12 months after implantation, respectively. At 24 months after implantation in the IVC, the MW of PLLA was about 2% of baseline, indicating complete loss of mechanical strength of PLLA yarn. On histological examination, many cracks were observed in the PLLA yarn at 24 months (Fig. 7C, Table 1), while the yarn at 1 month had a flat cross-sectional appearance (Fig. 7B). Interestingly, the PLLA degradation rate in the SHF was similar in the 2 implanted vessels. The MW of PET remained at the baseline level throughout the study period. Figure 7: Open in new tabDownload slide In vivo degradation of PLLA of the SHF. (A) Weight-averaged molecular weight measured by gel permeation chromatography. (B) PLLA yarn 1 month after implantation in the IVC. (C) PLLA yarn 24 months after implantation in the IVC. IVC: inferior vena cava; PLLA: poly-l-lactic acid; SHF: synthetic hybrid fabric. Figure 7: Open in new tabDownload slide In vivo degradation of PLLA of the SHF. (A) Weight-averaged molecular weight measured by gel permeation chromatography. (B) PLLA yarn 1 month after implantation in the IVC. (C) PLLA yarn 24 months after implantation in the IVC. IVC: inferior vena cava; PLLA: poly-l-lactic acid; SHF: synthetic hybrid fabric. DISCUSSION The key to successful in situ tissue engineering of a vascular wall is an appropriate selection of biodegradable material, structure design and fabrication of an implant as a scaffold and a combination thereof [15, 18, 20]. The degradation rate and mechanism are important factors in the selection of materials. Currently, synthetic polymers that rapidly degrade within a few months are the preferred choice [12–18]. Polyglycolic acid and a 1:1 blend of poly-ε-caprolactone and polylactic acid are used in classic in vitro tissue engineering. A sponge graft made from the polyglycolic acid/poly-ε-caprolactone and polylactic acid was implanted as an extracardiac conduit in patients who underwent a modified Fontan operation in Japan [13, 14]. However, critical graft stenosis occurred in the long term, probably due to mechanical loss after degradation of polymers before regenerated tissue maturation compensated the loss [13, 14]. The mechanical properties of the polyglycolic acid/poly-ε-caprolactone and polylactic acid sponge graft were improved by adding outer reinforcement with thick filaments [16] or with a PLLA knit laminated inside the sponge [17]. Therefore, slow biodegradable PLLA was chosen for the SHF to gain time for mechanical compensation by the matured regenerated tissue. The SHF sheet provided durable mechanical properties in vitro with satisfactory surgical handling and in vivo without stenosis or aneurysmal formation in both the aorta and the IVC. It is widely known that tissue regeneration using biodegradable materials is initiated by immune cells as a foreign body response [12, 14, 15]. However, sustained immune cell infiltration often causes tissue swelling and thickening with calcium deposition due to cell death, eventually leading to primary tissue failure of implants, as seen in xeno-bioprostheses [1, 10]. This is another reason for the preference for rapidly degradable polymers. Polycarbonate bis-urea (PC-BU) is rapidly degraded by immune cells and has been used to create an electrospinning scaffold for in situ tissue engineering of pulmonary valve leaflets [18]. However, its absorption occurs inhomogeneously with inflammatory cell infiltration sustained for more than 12 months, similar to chronic inflammation. We selected PLLA because it is not degraded by immune cells but by simple hydrolysis. In this study, PLLA degraded over 24 months, as measured by gel permeation chromatography, and only modest initial immune cell infiltration was detected in the regeneration process, which might have contributed to the tissue integration. The structure of an implant must have robust mechanical strength tolerating scar contraction and haemodynamic stress and sufficient porosity size to allow cell infiltration as a scaffold for subsequent tissue regeneration. Currently, 2 major structures have been used: a sponge and electrospun superfine fibres. The porous sponge structure is suitable for cell seeding in in vitro tissue engineering, but its mechanical strength is insufficient when combined with a rapidly degrading polymer [13, 14]. Electrospinning is a useful technique for fabricating various shapes of scaffolds (i.e. sheet, tube and valve leaflet). However, there are also unsolved intrinsic limitations of electrospun scaffolds, such as no porous structure, poor cell attachment and spreading, unpredictable and inhomogeneous absorption, sustained inflammatory reaction and significant bleeding from needle holes [18, 21]. To overcome the problems of sponge and electrospinning, the computational warp-knitting method was chosen to achieve sufficient mechanical strength and rich porosity for the SHF. In addition, both biodegradable and non-biodegradable yarns were used to maintain mechanical strength throughout the tissue restoration process, initially by PLLA and PET yarns and then later by the restored tissue and PET yarn. This mechanical stability might help to avoid deformity and promote regeneration of the aligned multilayered collagen and smooth muscle cells along both sides of the SHF. Serial histological examination clearly showed that numerous pores grew after the degradation of gelatin followed by PLLA and allowed bridging tissue development containing vasa vasorum as an oxygen and energy supply across the sheet. Limitations There are several limitations in this study. The sample size of the in vivo study was limited, especially in long-term follow-ups after implantation, and this made it difficult to draw a definite conclusion from the histological examination. However, since the SHF was produced under strict quality control with ISO standards, its mechanical properties and rate and extent of degradation were maintained within the consistent levels. Thus, we believe that the study shows, at least in part, the potential of SHF for in situ tissue regeneration in large vessels based on serial histological findings and regardless of individual differences among animals. The expandability of the non-biodegradable portion of the SHF after absorption of PLLA was not evaluated in this study. The expandable potential of the regenerated vascular tissue with body growth is crucial in paediatric congenital cardiac surgery and interventional dilation therapy, and further long-term follow-up is needed to examine this feature of the SHF. CONCLUSION In conclusion, we have developed a new warp-knitted fabric made of biodegradable and non-biodegradable yarns. The fabric exhibited sufficient mechanical properties for safe surgical implantation and avoiding deformity in canine aorta and IVC models. In situ vascular tissue regeneration was achieved with favourable integration. These results suggest that our new fabric may be an alternative to the current products used in paediatric congenital cardiac surgery. SUPPLEMENTARY MATERIAL Supplementary material is available at EJCTS online. Presented at the 31st Annual Meeting of the European Association for Cardio-Thoracic Surgery, Vienna, Austria, 7–10 October 2017. ACKNOWLEDGEMENTS We gratefully acknowledge Atsuko Ueda for coordinating the animal studies and data collection, Takafumi Ogawa (Kyodo Byori Inc., Kobe, Japan) for histological preparation and the Material Analysis Research Center of Teijin Ltd for gel permeation chromatography analysis. Funding This work was partially supported by a Development of Medical Devices Award through a Collaboration between Medicine and Industry from the Japan Agency for Medical Research and Development, AMED. Conflict of interest: none declared. REFERENCES 1 Hayabuchi Y , Mori K , Kitagawa T , Sakata M , Kagami S. Polytetrafluoroethylene graft calcification in patients with surgically repaired congenital heart disease: evaluation using multidetector-row computed tomography . Am Heart J 2007 ; 153 : 806.e1 – 8 . 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TI - In situ tissue regeneration using a warp-knitted fabric in the canine aorta and inferior vena cava
JF - European Journal of Cardio-Thoracic Surgery
DO - 10.1093/ejcts/ezy045
DA - 2018-08-01
UR - https://www.deepdyve.com/lp/oxford-university-press/in-situ-tissue-regeneration-using-a-warp-knitted-fabric-in-the-canine-BigLUoXELl
SP - 318
VL - 54
IS - 2
DP - DeepDyve
ER -