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Ann Thorac Surg 2009;88:1269-1276. doi:10.1016/j.athoracsur.2009.04.087
© 2009 The Society of Thoracic Surgeons
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Original Articles: Adult Cardiac

Newly Developed Tissue-Engineered Material for Reconstruction of Vascular Wall Without Cell Seeding

Hiroaki Takahashi, MD, PhDa, Takenori Yokota, MD, PhDb, Eiichiro Uchimura, PhDb, Shigeru Miyagawa, MD, PhDb, Takeyoshi Ota, MD, PhDa, Kei Torikai, MD, PhDb, Atsuhiro Saito, PhDb, Koichiro Hirakawa, MSc, Katsukiyo Kitabayashi, MDb, Kenji Okada, MD, PhDa, Yoshiki Sawa, MD, PhDb,*, Yutaka Okita, MD, PhDa

a Department of Surgery, Division of Cardiovascular Surgery, Kobe University Graduate School of Medicine, Kobe, Japan
b Department of Cardiovascular Surgery, Osaka University Graduate School of Medicine, Osaka, Japan
c Senko Medical Instrument Manufacturing Company Ltd, Tokyo, Japan

Accepted for publication April 16, 2009.

* Address correspondence to Dr Sawa, Department of Cardiovascular Surgery, Osaka University Graduate School of Medicine, 2-2 Yamada-oka, Suita, Osaka, 565-0871, Japan (Email: sawa{at}surg1.med.osaka-u.ac.jp).


Dr Hirakawa discloses a financial relationship with Senko Medical Instrument Manufacturing Co, Ltd.

 

   Abstract
Top
 Abstract
Introduction
 Material and Methods
 Results
 Comment
 Acknowledgments
 References
 
Background: We have developed a tissue-engineered patch for cardiovascular repair. Tissue-engineered patches facilitated site-specific in situ recellularization and required no pretreatment with cell seeding. This study evaluated the patches implanted into canine pulmonary arteries.

Methods: Tissue-engineered patches are biodegradable sheets woven with double-layer fibers. The fiber is composed of polyglycolic acid and poly-L-lactic acid, and compounding collagen microsponges. The patches (20- x 25-mm) were implanted into the canine pulmonary arterial trunks. At 1, 2, and 6 months after implantation (n = 4), they were explanted and characterized by histologic and biochemical analyses. Commercially available patches served as the control. No anticoagulant therapy was administered postoperatively.

Results: No aneurysm or thrombus was present within the patch area in all groups. The remodeled tissue predominantly consisted of elastic and collagen fibers, and the endoluminal surface was covered with a monolayer of endothelial cells and multilayers of smooth muscle cells beneath the endothelial layer. The elastic and collagen fibers and smooth muscle cells kept increasing with a maximum at 6 months, while a monolayer of endothelial cells was preserved. The expression levels of messenger RNA of several growth factors in the tissue-engineered patches were higher than those of native tissue at 1 and 2 months and decreased to normal level at 6 months. No regenerated tissue was found on the endoluminal surface in the control group.

Conclusions: The novel tissue-engineered patches showed in situ repopulation of host cells without prior ex vivo cell seeding. This is promising material for repair of the cardiovascular system.

Several commercially available patches have been clinically used for repair of the cardiovascular system, including surgical ventricular restorations for heart failure and right ventricular outflow reconstruction for congenital heart diseases [1, 2]. However, those synthetic materials, such as expanded polytetrafluoroethylene (ePTFE), Dacron (DuPont, Wilmington, DE), and glutaraldehyde-fixed equine pericardium, have well-known limitations, including infection, thrombogenicity, calcification, foreign body reaction, and the lack of growth potential.

Recently, biodegradable materials for cardiovascular operations have been developed and tested in preclinical and clinical studies. Shin'oka and colleagues [3] reported the first clinical application of a tissue-engineered vascular graft with pretreatment of human bone marrow cells. Likewise, tissue-engineered biodegradable materials with autologous cell seeding before implantation, where a bioreactor culture system is used, have been well documented as potential cardiovascular grafts [3–10]. However, the ex-vivo cell-seeding procedure is complicated, invasive, and can cause contaminations.

To overcome these problems, we had previously developed a tissue-engineered biodegradable patch that had sufficient durability to be used for vascular reconstruction. The design of our previous patch was composed of three layers: knitted polyglycolic acid (PGA), poly-{varepsilon}-caprolactone, and woven poly-L-lactic acid (PLLA). It showed acceptable in situ recellularization without prior cell seeding in an in vivo study [11, 12]. The previous patch still had problems, however, including insufficient cell repopulation, delamination of layers, and rigid material properties.

We developed the second generation of the tissue-engineered patch (TEP) by drastic structural modification. In this article, we described the characterization of remodeled tissue that was implanted into the wall of a canine pulmonary artery as a surgical patch to evaluate the efficiency of the newly developed TEP.


   Material and Methods
Top
 Abstract
 Introduction
 Material and Methods
Results
 Comment
 Acknowledgments
 References
 
Scaffold Design
The TEP was fabricated by compounding a collagen microsponge with a biodegradable polymeric scaffold that was woven with double-layer thread composed of PGA and PLLA. The double-layer thread was fabricated by an air-jet spinning technique with PLLA fibers as a core thread and PGA fibers as a sheath thread. They had a porous 3-dimensional structure. The core thread was composed of 30 PLLA fibers (each having a diameter of 20 µm), and the sheath thread was composed of 56 PGA fibers (each having a diameter of 15 µm; Fig 1A) [13].


Figure 1
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  		Fig 1. (A) Scaffold design of polyglycolic acid (PGA)/poly-L-lactic acid (PLAA) patch (B) Intraoperative view of the tissue-engineered patch (arrow) implanted to the canine pulmonary arterial trunk.

 
The TEPs were provided by Senko Medical Instrument Manufacturing Company Ltd (Tokyo, Japan). Five specimens of the current TEPs were tested to reveal the differences in the thickness and weight from the previous TEPs. The thickness of the current TEP is about 0.45 mm compared with 0.9 mm thickness of the previous TEP. The mean weight per a unit area in the current TEP is 27.4 ± 0.5 mg/cm2 compared with 36.2 ± 0.5 mg/cm2 for the previous TEP (Table 1).


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  		Table 1 In Vitro Study of Mechanical Properties
 
The TEPs with collagen microsponges were prepared using bovine type I collagen. The TEPs was immersed in a solution containing type I collagen (Koken Co Ltd, Tokyo, Japan), frozen at –80°C for 12 hours, and lyophilized in a vacuum of 0.2 mm Hg for an additional 24 hours. This process forms collagen microsponges. The collagen microsponges were cross-linked by glutaraldehyde vapor through exposure to a 25% glutaraldehyde aqueous solution at 37°C for 4 hours [14]. The TEPs were sterilized by exposure to ethylene oxide.

In Vitro Study of Mechanical Properties
Mechanical properties were evaluated by comparing the current TEP with the previous TEP [11, 15]. Five specimens were tested in each subgroup.

The pure bending property
The bending property under pure bending condition for patches was measured using KESFB2-AUTO-A (Kato Tech Co, Ltd, Kyoto, Japan). After the specimen is mounted, bending motion is performed automatically to measure stiffness and hysteresis (an arc of constant curvature: maximum curvature is ± 2.5 cm–1).

The tensile strength
Mechanical properties of the TEPs were evaluated using a tensile mechanical tester (Tensilon RTC-1150A, Orientec Co, Tokyo, Japan). The longitudinal and lateral strips (25 x 5 mm) were placed in the tensile tester and a maximum breaking point was measured.

The Expected Duration of Bioabsorption
In vitro study
The duration of bioabsorption was evaluated by degradation tests. After 0, 1, 3, 4, and 24 weeks of incubation in phosphate-buffered saline (PBS; Sigma, St Louis, MO) with shaking at 37°C, the tensile properties of 5 PGA and PLLA scaffolds were measured with a mechanical tester. The dimensions of the patches made of PGA or PLLA used for the tensile test were 5 x 20 mm, and the patches were pulled to failure at a rate of 50 mm/min. The tension strength was represented by N units (1 N = 1 MPa measured mm2).

In vivo study
In our other study [13], we investigated in situ tissue regeneration in small-diameter arteries using the same biodegradable vascular graft that did not require ex vivo cell seeding. For morphologic examination, the explanted grafts were observed with a scanning electron microscope.

Animal Operations
All animals received humane care in compliance with the Guide for the Care and Use of Laboratory Animals published by the National Institutes of Health (NIH publication No. 85–23, revised 1996).

An elliptical TEP (25 x 20 mm) or an expanded polytetrafluoroethylene patch (ePTFE) as a control was implanted into the canine pulmonary arterial trunk of 21 dogs (body weight: 19.2 ± 1.6 kg). The diameter of native canine pulmonary trunk was about 15 mm and the length was about 30 mm.

Anesthesia was induced with an intramuscular injection of 10 mg/kg ketamine (Sankyo Co, Tokyo, Japan) and 2 mg/kg of xylazine (Bayer Medical Co, Tokyo, Japan), and maintained by continuous intravenous infusion of propofol (AstraZeneca, Osaka, Japan) at a rate of 4 mg/kg/h throughout the operative procedure and additional injections of ketamine (5 mg/kg/h). Systemic anticoagulation was indicated with heparin (100 IU/kg).

The heart was exposed by a left anterolateral thoracotomy through the fourth intercostals space. After exposure and mobilization of the main and left pulmonary artery, the patch plasty was performed with side-bite partial clamp of the main pulmonary artery and total clamp of the left pulmonary artery, using a 5-0 monofilament running suture (Fig 1B). Heparin was reversed with protamine 100 IU/kg, followed by routine closure of the chest.

At 1, 2, and 6 months after implantation (N = 4 per each end point), the TEPs were explanted and evaluated by histologic and biological analyses or the control ePTFE patch was explanted (N = 3 per each end point). No anticoagulant was administered postoperatively.

Scanning Electron Microscopy
The explanted patches were incised longitudinally, and the middle of the patch area was examined with a scanning electron microscope (Model S-800; Hitachi, Tokyo, Japan).

Histologic and Immunohistological Examinations
The excised tissues were fixed in 10% phosphate-buffered formalin (pH 7.0) and embedded in paraffin. Each cross section included native tissues of proximal and distal ends, and patches. The sections were stained with hematoxylin and eosin and Victoria blue. A monoclonal antibody specific for {alpha}-smooth muscle actin (DAKO, Carpenteria, CA), and a polyclonal antibody against von Willebrand factor (DAKO) were used. We also measured the wall thickness of the neotissue microscopically, which was defined as regenerated tissue comprising newly organized autologous cells on the TEPs. Thickness was measured at three points: 5 mm from proximal and distal anastomoses and at the middle point of the patch. A random sampling technique was used.

Quantitative Determination of Collagen
A 4-hydroxyproline assay was used to measure collagen contents in the explanted TEPs 1, 2, and 6 months after implantation, as described previously [16]. The levels of collagen in the patches were compared with those of native canine pulmonary artery.

Mechanical Tensile Strength
Mechanical properties of the excised TEPs were evaluated using a tensile mechanical tester (Tensilon RTC-1150A, Orientec Co, Tokyo, Japan). The longitudinal tissue strips (25 x 5 mm) were placed in the tensile tester and a maximum breaking point was measured. As a control, canine native pulmonary arterial walls were used (N = 7).

Quantitative Reverse Transcription-Polymerase Chain Reaction
To quantify repopulated cells in the TEPs at 1, 2, and 6 months after implantation, the expression of vascular endothelial growth factor (VEGF) and smooth muscle 22{alpha} (SM22{alpha}), were determined with quantitative real-time reverse transcription-polymerase chain reaction (RT-PCR), as described previously [17]. Quantitative real-time RT-PCR was performed in quadruplicate using the TaqMan RT-PCR kit (ABI) in the Applied Biosystems 7700 sequence detector system (Applied Biosystems, Foster City, CA). Results were normalized to the level of glyceraldehyde-3-phosphate dehydrogenase (GAPDH) transcripts. The level of messenger RNA (mRNA) in the canine native pulmonary arterial wall was defined as 100%. Only TEP groups were evaluated because there was no regenerated tissue on the ePTFE patches. The sequences of the specific primers were as follows:

dog GAPDH: (forward) 5'-GTGATGCTGGTGCTGAGTATGTTG-3', (reverse) 5'-TGGCTAGAGGAGCCAAGCAGTT-3';
VEGF: (forward) 5'-GACGTCTACCAGCGCAGCTACT-3' (reverse) 5'-TTTGATCCGCATAATCTGCATG-3'; and
SM22{alpha}: (forward) 5'-AAGCTGGTCAACAGCCTGTATCC3', (reverse) 5'-ATAGAGGTCGACGGTCTGGAACA-3'.

Statistical Analysis
All values are expressed as the mean ± standard deviation. A standard t test was used to determine the significance of the difference between the two means, and the appropriate analysis of variance and Mann-Whitney test were used to compare the data among multiple groups. A value of p < 0.05 was considered to be statistically significant. Data analysis was performed with StatView 5.0 software (SAS Institute Inc, Cary, NC).


   Results
Top
 Abstract
 Introduction
 Material and Methods
 Results
Comment
 Acknowledgments
 References
 
In Vitro Study of Mechanical Properties
The pure bending property
The bending property of the current TEP was measured at 1.35 ± 0.17 gf x cm2/cm, in comparison with 2.5 ± 0.17 gf x cm2/cm measured for the previous TEP (p < 0.01). The result showed handling was improved in the current TEP compared with the previous TEP (Table 1).

The tensile strength
The mean maximal longitudinal tensile strength was 102.6 ± 14.6 N in the current TEP and 84.2 ± 2.7 N in the previous TEP (p = 0.01). The mean maximal lateral tensile strength was 50.4 ± 1.6 N in the current TEP and 61.1 ± 7.0 N in the previous TEP (p = 0.90). These results showed the current TEP was significantly stronger than the previous TEP in the longitudinal direction, although the differences were not statistically significant in the lateral direction (Table 1).

The Expected Duration of Bioabsorption
In vitro study
Maximal tensile strength was 47.2 ± 8.1 N after 0 week of incubation, 20.6 ± 4.5 N after 1 week, and 2.3 ± 0.2 N after 3 weeks in the PGA patch. This result indicated the PGA fibers were completely dissoluted up to 3 weeks. In the PLLA patch, maximal tensile strength was 50.3 ± 3.9 N after 0 week of incubation, 51.2 ± 4.3 N after 1 month, 47.4 ± 3.7 N after 6 months. Degradation tests (37°C) of the synthetic polymers, mainly involving hydrolytic reactions, demonstrated that the mechanical strength of PGA fibers were no longer adequate by 3 weeks, whereas PLLA fibers degraded so slowly that its strength was maintained through 6 months. Several reports showed it took longer than 1 year for PLLA fibers to be completely absorbed [18, 19].

In vivo study
A cross-section of the grafts showed that PGA fibers were completely absorbed as early as 2 months after implantation. However, PLLA fibers remained unabsorbed throughout the entire study (12 months after implantation) [13].

Macroscopic Findings
All animals tolerated the operation and survived without any postoperative complications until they were euthanized at 1, 2, and 6 months after implantation. There was no hematoma, seroma, aneurysm, or infection.

The endoluminal aspect of the TEPs at all time points was covered with a remodeled tissue with a thickness comparable to that of the adjacent pulmonary artery (Fig 2A, B). No thrombus was observed in the luminal surface of TEPs; however, thrombus formation was documented in some ePTFE patches (Fig 2C). Furthermore, there was always some endothelial defect into which blood had soaked, and the incidence of such defects was not associated with the duration of implantation. Longitudinal cross section of the explanted ePTFE patches showed intimal thickening at the proximal anastomosis (data not shown).


Figure 2
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  		Fig 2. Macroscopic observation of the tissue-engineered patches 1 and 6 months after implantation. There were no sign of thrombus or aneurysmal formation in any animals at (A) 1 month or (B) 6 months. (C) Macroscopic findings of expanded polytetrafluoroethylene patches at 6 months showed that endothelial defects were always detected. (D) Scanning electron microscopy image (x1000) shows luminal surface of the tissue-engineered patch 1 month after implantation.

 
Scanning Electron Microscope
The scanning electron microscope showed that the luminal surfaces of the TEPs were uniformly covered with confluent endothelial cells at 1 month after implantation (Fig 2D).

Histology and Immunohistochemistry
Histologic examination showed that the TEPs had a 3-layer structure that likely corresponded to intima, media, and adventitia (Figs 3E–G). In the TEPs, the remodeled tissue predominantly consisted of elastic and collagen fibers (Figs 3A–C), and the endoluminal surface was covered with a monolayer of von Willebrand factor–positive (endothelial cells) and multilayers of {alpha}-smooth muscle actin–positive cells beneath the endothelial layer (Fig 3E–G). These components tended to keep increasing, with a maximum at 6 months, preserving a monolayer of endothelial cells. The PLLA fibers still existed at 6 months after implantation (Figs 3C and G). These histologic examinations showed much better recellularization compared with the previous TEP [11].


Figure 3
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  		Fig 3. The upper side of the photomicrograph shows the luminal surface. Double staining with hematoxylin-eosin and Victoria blue of the tissue-engineered patches showed the presence of elastic fibers at (A) 1, (B) 2, and (C) 6 months after implantation (asterisks). Double staining with von Willebrand factor and {alpha}-smooth muscle actin showed the surface of the tissue-engineered patches were completely covered with monolayer of endothelial cells at (E) 1, (F) 2 and (G) 6 months (arrows), and {alpha}-smooth muscle actin was distributed throughout the regenerative tissue (asterisks). (D, H) In contrast, the surface of the expanded polytetrafluoroethylene patches was covered with thin layer that was not stained. The scale bar represents 300 µm.

 
Although the elastic components (blue in the hematoxylin and eosin and Victoria blue stain) and the cell density in the remodeled tissue of the TEPs were likely to increase, the thickness of the remodeled tissue was unchanged through 1 to 6 months. Specifically, the thickness of the regenerated tissue on the middle portion of TEP (excluding the TEP itself) was 1.37 ± 0.22 mm at 1 month, 1.54 ± 0.21 mm at 2 months, and 1.27 ± 0.21 mm at 6 months. In all groups, mean value at each measured point was approximately 1.4 mm. There were no significant differences in thickness at any point among the groups, indicating no tendency for wall thickness to increase in a time-dependent manner. Moreover, there were no significant differences when the three measured points were compared with each group (Table 2). In contrast, microscopic examination of ePTFE patches showed only partial regeneration of autologous tissue on the surfaces, although the surface of the ePTFE patches was covered with thin layer (Figs 3D and H).


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  		Table 2 The Wall Thickness (mm) of the Remodeled Tissue Formed in Implanted Polyglycolic Acid/Poly-L-Lactic Acid Patches
 
Quantitative Determination of Collagen
The 4-hydroxyproline assay demonstrated that collagen content in the TEPs at 1 month after implantation was already comparable with that in normal pulmonary artery. The collagen content in the TEPs at 2 and 6 months tended to increase compared with that in native pulmonary artery, although there was no significant difference among the groups. Collagen content was 238.8 ± 77.2 µg/mg (dry weight) in the native pulmonary artery, 338.1 ± 176.4 µg/mg in the TEPs at 1 month (p = 0.15 vs native pulmonary artery), and 361.0 ± 123.8 µg/mg in the TEPs at 2 months (p < 0.05 vs native pulmonary artery). The collagen content in the TEPs at 6 months was also significantly higher than that in the native pulmonary artery (375.8 ± 165.3 µg/mg, p < 0.05 vs native pulmonary artery; Table 3).


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  		Table 3 Biochemical Analysis
 
Mechanical Tensile Strength
The mechanical properties of the TEPs were compared with the native pulmonary artery. The TEPs were mechanically stronger at all time points, whereas mechanical properties of the TEPs approached those of the native pulmonary artery. Maximal tensile strength was 5.7 ± 3.6 N in native pulmonary artery, 58.0 ± 7.2 N in preimplant TEPs, 24.5 ± 14.7 N in the TEPs at 1 month, 16.4 ± 5.5 N in the TEPs at 2 months, and 12.1 ± 3.3 N in the TEPs at 6 months. The TEPs at 6 months after implantation were significantly stronger than the native pulmonary arteries (Table 3).

Reverse Transcription-Polymerase Chain Reaction
VEGF at 1 and 2 months was higher than that of the native tissue (1-month model: 178% ± 14% of normal, 2-month model: 232% ± 52% of normal). Expression of VEGF then dropped to a normal level at 6 months (108% ± 15% of normal; Table 4). The high expression of VEGF at 1 and 2 months appears to indicate that regenerative remodeling occurred in the TEPs at this period.


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  		Table 4 Quantitative Results of Reverse Transcription-Polymerase Chain Reaction a
 
There was an expression of SM22{alpha} at 1 month comparable with that of the native tissue (102% ± 55%). Expressions of SM22{alpha} at 2 and 6 months tended to be higher than that of the native tissue (2-month model: 136% ± 58%; 6-month model: 192% ± 34%). The SM22{alpha} expression in RT-PCR gradually increased in a time-dependent manner (Table 4). The expression of SM22{alpha} corresponds to increasing smooth muscle cells. On the other hand, RNA content of the tissue on the ePTFE patches was not sufficient, and the RT-PCR data in the control group could not be evaluated.


   Comment
Top
 Abstract
 Introduction
 Material and Methods
 Results
 Comment
Acknowledgments
 References
 
The previous patch had 3 layers where the PGA and the PLLA layers independently existed. In the current TEP, biodegradable fiber itself consists of PGA and PLLA. By changing the design of the polymer scaffold, the TEP was improved in the aspects of cell repopulation and material properties. This study showed that the TEP was capable of accommodating repopulated host cells. Histologic analyses showed that collagen and elastin was constructed in the early phase (ie, within 1 month) in conjunction with regeneration of a monolayer of endothelial cells. Smooth muscle cells beneath the endothelial layer then distributed toward the depth. Immunohistochemically, all the implanted TEPs showed complete endothelialization and a population of much more numerous SMCs with parallel accumulation in the subendothelial layers compared with the previous study [11]. Furthermore, handling in the current TEP was improved compared with the previous TEP. We speculated the improvement in the mechanical properties led to the promotion of elasticity in the neotissue.

On the other hand, the remodeled tissue on the TEPs was uniform in thickness in the entire patch and unchanged through 6 months. It seemed that this well-organized remodeling resulted from early regenerated endothelial cells. In addition, the early reendothelialization contributed to preventing thrombosis, although we did not use any postoperative antiplatelet or anticoagulant therapy [20]. Endothelial cells are known to produce various vasoactive factors and modulate vascular growth by secreting several antiproliferative factors, such as nitric oxide and vascular natriuretic peptides [21, 22]. Vascular smooth muscle cell growth is controlled by a balance of growth inhibitors and promoters [23]. Therefore, the presence of normal endothelial cells might have led to well-controlled vascular smooth muscle cell repopulation without intimal hyperplasia [24].

For the evaluation in the early phase, scanning electron micrographic examination of the luminal surfaces of the TEPs at 3 hours after implantation revealed cell adhesion and complete coverage on the scaffold surfaces (Fig 4A). On the other hand, scanning electron microscopy images at 3 hours after implantation showed that fibrils were directly exposed to the lumen of ePTFE without any endothelial coverage (Fig 4B). Although this experimental data by scanning electron microscopy was not enough to identify the mechanism of tissue regeneration in early phase, we speculated that the very efficient preclotting was what happened on the TEP in the early phase and it led to the good patency.


Figure 4
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  		Fig 4. (A) Scanning electron micrographs of luminal surfaces of tissue-engineered patches 3 hours after implantation showed cell adhesion and complete coverage on the scaffold surfaces. (B) Images at 3 hours after implantation showed that fibrils were directly exposed to the lumen of expanded polytetrafluoroethylene without any endothelial coverage. The scale bar represents 50 µm.

 
RT-PCR analyses showed expression of VEGF at 1 and 2 months after implantation in the TEPs was higher than that of the native tissue. This result suggests that the remodeling process is occurring in the TEPs during this period. Consequently, the number of smooth muscle cells tended to gradually increase, which was supported by the results of RT-PCR and histology. The expression of those growth factors reduced to the normal level of the native pulmonary artery at 6 months after implantation, which might indicate that the regeneration process led to complete constructive vascular remodeling.

The PGA component possesses 3-dimensional porous structures, which facilitates cell attachment and promotes in situ recellularization. The PLLA component reinforces the device and increases the durability. In our previous study, cell culture in vitro showed that proliferation of the cells seeded onto TEPs with the collagen microsponge was significantly higher than that onto TEPs without the collagen microsponge [13, 15]. The collagen microsponge is then incorporated to achieve complete sealing and enhance host cell repopulation. The PGA component is supposed to contribute to early site-specific remodeling and then degrade in an early phase (ie, less than 30 days), which was observed by the results of histology and scanning electron microscopy at the 1-month model in this study.

Although the residual PLLA fibers in the TEPs at 6 months after implantation did not contribute to the mechanical strength, the strength of the TEP even 6 months after implantation was greater than that of a native pulmonary artery. In addition, the collagen content in the TEPs at 2 and 6 months after implantation was significantly higher than in the native tissue. The increasing collagen content likely contributed enough to the strength of TEPs to withstand the venous blood pressure.

In the present study, we obtained excellent in situ regeneration. Bauer and colleagues [25] described that after trauma or tissue-level ischemia, elevated systemic concentrations of VEGF caused more endothelial progenitor cells to be released from the bone marrow into peripheral circulation. At the needed tissue-level site, the bone marrow-derived endothelial progenitor cells exit the circulation, invade or migrate through the extracellular matrix, and contribute to neovascularization by either direct vasculogenesis or by paracrine mechanisms [25, 26]. Although the mechanisms behind the in situ regeneration have not been clear in our study, we can speculate that circulating vascular progenitor cells attach to the TEP, and cellular and extracellular components from the native tissue also expand to the TEP [27]. Our goal of tissue engineering is to create scaffold constructs to direct tissue regeneration and restore function through the delivery of living elements that become integrated into the patient.

Despite the contributions of this study, several limitations have to be addressed. The first limitation is the choice of our animal model. The durability and growth potential of our TEP should be evaluated over a long-term follow-up period in another model, such as sheep, for the clinical approach in the repair of congenital heart diseases.

Second, although our results revealed site-specific regeneration of vascular tissue, more studies are needed to identify the mechanism of tissue regeneration, especially in an early phase.

Third, the TEP was tested in the low-pressure environment of pulmonary artery patch plasty. Further study in a high-pressure environment such as descending aorta replacement will be needed to verify the durability of the TEPs.

In conclusion, we developed a novel tissue-engineered patch that showed constructive in situ remodeling by site-specific host cells without prior ex vivo cell seeding. This study suggested our newly developed the tissue-engineered patch would be a promising surgical material for repair of cardiovascular system.


   Acknowledgments
Top
 Abstract
 Introduction
 Material and Methods
 Results
 Comment
 Acknowledgments
References
 
The PGA/PLLA patches were provided by Senko Medical Instrument Mfg Co, Ltd (Tokyo, Japan). Financial support was obtained from the Department of Cardiovascular Surgery, Osaka University Graduate School Of Medicine. We thank Dr Kawaguchi (Department of Molecular Pathology, Osaka University Graduate School of Medicine, Japan) for the histological analysis, Ms Masako Yokoyama (Department of Cardiovascular Surgery, Osaka University Graduate School of Medicine, Japan) for expert assistance with RT-PCR, and Mr Shigeru Matsumi (Department of Cardiovascular Surgery, Osaka University Graduate School of Medicine, Japan) for excellent technical support.


   References
Top
 Abstract
 Introduction
 Material and Methods
 Results
 Comment
 Acknowledgments
 References
 

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