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Ann Thorac Surg 2001;72:577-591
© 2001 The Society of Thoracic Surgeons

CCCETS basic science lecture

Tissue engineering: a 21st century solution to surgical reconstruction

Julie R. Fuchs, MDa, Boris A. Nasseri, MDa, Joseph P. Vacanti, MDa

a Massachusetts General Hospital and Harvard Medical School, Boston, Massachusetts, USA

Address reprint requests to Dr Vacanti, Department of Surgery, Massachusetts General Hospital, Warren 1157, 55 Fruit St, Boston, MA 02114
e-mail: jvacanti{at}partners.org

Coordinating Committee for Continuing Education in Thoracic Surgery (CCCETS) Basic Science Lecture, presented at the Thirty-fourth Postgraduate Program of The Society of Thoracic Surgeons, New Orleans, LA, Jan 28, 2001.


   Abstract
Top
 Abstract
Introduction
 Heart valves
 Blood vessels
 Myocardium
 Esophagus
 Trachea
 Future directions
 References
 
Tissue engineering has emerged as a rapidly expanding approach to address the organ shortage problem. It is an "interdisciplinary field that applies the principles and methods of engineering and the life sciences toward the development of biological substitutes that can restore, maintain, or improve tissue function." Much progress has been made in the tissue engineering of structures relevant to cardiothoracic surgery, including heart valves, blood vessels, myocardium, esophagus, and trachea.


   Introduction
Top
 Abstract
 Introduction
Heart valves
 Blood vessels
 Myocardium
 Esophagus
 Trachea
 Future directions
 References
 
End-stage organ failure or tissue loss is one of the most devastating and costly problems in medicine. Over 8 million surgical procedures are estimated to be performed each year to treat these disorders in the United States alone [1] (Table 1). This incurs a tremendous health care cost of more than $400 billion annually and millions of lost work days [1]. Over the last 50 years, transplantation of a wide variety of tissues, reconstructive surgical techniques, and replacement with mechanical devices have significantly improved patient outcomes. Murray and colleagues performed the first successful organ transplant in 1954 [2]. Since that historic accomplishment, the field of transplantation has evolved to include kidney, liver, split liver, pancreas, heart, lung, and small intestine at hundreds of transplant centers throughout the United States. In 1967, Barnard performed the first heart transplant for congestive heart failure [3]. These strides have been made possible because of the advances in transplantation biology and immunology leading to the development of a variety of immunosuppressive agents.


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  		Table 1. Cardiac Procedures Performed in the United States Annually

 
Unfortunately, organ and tissue transplantation are imperfect solutions because they are limited by a number of factors. Worsening donor shortages result in a discrepancy between the number of patients needing transplants and available organs. In 1989, 19,095 patients were awaiting transplantation. By February 2001, that number had increased to 74,800 [4] (Table 2). Additionally, transplantation recipients must follow lifelong immunosuppression regimens with their increased risks of infection, tumor development, and unwanted side effects. Surgical reconstruction also suffers from a lack of available donor tissue and donor site morbidity. Replacement with mechanical devices or artificial organs is limited by an increased risk of infection and thromboembolism and finite durability. Because of the above shortcomings, the field of tissue engineering and selective cell transplantation was born as a means to replace diseased tissue with living tissue that is "designed and constructed to meet the needs of each individual patient" [5].


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  		Table 2. The UNOS National Patient Waiting List for Organ Transplant (February, 2001)

 
Tissue engineering is "an interdisciplinary field that applies the principles and methods of engineering and the life sciences toward the development of biological substitutes that restore, maintain, or improve tissue function" [1, 5]. In 1933, this general concept was used when investigators encased mouse tumor cells in a polymer membrane and placed them in the abdominal cavity of a pig [6]. The cells were not destroyed by the immune system. In 1975, Chick and colleagues were the first to place pancreatic islet cells in semipermeable membranes to improve glucose control in diabetes [7]. Others created skin substitutes consisting of fibroblast cells seeded onto collagen scaffolds, which are currently used in clinical practice in the context of burns and diabetic ulcers [8]. In 2001, tissue engineering efforts are being undertaken for every type of tissue and organ (Fig 1). A recent survey revealed that more than 2,500 scientists were employed in these efforts with an annual expenditure of $450 million [9].



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  		Fig 1. The tissue engineering process. Reprinted with permission from Vacanti JP, Langer R. Tissue engineering: the design and fabrication of living replacement devices for surgical reconstruction and transplantation. Lancet 1999;354:SI32–4. © by The Lancet Ltd, 1999.

 
The goal of tissue engineering is to "restore function through the delivery of living elements which become integrated into the patient" [5]. This goal, which should lead to the fabrication of new, physiologic, functioning tissue, must involve the combined efforts of cell biologists, engineers, material scientists, mathematicians, geneticists, and clinicians to be successful. The three major approaches include guided tissue regeneration using engineered matrices alone, the injection of allogenic or xenogenic cells alone, or the use of cells placed on or within matrices [1, 5]. The latter two methods are the most common. Using isolated cells or cell substitutes avoids potential surgical complications and allows cell manipulation before injection/infusion (such as gene therapy), but has the drawbacks of possible rejection or loss of function [1]. The use of cell-matrix constructs, the most common method in tissue engineering, involves either an open or a closed system. An open system begins with the in vitro culture of isolated cells. The cells are then seeded onto a scaffold, either synthetic or natural. After appropriate cultivation time, the cell-matrix construct is implanted into the host (Fig 1). The matrix functions to guide the development of the new tissue and provides structural support. This approach is based on a number of biological observations. Firstly, all tissues undergo constant remodeling. Under appropriate environmental conditions, dissociated cells often reform their native structures [2]. Furthermore, normal parenchymal cells are anchorage dependent; they also require three-dimensional structure and an extracellular matrix to guide their growth. Lastly, the volume of tissue that can be implanted and survive is limited by the diffusion distance for nutritional molecules, gas exchange, and waste removal [2]. In a closed system, the cells are isolated from the body by a permeable membrane allowing exchange of nutrients and waste but protecting the cells from the immune response [1].

Polymer scaffolds can be produced from natural or synthetic materials. Natural materials may closely mimic the native cellular environment as they are often extracellular matrix components and include collagen, hydroxyapatite, Matrigel (Collaborative Biomedical, Madison, WI), and alginate, among others [5]. Synthetic materials have the advantage of being able to better control material properties such as strength, degradation time, porosity, and microstructure [5]. Cell attachment can be improved by modifying the polymer chemically or by coating it. Growth factors can also be incorporated into the matrix. Defined shapes and sizes can be fabricated readily and reproducibly. Ideally these polymers must be biocompatible and bioabsorbable, nonimmunogenic, support cell growth, and be able to induce angiogenesis to supply the newly formed tissue [5]. The most widely used polymers in tissue engineering fulfilling these criteria include the poly ({alpha}-hydroxy acids) of the aliphatic polyesters, which include polyglycolic acid (PGA), polylactic acid (PLA), and the copolymers (PLGA) of these materials [10]. Polyglycolic acid was first developed as Dexon (Davis and Geck, Danbury, CT), a synthetic absorbable suture. These polymers can be produced in a number of shapes and physical forms, including meshes (Fig 2), sponges, and films with a biodegradation rate that can be changed through copolymer proportions [10–12]. Polyhydroxyalkanoates (PHAs) are a class of natural polymers with thermoplastic properties. They are biocompatible, resorbable, extremely flexible, and induce only a minimal inflammatory response [13]. Because of these properties as well as their tensile strength, they are currently being used in the tissue engineering of heart valves and blood vessels. Hydrogels such as pluronic F-127 (a copolymer of 70% ethylene oxide and 30% propylene oxide) are also being used as cell delivery matrices.



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  		Fig 2. Scanning electron microscopy of a fiber-based polymer scaffold consisting of a nonwoven mesh of polyglycolic acid (PGA) fibers.

 
Cells used in tissue engineering can be derived from numerous sources, including primary tissues and cell lines. Cells may be allogenic, xenogenic, syngeneic, or autologous. Ideally, the cells should be nonimmunogenic, highly proliferative, easy to harvest, and have the ability to differentiate into a variety of cell types with specialized functions [10]. Autologous cells that can be easily isolated and expanded in vitro have been used in many tissue-engineering investigations, including skeletal muscle satellite cells, endothelial cells, and chondrocytes. However, certain cell types such as cardiomyocytes and hepatocytes proliferate poorly in culture or not at all. For this reason, more rapidly proliferating cell sources such as fetal, neonatal, genetically manipulated, or stem cells have been used and are an active area of investigation [5].

Over the last 15 years, our laboratory and others have applied the above tenets of tissue engineering to a wide variety of organs and tissues, many of which involve cardiac and thoracic structures. The structures currently under active investigation include heart valves, blood vessels, myocardium, esophagus, and trachea.


   Heart valves
Top
 Abstract
 Introduction
 Heart valves
Blood vessels
 Myocardium
 Esophagus
 Trachea
 Future directions
 References
 
The most common surgical therapy for end-stage valvular heart disease consists of valve replacement, with more than 60,000 implantations in the United States and 170,000 worldwide [14]. Additionally, each year in the United States approximately 15,000 infants are born with congenital cardiac malformations, many of which involve one or more cardiac valves [15]. Valve replacement surgery is effective and has been shown to significantly alter the course of valvular disease, but has numerous associated shortcomings. There are three main valvular replacements under current use, which include entirely prosthetic valves, bioprosthetic valves (combining heterograft tissue with prosthetic material), and homograft valves originating from cadaveric donors [15]. Each of these valves has limitations and associated complications. Mechanical valves are associated with a significant risk of thromboembolism and require lifelong anticoagulation. These valves are also susceptible to infection that is extremely difficult to overcome without removal of the device [16]. Bioprosthetic valves do not require anticoagulation, but they are far less durable and are subject to progressive tissue deterioration [17]. Homograft valves are believed to contain some viable cells and can therefore elicit an immunologic reaction because they are allografts [15]. None of the available valve replacements has the capacity for growth, which poses a future problem for pediatric patients, often necessitating reoperation [15].

Because of the above shortcomings associated with valvular replacements, efforts have been made by a number of investigators to use the techniques of tissue engineering to create a living valve substitute. The desirable characteristics of a heart valve substitute, as originally described by Harken and colleagues [18], must be taken into consideration when planning investigations. The ideal substitute would have prompt and complete closure and be nonobstructive, nonthrombogenic, infection resistant, chemically inert and nonhemolytic, durable, and easily and permanently inserted [18].

There are two basic approaches to valve substitution that have been pursued over the last 6 years. The first involves decellularizing xenogenic valves and then adding cells or allowing ingrowth of host cells to occur. Bader and colleagues [19] removed cells from porcine aortic valves by detergent cell extraction using Triton. They were able to show almost complete removal of the original cells while grossly maintaining the matrix. Then endothelial cells were isolated from human saphenous vein, expanded in vitro, and seeded onto the acellular matrix. By this method, the authors hoped to eliminate the immune-stimulating potential of xenogenic cells and avoid the calcification associated with glutaraldehyde. Additionally, they wished to further investigate whether the physiologic matrix proposed by their methodology will provide the microenvironment needed to sustain normal cellular differentiation in vivo in the future [19]. Another group developed a different decellularization process. After decellularization, a composite heart valve bioprosthesis was prepared by the "SynerGraft" process. The valve consisted of three size- and symmetry-matched decellularized porcine noncoronary cusp units, each with an aortic leaflet, anterior mitral leaflet, and aortic conduit sutured together. The valve composites were implanted as pulmonary valve replacements in sheep. All valves were hemodynamically functional at explant and histologic examination revealed in-growth of host fibroblastoid cells and no calcification [20].

The second approach involves traditional tissue engineering. The concept as it applies to heart valves is to transplant autologous cells onto a biodegradable scaffold in the shape of a heart valve [21]. After attachment, the cells form extracellular matrix material as the polymer simultaneously degrades. Theoretically, one can create a viable, autologous valve that can be transplanted into the same patient from whom the cells were harvested [21]. The tissue-engineered valve would have the potential to grow, remodel, and avoid infectious and thrombogenic complications. Mayer and associates have pioneered this approach over the last 6 years. The first study involved isolating mixed cell populations of endothelial cells and myofibroblasts from ovine arteries. Endothelial cells were labeled with an acetylated low-density lipoprotein marker and separated from myofibroblasts using fluorescence-activated cell sorting [22]. Polyglycolic acid scaffolds were seeded with myofibroblasts and then were subsequently seeded with endothelial cells. The constructs were implanted in place of the native right posterior leaflet of the pulmonary valve in sheep. In this preliminary study, postoperative echocardiography showed no stenosis and only trivial pulmonary regurgitation in autografts, but moderate regurgitation in allografts. Histology revealed appropriate cellular architecture and matrix that was less developed than native valves [22].

After this study, Mayer and colleagues further elaborated on their initial findings. They again implanted their autologous cell-polymer constructs as described above and the polymer leaflet alone without cells as a control [23]. They found that the acellular polymer leaflets were completely degraded at 8 weeks, whereas the tissue-engineered valve leaflet persisted. Collagen content increased progressively, and immunochemical staining demonstrated elastin in the matrix and factor VIII (endothelial cell specific) on the surface of the leaflet. Cell labeling revealed that the original cells persisted on the leaflet [23]. Echocardiography again demonstrated valve function without evidence of stenosis and only trivial regurgitation [24]. The group then investigated the possibility of using a different cell source to create leaflet constructs, namely, dermal fibroblasts. They found that leaflets created with arterial wall cells had a higher growth index than those containing dermal fibroblasts [25]. Elastin was more prominent in the leaflets with arterial cells, and those with dermal cells were thicker and contracted. There was no difference between the leaflets in terms of biomechanical testing or collagen content [25].

Zund and colleagues [26] further investigated the in vitro aspects of valve construction. They seeded human dermal fibroblasts and bovine aortic endothelial cells or sheep myofibroblasts and endothelial cells on sheets of PGA mesh and grew the constructs in vitro. Histologic analysis revealed an inner fibroblastic core and an outer endothelial monolayer positive for factor VIII [26]. They also used an MTT (3-[4,5-dimethylthiazol-2-yl]-2,5-diphenyltertra-zolium bromide) assay to determine optimal intervals for seeding human aortic myofibroblasts on PGA scaffolds [27] and further characterized the use of fluorescence-activated cell sorting as a method to gain pure autologous cell lines of endothelial cells and myofibroblasts from human mixed cells of the ascending aorta [28].

Further work by Mayer and colleagues has focused on different polymer scaffolds for use in tissue engineering of heart valves. In a recent study, three materials, PGA, PHA, and poly-4-hydroxybutyrate (P4HB) were compared. Trileaflet valves were constructed from each polymer and tested in a pulsatile flow bioreactor [29]. Additionally, cell attachment to each polymer was investigated by seeding the polymers with ovine vascular cells and culturing the constructs for 8 days. They found that it was not possible to create a trileaflet valve using PGA because of the material’s stiffness. Polyglycolic acid had the greatest cell attachment and collagen content, but PHA and P4HB also demonstrated a considerable amount of cell attachment and collagen development. The latter materials are thermoplastic, allowing them to be molded into the shape of a trileaflet valve [29].

Most recently, Mayer and colleagues have created a tissue-engineered trileaflet valve from porous PHA and vascular cells and have implanted this construct into the pulmonary position. The valves showed minimal regurgitation and no thrombus formation. Histology revealed laminated fibrous tissue and extracellular matrix [21]. Collagen content increased over time as well, indicating continued remodeling. Another study was performed using composite scaffolds of PGA and P4HB with similar results [30]. Lastly, after seeding, the composite scaffolds were grown for 14 days in a pulse duplicator in vitro system under gradually increasing flow and pressure conditions (Fig 3). The valve constructs were then implanted into the same lambs and explanted at different time points. The autologous tissue-engineered valves functioned up to 5 months and resembled normal heart valves histologically, mechanically, and in terms of extracellular matrix formation [17].



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  		Fig 3. Tissue-engineered heart valve after 14 days of pulsatile flow in a bioreactor. Reprinted with permission from Hoerstrup SP, Sodian R, Daebritz S, et al. Functional living trileaflet heart valves grown in vitro. Circulation 2000;102(Suppl III):III44–9. © 2000 American Heart Association, Inc.

 
Much progress has been made in the tissue engineering of heart valves, but further work is necessary. The optimal cell source, scaffold design, and in vitro conditions are still being investigated, and the ultimate goal may be to create a valve able to withstand the environment of higher hemodynamic stress in the systemic circulation [15, 17].


   Blood vessels
Top
 Abstract
 Introduction
 Heart valves
 Blood vessels
Myocardium
 Esophagus
 Trachea
 Future directions
 References
 
Atherosclerotic vascular disease, including peripheral vascular and coronary artery disease, is the major cause of mortality and morbidity in the United States, Europe, and other western nations [31, 32]. Current surgical therapy for diseased vessels less than 6 mm in diameter involves bypass grafting with autologous arteries or veins [31], and coronary artery bypass grafting procedures are performed approximately 600,000 times annually in the United States alone [33]. Although common surgical practice, vascular grafting has significant limitations and complications. Arterial conduits have restricted dimensions and are limited in supply. Venous conduits lack vasomotor tone and may have varicose degenerative alterations that can lead to aneurysm formation in the higher pressure arterial circulation [32, 33]. Allografts are problematic because of a high rate of rejection. Synthetic materials are excessively thrombotic when used to bypass arteries less than 6 mm in diameter, with thrombosis rates higher than 40% after 6 months [31, 33]. As a result of these complications, the need for a tissue-engineered vessel of small caliber composed of biological materials and autologous cells has arisen and has been an area of active investigation for more than 15 years [33].

Small vessel vascular tissue engineering can loosely be categorized into three groups. These include using nonbiodegradable grafts seeded with cells, using natural materials (such as collagen) as grafts with or without cells, and using biodegradable polymer matrices seeded with cells. Hybrid vascular grafts, consisting of synthetic material such as Dacron (C.R. Bard, Haverhill, PA) or polytetrafluoroethylene seeded with cultured endothelial cells before implantation, have been investigated by numerous groups [34, 35]. In 1978, the first successful isolation of endothelial cells from segments of vein and their subsequent transplantation onto synthetic vascular grafts was reported [36]. Miwa and Matsuda [34] created a graft of polyurethane, an artificial basement membrane composed of a complex gel of type I collagen and dermatan sulfate, and an autogenous endothelial cell monolayer. They implanted these grafts into the carotid arteries of dogs without anticoagulation and noted an overall patency rate of 75% [34].

Weinberg and Bell were the first to construct a tissue-engineered blood vessel composed of natural materials [37]. The media of the vessel consisted of smooth muscle cells and collagen gel. The adventitia consisted of fibroblasts, and after 2 weeks in culture, endothelial cells were added to the lumenal surface. Electron microscopy revealed endothelial cells lining the lumen, and staining revealed the presence of von Willebrand factor. These constructs were unable to attain adequate burst strengths for in vivo applications, despite their reinforcement with Dacron mesh [37]. L’Heureux and colleagues [33, 38] improved on the mechanical strength of these constructs by alterations in culture conditions. Human vascular smooth muscle cells were cultured with ascorbic acid and produced a cellular sheet, which was placed around a tubular support to produce the media of the vessel. A sheet of fibroblasts was then wrapped around the media to serve as the adventitia, and endothelial cells were seeded in the lumen after a period of culturing [38]. The tissue-engineered vessel exhibited a well-defined, three-layered organization as well as extracellular matrix proteins such as elastin on histologic analysis [38]. The vessel construct had a burst strength of more than 2,000 mm Hg, which is comparable to human vessels. These structures were implanted into the femoral arteries of mongrel dogs and remained patent in 50% at 1 week [38].

Campbell and associates [32] tested the hypothesis that cells from nontraditional sources can be used to create new grafts and achieve secondary function once in the arterial environment. They placed silastic tubing into the peritoneal cavity of rats and rabbits. After 2 weeks, the tubing was covered by multiple layers of myofibroblasts, collagen matrix, and a single layer of mesothelial cells. The tissue was removed from the tubing and everted to place the mesothelial layer within the lumen. The tube of tissue was then grafted into the carotid artery or abdominal aorta of the same animal and remained patent for at least 4 months [32]. This novel approach produced grafts that also responded to contractile agonists in a similar fashion to native vessels [32]. Teebken and colleagues [39] created decellularized matrix tubes by enzymatic cell extraction of porcine aortas and then seeded these natural grafts with human endothelial cells and myofibroblasts. The grafts were exposed to pulsatile flow conditions, and on histologic analysis, they resembled native vessels and had an intact endothelial cell monolayer [39]. Huynh and colleagues [40] constructed a 4-mm diameter graft from small intestinal submucosa and type I bovine collagen. Small intestinal submucosa is a biomaterial, composed primarily of type I collagen, that has shown good patency as a large diameter graft in the canine aorta. Their results revealed excellent patency up to 3 months; histologically, the grafts were remodeled into cellularized vessels that responded appropriately to vasoactive agents such as norepinephrine, serotonin, and bradykinin [40].

Niklason and colleagues [31] pioneered the approach of combining cells with biodegradable polymer. They seeded smooth muscle cells from bovine aorta onto tubular PGA polymer, and cultured the construct in a pulsatile flow bioreactor for 8 weeks [31]. Endothelial cells were then added to the constructs. Histology revealed elastin and collagen fibers, and thicker grafts when pulsed versus nonpulsed controls (Fig 4). A lumenal endothelial layer was present as shown by scanning electron microscopy and staining for platelet endothelial cell adhesion molecule (PECAM) antigen (Fig 4). Pulsed constructs had rupture strengths of more than 2,000 mm Hg, adequate suture retention strength, and collagen content of 50%. These tissue-engineered arteries were implanted into Yucatan pigs with patency demonstrated up to 4 weeks [31].



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  		Fig 4. Photomicrographs of engineered vessels. (A) Pulsed vessel cultured for 8 weeks. (Verhoff’s stain for elastin; x20 before 33% reduction.) (B) Pulsed vessel cultured for 8 weeks. #Cellular region, *polymer region. (Masson’s trichrome stain for collagen; x100 before 33% reduction.) (C) Nonpulsed vessel cultured for 8 weeks. (Verhoff’s stain; x20 before 33% reduction.) (D) Nonpulsed vessel cultured for 8 weeks. (Masson’s trichrome stain; x100 before 33% reduction.) (E) Pulsed vessel without medium supplementation. (Verhoff’s stain; x20 before 33% reduction.) (F) Pulsed vessel without medium supplementation. (Masson’s trichrome stain; x100 before 33% reduction.) (G) Scanning electron microscopy of the endothelial cell layer in an engineered vessel (Scale bar = 10 µm.) (H) Immunoperoxidase staining for platelet endothelial cell adhesion molecule (PECAM) antigen shows an endothelial monolayer on the vessel lumen. (x1,000 before 33% reduction.) Reprinted with permission from Niklason LE, Gao J, Abbott WM, et al. Functional arteries grown in vitro. Science 1999;284:489–93. © 1999 American Association for the Advancement of Science.

 
Ongoing work in our laboratory in this area involves (1) investigation of the ability of different polymer scaffolds to support attachment and growth of vascular endothelial and smooth muscle cells, (2) comparison of different dynamic seeding methods in the construction of the vascular conduit, (3) investigation of various cell-marking techniques of endothelial and smooth muscle cells, and (4) creation of a new pulsatile flow bioreactor capable of supporting fragile constructs during cellular maturation and architectural organization. Other groups are investigating the influence of mechanical stresses on vascular cells as this relates to the tissue engineering of blood vessels [35], the use of gene therapy in vivo to modify grafts (transfection with adenovirus expressing tissue plasminogen activator) [41], and the addition of peptide sequences to constructs to improve endothelial cell attachment [42], among others.

Large diameter conduits are necessary in the repair of many congenital cardiac defects and are often used to establish right ventricle to pulmonary artery continuity [43]. The drawback of homografts or prosthetic conduits is the lack of growth potential. Multiple operations are often necessary because of obstruction secondary to calcification or tissue ingrowth [43]. For these reasons, researchers have pursued the creation of tissue-engineered large diameter conduits. Shinóka and colleagues [43] seeded ovine arterial and venous cells onto synthetic biodegradable tubular scaffolds of polyglactin/PGA and cultured the constructs for 7 days. The constructs were used to replace a 2-cm segment of pulmonary artery in lambs. All tissue-engineered constructs were patent at harvest, whereas the acellular control showed thrombosis. Collagen, elastin, and a lumenal endothelial layer were found to be present on explant analysis [43] (Fig 5). Stock and associates [44] seeded ovine venous cells onto porous P4HB patches. After further culture of the construct, six autologously seeded patches were implanted into the proximal pulmonary artery in a patch augmentation procedure [44]. Echocardiography showed no dilatation or stenosis, and histology showed the formation of organized tissue. Biochemical assays revealed increasing collagen, elastin, and proteoglycans content [44]. The same group seeded ovine arterial cells onto a copolymer of PGA and PHA and used these constructs to replace 3- to 4-cm segments of the abdominal aorta in lambs [45]. All acellular control conduits became occluded, whereas the tissue-engineered conduits remained patent. The collagen and DNA content of the constructs approached the native aorta over time as did the mechanical strain-stress curve [45].



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  		Fig 5. Miller’s elastic staining shows the presence of elastic fibers (black) and collagen fibers in the conduit. (a) Tissue-engineered conduit 4 months after implantation, constructed using an arterial cell source. (x50 before 37% reduction.) (b) Junction between the tissue-engineered conduit and the native pulmonary artery (PA). (x50 before 37% reduction.) (c) Tissue-engineered conduit constructed using a venous cell source. (x100 before 37% reduction.) (d) Tissue-engineered conduit constructed using a venous cell source. (x200 before 37% reduction.) Reprinted with permission from Shinoka T, Shum-Tim D, Ma PX, et al. Creation of viable pulmonary artery autografts through tissue engineering. J Thorac Cardiovasc Surg 1998;115:536–46. © 1998 Mosby, Inc.

 
Recently, Shin’oka and colleagues [46] moved to the next level and implanted a tissue-engineered pulmonary artery into a human being. The patient was a 4-year-old girl who was born with a single right ventricle and pulmonary atresia. She had undergone pulmonary artery angioplasty and a Fontan procedure, but an angiogram revealed occlusion of her right intermediate pulmonary artery the following year. The group harvested a 2-cm segment of peripheral vein, isolated the cells, and expanded the cells in culture. A tubular scaffold of polycaprolactone-PLA copolymer reinforced with woven PGA was seeded with 12 million cells and transplanted after 10 days in culture. The tissue-engineered graft was used to reconstruct the pulmonary artery. Postoperatively, the graft was patent on angiography, and the patient was clinically well after 7 months [46]. This was the first reported human implant of a tissue-engineered blood vessel constructed from cells and polymer, truly an exciting accomplishment.


   Myocardium
Top
 Abstract
 Introduction
 Heart valves
 Blood vessels
 Myocardium
Esophagus
 Trachea
 Future directions
 References
 
Congestive heart failure affects an estimated 4.8 million patients in the United States and 400,000 new cases are diagnosed each year [47]. Mortality is greater than 50% within 5 years of this diagnosis [47]. Although 2,340 heart transplants were performed in 1998 in the United States, the utility of cardiac transplantation is limited by donor shortage, complications of immunosuppression regimens, and failure of the transplanted organ [47]. Congestive heart failure is characterized by loss or dysfunction of cardiomyocytes, whether because of ischemic heart disease, hypertensive heart disease, or idiopathic cardiomyopathy [47]. The heart lacks the ability to regenerate, as it lacks a population of proliferating cells [48, 49]. There are no stem cells within the myocardium and the mature cardiomyocyte is a terminally differentiated cell that does not enter the cell cycle. This is why injury to the myocardium, such as myocardial infarction, results in an irreversible loss of cells and replacement by fibroblasts [48]. Increasing the number of cardiomyocytes or "cardiomyocyte-like" cells in hearts afflicted by congestive heart failure could potentially improve heart function.

Additionally, congenital heart disease is a considerable problem worldwide, affecting approximately 1% of infants, and is associated with significant morbidity and mortality. Surgical correction of these anomalies often requires patch reconstruction of stenotic lesions or conduit placement for complex congenital defects [50]. Furthermore, conditions such as hypoplastic heart syndrome have few successful surgical options for clinical improvement other than transplantation. In general, most currently available grafts are subject to material-related failure and have the significant limitation of the lack of growth potential. Additionally, these materials lack the ability to contract and may be thrombogenic [50]. An autologous, contractile, tissue-engineered graft would be extremely useful in the surgical repair of congenital heart defects, either as a replacement of the wall of a cardiac chamber or as a means to augment the myocardium [50].

Because of these problems, significant efforts are being made in "myocardial tissue engineering." Investigators have taken two basic approaches toward this end. The first, cellular cardiomyoplasty, involves direct cell transplantation into the myocardium as a means of augmentation. Investigators have used numerous cell sources to achieve this goal. Soonpaa and associates [51] were the first to report that fetal cardiomyocytes could be engrafted and integrated with the normal myocardium of mice. Scorsin and colleagues [52] injected fetal cardiomyocytes into the left ventricle of rats in which a myocardial infarction had been created by proximal occlusion of the left coronary artery. They noted engrafted myocytes in 50% of the infarcted rats at the infarct border zone [52]. In a later study, the same group repeated the experiment, this time injecting male donor cells into female recipient myocardium. They were then able to detect the transplanted cells by fluorescent in situ hybridization using a DNA probe specific for the Y chromosome. They found that transplantation of fetal cardiomyocytes resulted in improved left ventricular function as demonstrated by ejection fraction and cardiac output [53]. Leor and associates [54] injected fragments of cultured fetal human or rat ventricles into postinfarction scar and found that the tissue survived up to 65 days. Li and colleagues [55] evaluated the viability and contractility of fetal cardiomyocytes transplanted into connective tissue of the rat. They found that these cells formed contractile cardiac tissue [55]. In a follow-up study, the same group [56] produced scar tissue in the left ventricular wall of rats by cryoinjury. After labeling fetal cardiomyocytes by transfecting them with a plasmid containing the ß-galactosidase gene, they injected the cells into the scar tissue. Two months after cryoinjury, they measured cardiac function using a Langendorff preparation and found improved heart function [56].

Unfortunately, allogenic cell transplantation is dependent on immunosuppression for success. Therefore, Li and colleagues injected autologous adult cardiomyocytes into cryoinjured hearts. They harvested and cultured atrial heart cells for this purpose. Again, myocardial function was improved as evaluated in a Langendorff preparation in comparison with control hearts [57]. The same group created myocardial infarctions in adult swine by occlusion of the distal left anterior descending artery with an intralumenal coil and harvested cardiomyocytes from the interventricular septum at the same time. The cells were then injected into the infarct zone after labeling with 5-bromo-2'deoxyuridine (BrDU) [58]. Technetium 99m-sestamibi single photon emission tomography demonstrated greater wall motion and an improvement in perfusion scores in those receiving transplantation [58]. Yokomuro and colleagues [59] also injected cryopreserved cardiomyocytes into infarcted hearts, noting viable engrafted cells. Reinecke and associates [60] compared fetal, neonatal, and adult cardiomyocyte survival after injection into normal and cryoinjured myocardium and noted that only fetal and neonatal cells formed viable grafts under all conditions.

Unlike heart muscle, skeletal muscle is able to regenerate because of the presence of immature, less differentiated "satellite cells" or myoblasts [61]. These cells, located between the plasma membrane and basal lamina of mature skeletal muscle, are able to fuse with surrounding myoblasts or damaged muscle fibers to regenerate viable skeletal muscle [48]. Because of their ability to proliferate (unlike adult cardiomyocytes) and the avoidance of immunosuppressive drugs, autologous myoblasts have been a recent choice for cellular cardiomyoplasty [48–50, 61]. Chiu and colleagues [61] isolated satellite cells from skeletal muscle of dogs, labeled the cells with tritiated thymidine or the lacZ reporter gene, and implanted them into acutely cryoinjured myocardium. New striated muscle was found at the implantation site within a dense scar showing histologic evidence of intercalated discs [61]. These results, along with a similar study in rats, support the hypothesis of "milieu-influenced" differentiation of myoblasts into cardiac-like muscle cells [48, 61]. Taylor and colleagues [49] transplanted autologous skeletal myoblasts into cryoinfarcted myocardium in rabbits and monitored cardiac function in vivo for 6 weeks. Striated cells that retained characteristics of both skeletal and cardiac cells were found in the infarct. Myocardial function was improved in animals with engrafted yoblasts [49, 60].

Bone marrow stromal cells or mesenchymal stem cells have been shown to have the potential of differentiating into cardiomyocytes in vitro after treatment with 5-azacytidine [62]. Because these cells can be harvested repeatedly by bone marrow aspiration, can be expanded significantly in vitro, and do not require immunosuppression, they are an attractive cell source for cellular cardiomyoplasty. In one study investigators harvested, expanded, and labeled marrow stromal cells in rats. They injected the cells into the myocardium and noted the expression of cardiomyogenic phenotypes in the new environment [62].

The second approach to myocardial tissue engineering involves seeding cells onto a biodegradable scaffold. Tissue-engineered constructs have a definitive structure and may be more apt to produce a significant myocardial augmentation when transplanted as opposed to a cell suspension alone. Furthermore, biodegradable polymers such as PGA and poly-L-lactic acid are well suited for the delivery of a large number of cells because of their high porosity and surface area which also allows for the vascularization and structural integration of the new tissue with surrounding native tissue after implantation. Li and colleagues [50] seeded fetal rat ventricular cardiomyocytes, stomach smooth muscle cells, skin fibroblasts, and adult human atrial and ventricular cardiomyocytes onto biodegradable gelatin meshes and studied their in vitro characteristics. All cell types except adult ventricular cardiomyocytes proliferated within the mesh. The same group then implanted meshes with fetal cardiomyocytes into the subcutaneous tissue or onto myocardial scar tissue. They found that the cells survived within the graft and formed junctions with the recipient heart cells, although a significant improvement in ventricular function was not seen [63]. Leor and colleagues [64] seeded porous alginate scaffolds with fetal cardiac cells and implanted the constructs onto myocardial scars. They noted extensive neovascularization of the grafts. Control animals developed left ventricular dilatation and deterioration of left ventricular function, but the grafted animals did not [64].

Ongoing investigations in our laboratory in this area include continued work on cellular cardiomyoplasty using a variety of cell sources including stem cells, improvement of cellular marking protocols, and formal myocardial tissue engineering using fetal skeletal myoblasts in combination with biodegradable polymers to address prenatally diagnosed cardiac defects such as hypoplastic left heart syndrome.


   Esophagus
Top
 Abstract
 Introduction
 Heart valves
 Blood vessels
 Myocardium
 Esophagus
Trachea
 Future directions
 References
 
Esophageal cancer is a difficult gastrointestinal malignancy to treat, and esophagectomy is often a highly morbid procedure with significant postoperative complications [65, 66]. The stomach, jejunum, or colon are often used in reconstruction procedures, and the patient may suffer from malnutrition postoperatively secondary to both poor functioning of the substitute or loss of length and function of the gastrointestinal tract being used [65]. Additionally, the treatment of long gap esophageal atresia often poses a challenge for pediatric surgeons. Current surgical therapies include stretching, circular myotomy, and colonic interposition. These procedures, although often successful, can be complicated by leakage, stricture, and elongation of the interposed colon. Over the last 80 years, development of an artificial, and more recently a tissue-engineered, esophageal replacement has been investigated as a means to treat these esophageal diseases.

Many attempts at creating a replacement esophagus using artificial materials have been investigated including rubber, polyethylene, teflon, and polypropylene, among others [67]. Most of these materials were eventually extruded, and the complications of leakage and stenosis long-term often developed. Additionally, autologous tissues and homografts have been used as esophageal substitutes [67]. The skin, trachea, and fascia have been used, but complications such as anastomotic leakage and stenosis again often ensued [67]. Two groups have tried to minimize the use of artificial materials; one by stimulating the regeneration of native esophagus, and the other by creating a tissue-engineered neoesophagus using cells and extracellular matrix materials.

Takimoto and colleagues [68] created a 5-cm cervical esophageal defect in dogs. The defect was then repaired using a two-layered tube consisting of a collagen sponge matrix (types I and III) and an inner silicone stent [68]. The inner silicone stent was removed endoscopically at weekly time points from 2 to 4 weeks. In the dogs in which the stent was removed at 2 or 3 weeks, constriction of the "regenerated esophagus" occurred and eventually the dogs were unable to swallow [68]. In the dogs in which the stent was removed at 4 weeks, regenerated esophageal tissue replaced the defect, and oral feeding was tolerated. Histologically, the regenerated tissue showed stratified epithelia, striated muscle tissue with an inner circular and an outer longitudinal layer, and esophageal glands [68]. In a follow-up study, Yamamoto and associates [69] replaced a 5-cm thoracic defect with the same prosthesis. In this group of animals, the silicone stent was removed after 4 weeks. At the time of stent removal, the host-regenerated tissue had replaced the gap in all dogs. The mucosa had fully regenerated at 3 months and the glands at 12 months. Some stenosis and shrinkage did occur, but there was no further progression after 3 months [69]. The lamina muscularis mucosae were seen as islets of smooth muscle by 12 months. Interestingly, skeletal muscle did not extend into the middle of the regenerated segment as had been seen in the cervical replacement. The authors hypothesized that the surrounding connective tissue is thinner and the blood supply poorer in the mediastinal region compared with the cervical region [69]. To test this theory, the same group attempted to promote tissue regeneration by omental pedicle wrapping. They also retained the silicone stent for a longer time to try to prevent stenosis [70]. Not only did most dogs with omental pedicle wrapping and longer stenting die, but only a thin epithelial and submucosal layer regenerated in this group [70].

Sato and colleagues have been investigating methods to create a tissue-engineered esophageal replacement over the last 8 years. Their goal is to harvest a small amount of normal esophageal tissue from a patient with esophageal cancer before operation, culture the cells, and create a natural esophageal replacement in vitro. Then they propose to graft the replacement into the same patient at the time of operation [66]. Toward these goals, Sato and colleagues isolated and cultured epithelial cells from normal human esophageal mucosa on collagen gels. Then the collagen sheets were transplanted onto the surface of the latissimus dorsi muscles of athymic mice [65]. The constructs were harvested at various time points and histology revealed neovascularization in the collagen layer and an epithelial cell layer similar to normal esophageal epithelium [65]. In the next study, Sato and associates [66] cultured esophageal epithelial cells on the surface of collagen gels in which biodegradable PGA mesh was embedded. They then sutured the mesh to create a tubular structure lined with cultured cells and wrapped these tubes in the latissimus dorsi muscle. The tubular structure was maintained in vivo, and appeared histologically similar to normal esophageal epithelium (Fig 6). A basement membrane was also identified by antilaminin staining [66]. The same group also then embedded fibroblasts into the collagen gels as a feeder layer and repeated the experiment. They noted that the fibroblasts accelerated the proliferation and differentiation of the epithelial cells both in vitro and in vivo [71].



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  		Fig 6. (a) Tissue-engineered esophagus, showing from top to bottom: squamous epithelial layer, collagen gel with fibroblasts, and skeletal muscle (latissimus dorsi). (Hematoxylin and eosin; x10 before 34% reduction.) (b) Tissue-engineered esophagus with a multicellular squamous epithelial layer. (Hematoxylin and eosin; x20 before 34% reduction.) (c) Native esophagus. (Hematoxylin and eosin; x10 before 35% reduction.) Unpublished data, courtesy of M. Sato.

 
Currently, continued efforts toward the creation of a tissue-engineered esophagus are ongoing in our laboratory. In the rat model, both the omentum and latissimus dorsi muscle are being used as an implantation bed for cultured rat esophageal keratinocytes on collagen gels. Furthermore, anastomosis of these tissue-engineered constructs with native small bowel and the cervical esophagus has begun.


   Trachea
Top
 Abstract
 Introduction
 Heart valves
 Blood vessels
 Myocardium
 Esophagus
 Trachea
Future directions
 References
 
The management of tracheal pathology, such as stenosis or cancer, often requires tracheal reconstruction. Primary anastomosis after resection is the method of choice and can usually be performed successfully for defects of up to 50% [72]. Unfortunately, there are cases in which primary anastomosis is not possible, such as after extensive burns, trauma, tumor resection, or postintubation injuries [72]. Additionally, the treatment of congenital tracheal atresia or stenosis can be hindered by the lack of sufficient tissue for surgical reconstruction, as the length of trachea involved may be extensive. Over the last 50 years, multiple approaches to tracheal reconstruction have been attempted both clinically and experimentally, including the use of prosthetics, autografts, homografts, and allografts [73]. Prostheses such as Dacron polyurethane mesh, polytetrafluoroethylene, polypropylene mesh, silicone rubber, and even glass tubes [74] have been tried and have often been fraught with complications such as infection, extrusion, and stenosis [72]. Autogenous and alloplastic tissues have also been used from sources such as fascia, skin, bone, periosteum, cartilage, perichondrium, tracheal allografts, muscle, esophagus, pericardium, dura mater, and small bowel [72, 75–78]. Again, in the majority of these grafts, the problem of late stenosis of the prosthesis is difficult to solve [79]. Because of these difficulties, investigators are still in search of the ideal tracheal replacement material.

Since 1988, tissue engineering of cartilage has been successfully pursued by a number of groups. During that year, Vacanti and associates [80, 81] were able to create new hyaline cartilage in athymic mice by using isolated bovine chondrocytes and biodegradable suture material (polyglactin 910 and PGA), which served as a temporary scaffold to which cells attached until they created their own matrix. Other polymers were used in subsequent studies with the same success, including nonwoven mesh of PGA and PGA/poly-L-lactic acid copolymers. These polymers were used to create cartilage constructs with predetermined shapes [80, 82]. Optimal cell concentrations were determined [80]. Other biomaterials have also been used in combination with chondrocytes to produce high-quality tissue-engineered cartilage. These biomaterials include calcium alginate gels, collagen gels, fibrin glue, and agarose gels, among others [80]. These techniques have been applied to create cartilage in the shape of the human ear [83], nasoseptal implants [84], and a temporomandibular joint disc [85], as well as for joint resurfacing [86] and meniscal replacement [87]. Tissue-engineered cartilage constructs have also been used in nonnative sites. PGA meshes were used to create ventriculoperitoneal shunts for hydrocephalus [88]. Chondrocytes in hydrogels have been used for nipple reconstruction in pigs [89] and as a potential treatment of vesiculoureteral reflux [90]. Recently, the in vitro modulation of chondrogenesis by dynamic cell seeding and bioreactors has been investigated [91]. Vunjak-Novakovic and colleagues [91] found that hydrodynamic conditions in convective-flow tissue culture bioreactors can modulate the composition (matrix components), morphology, mechanical properties, and electromechanical function of tissue-engineered cartilage.

The success of cartilage tissue engineering has led to the use of these techniques for tracheal replacement. Vacanti and colleagues [92] seeded chondrocytes obtained from newborn calf shoulders onto sheets of nonwoven PGA mesh after labeling with 5-bromo-2'deoxyuridine (BrDU). The cell-polymer constructs were wrapped around silastic tubes and implanted into nude mice for 1 month [92]. Specimens were then excised, examined, and found to be both grossly and histologically very similar to normal bovine cartilage. The cylinders were then sutured into large circumferential defects created in the cervical tracheas of nude rats. Four of 6 animals survived and were able to breathe without mechanical ventilation [92]. In a follow-up study by the same group, tracheal epithelial cells were isolated and injected into previously created cylindrical tubes of tissue-engineered cartilage [73]. Examination of these specimens revealed mature cartilage, and the lumens were lined with epithelial structures, including submucosal connective tissue [73]. The inner surface revealed different stages of pseudostratified columnar epithelial components, some cells showing cilia at 3 weeks after injection [73]. These initial results are encouraging for the development of a tissue-engineered tracheal replacement in humans. Current work in our laboratory involves dynamically seeding fetal chondrocytes on PGA to create high-quality cartilage constructs (Fig 7) for use in prenatally diagnosed tracheal defects in a large animal model.



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  		Fig 7. (a) Photomicrograph of tissue-engineered cartilage, using fetal chondrocytes from ear specimens. (Arrow = residual polymer. [Hematoxylin and eosin; x10 before 46% reduction.]) (b) Tissue-engineered cartilage. (Hematoxylin and eosin; x20 before 45% reduction.) (Unpublished data, J. R. Fuchs.)

 

   Future directions
Top
 Abstract
 Introduction
 Heart valves
 Blood vessels
 Myocardium
 Esophagus
 Trachea
 Future directions
References
 
The field of tissue engineering, although still in its infancy, has made significant progress toward organ and tissue replacement not only in the field of cardiothoracic surgery, but involving all tissues. Although there are hundreds of areas of current active research, three emerging concepts involve: the use and development of dynamic in vitro culture systems (bioreactors), the use and development of microfabrication technology to create vascularized tissues and organs, and the search for and use of an appropriate multipotent, undifferentiated stem cell in tissue engineering.

Producing a dynamic in vitro microenvironment for tissues may be an important aspect in guiding the formation of tissue with certain structural and functional characteristics [10]. The use of bioreactors (Fig 8) allows the investigator to control flow and mixing, which can enhance mass transfer of nutrients, wastes, and regulatory molecules [5]. Bioreactors can also provide mechanical cues to stimulate cells to produce specific components [5]. Vunjak-Novakovic and colleagues [5, 91], for example, have shown that stirred conditions improve the quality of cartilage produced in comparison to standard static culture conditions. Hydrodynamic stimulation of the cartilage resulted in the production of greater amounts of extracellular matrix components such as glycosaminoglycans and collagen, leading to improved mechanical properties [5, 91]. Additionally, pulsatile flow bioreactors have been used in the fabrication of trileaflet heart valves and blood vessels [17, 29]. Pulsed constructs had greater burst strengths, suture retention, and collagen content than nonpulsed constructs [31]. Additionally, hepatocytes cultured under continuous flow bioreactor conditions show increased organization (spheroid formation) and function (albumin production) as compared with static conditions [93, 94].



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  		Fig 8. Example of a bioreactor for use in tissue engineering.

 
One of the major challenges of the field is to be able to fabricate larger tissues and organs such as liver, kidney, and heart. Most early approaches in tissue engineering relied on blood vessel in-growth into tissue-engineered structures from the host to achieve long-term vascularization [95]. This method has been successful for some tissues, but is inadequate for thick, complex organs [95]. With these organs, there has been an insufficient survival of an adequate mass of transplanted cells to provide function. One of the first attempts to solve this problem involved designing a complex three-dimensional synthetic biodegradable polymer scaffold with an intrinsic network of interconnected branching channels using the three-dimensional printing technique [94]. Hepatocytes were shown to attach to and survive on these scaffolds in both static and flow conditions [94]. The limitation of this technology is that it lacks the resolution needed to form templates containing the microcirculation (the capillaries), which are on the order of 10 microns in diameter. Microfabrication technology has emerged as an approach to this hurdle and arises from the field of microelectromechanical systems (MEMS) [95]. MEMS is a rapidly emerging science with many applications including biomedical engineering. It was originally developed for the circuit industry, and allows complex branching patterns at the 1-micron level to be imprinted onto silicon or Pyrex wafers using photolithography and a series of subtractive etching techniques [95]. These templates can be used as molds to replicate their branching patterns onto bioresorbable polymer films, such as poly-DL-lactic-coglycolic acid (PLGA) (Fig 9), which can then be seeded with endothelial cells to form a microvasculature. Our laboratory has adapted this technology to engineering branched vascular channels for liver fabrication [95]. Currently, our laboratory, in collaboration with Draper Laboratories (Cambridge, MA), is developing prototype vascular networks on silicon wafers and other materials using mathematical modeling and the principles of fluid dynamics/microfluidics as a guide to optimize microvascular networks, to provide an even distribution of blood and maximal mass transfer of oxygen and nutrients. Additionally, micromagnetic resonance imaging and microcomputerized tomography may provide detailed three-dimensional images of the microvasculature for mathematical modeling of flow, branching angles, and resistance to optimize template design. Studies are underway to characterize endothelial cells in a dynamic flow model to determine the optimal conditions for producing endothelialized microvascular networks in vitro.



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  		Fig 9. Scanning electron microscopy of a biodegradable polymer film composed of poly-DL-lactic-coglycolic acid with channels created by a silicon wafer template. The silicon template was created using microfabrication technology. Unpublished data, courtesy of H. Terai.

 
Although autologous cells are an excellent source for use in tissue engineering, they are fully differentiated cells and as such have a decreased ability to proliferate (hepatocytes, cardiomyocytes, and others). A pluripotent, highly proliferative cell source from which a variety of cell types could be derived would be extremely useful in tissue engineering [5]. The field of stem cell biology is rapidly expanding. Recently, pluripotent stem cells have been isolated from human fetal tissue and have shown the ability to differentiate into a variety of cell types found in embryonic germ layers [96]. Many adult tissues contain stem cell populations with the ability to repair tissue after disease or trauma. One example is the oval cell in the liver and now the small hepatocyte, each of which can differentiate into either mature liver cells or bile ducts in culture and are considered committed progenitor cells [5, 97, 98]. Another example is the skeletal muscle satellite cell. Additionally, a cell population isolated from human bone marrow, "the mesenchymal stem cell," has been induced to differentiate into adipocytic, chondrocytic, and osteocytic lineages under different culture conditions [99]. Characterization of these and other potential "stem cell" populations is an area of intensive research by many groups. In the future, these cells may prove extremely important in tissue engineering.

Tissue engineering has emerged as a rapidly expanding approach to address the organ shortage problem. Much progress has been made in areas relevant to cardiothoracic surgery as well as in other tissues. There are still many questions to be answered that will require the close, interdisciplinary collaboration of surgeons, engineers, chemists, and biologists, with the ultimate goal of functional neoorganogenesis for human use.


   References
Top
 Abstract
 Introduction
 Heart valves
 Blood vessels
 Myocardium
 Esophagus
 Trachea
 Future directions
 References
 

  1. Langer R., Vacanti J.P. Tissue engineering. Science 1993;260:920-926.[Abstract/Free Full Text]
  2. Vacanti J.P. Beyond transplantation. Third annual Samuel Jason Mixter lecture. Arch Surg 1988;123:545-549.[Abstract/Free Full Text]
  3. Barnard C.N. Human cardiac transplantation. An evaluation of the first two operations performed at the Groote Schuur Hospital, Cape Town. Am J Cardiol 1968;22:584-596.[Medline]
  4. United Network for Organ Sharing (UNOS). Available from: URL: http://www.unos.org/frame_Default.asp?Category=Newsdata.
  5. Vacanti J.P., Langer R. Tissue engineering: the design and fabrication of living replacement devices for surgical reconstruction and transplantation. Lancet 1999;354(Suppl 1):SI32-SI34.
  6. Bisceglie V. Über die antineoplastische Immunität: heterologe Einpflanzung von Tumoren in Hühnerembryonen. Ztschr f Krebsforsch 1933;40:122-140.
  7. Chick W.L., Like A.A., Lauris V. Beta cell culture on synthetic capillaries: an artificial endocrine pancreas. Science 1975;187:847-849.[Abstract/Free Full Text]
  8. Burke J.F., Yannas I.V., Quinby W.C., Jr, Bondoc C.C., Jung W.K. Successful use of a physiologically acceptable artificial skin in the treatment of extensive burn injury. Ann Surg 1981;194:413-428.[Medline]
  9. Lysaght M.J., Nguy N.A., Sullivan K. An economic survey of the emerging tissue engineering industry. Tissue Eng 1998;4:231-238.[Medline]
  10. Marler J.J., Upton J., Langer R., Vacanti J.P. Transplantation of cells in matrices for tissue regeneration. Adv Drug Del Rev 1998;33:165-182.[Medline]
  11. Cima L.G., Langer R., Vacanti J.P. Polymers for tissue and organ culture. J Bioactive Compat Polymers 1991;6:232-240.
  12. Langer R.S., Vacanti J.P. Artificial organs. Sci Am 1995;273:100-103.
  13. Williams S.F., Martin D.P., Horowitz D.M., Peoples O.P. PHA applications: addressing the price performance issue: I. Tissue engineering. Int J Biol Macromol 1999;25:111-121.[Medline]
  14. Schoen F.J., Levy R.J. Founder’s Award, 25th Annual Meeting of the Society for Biomaterials, perspectives. Providence, RI, April 28–May 2, 1999. Tissue heart valves: current challenges and future research perspectives. J Biomed Mater Res 1999;47:439-465.[Medline]
  15. Mayer J.E., Jr, Shin’oka T., Shum-Tim D. Tissue engineering of cardiovascular structures. Curr Opin Cardiol 1997;12:528-532.[Medline]
  16. Hammermeister K.E., Sethi G.K., Henderson W.G., Oprian C., Kim T., Rahimtoola S. A comparison of outcomes in men 11 years after heart-valve replacement with a mechanical valve or bioprosthesis. Veterans Affairs Cooperative Study on Valvular Heart Disease. N Engl J Med 1993;328:1289-1296.[Abstract/Free Full Text]
  17. Hoerstrup S.P., Sodian R., Daebritz S., et al. Functional living trileaflet heart valves grown in vitro. Circulation 2000;102(Suppl 3):III44-III49.[Abstract/Free Full Text]
  18. Harken D.F., Taylor W.J., LeFemine A.A., et al. Aortic valve replacement with a caged ball valve. Am J Cardiol 1962;9:292-299.[Medline]
  19. Bader A., Schilling T., Teebken O.E., et al. Tissue engineering of heart valves—human endothelial cell seeding of detergent acellularized porcine valves. Eur J Cardiothorac Surg 1998;14:279-284.[Abstract/Free Full Text]
  20. O’Brien M.F., Goldstein S., Walsh S., Black K.S., Elkins R., Clarke D. The SynerGraft valve: a new acellular (nonglutaraldehyde-fixed) tissue heart valve for autologous recellularization first experimental studies before clinical implantation. Semin Thorac Cardiovasc Surg 1999;11(Suppl 1):194-200.[Medline]
  21. Sodian R., Hoerstrup S.P., Sperling J.S., et al. Early in vivo experience with tissue-engineered trileaflet heart valves. Circulation 2000;102(Suppl 3):III22-III29.[Medline]
  22. Shinoka T., Breuer C.K., Tanel R.E., et al. Tissue engineering heart valves: valve leaflet replacement study in a lamb model. Ann Thorac Surg 1995;60:S513-S516.
  23. Shinoka T., Ma P.X., Shum-Tim D., et al. Tissue-engineered heart valves. Autologous valve leaflet replacement study in a lamb model. Circulation 1996;94(Suppl):II164-II168.
  24. Breuer C.K., Shin’oka T., Tanel R.E., et al. Tissue engineering lamb heart valve leaflets. Biotech Bioeng 1996;50:562-567.
  25. Shinoka T., Shum-Tim D., Ma P.X., et al. Tissue-engineered heart valve leaflets: does cell origin affect outcome?. Circulation 1997;96(Suppl):II102-II107.
  26. Zund G., Breuer C.K., Shinoka T., et al. The in vitro construction of a tissue engineered bioprosthetic heart valve. Eur J Cardiothor Surg 1997;11:493-497.[Abstract]
  27. Zund G., Ye Q., Hoerstrup S.P., et al. Tissue engineering in cardiovascular surgery: MTT, a rapid and reliable quantitative method to assess the optimal human cell seeding on polymeric meshes. Eur J Cardiothorac Surg 1999;15:519-524.[Abstract/Free Full Text]
  28. Hoerstrup S.P., Zund G., Schoeberlein A., Ye Q., Vogt P.R., Turina M.I. Fluorescence activated cell sorting: a reliable method in tissue engineering of a bioprosthetic heart valve. Ann Thorac Surg 1998;66:1653-1657.[Abstract/Free Full Text]
  29. Sodian R., Hoerstrup S.P., Sperling J.S., et al. Evaluation of biodegradable, three-dimensional matrices for tissue engineering of heart valves. ASAIO J 2000;46:107-110.[Medline]
  30. Stock U.A., Nagashima M., Khalil P.N., et al. Tissue-engineered valved conduits in the pulmonary circulation. J Thorac Cardiovasc Surg 2000;119:732-740.[Abstract/Free Full Text]
  31. Niklason L.E., Gao J., Abbott W.M., et al. Functional arteries grown in vitro. Science 1999;284:489-493.[Abstract/Free Full Text]
  32. Campbell J.H., Efendy J.L., Campbell G.R. Novel vascular graft grown within recipient’s own peritoneal cavity. Circ Res 1999;85:1173-1178.[Abstract/Free Full Text]
  33. Edelman E.R. Vascular tissue engineering: designer arteries. Circ Res 1999;85:1115-1117.[Free Full Text]
  34. Miwa H., Matsuda T. An integrated approach to the design and engineering of hybrid arterial prostheses. J Vasc Surg 1994;19:658-667.[Medline]
  35. Nerem R.M. Tissue engineering in the USA. Med Biol Eng Comput 1992;30:CE8-CE12.[Medline]
  36. Williams S.K. Endothelial cell transplantation. Cell Transplant 1995;4:401-410.[Medline]
  37. Weinberg C.B., Bell E. A blood vessel model constructed from collagen and cultured vascular cells. Science 1986;231:397-400.[Abstract/Free Full Text]
  38. L’Heureux N., Paquet S., Labbe R., Germain L., Auger F.A. A completely biological tissue-engineered human blood vessel. FASEB J 1998;12:47-56.[Abstract/Free Full Text]
  39. Teebken O.E., Bader A., Steinhoff G., Haverich A. Tissue engineering of vascular grafts: human cell seeding of decellularised porcine matrix. Eur J Vasc Endovasc Surg 2000;19:381-386.[Medline]
  40. Huynh T., Abraham G., Murray J., Brockbank K., Hagen P.O., Sullivan S. Remodeling of an acellular collagen graft into a physiologically responsive neovessel. Nat Biotechnol 1999;17:1083-1086.[Medline]
  41. Kuo M.D., Waugh J.M., Yuksel E., Weinfeld A.B., Yuksel M., Dake M.D. 1998 ARRS President’s Award. The potential of in vivo vascular tissue engineering for the treatment of vascular thrombosis: a preliminary report. American Roentgen Ray Society. AJR Am J Roentgenol 1998;171:553-558.[Abstract/Free Full Text]
  42. Hubbell J.A., Massia S.P., Desai N.P., Drumheller P.D. Endothelial cell-selective materials for tissue engineering in the vascular graft via a new receptor. Biotechnology (NY) 1991;9:568-572.[Medline]
  43. Shinoka T., Shum-Tim D., Ma P.X., et al. Creation of viable pulmonary artery autografts through tissue engineering. J Thorac Cardiovasc Surg 1998;115:536-546.[Abstract/Free Full Text]
  44. Stock U.A., Sakamoto T., Hatsuoka S., et al. Patch augmentation of the pulmonary artery with bioabsorbable polymers and autologous cell seeding. J Thorac Cardiovasc Surg 2000;120:1158-1168.[Abstract/Free Full Text]
  45. Shum-Tim D., Stock U., Hrkach J., et al. Tissue engineering of autologous aorta using a new biodegradable polymer. Ann Thorac Surg 1999;68:2298-2305.[Abstract/Free Full Text]
  46. Shin’oka T., Imai Y., Ikada Y. Transplantation of a tissue-engineered pulmonary artery. N Engl J Med 2001;344:532-533.[Free Full Text]
  47. Kessler P.D., Byrne B.J. Myoblast cell grafting into heart muscle: cellular biology and potential applications. Annu Rev Physiol 1999;61:219-242.[Medline]
  48. Dorfman J., Duong M., Zibaitis A., et al. Myocardial tissue engineering with autologous myoblast implantation. J Thorac Cardiovasc Surg 1998;116:744-751.[Abstract/Free Full Text]
  49. Taylor D.A., Atkins B.Z., Hungspreugs P., et al. Regenerating functional myocardium: improved performance after skeletal myoblast transplantation. Nat Med 1998;4:929-933.[Medline]
  50. Li R.K., Yau T.M., Weisel R.D., et al. Construction of a bioengineered cardiac graft. J Thorac Cardiovasc Surg 2000;119:368-375.[Abstract/Free Full Text]
  51. Soonpaa M.H., Koh G.Y., Klug M.G., Field L.J. Formation of nascent intercalated disks between grafted fetal cardiomyocytes and host myocardium. Science 1994;264:98-101.[Abstract/Free Full Text]
  52. Scorsin M., Marotte F., Sabri A., et al. Can grafted cardiomyocytes colonize peri-infarct myocardial areas?. Circulation 1996;94(Suppl):II337-II340.
  53. Scorsin M., Hagege A.A., Marotte F., et al. Does transplantation of cardiomyocytes improve function of infarcted myocardium?. Circulation 1997;96(Suppl):II188-II193.
  54. Leor J., Patterson M., Quinones M.J., Kedes L.H., Kloner R.A. Transplantation of fetal myocardial tissue into the infarcted myocardium of rat. A potential method for repair of infarcted myocardium?. Circulation 1996;94(Suppl):II332-II336.
  55. Li R.K., Mickle D.A., Weisel R.D., Zhang J., Mohabeer M.K. In vivo survival and function of transplanted rat cardiomyocytes. Circ Res 1996;78:283-288.[Abstract/Free Full Text]
  56. Li R.K., Jia Z.Q., Weisel R.D., et al. Cardiomyocyte transplantation improves heart function. Ann Thorac Surg 1996;62:654-661.[Abstract/Free Full Text]
  57. Sakai T., Li R.K., Weisel R.D., et al. Autologous heart cell transplantation improves cardiac function after myocardial injury. Ann Thorac Surg 1999;68:2074-2081.[Abstract/Free Full Text]
  58. Li R.K., Weisel R.D., Mickle D.A., et al. Autologous porcine heart cell transplantation improved heart function after a myocardial infarction. J Thorac Cardiovasc Surg 2000;119:62-68.[Abstract/Free Full Text]
  59. Yokomuro H., Li R.K., Mickle D.A., Weisel R.D., Verma S., Yau T.M. Transplantation of cryopreserved cardiomyocytes. J Thorac Cardiovasc Surg 2001;121:98-107.
  60. Reinecke H., Zhang M., Bartosek T., Murry C.E. Survival, integration, and differentiation of cardiomyocyte grafts: a study in normal and injured rat hearts. Circulation 1999;100:193-202.[Abstract/Free Full Text]
  61. Chiu R.C., Zibaitis A., Kao R.L. Cellular cardiomyoplasty: myocardial regeneration with satellite cell implantation. Ann Thorac Surg 1995;60:12-18.[Abstract/Free Full Text]
  62. Wang J.S., Shum-Tim D., Galipeau J., Chedrawy E., Eliopoulos N., Chiu R.C. Marrow stromal cells for cellular cardiomyoplasty: feasibility and potential clinical advantages. J Thorac Cardiovasc Surg 2000;120:999-1005.[Abstract/Free Full Text]
  63. Li R.K., Jia Z.Q., Weisel R.D., Mickle D.A., Choi A., Yau T.M. Survival and function of bioengineered cardiac grafts. Circulation 1999;100(Suppl):II63-II69.
  64. Leor J., Aboulafia-Etzion S., Dar A., et al. Bioengineered cardiac grafts: a new approach to repair the infarcted myocardium?. Circulation 2000;102(Suppl):III56-III61.[Abstract/Free Full Text]
  65. Sato M., Ando N., Ozawa S., Nagashima A., Kitajima M. A hybrid artificial esophagus using cultured human esophageal epithelial cells. ASAIO J 1993;39:M554-M557.[Medline]
  66. Sato M., Ando N., Ozawa S., Miki H., Kitajima M. An artificial esophagus consisting of cultured human esophageal epithelial cells, polyglycolic acid mesh, and collagen. ASAIO J 1994;40:M389-M392.[Medline]
  67. Sato M., Ando N., Ozawa S., Miki H., Hayashi K., Kitajima M. Artificial esophagus. Mater Sci Forum 1997;250:105-114.
  68. Takimoto Y., Nakamura T., Yamamoto Y., Kiyotani T., Teramachi M., Shimizu Y. The experimental replacement of a cervical esophageal segment with an artificial prosthesis with the use of collagen matrix and a silicone stent. J Thorac Cardiovasc Surg 1998;116:98-106.[Abstract/Free Full Text]
  69. Yamamoto Y., Nakamura T., Shimizu Y., et al. Intrathoracic esophageal replacement in the dog with the use of an artificial esophagus composed of a collagen sponge with a double-layered silicone tube. J Thorac Cardiovasc Surg 1999;118:276-286.[Abstract/Free Full Text]
  70. Yamamoto Y., Nakamura T., Shimizu Y., et al. Intrathoracic esophageal replacement with a collagen sponge–silicone double layer tube: evaluation of omental-pedicle wrapping and prolonged placement of an inner stent. ASAIO J 2000;46:734-739.[Medline]
  71. Miki H., Ando N., Ozawa S., Sato M., Hayashi K., Kitajima M. An artificial esophagus constructed of cultured human esophageal epithelial cells, fibroblasts, polyglycolic acid mesh, and collagen. ASAIO J 1999;45:502-508.[Medline]
  72. Kon M., van den Hooff A. Cartilage tube formation by perichondrium: a new concept for tracheal reconstruction. Plast Reconstr Surg 1983;72:791-797.[Medline]
  73. Sakata J., Vacanti C.A., Schloo B., Healy G.B., Langer R., Vacanti J.P. Tracheal composites tissue engineered from chondrocytes, tracheal epithelial cells, and synthetic degradable scaffolding. Transplant Proc 1994;26:3309-3310.[Medline]
  74. Jacobs J.R. Investigations into tracheal prosthetic reconstruction. Laryngoscope 1988;98:1239-1245.[Medline]
  75. Fonkalsrud E.W., Sumida S. Tracheal replacement with autologous esophagus for tracheal stricture. Arch Surg 1971;102:139-142.[Abstract/Free Full Text]
  76. Letang E., Sanchez-Lloret J., Gimferrer J.M., Ramirez J., Vicens A. Experimental reconstruction of the canine trachea with a free revascularized small bowel graft. Ann Thorac Surg 1990;49:955-958.[Abstract]
  77. Har-El G., Krespi Y.P., Goldsher M. The combined use of muscle flaps and alloplasts for tracheal reconstruction. Arch Otolaryngol Head Neck Surg 1989;115:1310-1313.
  78. Cohen R.C., Filler R.M., Konuma K., Bahoric A., Kent G., Smith C. The successful reconstruction of thoracic tracheal defects with free periosteal grafts. J Pediatr Surg 1985;20:852-858.[Medline]
  79. Sabas A.A., Uez J.B., Rojas O., Inones A., Aranguren J.A. Replacement of the trachea with dura mater. Experimental work. J Thorac Cardiovasc Surg 1977;74:761-765.[Abstract]
  80. Ibarra C, Langer R, Vacanti JP. Tissue engineering: cartilage, bone, and muscle. In: Lanza RP, Chick WL, eds. Yearbook of cell and tissue transplantation. Kluwer Academic Publishers: Amsterdam, The Netherlands 1996/1997:235–45.
  81. Vacanti C.A., Langer R., Schloo B., Vacanti J.P. Synthetic polymers seeded with chondrocytes provide a template for new cartilage formation. Plast Reconstr Surg 1991;88:753-759.[Medline]
  82. Kim W.S., Vacanti J.P., Cima L., et al. Cartilage engineered in predetermined shapes employing cell transplantation on synthetic biodegradable polymers. Plast Reconstr Surg 1994;94:233-240.[Medline]
  83. Vacanti C.A., Cima L.G., Ratkowski D., et al. Tissue engineered growth of new cartilage in the shape of a human ear using synthetic polymers seeded with chondrocytes. Mater Res Soc Symp Proc 1992;252:367-373.
  84. Puelacher W.C., Mooney D., Langer R., Upton J., Vacanti J.P., Vacanti C.A. Design of nasoseptal cartilage replacements synthesized from biodegradable polymers and chondrocytes. Biomaterials 1994;15:774-778.[Medline]
  85. Puelacher W.C., Wisser J., Vacanti C.A., Ferraro N.F., Jaramillo D., Vacanti J.P. Temporomandibular joint disc replacement made by tissue-engineered growth of cartilage. J Oral Maxillofac Surg 1994;52:1172-1178.[Medline]
  86. Vacanti C.A., Kim W., Schloo B., Upton J., Vacanti J.P. Joint resurfacing with cartilage grown in situ from cell-polymer structures. Am J Sports Med 1994;22:485-488.[Abstract/Free Full Text]
  87. Ibarra C., Jannetta C., Vacanti C.A., et al. Tissue engineered meniscus: a potential new alternative to allogeneic meniscus transplantation. Transplant Proc 1997;29:986-988.[Medline]
  88. Lee I.W., Vacanti J.P., Taylor G.A., Madsen J.R. The living shunt: a tissue engineering approach in the treatment of hydrocephalus. Neurol Res 2000;22:105-110.[Medline]
  89. Cao Y.L., Lach E., Kim T.H., Rodriguez A., Arevalo C.A., Vacanti C.A. Tissue-engineered nipple reconstruction. Plast Reconstr Surg 1998;102:2293-2298.[Medline]
  90. Atala A., Cima L.G., Kim W., et al. Injectable alginate seeded with chondrocytes as a potential treatment for vesicoureteral reflux. J Urol 1993;150:745-747.[Medline]
  91. Vunjak-Novakovic G., Martin I., Obradovic B., et al. Bioreactor cultivation conditions modulate the composition and mechanical properties of tissue-engineered cartilage. J Orthop Res 1999;17:130-138.[Medline]
  92. Vacanti C.A., Paige K.T., Kim W.S., Sakata J., Upton J., Vacanti J.P. Experimental tracheal replacement using tissue-engineered cartilage. J Pediatr Surg 1994;29:201-205.[Medline]
  93. Pollok J.M., Kluth D., Cusick R.A., et al. Formation of spheroidal aggregates of hepatocytes on biodegradable polymers under continuous-flow bioreactor conditions. Eur J Pediatr Surg 1998;8:195-199.[Medline]
  94. Kim S.S., Utsunomiya H., Koski J.A., et al. Survival and function of hepatocytes on a novel three-dimensional synthetic biodegradable polymer scaffold with an intrinsic network of channels. Ann Surg 1998;228:8-13.[Medline]
  95. Kaihara S., Borenstein J., Koka R., et al. Silicon micromachining to tissue engineer branched vascular channels for liver fabrication. Tissue Eng 2000;6:105-117.[Medline]
  96. Thomson J.A., Itskovitz-Eldor J., Shapiro S.S., et al. Embryonic stem cell lines derived from human blastocysts. Science 1998;282:1145-1147.[Abstract/Free Full Text]
  97. Mitaka T., Sato F., Mizuguchi T., Yokono T., Mochizuki Y. Reconstruction of hepatic organoid by rat small hepatocytes and hepatic nonparenchymal cells. Hepatology 1999;29:111-125.[Medline]
  98. Lazaro C.A., Rhim J.A., Yamada Y., Fausto N. Generation of hepatocytes from oval cell precursors in culture. Cancer Res 1998;58:5514-5522.[Abstract/Free Full Text]
  99. Pittenger M.F., Mackay A.M., Beck S.C., et al. Multilineage potential of adult human mesenchymal stem cells. Science 1999;284:143-147.[Abstract/Free Full Text]



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