Ann Thorac Surg 2006;81:57-63
© 2006 The Society of Thoracic Surgeons
Original article: Cardiovascular
In Vitro Fabrication of a Tissue Engineered Human Cardiovascular Patch for Future Use in Cardiovascular Surgery
Chao Yang, MD
a
,
Ralf Sodian, MD
a
,
*
,
Ping Fu, MD
a
,
Cora Lüders, PhD
a
,
Thees Lemke, MD
a
,
Jing Du, MD
b
,
Michael Hübler, MD
a
,
Yuguo Weng, MD
a
,
Rudolf Meyer, MD, PhD
a
,
Roland Hetzer, MD, PhD
a
a Department of Cardiothoracic and Vascular Surgery, Laboratory for Tissue Engineering, Deutsches Herzzentrum Berlin
b Department of Cardiology, Charité-Universitätsmedizin Berlin, Campus Benjamin Franklin, Berlin, Germany
Accepted for publication July 5, 2005.
* Address correspondence to Dr Sodian, Department of Cardiothoracic and Vascular Surgery, Laboratory for Tissue Engineering, German Heart Institute Berlin, Augustenburger Platz 1, 13353 Berlin, Germany (Email: sodian{at}dhzb.de).
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Abstract
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BACKGROUND: One approach to tissue engineering has been the development of in vitro conditions for the fabrication of functional cardiovascular structures intended for implantation. In this experiment, we developed a pulsatile flow system that provides biochemical and biomechanical signals in order to regulate autologous, human patch-tissue development in vitro.
METHODS: We constructed a biodegradable patch scaffold from porous poly-4-hydroxy-butyrate (P4HB; pore size 80 to 150 µm). The scaffold was seeded with pediatric aortic cells. The cell-seeded patch constructs were placed in a self-developed bioreactor for 7 days to observe potential tissue formation under dynamic cell culture conditions. As a control, cell-seeded scaffolds were not conditioned in the bioreactor system. After maturation in vitro, the analysis of the tissue engineered constructs included biochemical, biomechanical, morphologic, and immunohistochemical examination.
RESULTS: Macroscopically, all tissue engineered constructs were covered by cells. After conditioning in the bioreactor, the cells were mostly viable, had grown into the pores, and had formed tissue on the patch construct. Electron microscopy showed confluent smooth surfaces. Additionally, we demonstrated the capacity to generate collagen and elastin under in vitro pulsatile flow conditions in biochemical examination. Biomechanical tesing showed mechanical properties of the tissue engineered human patch tissue without any statistical differences in strength or resistance to stretch between the static controls and the conditioned patches. Immunohistochemical examination stained positive for alpha smooth muscle actin, collagen type I, and fibronectin. There was minor tissue formation in the nonconditioned control samples.
CONCLUSIONS: Porous P4HB may be used to fabricate a biodegradable patch scaffold. Human vascular cells attached themselves to the polymeric scaffold, and extracellular matrix formation was induced under controlled biomechanical and biodynamic stimuli in a self-developed pulsatile bioreactor system.
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Introduction
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The application of patches or grafts in cardiovascular surgery, and particularly in pediatric cardiac surgery, is a widely accepted surgical technique for repair or reconstruction of cardiovascular structures [1, 2]. Currently, the patch materials used clinically are limited to prosthetic materials, autologous pericardium, and allogenic or xenogenic (glutaraldehyde-fixed) pericardium [2]. All materials used for patch repair or reconstruction have limitations such as their inability to grow, repair, and remodel. Aneurysm formation and the inability of patches to grow or remodel are important sources of morbidity and mortality after repair or reconstruction of cardiovascular structures, especially in children and young adults. Tissue engineering is proposed as a solution to these problems by transplanting autologous cells onto biodegradable scaffolds to ultimately form new functional autologous tissue.
To fabricate an implantable human autologous and viable tissue patch, our laboratory focused on applying the principles of tissue engineering to fabricating human cardiovascular structures for implantation. We have now developed a tissue-engineered (TE) patch, which was seeded with human vascular cells from pediatric patients and transferred into a self-developed pulsatile patch bioreactor to induce the formation of a viable and surgically feasible patch to be used in cardiovascular surgery. This study was designed to evaluate the new, dynamic in vitro conditions for tissue engineering of human cardiovascular patches for potential reconstructive surgery of congenital defects.
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Material and Methods
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Scaffold Materials
Biodegradable polymer poly-4-hydroxybutyrate (P4HB; Tepha, Cambridge, MA) was used for cell seeding as a round sheet (pore size 80 to 150 µm, diameter 45 mm, thickness 0.5 mm). The P4HB is a rapidly absorbable biopolymer biologically derived from bacteria that is strong and pliable. It is a semicrystalline, thermoplastic elastomer with a melting point of about 60°C and a glass transition temperature of –50°C. The polymer scaffold was sterilized with cold gas (ethylene oxide).
Cell Harvesting
Cell harvesting was approved by the Ethics Committee of the Charité Medical School, Humboldt University, Berlin, Germany. Tissue from the human ascending aorta was harvested from pediatric patients undergoing heart transplantation or corrective or reconstructive aortic surgery. After harvesting, the arterial tissue was rinsed of blood with phosphate buffered saline (PBS [Gibco Life Technologies, Gaithersburg, MD]) and the tunica adventitia of the aorta was peeled off. Using a scalpel the tissue was then minced into 1 mm x 1 mm pieces, which were distributed over the bottom of 10 cm2 Petri dishes for primary culture. The cells were cultured in Dulbecco's Modified Eagle Medium (DMEM [Gibco Life Technologies]) supplemented with 10% fetal calf serum (FCS [Sigma-Aldrich, Taufkirchen, Germany]) and 1% penicillin-streptomycin-glutamine solution (Gibco Life Technologies). The explants were placed in a humidified incubator at 37°C with 5% CO2. After 2 weeks, when the cell population had grown into a 80% to 100% confluent monolayer, the primary culture was passaged. Sufficient cell numbers for cell seeding on bioabsorbable polymer scaffolds were obtained after 21 to 28 days (passage 3 to 4). The medium was changed every 4 days. Before cell seeding, cellular phenotype was clarified using immunofluorescence methods with monoclonal antibodies for alpha smooth muscle actin (DAKO Corporation, San Diego, CA), smooth muscle myosin heavy chain (Sigma-Aldrich), CD31 (DAKO Corporation), and fibronectin (Novocastra Lab, Bagsvaerd, UK).
Scaffold Seeding
The sterilized polymer scaffolds were soaked in PBS solution at 37°C for 4 hours, subsequently precoated with DMEM with 10% FCS and 1% penicillin-streptomycin-glutamine solution for 24 hours before cell seeding. The scaffolds were seeded with 8 million vascular cells each day on 3 consecutive days, and then were cultured in a humidified incubator at 37°C with 5% CO2 for 4 days. From the 8th to the 10th day, the other side of the scaffolds was seeded with cells in the same way. We developed a new seeding technique using two metal rings to seed both sides of the patch and provide an optimal amount of cell culture medium to both sides of the patches (Fig 1). The cell-polymer constructs were further incubated in the bioreactor or in static nutrient media as a control.
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Fig 1. Schema illustrating the principle of metal rings used in cell seeding. The scaffold was positioned between two metal rings to concentrate the cell suspension on the seeded area of patch. The lower metal ring had a gap, which was used for deairing of the space under the patch and allowed the cell medium subsequently to enter through it because two sides of the patch were seeded. The upper metal ring with no gap fixed the polymer and circumscribed the cell suspension within the seeding area.
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Cell Seeded Patch Cultivation
From the 15th day, four cell-seeded polymer constructs were incubated in static nutrient media in a humidified incubator at 37°C with 5% CO2 for 7 days. On the 22nd day, the polymer constructs were transferred into the bioreactor under sterile conditions and incubated for 7 days more. The patch bioreactor system used in this experiment is self-developed and has been described in detail [3]. The whole system is made of acrylic glass (Berlin Heart AG, Berlin, Germany) and is a highly isolated cell culture setting, which is designed for long-term patch tissue conditioning. It combines continuous, pulsatile perfusion and mechanical stimulation by periodically stretching the TE patch constructs under controlled conditions in a humidified incubator. The bioreactor was adjusted to a low flow of 250 mL/min and a systolic pressure of 8 mm Hg. As a control, four other cell-seeded constructs were incubated in static nutrient media in the humidified incubator at 37°C with 5% CO2 for 14 days. The cell culture medium was changed every 4 days.
Structure Assessment
Sections of the TE patches were fixed in 8% phosphate-buffered formalin, embedded in paraffin, and stained with hematoxylin and eosin. For immunohistochemical staining the alkaline phosphatase/antialkaline-phosphatase method was utilized with monoclonal antibodies for collagen I (Novocastra Lab), alpha smooth muscle actin (DAKO Corporation), fibronectin (Novocastra Lab), and CD31 and CD34 (DAKO Corporation).
For scanning electron microscopic examination, a representative portion of the TE patch was fixed in 2.5% glutaraldehyde, followed by gradual dehydration. Afterward, a piece of the specimen was critical point dried in CO2 using a Hitachi HCP-2 critical point dryer (Hitachi, Tokyo, Japan) and was subsequently sputter-coated in a Jeol JEC-1100 ion sputter (Hitachi). The samples were examined with Philips XL-30 scanning electron microscopes (FEI; Philips, Hamburg, Germany).
For DNA assay, the QIAamp DNA Minikit 50 (Qiagen, Duesseldorf, Germany) was used. Samples were gently rinsed with PBS twice and were then cut into small pieces. The cells in the samples were lysed using proteinase K at 56°C for 12 hours in a shaking water bath. The sample lysate was loaded onto the mini-column and then centrifuged at 6,000g for 1 minute. The DNA that was bound to the membrane of the Spin column was eluted in buffer after washing. Total DNA was quantified by spectrophotometry with absorbance at 260 nm.
To assess the presence of extracellular matrix formation biochemically, collagen assay and elastin assay were performed using a collagen assay kit (Biocolor, Newtownabbey, Northern Ireland) and an elastin assay kit (Biocolor), respectively, according to the product recommendations. Samples were lyophilized and weighed. For collagen and elastin measurement, the absorbance of the spectrophotometer (Shimadzu Deutschland GmbH, Duisburg, Germany) was set at 540 nm and 513 nm, respectively.
Biomechanical uniaxial tension testing of all tissue engineered patches was performed with the Instron 4465 materials testing machine (Instron, Darmstadt, Germany). The testing was carried out at room temperature (22°C). A rectangular shaped specimen was cut from the TE patch (5 x 6 mm) using a scalpel. The specimen was attached at opposite ends to the test apparatus. One arm of the test apparatus progressively stretched the specimen (at 2.0 mm per minute) until failure (complete tear). The passive tensile strength of each specimen was continuously recorded during the displacement. These data were plotted for each specimen as a displacement versus force graph and compared between the two groups. Maximal patch stress (index of maximal tensile strength) was determined by the peak of the curve whereas tissue resistance (inverse of elastance) to stretch was determined by the slope of the curve. Engineering strain (e) of the specimen was expressed as the ratio of the change in length (
L) to the original length (L), e = L/L.
Statistics
Data results were expressed as mean ± SD. Comparisons between the conditioned TE patch group and the static control group were performed with paired t test (Student's t test). Statistical significance was set at p less than 0.05. Data and graphs were produced with MS Excel 2000 (Microsoft Corporation).
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Results
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Characterization of Cell Cultivation
Before seeding of the polymer patch, the cell culture showed elongated cells with fibroblast-like morphology (Fig 2A). Fluorescence immunohistochemistry of the cultivated cells showed positive staining for alpha smooth muscle actin (Fig 2B) and smooth muscle myosin heavy chain (Fig 2C). Immunofluorescent staining of cultivated cells showed negative staining of CD31 (Fig 2D). The cellular morphology did not change over the period of cell cultivation.
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Fig 2. (A) Elongated cells with fibroblast-like morphology reached confluence after 10 to 14 days. The cultivated cells are bipolar in shape and look like a fish shoal. Fluorescence immunohistochemistry of the cultivated cells showed positive staining for (B) alpha smooth muscle actin and (C) myosin heavy chain. (D) Cultivated cell cultures showed negative immunofluorescent staining of CD31; cell nuclei are stained with 6-diamidino-2-phenylindole. (A: magnification 100x; B, C, D: magnification 400x.)
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Dynamic Cell Culture Conditioning
The bioreactor was easily assembled and placed in the humidified incubator. During the dynamic conditioning period, the whole system was compact without any sign of leakage or contamination and functioned stably.
Morphology
Macroscopically, all cell-seeded patches were covered by cells. The gross appearance showed that all TE constructs were intact without any rupture of structure, and both sides of patches exhibited a confluent and smooth surface.
Histology
Microscopically, all TE patch constructs were covered by cells. Hematoxylin and eosin staining of bioreactor conditioned TE patch sections demonstrated cellular tissue organized in a layered fashion with a dense outer layer and lesser cellularity in the deeper portions (Fig 3). The histologic examination showed mutiple, confluent cell layers on both sides of the P4HB patch scaffold. Additionally, after conditioning in the pulsatile flow bioreactor, the cells were mostly viable and had grown into the pores and formed tissue inside the scaffold. Static controls showed a loose, less organized tissue formation with irregular cellular ingrowth (Fig 4).
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Fig 3. Hematoxylin and eosin staining of bioreactor conditioned tissue-engineered patch. (Magnification 200x.)
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Immunohistochemistry
Immunohistochemistry showed positive staining for collagen type I (Fig 5A), alpha smooth muscle actin (Fig 5B), and fibronectin (Fig 5C) throughout the TE constructs, whereas staining for CD31 and CD34 was negative, as expected.
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Fig 5. (A) Positive immunohistochemical staining for collagen I of a tissue-engineered (TE) patch seeded with human vascular cells. Open arrow shows stained collagen I. Polymeric scaffold (solid arrow) was partially absorbed. (B) Immunohistochemical staining for alpha smooth muscle actin (open arrow) of TE patch was positive. (C) Staining for fibronectin (open arrow) of TE patch was positive. (Magnification 200x.)
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Scanning Electron Microscopy
The scanning electron microscopic examination showed that cells attached well to the surface of the P4HB scaffold and formed a homogeneous confluent smooth surface after conditioning in the bioreactor system (Fig 6A), whereas the static controls showed a less confluent, inhomogeneous surface (Fig 6B).
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Fig 6. (A) Scanning electron microscopic examination of the conditioned tissue-engineered patch demonstrated homogenous tissue and confluent smooth surfaces with cell orientation in the direction of flow exposition. (B) The static controls showed a less confluent, inhomogenous surface without cellular orientation. (Acc.V = accelerating voltage; WD = working distance.) (Magnification 500x.)
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Cell Proliferation (DNA Assay)
The DNA assay demonstrated cell proliferation among groups. The DNA content of the bioreactor conditioned patch was significantly higher than that of the static controls (p < 0.05). Comparison of DNA content (DNA/tissue dry weight) between the two groups is illustrated in Figure 7. The results showed that the conditioned patches underwent more cell proliferation than the nonconditioned patches.
Extracellular Matrix Formation
Additionally, we demonstrated the capacity to generate collagen and elastin under in vitro pulsatile flow conditions in our biochemical examination. Collagen and elastin formation occurred under both static and flow conditions. The conditioned patch showed significantly higher collagen values than the static patch (p < 0.05; Fig 8A). The concentration of elastin was not significantly different between the two groups (p > 0.05; Fig 8B).
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Fig 8. (A) Collagen concentration of the dynamic versus the static tissue-engineered patch. (B) Quantitative biochemical assays of elastin of the tissue-engineered patch.
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Biomechanical Testing
The testing showed appropriate mechanical properties of the TE human patch tissue without any statistical differences in strength or resistance to stretch between the static controls and the conditioned patches (Table 1).
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Comment
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Cardiovascular diseases are the most common causes of death and serious morbidity in the world [4–7]. Additionally, congenital heart disease is a considerable problem worldwide, affecting approximately 1% of infants, and is associated with significant morbidity and mortality [8]. Congenital heart defects, such as atrial septal defect, ventricular septal defect, double outlet ventricles and hypoplastic left heart syndrome, and ischemic heart disease are associated with aplastic, defective, or necrotic myocardial structures. In many of these instances, surgical treatment is necessary and patch closure, reconstruction of the defect, or revascularization is required [9]. The patch materials currently used are limited to prosthetic materials, autologous pericardium, and allogenic or xenogenic (glutaraldehyde-fixed) pericardium [2]. All of these materials are unable to grow, repair, and remodel, which leads to aneurysm formation in patch aortoplasty, inelasticity of the prosthetic materials, and an increased risk of hemolysis induced by contact of the blood with the prosthetic materials [10, 11]. The optimal cardiovascular patch materials should be characterized by long durability, absence of thrombogenicity, resistance to infections, lack of antigenicity, potential for growth, and ability to prevent patch dilatation and thinning. To meet these requirements, our laboratory applies the principles of tissue engineering in an attempt to develop viable, autologous replacements for deficient cardiovascular structures.
One approach in tissue engineering is to seed autologous cells onto biodegradable polymers and to establish optimal in vitro conditions to guide tissue development and to create cell-polymer constructs with a high degree of maturity before implantation. To fabricate such constructs in vitro, it has been widely confirmed that a dynamic tissue environment significantly enhances tissue maturation and mechanical properties [12–16]. Pulsatile flow or fluid dynamics have a well-known impact on cell morphology [17], proliferation [18], and composition of extracellular matrix that may lead to TE constructs suitable for implantation [19–21]. Therefore, in our laboratory we have developed a new pulsatile flow system that provides biochemical and biomechanical signals to regulate ovine patch tissue development in vitro.
Since our early testing of the concept appeared promising, in the current study we modified the whole system to create human tissue. Additionally, in this experiment we were able to show that it is possible to fabricate appropriate human cardiovascular patch constructs using this concept. In our experiment, we started with a relatively low systolic pressure to improve cellular attachment to the polymeric scaffold. From previous experiments we learned that systemic pressure in the early phase of conditioning washes a significant number of cells off the scaffold. After successful conditioning in vitro, which includes the ingrowth of cells into the polymer, systemic pressure conditions do not wash off the newly formed neotissue. In addition, our new pulsatile patch bioreactor is a compact, stable dynamic flow system. It can be easily placed into a standard cell incubator, representing a highly isolated dynamic cell culture setting with maximum sterility, optimal gas supply, and stable temperature conditions especially suited for long-term experiments in human cardiovascular tissue engineering.
Moreover, we describe a whole tissue engineering concept for potential human application. We chose aortic derived cells because we are working in a human model, and this was a source of cells that we were able to obtain during cardiac surgery without sacrificing a healthy vessel. Therefore, there are mainly ethical reasons to choose this cell source, although it is probably not very practicable for clinical application, and not to sacrifice an intact peripheral vessel or even perform a cell harvest from bone marrow. Furthermore, experience with these cell types shows that the results are easy to transfer to cells from peripheral vessels.
Our results demonstrated positive evidence that the seeded vascular cells adhered to, migrated into, proliferated in, and differentiated in the porous P4HB scaffold to form viable, oriented, and confluently layered tissue without any signs of contamination in our bioreactor system. The cells appeared viable on the polymeric patch scaffold. In addition, after conditioning in our bioreactor, each polymeric scaffold was partially absorbed and replaced with cells and extracellular matrix. In contrast to the static controls, the conditioned patch constructs were covered completely with organized tissue in a layered fashion and did not form unorganized tissue particles. Additionally, more cells penetrated and grew into the deeper layer of the conditioned scaffolds. The DNA assay demonstrated different cell proliferation between the groups. The bioreactor-conditioned patches showed an increase of cell proliferation compared with static controls.
Moreover, the scanning electron microscope examination of the TE patches demonstrated that the vascular cells in conditioned patches attached well to the surface of the polymeric scaffold, exhibited a characteristic elongated bipolar spindle shape, and were oriented parallel to flow direction, whereas those in static controls were polygonal shaped and randomly oriented. This finding emphasizes that the biomimetic in vitro environment influences the degree of tissue organization and the cell orientation in human tissue constructs.
Additionally, our results further demonstrated formation of extracelluar matrix proteins (collagen, elastin) under pulsatile flow conditions as measured by Biocolor assays and stained positive for collagen and fibronectin. These findings indicated that the new bioreactor system induces not only cell proliferation but also has a beneficial impact on the composition of extracellular matrix, and facilitates the development of new tissue and enhances tissue maturation. One limitation of our study was that we found no difference in the distribution of the overall small amount of elastin in our constructs. Moreover, we could not prove the presence of certain elastin layers. These findings are not very surprising because the lack of elastin is a well-known problem in the fabrication of in vitro conditioned constructs [15]. One potential reason for this lack may be that, during conditioning, the elastic polymeric scaffold overtakes the role of elastin in native tissue, so there are no adequate stimuli for the cells to synthesize elastin. Nevertheless, this is an important issue and is currently under investigation in our laboratory. Moreover, there was no positive staining for CD31 and CD34, indicating the absence of endothelial cells in the TE patch. This result was expected, because no endothelial cells were electively seeded on the patch.
The current experiment is one of the early studies that attempt to create human cardiovascular tissue in vitro. Although all currently used cardiovascular prostheses (namely, polytetrafluoroethylene [PTFE], Dacron, fixed pericardium, xenografts, mechanical heart valves, and homografts) do not have any endothelial coverage but function well in thousands of patients, there is no question that a functional endothelial coverage would be beneficial in a tissue engineered cardiovascular construct. Therefore, to develop a feasible seeding and in vitro conditioning method is one of our ultimate goals in the experiments to follow.
Although our early results with tissue engineering of autologous human cardiovascular patches appear promising, issues such as endothelial coverage and control of extracellular matrix formation have to be addressed in future experiments before a final human implantation.
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Acknowledgments
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This study was partly supported by a grant from the Deutsche Forschungsgemeinschaft (462/1-1). We thank David P. Martin, PhD (Tepha, Cambridge, Massachusetts), for his generous gifts of P4HB and his advice concerning the polymeric scaffold. We thank Professor Dr Monica Bauer (Fraunhofer Institute, Teltow, Berlin) for her help and advice concerning the biomechanical testing of the constructs. Furthermore, we thank Anne Gale (editorial), Annette Gaussmann (graphics) and Carla Weber (graphics) for their assistance in the preparation of the manuscript.
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Abstract
Introduction
