- Email: [email protected]
Contractile cardiac grafts using a novel nanofibrous mesh
ARTICLE IN PRESS
Biomaterials 25 (2004) 3717–3723
Contractile cardiac grafts using a novel nanofibrous mesh$ M. Shina,*,1, O. Ishiia,b,1, T. Suedab, J.P. Vacantia a
Department of Surgery, Massachusetts General Hospital and Harvard Medical School, Wellman 627, 55 Fruit Street, Boston, MA 02114, USA b First Department of Surgery, Hiroshima University School of Medicine, 1-2-3 Kasumi, Minami-ku, Hiroshima 734-8551, Japan Received 25 September 2003; accepted 21 October 2003
Abstract Cardiomyoctes are terminally differentiated cells and therefore unable to regenerate after infarction. The use of autologous bioengineered cardiac grafts has been suggested to replace infarcted myocardium and enhance cardiac function. Here we report the development of an in vitro system for engineered myocardium. Cardiac nanofibrous meshes (CNM) were developed by culturing cardiomyocytes from neonatal Lewis rats on electrospun, nanofibrous polycaprolactone (PCL) meshes. The mesh had an ECM-like topography and was suspended across a wire ring that acted as a passive load to contracting cardiomyocytes. The cardiomyocytes started beating after 3 days and were cultured in vitro for 14 days. The cardiomyocytes attached well on the PCL meshes and expressed cardiac-specific proteins such as a-myosin heavy chain, connexin43 and cardiac troponin I. The results demonstrate the formation of contractile cardiac grafts in vitro. Using this technique, cardiac grafts can be matured in vitro to obtain sufficient function prior to implantation. It is conjectured that cardiac grafts with clinically relevant dimensions can be obtained by stacking CNMs and inducing vascularization with angiogenic factors. r 2003 Elsevier Ltd. All rights reserved. Keywords: Cardiac tissue engineering; Polycaprolactone; Heart; Electrospinning
1. Introduction The limited ability of myocardium to regenerate after extensive infarction has led to the development of new treatment modalities to regenerate functional myocardium. Autologous cell transplantation into damaged hearts has received significant attention, and two general approaches have been studied. The first approach is to inject cell suspensions directly into the heart. A number of studies have been carried out in animals and humans using a variety of cell types. The results are promising and demonstrate that implanted cells can incorporate themselves into the host hearts [1,2]. The second approach combines cells and biodegradable scaffolds to provide temporal structural support and guide tissue regeneration. To this end, different scaffold materials, cell sources and bioreactors have been assessed to obtain viable and functional tissue. $ Supplementary data associated with this article can be found, in the online version, at doi: 10.1016/j.biomaterials.2003.10.055 *Corresponding author. Tel.: +1-617-724-0832; fax: +1-617-7265057. E-mail address: [email protected] (M. Shin). 1 M. Shin and O. Ishii contributed equally to this study.
0142-9612/$ - see front matter r 2003 Elsevier Ltd. All rights reserved. doi:10.1016/j.biomaterials.2003.10.055
Li et al. produced three-dimensional (3-D) contractile cardiac grafts using gelatin sponges. These grafts were used to replace myocardial scar tissue and right ventricular outflow tract defects in rats [3,4]. It was found that mechanical stretching enhanced the formation of tissue-engineered grafts [5]. In addition, poly(e caprolactone-co-l-lactide) (PCLA) sponge reinforced with knitted poly-l-lactide fabric (PLLA), and polyglycolic acid (PGA) were assessed [6]. Leor et al. reported the formation of bioengineered cardiac grafts with 3-D alginate scaffolds. These grafts stimulated neovascularization and attenuated left ventricular dilatation and heart failure in a rat infarction model [7]. Eschenhagen et al. engineered 3-D heart tissue by gelling a mixture of rat cardiomyocytes and collagen [8]. These constructs were cast in circular shapes and subjected to a cyclical mechanical stress. The morphology, electrophysiology and contractile properties of these grafts were characterized [9]. The engineered heart tissue survived and matured after implantation on uninjured hearts in a rat model [10]. Recently, Okano’s group has developed cell sheet engineering without scaffolds. Using a novel temperature-
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responsive polymer, poly(N-isopropylacrylamide) (PIPAAm), cultured rat cardiomyocytes could be detached as a continuous sheet without destroying cellcell junction proteins and extracellular matrix (ECM) [11]. To create thicker grafts, several cell sheets were layered on top of each other. Electrical and morphological connections between the layers were verified both in vitro and in vivo [12,13]. Despite these advances, several challenges remain. In addition to myocardial cell sourcing, cell survival within the graft is of importance. Typically, the cells are dense in the graft periphery but sparse in interior due to insufficient oxygen and nutrient perfusion [14]. To avoid core ischemia in grafts of clinically relevant dimensions, either new methods to accelerate blood vessel formation or connected layers of thinner grafts are required. Furthermore, cardiomyocytes in the native heart are elongated, and hypertrophied by a mechanical load. This indicates the importance of the culture conditions, and the development of bioreactors remains an active area of research [15–17]. The present study introduces an in vitro system for engineered myocardium using a degradable, nanofibrous scaffold made by electrostatic fiber spinning. Electrostatic fiber spinning, or electrospinning, is a process that produces ultrafine fibers in the form of a non-woven mesh through the action of a high electric field [18]. The resulting fiber diameters are in the submicron range, and are unattainable by other fiber spinning techniques. Due to the small fiber diameters, the meshes have a high specific surface area conjectured to be beneficial for cell attachment and proliferation [19–21]. Qualitatively, the topography of the scaffold resembles that of the ECM. The scaffold used in this study is a thin, highly porous, non-woven mesh stretched across a wire ring. The wire ring acts as a passive stretching device to condition and mature the cardiomyocytes. To assess its potential as a cardiac graft, neonatal rat cardiomyocytes were seeded on electrospun polycaprolactone (PCL) meshes, henceforth referred to as cardiac nanofibrous mesh (CNM). After two weeks of in vitro culture, the CNMs were characterized using histology, immunohistochemistry and scanning electron microscopy (SEM).
2. Materials and methods 2.1. Scaffold fabrication PCL with an average molecular weight of 80 kDa (Aldrich, Milwaukee, WI) was dissolved in a 1:1 mixture of chloroform and methanol (Sigma, St. Louis, MO) to obtain a 10 wt% solution. The polymer solution was delivered with a syringe pump (Harvard Apparatus, Holliston, MA) at a flow rate of 0.1 ml/min to a stainless
Fig. 1. Experimental set-up for scaffold fabrication. A high voltage power supply (1) generates a high electric field between a metal capillary (2) and a grounded collector (3). The polymer fluid is delivered to the capillary via tubing at a constant flow rate with a syringe pump (not shown). The instability region is indicated by the shaded area. The distance between the capillary and the collector is 30 cm.
steel capillary (inner diameter=1 mm; Cole-Parmer, Vernon Hills, IL) connected to a high voltage power supply (Gamma High Voltage Research, Ormond Beach, FL). A schematic diagram of the fiber spinner is shown in Fig. 1. At a voltage of 12 kV, a fluid jet was ejected from the capillary, and a non-woven fibrous mesh was collected on nickel-chrome wire rings (ring diameter=15 mm, wire thickness=0.08 mm; McMasterCarr, Atlanta, GA). The scaffolds were stored in a dessicator for several days. For sterilization, the scaffolds were placed in 70% ethanol overnight. To promote cell attachment, the scaffolds were coated with a purified type 1 collagen solution (Cohesion Technologies, Palo Alto, CA). 2.2. Cell source and seeding 3-day old Lewis rats (Charles River Laboratories, Wilmington, MA) were used as donors. Following euthanasia by isoflurane inhalation, the ventricles were harvested and placed in ice-cold phosphate-buffered saline (PBS; Gibco, Grand Island, NY). The ventricles were minced into pieces o1 mm3, and the tissue fragments were digested in a mixed enzyme solution for 10 min in a 37 C water bath. The enzyme solution consisted of 15 ml of 0.25% trypsin (Gibco) and 90 units/ml of collagenase 2 (Worthington, Lakewood, NJ). Cell clumps were dispersed by filtering through a 100 mm cell strainer (Fisher Scientific, Pittsburgh, PA). The cell suspension was centrifuged at 1200 rpm for 8 min. The supernatant was removed, and the remaining cell pellet was resuspended in culture medium. The culture medium consisted of 40% M–199 medium (Gibco), 54% Earle’s Balanced Salt Solution (Sigma) and 6% fetal bovine serum (FBS; Gibco), supplemented
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with 100 U/ml penicillin and streptomycin (Gibco). To increase the fraction of cardiomyocytes, the cells were preplated for 20 min. Non-adherent cells were removed and counted with a cell counter. The cells were seeded on the scaffolds at a concentration of 4–5 million cells/ scaffold and incubated under standard cell culture conditions. Assessment of cell viability by light microscopy and videomicroscopy, and exchange of culture medium were performed daily. 2.3. Histology The surfaces and cross-sections of the CNMs were analyzed after 14 days of in vitro culture to assess the character of the cells and to ensure that the cardiomyocytes had not been overgrown by fibroblasts. Following dehydration according to standard protocols, the CNMs were embedded in paraffin. Plane and cross sections (5 mm) were cut and stained with hematoxyline and eosin (H&E). For cell F-actin and nuclei staining, the CNMs were fixed with 4% paraformaldehyde. The CNMs were stained with a 1:20 dilution of rhodamine-conjugated phalloidin (Sigma) for 2 h and counterstained with 2 mg/ml 40 ,60 -diamidino-2-phenylindole hydrochloride (DAPI; Sigma) for 2 min. 2.4. Immunohistochemistry Immunohistochemical staining for tropomyosin was used to assess the fraction of cardiomyocytes in the CNMs. Following fixation with 4% paraformaldehyde, the CNMs were incubated overnight with a 1:100 dilution of anti-sarcomeric tropomyosin antibody (MF20; Developmental Studies Hybridoma Bank, University of Iowa, Iowa City, IA), and then with a 1:200 dilution of alkaline phosphatase-conjugated antimouse IgG antibody (DAKO, Carpinteria, CA) for 2 h. The specimens were then incubated with avidinbiotin complex reagent and 3,30 -diaminobenzidine (Sigma). Cardiac troponin I was identified on the surface and the cross-sections of the CNMs. For surface assessment, the CNMs were incubated overnight with a 1:100 dilution of anti-cardiac troponin I antibody (Spectral Diagnostics, Whitestone, VA), followed by a 1:200 dilution of alkaline phosphatase-conjugated anti-mouse IgG antibody (DAKO) for 2 h. For cross-sectional analysis, incubation of the primary antibody was followed by a 1:200 dilution of biotin-conjugated antimouse IgG antibody (DAKO) for 2 h and 1:200 dilution of horse-radish peroxidase conjugated streptavidine (DAKO). To examine the presence of gap junctions, the CNMs were fixed with acetone, followed by incubation with a 1:50 dilution of anti-connexin43 antibody (Chemicon, Temecula, CA) overnight. Subsequently, the CNMs
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were incubated with a 1:400 dilution of Texas-red conjugated anti-mouse IgG antibody (Molecular Probes, Eugene, OR) for 2 h. Images were photographed with a digital camera in a Nikon fluorescence microscope. 2.5. Scanning electron microscopy The fixed CNMs were rinsed in 0.1 m PBS and dehydrated in increasing concentrations of ethanol. Dehydrated CNMs were immersed in hexamethyldisilazane (HMDS; Fluka Chemical Corp., Milwaukee, WI) for 15 min and dried overnight in a dessicator. After drying, the samples were mounted on aluminum stubs, sputter-coated with gold-palladium (AuPd) and viewed in a scanning electron microscope (SEM; Philips XL-20, Eindhoven, The Netherlands) with an accelerating voltage of 10–15 kV.
3. Results Fig. 2a shows the gross morphology of the scaffold. The thickness of the scaffold is 10 mm. As shown in Fig. 2b, the scaffold consists of a 3-D mesh of ultrafine fibers with an ECM-like topography. Typical fiber diameters range from 100 nm to 5 mm. The average fiber diameter is 250 nm. Occasional outliers of up to 10 mm are observed. Some of the fibers are fused at intersections and form physical crosslinks. The mesh has an interconnected, open structure with pore sizes well above that of the fibers. Neonatal rat cardiomyocytes were cultured on electrospun PCL scaffolds for 14 days. The cardiomyocytes attached well to the scaffold. Due to the high porosity and low thickness, cells could be sustained throughout the entire thickness of the scaffold. Three days after seeding, the cardiomyocytes started to contract. Contractions were ubiquitous and synchronized. All CNMs displayed continuous contractility for the duration of the experiment. The fibrous mesh is sufficiently soft to allow spontaneous contraction of the cardiomyocytes. The contractions of the CNMs are shown in Movie 1. Fig. 3a shows the actin filament organization on the CNM surfaces. Muscle-like actin is abundantly present. At higher magnifications, cross-striations are also visible as shown in Fig. 3b. The cardiomyocytes stain positively for tropomyosin and show spindle, round, and multiangular morphologies with striations that are indicative of the sarcomeric structures of muscle cells as seen in Figs. 3c. Fig. 3d shows that the cardiomyocytes stain positively for cardiac troponin-I. Diffuse gap junctions within the CNMs are shown in Fig. 3e. Fig. 4a shows the cross-sectional view of the CNMs. The cardiomyocytes have formed a tight arrangement and intercellular contacts throughout the entire mesh.
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Fig. 2. (a) Gross view of culture system. A thin, non-woven fibrous mesh is suspended across a wire ring. The thickness of the mesh is 10 mm. The wire ring acts as a passive load to condition cardiomyocytes during in vitro culture. Scalebar=15 mm. (b) SEM micrograph of mesh. The electrospun fibers have an average diameter of 250 nm, and observed fiber diameters range from 100 nm to 5 mm. The pores are interconnected and are much larger than the fibers. The topography of the non-woven mesh resembles that of an ECM and facilitates cell attachment. Scalebar=200 mm.
Fig. 4b shows that muscle-like actin filaments traverse the entire thickness of the mesh. The CNMs stain positively for cardiac troponin-I in the interior of the scaffolds as seen in Fig. 4c. Fig. 5a shows an SEM micrograph of the surface. The CNM surface is abundantly covered with multilayers of cells. Fig. 5b shows a cross-sectional view. Although the cell density is higher on the surfaces, cells are present through the entire mesh, and non-degraded PCL fibers are also visible.
4. Discussion Here we present an initial description of a new in vitro culture system that shows potential as a myocardial patch. Neonatal rat cardiomyocytes were harvested and seeded on electrospun nanofibrous PCL meshes suspended on wire rings (CNMs). The cardiomyocytes attached well to the scaffolds and were cultured in vitro for 14 days. After 3 days, the cardiomyocytes started contracting, and the wire ring acted as a passive load. The cardiomyocytes penetrated the entire scaffold and stained positively for cardiotypical proteins, i.e., actin, tropomyosin and cardiac troponin-I. In addition, diffuse gap junctions were observed. The stains verify that the cardiomyocytes retained their character and were not overgrown by fibroblasts. Several studies on creating beating bioengineered patches have been reported recently. Studies that used a synthetic polymer scaffold typically found that the stiffness of the scaffold impeded the contractions. More promising results were obtained with ECM-based scaffolds, e.g., collagen or cell sheets. In these studies, good contractions and function were observed. How-
ever, handling of these constructs remains difficult. The work presented here provides an alternative approach to engineered myocardium by relying on a synthetic polymer that provides sufficient stability and low resistance to contractions. A key feature of our approach is the topography and porosity of the mesh. Through the electrospinning process, fibers with diameters smaller than those obtained through conventional fiber spinning processes were obtained. Fiber diameters from conventional fiber spinning operations, e.g., melt spinning or wet spinning, are at least one order of magnitude larger than those obtained in the electrospinning process. The small fiber diameters result in a high specific surface area that is beneficial for cell attachment and proliferation. The end product is a highly porous, non-woven mesh with interconnected pores that can act as a temporary ECM. The variation in fiber diameters is primarily due to fluid instabilities during the process. Electrospun PCL meshes were sufficiently soft so as to not restrict contractions of the cardiomyocytes but also sufficiently stable for handling. Electrospinning is a versatile process and can be applied to a wide range of synthetic or natural polymers. Through judicious choice of polymer and processing conditions, the properties of the scaffold can be tailored for specific applications. Another advantage is the suspension of the mesh across a wire ring that provides tension. The wire acts as a passive load for beating cardiomyocytes and permits contractions at their natural frequency. For successful implantation, the cardiomyocytes must be sufficiently strong and mature. The mechanical load is conjectured to mature the cardiomyocytes in vitro prior to implantation and longer term studies are currently
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Fig. 3. (a) Immunohistochemical staining for F-actin. Actin filaments cover the entire surface of the CNM (100 ). Scalebar=50 mm. (b) At high magnifications, cross-striations can be seen that are indicative of maturing (400 ). Scalebar=50 mm. (c) Immunohistochemical staining for tropomyosin. The surface is covered with tropomyosin-positive cells. Some fiber segments are also visible (100 ). Scalebar=50 mm. (d) Immunohistochemical staining for cardiac troponin-I. The surface is covered with cardiac troponin-I-positive cells. The nuclei are also shown (200 ). Scalebar=50 mm. (e) Immunofluorescent staining for connexin43. Diffuse gap junctions between the cells can be seen (400 ). Scalebar=50 mm.
underway to assess this with the CNMs. This approach also prevents the CNMs from damage due to excessive loading. It is anticipated that the wire ring will improve handling of the CNMs for in vivo applications, i.e., placing the CNM on the heart, securing the mesh with suture and removing the ring.
5. Conclusion We have established a versatile in vitro system for engineered myocardium combining a nanofibrous, biodegradable mesh and a passive stretching device. To fabricate a functional myocardial patch, sufficient
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Fig. 4. (a) Hematoxylin and eosin staining. Cells have attached through the entire mesh (400 ). Scalebar=50 mm. (b) Immunohistochemical staining for F-actin. Actin fibers traverse the entire thickness of the CNM (400 ). Scalebar=50 mm. (c) Immunohistochemical staining for cardiac troponin-I. Troponin-positive cells are found in the interior of the CNM. Scalebar=50 mm.
Fig. 5. (a) SEM micrograph of the CNM surface. The mesh is covered with multilayers of cells. Scalebar=100 mm. (b) SEM micrograph of the CNM cross-section. Cells have attached to the fibers through the entire mesh. Scalebar=50 mm.
cardiomyocyte function at the time of implantation is required. The results presented here indicate that cardiomyocyte sheets can be matured in vitro. In addition to the question of cell source, the key challenge in myocardial tissue engineering is the growth of 3-D structures that contain more than a few layers of
muscle cells. The innermost cells, beyond 100 mm from the surface, are too far from the supply of fresh growth medium to thrive. While an individual CMN is not sufficiently thick to be of clinical relevance, e.g., left ventricle free wall with a thickness >1 cm, current experiments are underway to assess the in vivo
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performance and to create thicker patches by stacking CNMs and incorporating angiogenic factors to induce vascularization. Depending on future progress in stem cell technology and the discovery of appropriate cell sources, autologous CNMs may be used for regeneration of infarcted myocardium and treatment of other cardiac defects. Acknowledgements This work was supported by a grant from the Center for Integration of Medicine and Innovative Technology (CIMIT).
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References [13] [1] Kocher AA, Schuster MD, Szabolcs MJ, Takuma S, Burkhoff D, Wang J, Homma S, Edwards NM, Itescu S. Neovascularization of ischemic myocardium by human bone-marrow-derived angioblasts prevents cardiomyocyte apoptosis, reduces remodeling and improves cardiac function. Nat Med 2001;7:430–6. [2] Penn MS, Francis GS, Ellis SG, Young JB, McCarthy PM, Topol EJ. Autologous cell transplantation for the treatment of damaged myocardium. Prog Cardiovasc Dis 2002;45:21–32. [3] Li RK, Jia ZQ, Weisel RD, Mickle DA, Choi A, Yau TM. Survival and function of bioengineered cardiac grafts. Circulation 1999;100:II63–9. [4] Sakai T, Li RK, Weisel RD, Mickle DA, Kim ET, Jia ZQ, Yau TM. The fate of a tissue-engineered cardiac graft in the right ventricular outflow tract of the rat. J Thorac Cardiovasc Surg 2001;121:932–42. [5] Akhyari P, Fedak PW, Weisel RD, Lee TY, Verma S, Mickle DA, Li RK. Mechanical stretch regimen enhances the formation of bioengineered autologous cardiac muscle grafts. Circulation 2002;106(12 Suppl 1):I137–42. [6] Ozawa T, Mickle DA, Weisel RD, Koyama N, Ozawa S, Li RK. Optimal biomaterial for creation of autologous cardiac grafts. Circulation 2002;106(12 Suppl 1):I176–82. [7] Leor J, Aboulafia-Etzion S, Dar A, Shapiro L, Barbash IM, Battler A, Granot Y, Cohen S. Bioengineered cardiac grafts: A new approach to repair the infarcted myocardium? Circulation 2000;102(19 Suppl 3):III56–61. [8] Eschenhagen T, Fink C, Remmers U, Scholz H, Wattchow J, Weil J, Zimmermann W, Dohmen HH, Schafer H, Bishopric N,
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Biomaterials 25 (2004) 3717–3723
Contractile cardiac grafts using a novel nanofibrous mesh$ M. Shina,*,1, O. Ishiia,b,1, T. Suedab, J.P. Vacantia a
Department of Surgery, Massachusetts General Hospital and Harvard Medical School, Wellman 627, 55 Fruit Street, Boston, MA 02114, USA b First Department of Surgery, Hiroshima University School of Medicine, 1-2-3 Kasumi, Minami-ku, Hiroshima 734-8551, Japan Received 25 September 2003; accepted 21 October 2003
Abstract Cardiomyoctes are terminally differentiated cells and therefore unable to regenerate after infarction. The use of autologous bioengineered cardiac grafts has been suggested to replace infarcted myocardium and enhance cardiac function. Here we report the development of an in vitro system for engineered myocardium. Cardiac nanofibrous meshes (CNM) were developed by culturing cardiomyocytes from neonatal Lewis rats on electrospun, nanofibrous polycaprolactone (PCL) meshes. The mesh had an ECM-like topography and was suspended across a wire ring that acted as a passive load to contracting cardiomyocytes. The cardiomyocytes started beating after 3 days and were cultured in vitro for 14 days. The cardiomyocytes attached well on the PCL meshes and expressed cardiac-specific proteins such as a-myosin heavy chain, connexin43 and cardiac troponin I. The results demonstrate the formation of contractile cardiac grafts in vitro. Using this technique, cardiac grafts can be matured in vitro to obtain sufficient function prior to implantation. It is conjectured that cardiac grafts with clinically relevant dimensions can be obtained by stacking CNMs and inducing vascularization with angiogenic factors. r 2003 Elsevier Ltd. All rights reserved. Keywords: Cardiac tissue engineering; Polycaprolactone; Heart; Electrospinning
1. Introduction The limited ability of myocardium to regenerate after extensive infarction has led to the development of new treatment modalities to regenerate functional myocardium. Autologous cell transplantation into damaged hearts has received significant attention, and two general approaches have been studied. The first approach is to inject cell suspensions directly into the heart. A number of studies have been carried out in animals and humans using a variety of cell types. The results are promising and demonstrate that implanted cells can incorporate themselves into the host hearts [1,2]. The second approach combines cells and biodegradable scaffolds to provide temporal structural support and guide tissue regeneration. To this end, different scaffold materials, cell sources and bioreactors have been assessed to obtain viable and functional tissue. $ Supplementary data associated with this article can be found, in the online version, at doi: 10.1016/j.biomaterials.2003.10.055 *Corresponding author. Tel.: +1-617-724-0832; fax: +1-617-7265057. E-mail address: [email protected] (M. Shin). 1 M. Shin and O. Ishii contributed equally to this study.
0142-9612/$ - see front matter r 2003 Elsevier Ltd. All rights reserved. doi:10.1016/j.biomaterials.2003.10.055
Li et al. produced three-dimensional (3-D) contractile cardiac grafts using gelatin sponges. These grafts were used to replace myocardial scar tissue and right ventricular outflow tract defects in rats [3,4]. It was found that mechanical stretching enhanced the formation of tissue-engineered grafts [5]. In addition, poly(e caprolactone-co-l-lactide) (PCLA) sponge reinforced with knitted poly-l-lactide fabric (PLLA), and polyglycolic acid (PGA) were assessed [6]. Leor et al. reported the formation of bioengineered cardiac grafts with 3-D alginate scaffolds. These grafts stimulated neovascularization and attenuated left ventricular dilatation and heart failure in a rat infarction model [7]. Eschenhagen et al. engineered 3-D heart tissue by gelling a mixture of rat cardiomyocytes and collagen [8]. These constructs were cast in circular shapes and subjected to a cyclical mechanical stress. The morphology, electrophysiology and contractile properties of these grafts were characterized [9]. The engineered heart tissue survived and matured after implantation on uninjured hearts in a rat model [10]. Recently, Okano’s group has developed cell sheet engineering without scaffolds. Using a novel temperature-
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responsive polymer, poly(N-isopropylacrylamide) (PIPAAm), cultured rat cardiomyocytes could be detached as a continuous sheet without destroying cellcell junction proteins and extracellular matrix (ECM) [11]. To create thicker grafts, several cell sheets were layered on top of each other. Electrical and morphological connections between the layers were verified both in vitro and in vivo [12,13]. Despite these advances, several challenges remain. In addition to myocardial cell sourcing, cell survival within the graft is of importance. Typically, the cells are dense in the graft periphery but sparse in interior due to insufficient oxygen and nutrient perfusion [14]. To avoid core ischemia in grafts of clinically relevant dimensions, either new methods to accelerate blood vessel formation or connected layers of thinner grafts are required. Furthermore, cardiomyocytes in the native heart are elongated, and hypertrophied by a mechanical load. This indicates the importance of the culture conditions, and the development of bioreactors remains an active area of research [15–17]. The present study introduces an in vitro system for engineered myocardium using a degradable, nanofibrous scaffold made by electrostatic fiber spinning. Electrostatic fiber spinning, or electrospinning, is a process that produces ultrafine fibers in the form of a non-woven mesh through the action of a high electric field [18]. The resulting fiber diameters are in the submicron range, and are unattainable by other fiber spinning techniques. Due to the small fiber diameters, the meshes have a high specific surface area conjectured to be beneficial for cell attachment and proliferation [19–21]. Qualitatively, the topography of the scaffold resembles that of the ECM. The scaffold used in this study is a thin, highly porous, non-woven mesh stretched across a wire ring. The wire ring acts as a passive stretching device to condition and mature the cardiomyocytes. To assess its potential as a cardiac graft, neonatal rat cardiomyocytes were seeded on electrospun polycaprolactone (PCL) meshes, henceforth referred to as cardiac nanofibrous mesh (CNM). After two weeks of in vitro culture, the CNMs were characterized using histology, immunohistochemistry and scanning electron microscopy (SEM).
2. Materials and methods 2.1. Scaffold fabrication PCL with an average molecular weight of 80 kDa (Aldrich, Milwaukee, WI) was dissolved in a 1:1 mixture of chloroform and methanol (Sigma, St. Louis, MO) to obtain a 10 wt% solution. The polymer solution was delivered with a syringe pump (Harvard Apparatus, Holliston, MA) at a flow rate of 0.1 ml/min to a stainless
Fig. 1. Experimental set-up for scaffold fabrication. A high voltage power supply (1) generates a high electric field between a metal capillary (2) and a grounded collector (3). The polymer fluid is delivered to the capillary via tubing at a constant flow rate with a syringe pump (not shown). The instability region is indicated by the shaded area. The distance between the capillary and the collector is 30 cm.
steel capillary (inner diameter=1 mm; Cole-Parmer, Vernon Hills, IL) connected to a high voltage power supply (Gamma High Voltage Research, Ormond Beach, FL). A schematic diagram of the fiber spinner is shown in Fig. 1. At a voltage of 12 kV, a fluid jet was ejected from the capillary, and a non-woven fibrous mesh was collected on nickel-chrome wire rings (ring diameter=15 mm, wire thickness=0.08 mm; McMasterCarr, Atlanta, GA). The scaffolds were stored in a dessicator for several days. For sterilization, the scaffolds were placed in 70% ethanol overnight. To promote cell attachment, the scaffolds were coated with a purified type 1 collagen solution (Cohesion Technologies, Palo Alto, CA). 2.2. Cell source and seeding 3-day old Lewis rats (Charles River Laboratories, Wilmington, MA) were used as donors. Following euthanasia by isoflurane inhalation, the ventricles were harvested and placed in ice-cold phosphate-buffered saline (PBS; Gibco, Grand Island, NY). The ventricles were minced into pieces o1 mm3, and the tissue fragments were digested in a mixed enzyme solution for 10 min in a 37 C water bath. The enzyme solution consisted of 15 ml of 0.25% trypsin (Gibco) and 90 units/ml of collagenase 2 (Worthington, Lakewood, NJ). Cell clumps were dispersed by filtering through a 100 mm cell strainer (Fisher Scientific, Pittsburgh, PA). The cell suspension was centrifuged at 1200 rpm for 8 min. The supernatant was removed, and the remaining cell pellet was resuspended in culture medium. The culture medium consisted of 40% M–199 medium (Gibco), 54% Earle’s Balanced Salt Solution (Sigma) and 6% fetal bovine serum (FBS; Gibco), supplemented
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with 100 U/ml penicillin and streptomycin (Gibco). To increase the fraction of cardiomyocytes, the cells were preplated for 20 min. Non-adherent cells were removed and counted with a cell counter. The cells were seeded on the scaffolds at a concentration of 4–5 million cells/ scaffold and incubated under standard cell culture conditions. Assessment of cell viability by light microscopy and videomicroscopy, and exchange of culture medium were performed daily. 2.3. Histology The surfaces and cross-sections of the CNMs were analyzed after 14 days of in vitro culture to assess the character of the cells and to ensure that the cardiomyocytes had not been overgrown by fibroblasts. Following dehydration according to standard protocols, the CNMs were embedded in paraffin. Plane and cross sections (5 mm) were cut and stained with hematoxyline and eosin (H&E). For cell F-actin and nuclei staining, the CNMs were fixed with 4% paraformaldehyde. The CNMs were stained with a 1:20 dilution of rhodamine-conjugated phalloidin (Sigma) for 2 h and counterstained with 2 mg/ml 40 ,60 -diamidino-2-phenylindole hydrochloride (DAPI; Sigma) for 2 min. 2.4. Immunohistochemistry Immunohistochemical staining for tropomyosin was used to assess the fraction of cardiomyocytes in the CNMs. Following fixation with 4% paraformaldehyde, the CNMs were incubated overnight with a 1:100 dilution of anti-sarcomeric tropomyosin antibody (MF20; Developmental Studies Hybridoma Bank, University of Iowa, Iowa City, IA), and then with a 1:200 dilution of alkaline phosphatase-conjugated antimouse IgG antibody (DAKO, Carpinteria, CA) for 2 h. The specimens were then incubated with avidinbiotin complex reagent and 3,30 -diaminobenzidine (Sigma). Cardiac troponin I was identified on the surface and the cross-sections of the CNMs. For surface assessment, the CNMs were incubated overnight with a 1:100 dilution of anti-cardiac troponin I antibody (Spectral Diagnostics, Whitestone, VA), followed by a 1:200 dilution of alkaline phosphatase-conjugated anti-mouse IgG antibody (DAKO) for 2 h. For cross-sectional analysis, incubation of the primary antibody was followed by a 1:200 dilution of biotin-conjugated antimouse IgG antibody (DAKO) for 2 h and 1:200 dilution of horse-radish peroxidase conjugated streptavidine (DAKO). To examine the presence of gap junctions, the CNMs were fixed with acetone, followed by incubation with a 1:50 dilution of anti-connexin43 antibody (Chemicon, Temecula, CA) overnight. Subsequently, the CNMs
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were incubated with a 1:400 dilution of Texas-red conjugated anti-mouse IgG antibody (Molecular Probes, Eugene, OR) for 2 h. Images were photographed with a digital camera in a Nikon fluorescence microscope. 2.5. Scanning electron microscopy The fixed CNMs were rinsed in 0.1 m PBS and dehydrated in increasing concentrations of ethanol. Dehydrated CNMs were immersed in hexamethyldisilazane (HMDS; Fluka Chemical Corp., Milwaukee, WI) for 15 min and dried overnight in a dessicator. After drying, the samples were mounted on aluminum stubs, sputter-coated with gold-palladium (AuPd) and viewed in a scanning electron microscope (SEM; Philips XL-20, Eindhoven, The Netherlands) with an accelerating voltage of 10–15 kV.
3. Results Fig. 2a shows the gross morphology of the scaffold. The thickness of the scaffold is 10 mm. As shown in Fig. 2b, the scaffold consists of a 3-D mesh of ultrafine fibers with an ECM-like topography. Typical fiber diameters range from 100 nm to 5 mm. The average fiber diameter is 250 nm. Occasional outliers of up to 10 mm are observed. Some of the fibers are fused at intersections and form physical crosslinks. The mesh has an interconnected, open structure with pore sizes well above that of the fibers. Neonatal rat cardiomyocytes were cultured on electrospun PCL scaffolds for 14 days. The cardiomyocytes attached well to the scaffold. Due to the high porosity and low thickness, cells could be sustained throughout the entire thickness of the scaffold. Three days after seeding, the cardiomyocytes started to contract. Contractions were ubiquitous and synchronized. All CNMs displayed continuous contractility for the duration of the experiment. The fibrous mesh is sufficiently soft to allow spontaneous contraction of the cardiomyocytes. The contractions of the CNMs are shown in Movie 1. Fig. 3a shows the actin filament organization on the CNM surfaces. Muscle-like actin is abundantly present. At higher magnifications, cross-striations are also visible as shown in Fig. 3b. The cardiomyocytes stain positively for tropomyosin and show spindle, round, and multiangular morphologies with striations that are indicative of the sarcomeric structures of muscle cells as seen in Figs. 3c. Fig. 3d shows that the cardiomyocytes stain positively for cardiac troponin-I. Diffuse gap junctions within the CNMs are shown in Fig. 3e. Fig. 4a shows the cross-sectional view of the CNMs. The cardiomyocytes have formed a tight arrangement and intercellular contacts throughout the entire mesh.
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Fig. 2. (a) Gross view of culture system. A thin, non-woven fibrous mesh is suspended across a wire ring. The thickness of the mesh is 10 mm. The wire ring acts as a passive load to condition cardiomyocytes during in vitro culture. Scalebar=15 mm. (b) SEM micrograph of mesh. The electrospun fibers have an average diameter of 250 nm, and observed fiber diameters range from 100 nm to 5 mm. The pores are interconnected and are much larger than the fibers. The topography of the non-woven mesh resembles that of an ECM and facilitates cell attachment. Scalebar=200 mm.
Fig. 4b shows that muscle-like actin filaments traverse the entire thickness of the mesh. The CNMs stain positively for cardiac troponin-I in the interior of the scaffolds as seen in Fig. 4c. Fig. 5a shows an SEM micrograph of the surface. The CNM surface is abundantly covered with multilayers of cells. Fig. 5b shows a cross-sectional view. Although the cell density is higher on the surfaces, cells are present through the entire mesh, and non-degraded PCL fibers are also visible.
4. Discussion Here we present an initial description of a new in vitro culture system that shows potential as a myocardial patch. Neonatal rat cardiomyocytes were harvested and seeded on electrospun nanofibrous PCL meshes suspended on wire rings (CNMs). The cardiomyocytes attached well to the scaffolds and were cultured in vitro for 14 days. After 3 days, the cardiomyocytes started contracting, and the wire ring acted as a passive load. The cardiomyocytes penetrated the entire scaffold and stained positively for cardiotypical proteins, i.e., actin, tropomyosin and cardiac troponin-I. In addition, diffuse gap junctions were observed. The stains verify that the cardiomyocytes retained their character and were not overgrown by fibroblasts. Several studies on creating beating bioengineered patches have been reported recently. Studies that used a synthetic polymer scaffold typically found that the stiffness of the scaffold impeded the contractions. More promising results were obtained with ECM-based scaffolds, e.g., collagen or cell sheets. In these studies, good contractions and function were observed. How-
ever, handling of these constructs remains difficult. The work presented here provides an alternative approach to engineered myocardium by relying on a synthetic polymer that provides sufficient stability and low resistance to contractions. A key feature of our approach is the topography and porosity of the mesh. Through the electrospinning process, fibers with diameters smaller than those obtained through conventional fiber spinning processes were obtained. Fiber diameters from conventional fiber spinning operations, e.g., melt spinning or wet spinning, are at least one order of magnitude larger than those obtained in the electrospinning process. The small fiber diameters result in a high specific surface area that is beneficial for cell attachment and proliferation. The end product is a highly porous, non-woven mesh with interconnected pores that can act as a temporary ECM. The variation in fiber diameters is primarily due to fluid instabilities during the process. Electrospun PCL meshes were sufficiently soft so as to not restrict contractions of the cardiomyocytes but also sufficiently stable for handling. Electrospinning is a versatile process and can be applied to a wide range of synthetic or natural polymers. Through judicious choice of polymer and processing conditions, the properties of the scaffold can be tailored for specific applications. Another advantage is the suspension of the mesh across a wire ring that provides tension. The wire acts as a passive load for beating cardiomyocytes and permits contractions at their natural frequency. For successful implantation, the cardiomyocytes must be sufficiently strong and mature. The mechanical load is conjectured to mature the cardiomyocytes in vitro prior to implantation and longer term studies are currently
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Fig. 3. (a) Immunohistochemical staining for F-actin. Actin filaments cover the entire surface of the CNM (100 ). Scalebar=50 mm. (b) At high magnifications, cross-striations can be seen that are indicative of maturing (400 ). Scalebar=50 mm. (c) Immunohistochemical staining for tropomyosin. The surface is covered with tropomyosin-positive cells. Some fiber segments are also visible (100 ). Scalebar=50 mm. (d) Immunohistochemical staining for cardiac troponin-I. The surface is covered with cardiac troponin-I-positive cells. The nuclei are also shown (200 ). Scalebar=50 mm. (e) Immunofluorescent staining for connexin43. Diffuse gap junctions between the cells can be seen (400 ). Scalebar=50 mm.
underway to assess this with the CNMs. This approach also prevents the CNMs from damage due to excessive loading. It is anticipated that the wire ring will improve handling of the CNMs for in vivo applications, i.e., placing the CNM on the heart, securing the mesh with suture and removing the ring.
5. Conclusion We have established a versatile in vitro system for engineered myocardium combining a nanofibrous, biodegradable mesh and a passive stretching device. To fabricate a functional myocardial patch, sufficient
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Fig. 4. (a) Hematoxylin and eosin staining. Cells have attached through the entire mesh (400 ). Scalebar=50 mm. (b) Immunohistochemical staining for F-actin. Actin fibers traverse the entire thickness of the CNM (400 ). Scalebar=50 mm. (c) Immunohistochemical staining for cardiac troponin-I. Troponin-positive cells are found in the interior of the CNM. Scalebar=50 mm.
Fig. 5. (a) SEM micrograph of the CNM surface. The mesh is covered with multilayers of cells. Scalebar=100 mm. (b) SEM micrograph of the CNM cross-section. Cells have attached to the fibers through the entire mesh. Scalebar=50 mm.
cardiomyocyte function at the time of implantation is required. The results presented here indicate that cardiomyocyte sheets can be matured in vitro. In addition to the question of cell source, the key challenge in myocardial tissue engineering is the growth of 3-D structures that contain more than a few layers of
muscle cells. The innermost cells, beyond 100 mm from the surface, are too far from the supply of fresh growth medium to thrive. While an individual CMN is not sufficiently thick to be of clinical relevance, e.g., left ventricle free wall with a thickness >1 cm, current experiments are underway to assess the in vivo
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performance and to create thicker patches by stacking CNMs and incorporating angiogenic factors to induce vascularization. Depending on future progress in stem cell technology and the discovery of appropriate cell sources, autologous CNMs may be used for regeneration of infarcted myocardium and treatment of other cardiac defects. Acknowledgements This work was supported by a grant from the Center for Integration of Medicine and Innovative Technology (CIMIT).
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