Tissue Engineering of Organ Systems

C H AP TER 1 9

Tissue Engineering of Organ Systems Jason M. Sandberg, Anthony Atala Department of Urology, Wake Forest University School of Medicine, Winston–Salem, NC, USA

LEARNING OBJECTIVES To understand the basic anatomy and physiology of the organs involved in tissue engineering n To understand organ failure and the respective diseases causing it n To understand the importance of cell–cell and cell–matrix interaction in tissue-engineered organs n To be aware of the importance of adequate mechanical properties of the matrix in organ tissue engineering n To understand the importance of using the human body as a “bioreactor” to allow final tissue differentiation n To know the limiting factors for successful in vivo implantation of in vitro engineered organs n To take note of the most recent biologic constructs in development and their current clinical applications n

19.1 INTRODUCTION Organ replacement, repair, or restoration of organ function is classically treated by organ or tissue transplantation to replace diseased, missing, or malformed organs. In cases where congenital malformation can lead to severe organ agenesis; dysplasia and hypoplasia; or anatomical, structural, and functional deficiencies, it might be necessary to replace the entire or part of the organ. In other cases, cancer, trauma, infection, inflammation, or iatrogenic injuries may lead to organ damage or loss and dictate eventual replacement. Yet three major problems are linked to human organ donation: tissue compatibility, tissue rejection, and donor shortage. Depending on the type of organ, donors and receivers need to be immunologically matched to minimize the risk of rejection, and the donor tissue needs to be of the highest quality to provide the best chance of obtaining long-term function, while for some vital and delicate organs such as pancreas, the number of available donor organs is scarce. Tissue Engineering. http://dx.doi.org/10.1016/B978-0-12-420145-3.00019-5 Copyright © 2015 Elsevier Inc. All rights reserved.

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In some organ systems, these problems can be minimized by reconstruction with tissue from other organ systems when native tissue is lacking. For example, in the genitourinary tract, surgeons can use nonurologic tissues, such as intestinal or gastric segments, mucosa from several body sites, and skin with homologous tissues (e.g., cadaver fascia, cadaver, or donor kidney) to reconstruct vital structures. Other options include biopolymers (e.g., bovine collagen) or synthetic polymers (e.g., silicone, polyurethane or polytetrafluoroethylene (PTFE), poly(glycolic acid) (PGA), poly(lactic acid), and ­poly(lactic-co-glycolic acid)), among others. These approaches are, however, often linked to complications, such as metabolic disturbances and even cancer, and rarely replace the entire function of the original tissues because of their inherently different functional parameters. Although engineering of two-dimensional (2D) and very small three-­ dimensional (3D) tissue structures has been achieved, construction of true 3D functional tissue structures, acting as organ systems, is still a big challenge. This is even more so if the organ consists of a multitude of different cell types, each having a particular function and influencing each other. Such engineered constructs depend on very complex cell carrier matrices, often combined with biological factors, such as morphogens and extracellular matrix (ECM) components. This supports the respective cell types by promoting vascularization and allowing adequate nutritional support—a key issue in large 3D constructs. In tissue engineering, biomaterials should act as artificial ECMs, and mimic biologic and mechanical functions of native ECM as much as possible. Biomaterials should also allow controlled delivery of bioactive factors (e.g., cell adhesion peptides, growth factors) and should offer a 3D space for the development of new tissues with appropriate structure and function (Kim and Mooney, 1998). In other words, the optimal biomaterial should be biocompatible, promote cellular interaction and tissue development, and possess proper mechanical and physical properties and act as a delivery vehicle to control the localization of transplanted cells (Brittberg et al., 1994). Three main types of biomaterials can be used for the engineering of bioartificial organs: naturally derived matrices (e.g., collagen and alginate), acellular tissue matrices (e.g., acellular bladder submucosa and small intestinal submucosa (SIS)), and synthetic polymers. Naturally derived materials and acellular tissue matrices have the advantage of being biologically recognized, but their production needs to be closely controlled and monitored to maintain good and reproducible quality. Synthetic polymers, however, can be reproducibly manufactured in large quantities with good quality and standardized physical properties. In many cases, a cell source is necessary to fully replace the organ’s metabolic, exocrine, or endocrine functions. The source of donor tissue can be xenogeneic

19.2  Urogenital Tissue Engineering

(from another species), allogeneic (same species, different individual), or autologous (same individual). Autologous cells are the preferred source as they avoid rejection from the host and preclude the use of immunosuppressive drugs. However, many patients with a specific organ disease no longer posses the specific healthy tissues or cells that should be used for culture expansion and subsequent transplantation. In these instances, donor cells or tissues are required to serve as an appropriate cell source. Another cell type with great potential applicable to many different organs could be the use of stem cells. It is now known that human embryonic stem cells (hESCs), for example, are capable of self-renewal, proliferating in an undifferentiated state while maintaining their pluripotent properties. We also know that we can differentiate these cells into almost any specialized human tissue cell type when providing the correct culture conditions. Researchers are making progress in directing these and other stem cells toward specific cell lineages with the goal of regenerating healthy organs that can replace diseased ones without depending on the availability of donor cells. According to Folkman and Hochberg (1973), engineered tissues exceeding a volume of 3 mm3 cannot survive adequately because of impaired nutrition and gas exchange. Vascularization of the regenerating cells is mandatory to engineer large complex tissues. Progress toward this goal is being achieved by incorporation of angiogenic growth factors into the engineered tissue, by seeding endothelial cells in combination with the other cell types (including stem cells) and by prevascularization prior to cell seeding. There is no doubt that the problem of tissue construct vascularization has to be solved before large entire solid organs can be regenerated.

19.2 UROGENITAL TISSUE ENGINEERING The urogenital system comprises the urinary tract and the internal and external genital apparatus. The urinary tract is subdivided into the upper tract (comprising the kidney and ureter) and the lower tract (comprising the bladder and urethra). In the female, the ovaries, the uterus, and the vagina belong to the internal genital apparatus. The external apparatus comprises the introitus, in the female the minor and major labia, and in the male the testicles and the penis, and, in particular, the corporal structures (Figure 19.1).

19.2.1 Bladder Intestinal and, less frequently, gastric segments are normally used as tissues for bladder replacement or repair. However, intestinal tissues are designed to absorb, and when they come into contact with the urinary tract, they also absorb toxic substances normally eliminated with the urine. Therefore,

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FIGURE 19.1 Schematic representation of the human urogenital system. (a) Vascular anatomy (b) Upper urinary tract.

multiple complications may ensue, such as infection, metabolic disturbances, increased mucus production, stone formation, and even malignancies (Kaefer et al., 1997). The use of gastric tissue may induce metabolic disturbances or even perforation. Owing to the problems encountered with the use of gastrointestinal segments, numerous investigators have attempted alternative reconstructive procedures for bladder replacement or repair, such as the use of seromuscular grafts, matrices for tissue regeneration, and tissue engineering with cell transplantation. Porcine SIS, a biodegradable, acellular, xenogeneic predominantly ­collagen-based tissue-matrix graft, was first used in the early 1980s as a matrix for tissue replacement in the vascular field. It has been shown to promote the regeneration of a variety of host tissues, including blood vessels and ligaments (Badylak et al., 1989). Animal experiments have shown that the noncell-seeded SIS matrix used for bladder augmentation is able to regenerate a bladder wall in vivo (Kropp et al., 2004). Allogeneic acellular bladder matrices can also serve as scaffolds for host cellular bladder wall components. The matrices are prepared by mechanically and chemically removing all cellular components from the donor bladder tissue a method called “decellularization” (Yoo et al., 1998). The ECMs serve as vehicles for partial bladder regeneration and relevant graft versus host reaction is not evident. In several studies using various materials as nonseeded grafts for surgical repair of the bladder, the urothelial layer was able to regenerate normally, but the muscle layer, although present, was not adequately developed (Kropp et al., 2004; Yoo et al., 1998). The grafts contracted to 60–70% of their original size

19.2  Urogenital Tissue Engineering

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FIGURE 19.2 Scanning electron microscopy of native equine collagen matrix (TissuFleece®) (a) and with human smooth muscle cells cultured for 10 days (b).

(Portis et al., 2000) with little increase in bladder capacity or compliance (Landman et al., 2004). Several studies of acellular matrices indicated that they might provide the necessary environment to promote cell migration, adherence, growth, and differentiation (Hubschmid et al., 2005; Danielsson et al., 2006); with continued research, these matrices may have a clinical role in bladder replacement in the future (Figure 19.2). In the bladder, the trigone is a smooth triangular region of the inner urinary bladder formed by the two ureteral orifices and the internal urethral orifice. Cell-seeded allogeneic acellular bladder matrices have been used for bladder augmentation in dogs. A group of experimental dogs underwent a trigone-­ sparing cystectomy and were randomly assigned to one of three groups. One group underwent closure of the trigone without a reconstructive procedure, another underwent reconstruction with a nonseeded bladder-shaped ­biodegradable scaffold, and the last underwent reconstruction using a bladdershaped biodegradable scaffold that delivered seeded autologous urothelial cells and smooth muscle cells (Oberpenning et al., 1999). The cystectomy-only and nonseeded controls maintained average capacities of 22% and 46% of preoperative values, respectively. However, in the cell-seeded tissue-engineered bladder replacements, an average bladder capacity of 95% of the original precystectomy volume was achieved. The compliance of the cell-seeded tissue-engineered bladders was also superior. Histologically, the nonseeded scaffold bladders presented a pattern of normal urothelial cells, however, with a thickened fibrotic submucosa and a thin layer of smooth muscle fibers. The retrieved tissue-engineered bladders showed a normal cellular

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organization, consisting of a trilayer of urothelium, submucosa, and smooth muscle (Oberpenning et al., 1999). Applying this technology, Atala et al. (2006) performed cystoplasty in seven patients with myelomeningocele with high-pressure, or poorly compliant bladders (State-of-the-Art Experiment). Researchers followed this study with a multicenter phase 2 clinical trial to evaluate bladder capacity, compliance, and safety in 10 pediatric patients with neurogenic bladder secondary to spina bifida (Joseph et al., 2014). Unfortunately, no statistically significant improvement in bladder capacity or compliance was observed, and four patients sustained significant adverse events including bowel obstruction and/or bladder rupture.

19.2.2 Urethra Bladder-derived acellular collagen matrix has also proven to be a suitable graft for repair of urethral defects, and newly created neourethras have shown normal urothelial lining and organized muscle bundles in a rabbit model (Chen et al., 1999). These results were confirmed clinically in a series of patients with a history of failed hypospadias repair wherein the urethral defects were repaired with human bladder acellular collagen matrices (Atala et al., 1999). The neourethral tissue was created by connecting the matrix to the urethral plate in an onlay fashion. The size of the created neourethra ranged from 5 to 15 cm. After a 3-year follow-up, three of the four patients had a successful outcome in regard to cosmetic appearance and function. One patient who had a 15-cm neourethra created developed a subglandular fistula. Similar results were obtained in pediatric and adult patients with primary urethral stricture disease using the same collagen matrices (El-Kassaby et al., 2003) (Figure 19.3). The described techniques, using nonseeded acellular matrices, have been applied experimentally and clinically with varying success for only urethral repairs by using SIS as a ventral onlay graft for small defects. However, when tubularized urethral repairs were attempted experimentally, adequate functional urethral tissue regeneration was not achieved, and complications, predominantly graft contracture, and stricture formation, were noticed (Le Roux, 2005). On the contrary, seeded tubularized collagen matrices have performed better in animal studies. In a rabbit model, urethroplasty was performed with tubularized collagen matrices either seeded with cells or without cells. The matrices seeded with autologous cells formed new tissue, which was histologically comparable to native urethra. However, the tubularized collagen matrices without cells led to limited tissue development, fibrosis, and stricture formation. Further, reliable regeneration and repair in subsequent studies could only be achieved at up to 0.5 cm (Dorin et al., 2008). Still, the ultimate goal remains to provide fully engineered tissues to patients requiring complete substitution urethroplasties, such as those with long strictures or significant penile trauma. Poor success was reported after initial

19.2  Urogenital Tissue Engineering

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FIGURE 19.3 (a) Representative case of a patient with a bulbar stricture. (b) During the urethral repair surgery, strictured tissue is excised; this preserves the urethral plate on the left side, and the matrix is anastomosed to the urethral plate in an onlay fashion on the right side. Both photographs indicate the area of interest, including the urethra, which appears white in the left photograph. The left photograph, indicates the area of stricture in the urethra. The right photograph indicates the repaired stricture. Note that the engineered tissue now obscures the native white urethral tissue in an onlay fashion in the right photograph. (c) Urethrogram 6 months after repair. (d) Cystoscopic view of the urethra before surgery on the left side, and 4 months after repair on the right side.

attempts in five patients with significant stricture secondary to lichen sclerosis. Keratinocytes and fibroblasts isolated from buccal mucosal biopsies were seeded on acellular dermal grafts used for complete substitution tubularized urethral replacement (Bhargava et al., 2008). Following this attempt, a study reported the successful implantation of bioengineered urethras into five boys suffering severe strictures (Raya-Rivera et al., 2011). Open bladder biopsies harvested 1-cm2 full-thickness tissues that were divided into urothelium and smooth muscle, and then cultured separately in the laboratory. These autologous tissues were then seeded onto the intraluminal and outer surfaces of a tubularized PGA acellular matrix for treatment of defects 4–6 cm in length. Of the patients, 100% demonstrated continence with normal flow rates and continence at a median follow-up of nearly 5 years. Further, cellularized implants did not exhibit severe fibrosis, inflammation, or stricture formation, and demonstrated histologically normal tissue on ­follow-up biopsies. Work in this area continues; recent success demonstrated that lengths of 6 cm were achievable in a canine model by using autologous bladder epithelium on which smooth muscle cells were seeded, and that

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selecting a combination of the appropriate cell type and biomaterial could lead to a successful clinical outcome (Orabi et al., 2013).

19.2.3 Kidney Kidney transplantation is superior to dialysis in treating patients with end-stage renal disease (ESRD). However, transplants expose patients to the risk of developing infection, rejection, and on-going costs for medical care. The current demand for kidney transplantation is also substantially greater than that of organ supply; this puts transplantation out of reach for many patients who require it. The alternative is the commonly used dialysis to replace absent renal function, but this is associated with a relatively high morbidity and mortality, and only partially replaces healthy kidney function. The kidney is surely the most challenging organ in the urinary system to be reconstructed by tissue engineering methods because of its complex structure and function (Figure 19.4). Unlike hollow or viscous organs, the kidney’s solid architecture presents the unique challenge of engineering several structural and functional capabilities into a single organ. The challenge is to include the normal kidney’s one million highly specialized nephrons working in concert to drain into a single renal pelvis. Two approaches have flourished in creating a bioartificial kidney: some researchers are pursuing the replacement of isolated kidney function parameters with the use of extracorporeal cellular units, while others are aiming at the replacement of total renal function by tissue-engineered bioartificial structures. Some investigators have sought alternative solutions involving ex vivo cellular systems, in an attempt to assess the viability and physiologic functionality of a ­cell-seeded device to replace the filtration, transport, and metabolic and endocrinologic functions of the kidney. The initial clinical experience suggests that

Cortex

Basement membrane Fibrous capsule Veseral epithelium

Minor calyces Renal sinus Major calyces

Afferent arteriole Smooth muscle Endothelium

Renal pelvis

Pseudofenestrations

Proximal tubule Distal convoluted tubule

Ureter Efferent arteriole

FIGURE 19.4 The complex renal structure. A cross-section of the whole kidney (left) and the glomerular apparatus (right).

19.3  Liver Tissue Engineering

renal tubule cell therapy may provide a dynamic and individualized treatment program as assessed by acute physiologic and biochemical indices (Humes et al., 2003). Two recent studies have demonstrated the possibility of kidney decellularization–recellularization technology for the creation of an implantable bioartificial organ. The first study demonstrated the ability to decellularize intact rat kidneys in a manner that preserved the intricate architecture and seeded them with pluripotent murine embryonic stem cells (ESs) antegrade through the artery or retrograde through the ureter (Ross et al., 2009). In this study, cells seeded onto the scaffolds demonstrated the immunohistochemical markers of ­differentiation including epithelization, while cells that did not contact the scaffold became apoptotic and formed a potential lumen. The second study in a rhesus monkey model demonstrated the capacity of the scaffold to support Pax2+/vimentin+ cell attachment and migration to recellularize the scaffold after layering with fetal kidney explants (Nakayama et al., 2010). Follow-up studies using scaffolds and cell explants from different age groups (fetal, juvenile, and adult) showed the extent of cellular repopulation was the greatest with scaffolds from the youngest donors, and with seeding of mixed fetal renal aggregates that formed tubular structures within the kidney scaffold (Nakayama et al., 2011). Even further investigation into both kidney and lung scaffolds using proteomic analysis and seeding with human embryonic stem cells (hESCs) suggests that decellularized scaffolds have an intrinsic spatial ability to influence hESC differentiation by physically shaping cells into tissue-appropriate structures and phenotypes (Nakayama et al., 2013). Future applications will include the evaluation of bioengineered kidneys in humans with ESRD. In addition, planning for a phase 1 clinical trial to evaluate safety and efficacy of a neokidney composed of autologous renal cells formulated in a gelatin-based hydrogel is underway.

19.3 LIVER TISSUE ENGINEERING The liver is a very complex organ that has the ability to regrow and regenerate (Figure 19.5). However, liver failure can still occur due to many different causes, including acquired liver disease, inborn errors of metabolism, or drug toxicity. Currently, when a liver fails, the lost function can only be replaced fully by liver transplantation. The options include transplantation of the whole liver or parts of the liver from cadaveric donors, portions of the liver from living donors, or xenotransplantation, typically porcine. Unfortunately, there are far too few livers available for the need worldwide, with approximately five times as many recipients as donors (Pruett, 2002), and xenotransplantation is limited by immune and infectious complications.

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Right triangular ligament

Sinusoids

Coronary ligament Left triangular ligament

Hepatocyte cords Central vein (systemic)

Lobule

Fibrous appendix of liver Left lobe Right lobe Falciform ligament Round ligament (ligamentum teres) of liver (obliterated umbilical vein) forming free border of falciform ligament Inferior border of liver

Portal vein branch Interlobular bile duct

Gall bladder (fundus)

Hepatic artery branch

Portal triad

FIGURE 19.5 Liver structure. (a) Surfaces and bed of liver (b) Hepatic architecture.

The simplest method of liver failure treatment using hepatocytes is cell transplantation. The first liver cell transplantation was described in 1979, with the successful transplant of viable rat hepatocytes into the spleen in an animal study (Mito et al., 1979). This and other studies taught us valuable lessons about the importance of the microenvironment to cell survival, as the cells required conditions resembling those of the liver, including a matrix to grow on and a venous blood supply similar physiologically to the hepatic sinusoid to grow and function normally. They also demonstrated the role of nonhepatocyte cells in hepatocyte survival, as coculture with nonparenchymal cells was often required to ensure cell viability (Selden and Hodgson, 2004). Unfortunately, clinical experience shows that many patients treated with hepatocytes for acute liver failure die of sepsis and most trials investigating the treatment of fulminate liver failure with hepatocyte transplantation have given equivocal results (Selden and Hodgson, 2004). For cells to be useful in transplantation, they must have a high proliferative capacity without loss of function and they must be economically feasible. There are several groups of possible cell sources for liver replacement. These include primary hepatocytes, immortalized cell lines, stem cells, and transdifferentiated cells (Selden and Hodgson, 2004). Primary hepatocytes are the most desirable and commonly used cells. Unfortunately, hepatocyte function is closely tied to the orientation, location, and surrounding environment of the cells. These environmental characteristics are lost in culture, so these cells

19.3  Liver Tissue Engineering

rapidly and irreversibly lose their function. Therefore, novel methods of cell culture, or alternative sources of cells, must be developed to obtain the necessary hepatocytes by perhaps mimicking the appropriate microenvironment (Selden and Hodgson, 2004). Human hepatocytes would be the best choice for cell transplantation, but just as liver donors are limited, so are human hepatocytes. Fetal or neonatal hepatocytes may also provide a source of cells. They have finite proliferation capacity, but may be immortalized (Selden and Hodgson, 2004). Another source of hepatocytes is stem cells. Stem cells have an unlimited replication potential and can differentiate into various different cell types. Potential sources of stem cells for hepatocyte generation include ES cells, adult liver progenitor cells, and transdifferentiated nonhepatic cells. ES cells offer great hope, but of course, there are ethical and legal concerns with ES cell use. Also, transdifferentiation of adult nonhepatic cells may also provide a novel cell source for liver replacement therapy (Baccarani et al., 2005). To this end, a small clinical trial investigating autologous transplantation of adult mesenchymal stem cells (MSCs) in patients with cirrhosis has demonstrated an improvement in liver and kidney functions (Kharaziha et al., 2009). As a practical treatment, however, this method has limitations. The small scale of this study (eight patients) and short-term follow-up did not allow for the analysis of a survival benefit. Further, preparation of harvested MSCs required greater than two months in culture, during which time many patients to potentially benefit may die. Another trial using hematopoietic stem cells in patients with viral or autoimmune hepatitis demonstrated similar results (Salama et al., 2010). Alternatives to human cells include primary porcine hepatocytes. These cells are plentiful and readily available, and have similar functional characteristics as those of human cells (Barshes et al., 2005). This strategy is not without its limitations, as there are possibilities of rejection and infectious complications such as xenozoonosis (e.g., porcine endogenous retroviral infection). To ensure a clinically safe method of transplantation of such cells, immunoprotective strategies are necessary. Cell encapsulation is one strategy to prevent rejection; it involves the encapsulation of single cells or groups of cells by using specific hydrogels, ­nanometer-thick coatings, or thin membranes of a nondegradable biomaterial. It does not prevent the development of antibodies to secreted proteins that cross the encapsulation barrier, but does prevent access of host immune cells to the transplanted cells. Systemic immunosuppressants are also useful if needed, but cotransplanting cells with Sertoli cells, or MSCs, can also provide additional immunosuppression locally, or at least dampen the immune response. Sertoli cells are interesting since they are central to the immune privileged environment in the testes. Therefore, coculture of cells with Sertoli cells may provide a level of protection from the host immune response. Preliminary

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experiments have been performed in which encapsulation of Sertoli cells and hepatocytes in microporous polypropylene, acetonitrile methallyl sulfonate copolymers, and alginate with poly-l-lysine demonstrated improvements in cell survival time and hepatic cellular function (Zheng et al., 2009).

19.3.1 The Bioartificial Liver Another strategy of liver replacement in the setting of liver failure is bioartificial liver (BAL) replacement. As an evolution of the artificial liver devices, BAL constructs were developed to harness the function of hepatocytes in an extracorporeal system. BAL devices have been under development for >40 years and usually consist of nearly 15 billion isolated hepatocytes with a membrane through which the plasma is circulated (Park and Lee, 2005). There are already several examples of BALs currently undergoing clinical trials, which include ELAD, HepatAssist®, BLSS, AMC-BAL, MELS, RFB, and HBAL/TECA-HALSS. Unfortunately, only two clinical trials that explored the effectiveness of BALs have been reported, and it is clear that many challenges remain before widespread dissemination for clinical use can be achieved (Zhao et al., 2012). The creation of a complete, implantable bioengineered liver remains the ultimate goal of tissue engineering for liver replacement. Many strategies have been explored as a means to achieve this goal, such as the injection of hepatocytes into vascular beds or the use of implantable seeded scaffolds. One of the main limitations of bioengineering large organs such as the liver is vascularization to provide cells embedded in such constructs with adequate nutrient supply ensuring a proper physiological response. Prior to transplantation, researchers must have a fully functional bioengineered liver, as it must immediately perform all hepatic functions. Vacanti’s group used a prevascularized sponge device that was fabricated and seeded with adequate cells in vitro, and then implanted in vivo. Another strategy is to use an implantable scaffold with direct access to the portal blood. One example used SIS grafted between the portal vein and inferior vena cava as a venous conduit, which could be seeded with hepatocytes. Microelectromechanical systems are another strategy, using silicon microfabrication technology to form large sheets of liver tissue, which can then be folded into 3D structures (Kulig and Vacanti, 2004). Animal studies have shown that the liver architecture remains viable after stripping and can be repopulated with hepatocytes. Hepatic functional data using this strategy was demonstrated for the first time in rats (Uygun et al., 2010). As early as 2011, Soker and Atala were able to generate a fully vascularized liver organoid after decellularization of livers across multiple species (Baptista et al., 2011). After rescue with human fetal liver and endothelial cells, these in vitro constructs were able to withstand fluid flow through the vascular network and generated significantly higher concentrations of urea and albumin than equivalent hepatic cells grown in culture dishes. Badylak’s group subsequently endorsed similar results, and provided encouragement that these methods may one day provide

19.4  Gastrointestinal Tissue Engineering

real therapeutic benefit to patients (Soto-Gutierrez et al., 2011). Other researchers have been able to successfully generated nonimmunogenic, large-scale scaffolds of clinically significant size from porcine livers (Mirmalek-Sani et al., 2013). As a conglomerate, these studies and others like them are ushering in the next phase of research in regenerating a viable liver construct for therapeutic use in humans by overcoming the challenges of oxygen and nutrient delivery, waste removal, and immunologic rejection. Researchers still must overcome challenges associated with establishing an ideal cell source, maintaining in vivo-like oxygen conditions, achieving bile secretory function, increasing convenience for the patient, and establishing efficacy of the devices overall.

19.4 GASTROINTESTINAL TISSUE ENGINEERING The current best therapy for replacing diseased or absent intestine is organ transplantation, aimed at restoring intestinal function for adequate nutrient absorption, fluid and electrolyte transport, and restoration of metabolic homeostasis. This involves an extensive operative procedure with its own morbidity and mortality, along with the well-described side effects of chronic systemic immunosuppression. There are also the concerns of inadequate numbers of organ donors and rejection of the transplanted tissue. An excellent alternative to current therapies would be to regenerate intestine using tissue engineering techniques that have been successfully applied to other organs (Chen and Badylak, 2001; Chen and Beierle, 2004). Early attempts at tissue engineering bowel segments involved patch repairs of intestinal defects. As far back as Kobold and Thal (1963), who used jejunal serosa to repair duodenal defects in dogs. These patches were covered by duodenal epithelium in 2 months. Binnington et al. (1973) repaired jejunal defects in animals with the serosa of the descending colon. The new mucosal covering was noted to grow in from the edges, which covered the serosa. Expanding their work, they used a model of “short bowel syndrome” in pigs to show that the jejunal patch was able to induce increased weight gain (Chen and Badylak, 2001). These “patch” experiments were repeated using different models and strategies, such as colonic serosa or abdominal wall pedicle flap patches in rabbits, with similar results (Chen and Beierle, 2004). The studies concluded that these patches were all able to support the growth of functional neomucosa.

19.4.1 The Intestinal Mucosa All these studies share the need for intestinal mucosa to grow and spread across the patch surface. Therefore, the regulation of intestinal growth is an area of research that is crucial to the advancement of bowel tissue engineering. Numerous biomolecules have been studied and proven to have an effect on intestinal mucosal growth. Basic fibroblast growth factor (bFGF) delivered to

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smooth muscle in a mouse model was shown to maintain the viability of small intestinal implanted sphincteric smooth muscle (Orlando et al., 2012). Similar results were reported with the use of platelet-derived growth factor as applied to bioengineered internal anal sphincter constructs (Miyasaka et al., 2011). Early work involving biomaterials focused on synthetic scaffolds such as Dacron and PTFE to repair intestinal defects in animals. However, these strategies suffered from extrusion into the bowel wall and incomplete neomucosal coverage (Harmon et al., 1979; Watson et al., 1980). The studies demonstrated that selecting the ideal biomaterial is important and that biocompatibility and biodegradability are crucial for a successful in vivo outcome.

19.4.2 Natural Biomaterials for Intestinal Repair Acellular and natural scaffolds have recently been found to be more appropriate for bowel repair. An example of a widely used acellular biomaterial is SIS. SIS is primarily composed of ECM material, following the removal of the mucosa, external muscle layers, and other cells from the small intestine of pigs. The naturally occurring ECM is rich in components that support angiogenesis (i.e., fibronectin; glycosaminoglycans; collagen types I, III, IV, and V; bFGF; and vascular endothelial growth factor). SIS has also been shown to be nonimmunogenic (Demirbilek et al., 2003). SIS has been extensively investigated as a scaffold for in numerous tissue engineering applications, and is now processed in a standardized fashion in mass quantities; this makes it readily available for tissue engineering studies (Chen and Badylak, 2001). SIS has been used to cover defects in the esophagus and small bowel in dogs. In these studies, the SIS patch allowed the mucosa to grow over the scaffold, and the histology of the patch was similar to the native intestine (Chen and Badylak, 2001). In another study, tubular noncell containing SIS grafts from Sprague–Dawley rat donors were used to create an ileostomy (Figures 19.6). These grafts were initially covered by inflammatory cells, but by 3 months, they were completely covered by mucosa. Also, the outsides of the tubular grafts were covered with bundles of smooth muscle-like cells and fibrovascular tissue. A similar study performed in the stomach showed similar results. In this study, rats underwent the removal of a full-thickness 1-cm segment of the stomach, with a patch repair using a double-layer patch of porcine-derived SIS. Three weeks postoperatively, the site was covered by granulation tissue and fibrosis, with some growth of normal gastric mucosa at the edges of the graft (de la Fuente et al., 2003). These studies demonstrated that SIS can develop structural features of the normal intestine and may provide an off-the-shelf option for replacement patches for intestinal tissue (Wang et al., 2005). Other studies using various other novel scaffold materials, such as fibrin glue and PGA as a framework for tissue engineering neointestine, have had promising results as well. Hori et al. (2001) used a collagen sponge enzymatically

19.4  Gastrointestinal Tissue Engineering

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FIGURE 19.6 Histologic photomicrographs of SIS-regenerated rat small bowel tissue at 12 weeks. (a) Well-organized mucosal epithelial layer with smooth muscle cells regeneration. (b) Immunohistochemically stained intestinal tissue suggesting bundles of well-formed smooth muscle fibers with less regularity.

processed from porcine skin as a scaffold for regeneration of stomach and small bowel with good success in a beagle model. Isch et al. (2001) have also performed esophagoplasty by patching defects in the cervical esophagus in a dog with AlloDerm® (human decellularized skin).

19.4.3 Combining Biomaterials with Cells for Intestinal Repair The key components for this process include a cell source, and an appropriate biocompatible and biodegradable scaffold. Early work in intestinal tissue engineering by Vacanti used just such a strategy. Cells were obtained from an intestinal organoid unit; a scaffold of sheets of collagen-coated, microporous, nonwoven PGA was formed into a tubular shape, and these tubes once seeded with cells were implanted into the omenta of animals (Vacanti, 2003). Vacanti’s group harvested cells from intestinal organoids, seeded these on the constructs, and after subsequent implantation, they found it decreased morbidity associated with bowel resections in rats (Grikscheit et al., 2004). Before this study was conducted, Tait et al. (1994) had already described these organoids to be excellent for intestinal tissue engineering because of their unique composition, which includes progenitor cells, epithelium, and mesenchymal stroma. What was very interesting to see is that the implanted devices demonstrated angiogenesis from the omental vessels, crucial for the long-term survival of the constructs. After additional examination in animal studies, the constructs were found to be quite similar to normal small bowel and were anastomosed to the native bowel without causing feeding problems (Chen and Beierle, 2004; Rocha and Whang, 2004). More recently, smooth muscle cells on PGA sheets, a synthetic approach, in combination with fibroblastic keratinocytes have been

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used to successfully replace portions of the esophagus (Nakase et al., 2008). However, such constructs have not demonstrated physiologic functionality; this shows again that a proper combination of cells and biomaterials is critical for a good outcome. Although these implants have now demonstrated basic metabolic function, one important aspect is still missing, namely, integration with the enteric nervous system that is crucial for a normal bowel motility. There are however indications that nerve innervation, or at least the stimulation of nerve ingrowth, is possible by using the right cell type. Notable success has been achieved with the implantation and subsequent neovascularization of bioengineered internal anal sphincters in murine models. Implanted constructs have demonstrated maintenance of basal tone and sphincteric relaxation similar to native sphincters (Raghavan et al., 2011). The group described their method to bioengineer functionally innervated gastrointestinal smooth muscle constructs using neuronal progenitor cells and smooth muscle cells that were isolated and cultured from intestinal tissues of adult human donors (Gilmont et al., 2014). This breakthrough in bioengineering functional gastrointestinal tissue in vitro provides not only a useful construct for the examination of the disease processes of the bowel but also a possible “off-the-shelf” source of intestinal tissues with physiological and metabolic functions for clinical use.

19.5 PANCREAS TISSUE ENGINEERING Diabetes mellitus is a common disease, which affects 4% of the world, and causes significant morbidity and mortality through its effects on vascular/microvascular disease. Diabetes is due to systemic hyperglycemia, resulting from inadequate insulin production by the pancreas and/or resistance to insulin (Kahn, 2004). The insulin-producing β cells can be found in the pancreatic islets, or islets of Langerhans. The pancreas is composed of acinar cells, ducts, and islets of endocrine cells. Pancreatic cells develop from endodermal progenitor cells. The islets contain a specific set of endocrine cells comprising the β cells, α cells, which produce glucagon, δ cells, which produce somatostatin, ε cells, which produce ghrelin, and PP cells, which produce pancreatic polypeptide. These endocrine cells all work in unison and stimulate, or inhibit, each other through paracrine effects. Although the islets only take up a very small portion (1–2%) of the total pancreas volume, they receive 10–15% of the blood flow that increases when blood glucose levels are elevated. Each islet therefore is densely vascularized, and islets cells are never further away than 100 μm from a blood vessel to provide them with sufficient amounts of nutrients and oxygen and to efficiently release insulin into the bloodstream (Figure 19.7). There are two main types (a third has been described) of diabetes, types I and II, and they are caused by islet β cell dysfunction, with an absolute or relative

19.5  Pancreas Tissue Engineering

Inferior vena cava Portal vein Common bile duct

Stomach Spleen Aorta

Pancreas body Right kidney

Colon

Duodenum Left kidney Jejunum

Root of mesentery

Common bile duct

FIGURE 19.7 The pancreas anatomy.

decrease in insulin production. Type I diabetes results from the autoimmune destruction of β islet cells. Type II diabetes is characterized by systemic insulin resistance, an inability of the β cell to rapidly release sufficient insulin in response to oral nutrient stimulation, and to impaired glucose tolerance. Also, β cells may not be able to convert proinsulin, which is the precursor in the cell for forming insulin (Kahn, 2004). The most effective therapy for replacing lost β cell function, since the discovery of insulin by Banting and Best in 1921, has been insulin replacement therapy (Bliss, 1982). The ability to monitor blood glucose levels and provide patients with insulin on a daily basis multiple times a day works for many patients today. However, it is still very difficult to match insulin levels appropriately to the blood glucose concentration in real time. Scientists therefore continue to search for new therapies to improve the treatment of diabetes mellitus, such as β cell replacement, which seeks to replace the deficient endogenous insulin production especially for type I diabetes patients. Designing of new β cells involves two approaches: gene therapy and cell replacement strategies. Engineered β cells need to be able to modulate basal insulin release for normal events such as exercise or infection, and respond

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to continuous changing blood glucose levels. Keep in mind that the islets of Langerhans are innervated by the sympathetic, parasympathetic, and sensory nerve fibers, so that engineered β cells need to respond to neural input. Over life, there are changing demands on insulin release that are the result of changes in insulin sensitivity, such as in puberty, pregnancy, and old age. There is quite a challenging list of criteria an in vitro engineered β cell should obey. The cells should be able to increase their replicative capacity and involute when no longer needed. Cells should also produce insulin and appropriately process proinsulin to insulin to function properly, which is highly dependent on adequate oxygen supply. Preferably, there has to be a way to recreate the influence of gastrointestinal derived factors in the modulation of insulin release (Kahn, 2004). β Cells normally release the hormone insulin in a pulsatile fashion, so the newly generated β cells should mimic this behavior. A major concern for scientists in β cell engineering is the recurrence of disease. Type I diabetes has an immunological cause, so it may recur with in vitro differentiated cells, if they express the antigens leading to an autoimmune reaction. The causes of type II diabetes are not as clear, so recurrence could be a possibility here as well (Kahn, 2004). Therefore, in vitro generated β cells may require immunosuppression, and these agents (specifically glucocorticoids) can cause insulin resistance, suppress insulin release, and prevent insulin processing. They also have other more severe risks, such as malignancy. Also, proinsulin may pose a cardiovascular risk, so proper processing of the hormones is essential (Kahn, 2004).

19.5.1 Creating New β Cells for Cell Therapy in Type I Diabetes Based on the aforementioned challenging checklist, the cells developed must have certain physiologic functions, such as insulin gene expression, appropriate posttranslational modification of insulin, glucose-sensitive insulin production, and, ideally, be available in an unlimited supply to be used as cell therapy for diabetes mellitus (Kahn, 2004). All these characteristics are typical of β cells, but this does not mean that one needs β cells to replace them; just cells that act like them would suffice. For example, modification of genes in cells that have similar origins to pancreatic cells may produce β cell-like qualities, such as the transfer of the Pdx-1 gene into mouse liver cells. These modified liver cells can activate gene expressing of insulin 1 and 2 and the prohormone convertase, but unfortunately are not sensitive to glucose in the environment. It has also been demonstrated that ductal cells also, which comprise the ducts in the pancreas through which digestive enzymes are secreted, can be modified to produce insulin if modified with gene transfection (Kahn, 2004). Zhou et al. have demonstrated the ability to reprogram fully differentiated adult exocrine cells into insulin secreting β-like cells (Zhou et al., 2008). In this study, developmental transcription factors including Pdx-1, Ngn3, and MafA were delivered

19.5  Pancreas Tissue Engineering

using a viral vector via injection into the pancreas, and over time, acinar cells began to exhibit an endocrine rather than exocrine phenotype. An alternative approach is to start with cells that already have some β cell characteristics and to modify these as needed to obtain the desired function. For example, Cheung et al. modified intestinal K cells, which normally secrete peptides in response to insulin secretion, using a transgenic model to express insulin in response to glucose sensation and these cells were able to restore normoglycemia to streptozotocin-induced diabetic mice (Cheung et al., 2000) (Figure 19.8). But what about a stem cell approach for replacement of exogenous insulin therapy? The goal of stem cell therapy, as with all studies of multipotent cells in regenerative medicine, is the directed differentiation of these cells and the (a) Control, STZ Transgenic, STZ Control

Glucose (mM)

40 30 20 10 0

20

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Control

Control

+STZ

FIGURE 19.8 Production of human insulin from K cells protects transgenic mice from diabetes induced by destruction of pancreatic cells. (a) Oral glucose tolerance tests. Mice were given intraperitoneal injection of streptozotocin (STZ, 200 mg/kg) or an equal volume of saline. On the fifth day after treatment, after overnight food deprivation, glucose (1.5 g/kg body weight) was administered orally by feeding tube at 0 min. Results are means (6SEM) from at least three animals in each group. (b) Immunohistochemical staining for mouse insulin in pancreatic sections from control mice and an STZ-treated transgenic mouse. Arrows indicate mouse islets.

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introduction of specific characteristics to these cells via genetic manipulation. There are several different choices of cells with which to work in this field of study, such as ES cells and adult stem cells. However, it is still somewhat unclear if there is a pool of adult stem cells in the pancreas and whether the duct cells can act as islet progenitors, or even if β cells can renew themselves (Kahn, 2004). But perhaps an answer will be found in the near future; a study by Seaberg et al. (2004) described the isolation of pancreas-derived multipotent precursor cells. They isolated these cells by dissociating and plating pancreatic ductal and islet tissue. These tissues formed floating colonies typical of neurospheres. These colonies expressed markers such as PDX-1 and Nestin, characteristic of both pancreatic and neural precursors (Seaberg et al., 2004). It was discovered that these cells could not only develop into neurons, astrocytes, and oligodendrocytes but they also produce insulin-producing β-like pancreatic cells. They were found to be glucose responsive, with increased insulin release in response to insulin. These findings are very encouraging, but require considerably more research to prove clinical utility. Is there a role for stem cells in diabetes cell therapy? Perhaps, there has been some recent success in exploring strategies using ES cells and other stem cell types that can now be fully differentiated into cells with functional characteristics of native β cells. For example, some researchers have been using precursors, such as adipose-derived stem cells, capable of glucose regulation in mice (Chandra et al., 2009). Yet, it has been considerably challenging to achieve a full β cell phenotype in vitro, and multiple studies have suggested the transplant site milieu to be an important factor necessary for terminal differentiation, including the effect of in situ signals (Orlando et al., 2012). Notable success was achieved when multipotent cells of both adult progenitor (bone marrow) and blastocyst (ES) origin showed β-like potential (Kumar et al., 2013). These rat stem cells were cultured and differentiated ex vivo into β-like cells. On day 21, 20% of progeny expressed Pdx-1 and C-peptide and secreted C-peptide under the stimulus of insulin agonist carbachol. When implanted into diabetic nude mice for 4 weeks, hypoglycemia was reversed in roughly half of the subjects with hyperglycemic recurrence within 24 h of graft removal, which showed that at least to a certain extent differentiated stem cells could be used to treat type 1 diabetes. Similar to tissue engineering of other organ systems, there is much interest in using scaffolds and/or tissue constructs to regenerate a 3D organ that mimics the natural environment of the native endocrine pancreas compartment. It may even prove necessary for efficient terminal differentiation (Wang and Ye, 2009). Multiple groups have explored this option in rodents with some success, including the constitution of viable vascular networks and insulin secretion for several weeks (Orlando et al., 2012). Although relatively new in the field of tissue engineering, there is clearly an increasing trend toward using regenerative medicine- and bioengineering-based strategies for the creation of

19.6  Lung Tissue Engineering

a bioartificial organ that mimics the endocrine function of the native pancreas (Forster et al., 2011; Dufour et al., 2005; Buitinga et al., 2013; Yap et al., 2013; Pedraza et al., 2012; Gibly et al., 2011; Blomeier et al., 2006).

19.6 LUNG TISSUE ENGINEERING As in other fields currently being investigated by tissue engineers, organ transplantation is the preferred treatment and standard of care for organ failure for the lung. The lungs might seem to comprise a relative simple anatomy, but in fact contain an elaborate network of branching bronchi, and ultimately end in very small alveoli that are covered by a thin layer of squamous epithelium that permits diffusion of oxygen into the dense microvascular network lining the epithelium on the inside. The trachea is lined by rings of cartilage to support the tubular structure that otherwise would collapse on itself (Figure 19.9). The demand for human organs for transplantation is large and growing, and the supply of donor lungs (as with most transplanted organs) does not meet the demand. Developing technologies for lung replacement include stem cell biology, therapeutic cloning, tissue and organ bioengineering, and ­xenotransplantation (Ogle et al., 2004).

(a)

(b) Apex

Smooth muscle Alveolus

Eparterial bronchus Pulmonary artery Pulmonary veins Cardiac impression Pulmonary ligament

Right lung

FIGURE 19.9 Anatomy of the lung.

Pores of Kohn

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Studies have demonstrated that ES cells can be induced to differentiate in vitro into pulmonary epithelial cells, type II pneumocytes, using a specific medium designed for the maintenance of lung epithelial cells in culture (Rippon et al., 2004). The generation of differentiated airway epithelial tissue had not been reported, however, until Coraux et al. also demonstrated that type II alveolar epithelial cells could be derived from ES cells in a murine model (Coraux et al., 2005). They went on to show that murine ES cells were able to differentiate into nonciliated secretory Clara cells and that type I collagen induced this development. In addition, when these cells were cultured at an air–liquid interface, ES cells were able to form fully differentiated airway epithelium. The cells were composed of basal, ciliated, intermediate, and Clara cells, such as found in the tracheobronchial airway epithelium (Coraux et al., 2005). Another method of inducing the development of ES cells along the pulmonary epithelial lineage is the use of coculture techniques. Coculture of ES cells from a murine model with embryonic mesenchyme from the distal lung promoted the development of pneumocytes. The technique involved differentiating the stem cells first into embryoid bodies (EBs) and then culturing these EBs for one to two weeks with embryonic pulmonary mesenchyme. The EBs demonstrated the development of channels lined with pulmonary epithelial cells including type II pneumocytes. Even more exciting was the observation in some animal studies that ES cells delivered systemically to animals with pulmonary damage seeded at the site of the diseased tissue and developed pulmonary phenotypes. These results encourage the search for an ideal use of ES cells as some form of cell therapy for pulmonary disease (Van Vranken et al., 2005). The final goal of all this research is the ability to regenerate portions of the lung for replacement of damaged tissue. Clinically, the experience of this scale of bioengineering, while limited, is progressing. In one reported case, a bioengineered airway patch was created for a patient who developed a dehiscence of his tracheobronchial anastomosis following a right completion carinal pneumonectomy. The patch was created by harvesting autologous muscle cells and fibroblasts from the patient, culturing and expanding these cells, and plating them on a collagen matrix derived from decellularized porcine jejunum. The patch was placed over an airway defect, satisfactorily closed the opening, and was immediately airtight. The graft demonstrated neovascularization and excellent coverage with functional and viable ciliated respiratory epithelium. The cellularized patch worked well in this instance, although it is possible that an acellular patch could be used, as there was ingrowth of the normal respiratory epithelium over the patch (Macchiarini et al., 2004) (Figure 19.10). The above-mentioned procedure has been improved over the last decade, and the investigators have extended their reach to multiple patients. They have shortened the time to obtain tracheal decellularization by using an alternative cell technological approach (Bader and Macchiarini, 2010). The improved approach involves the decellularized human tracheal graft being reseeded

19.7  Future Developments

(1)

(2) Human Biopsy

Transplantation of the tissue-engineered patch

(50 ml blood & muscle cells)

Cellular isolation

(a)

(b)

(c)

(d)

Human cells are seeded on the matrix

Isolation of porcine jejunum segment Decellularization of the matrix

FIGURE 19.10 Engineering process (1) and histologic composition (2) of the bioartificial patch. Muscle cells and fibroblasts are isolated from a biopsy specimen obtained from the patient. These cells are seeded on a biologic matrix representing a collagen network generated from a decellularized porcine jejunal segment. During the incubation period, the cells start to remodel the xenogeneic matrix and replace it with autologous connective tissue. Within 4 weeks, this autologous bioartificial implant can be clinically used. The computed tomographic scan of the chest 6 weeks after graft implantation shows the site (arrow) where the tissue-engineered patch was transplanted. (a) Acellular porcine collagen matrix (hematoxylin and eosin staining). Notice the absence of cellular components and collagen fibers. (b) Reseeded patch before implantation. Antivimentin immunohistology depicting fibroblast and smooth muscle cell seeding. (c) Brush cytology from the patch surface showing the respiratory epithelium (12th week). The arrows indicate epithelial cilia. (d) Immunohistochemical staining for respiratory epithelium (12th week) (Villin antibody).

intraoperatively with autologous cells (bone marrow MSCs and respiratory cells) and conditioned with growing, differentiation, and “boosting” factors. This in vivo tissue-engineered approach was used in five patients with benign tracheal diseases and in two patients with primary tracheal cancers involving the entire trachea (Kalathur et al., 2010). The aforementioned therapy has improved not only lung function for these patients many months after surgery but also the overall quality of life.

19.7 FUTURE DEVELOPMENTS Today’s strategy for tissue engineering predominantly depends upon autologous cells from the diseased organ of the patient. In the case of extensive end-stage organ failure, a tissue biopsy may not yield enough normal cells for culture and ultimate transplantation. Primary autologous human cells cannot always be expanded from a particular organ, as is the case with the pancreas. In these cases, pluripotent human ES cells could become a source of cells from which the desired tissue can be derived. Embryonic stem cells exhibit two remarkable properties: the

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ability to proliferate in an undifferentiated, but pluripotent state and the ability to differentiate into many specialized cell types. Although these cells have great therapeutic potential, their use is currently limited by biological and ethical factors alike. ES cells are an area of continued fertile research and one day may prove to be an extremely powerful cell source for tissue engineering. Adult stem cells are ethically more acceptable and have the advantage of not transdifferentiating spontaneously into a malignant phenotype. Because some adult stem cells are notoriously difficult to expand into large quantities, their use for clinical therapy in these cases is limited. A relatively recent development, induced pluripotent stem cells (iPSCs) represent an exciting alternative to both ES cells and adult stem cells. These cells are derived from adult cells, but can be reprogrammed via genetic engineering to exhibit the pluripotency that was at one time unique to ES cells. For this reason, iPSCs are considered more ethically acceptable than ES cells and hold potential to overcome many of their biological limitations, including those associated with immunogenicity. The development of iPSCs for the use in ­tissue engineering is an area of fertile research. Fetal stem cells, harvested from amniotic fluid and placentas, may represent another novel source of stem cells. These cells express markers consistent with human ES cells, such as octamer-binding transcription factor 4 and stage-specific embryonic antigen 4. They are multipotent and are able to differentiate into all three germ layers, and they do not form teratomas, nor do they spontaneously transdifferentiate into a malignant phenotype. A further advantage is that the cells have a high replicative potential and they could be stored for future possible use in the patient, without the risks of rejection and ethical concerns (Siddiqui and Atala, 2004). Further, growth factors and other signaling molecules alone or in combination with biological or synthetic matrices may be used for cell proliferation and differentiation, tissue guidance, and regeneration, and thus replace cell-based tissue engineering. Once the growth factors to be delivered have been identified, there are several important criteria to take into consideration. First, the delivery must be controlled, and allow sustained and localized delivery of small amounts of growth factors to the tissue repair site, as systemic administration can cause toxicity in other organs. Further, the release profile must be controlled in ensuring frequent exposure to these factors for a relatively long time to obtain the desired effect. Growth factors typically have short half-lives once they are introduced into the body and are eliminated rapidly. Finally, the process of incorporating the factors must not involve harsh solvents or high temperatures that might denature or deactivate the proteins. To meet these criteria, the growth factors could be entrapped within or covalently linked to a polymer matrix and released by matrix degradation. As the prenatal diagnosis of fetal abnormalities has become extremely accurate, theoretically, a small tissue biopsy could be obtained under ultrasound

19.8 Summary

guidance. These tissues could then be processed and the various cell types could be cultured in vitro. Perhaps an interesting approach to prevent the onset of certain life-long diseases at a very early age would be to apply tissue engineering strategies, to engineer in vitro reconstituted structures of tissues, or organs, which could then be available for reconstruction or transplantation at the time of birth.

19.8 SUMMARY The engineering of simple hollow organs (e.g., the bladder) has been achieved, but the construction of more complex hollow organs (e.g., blood vessels) and functional solid organs (e.g., kidney, liver) is the main challenge facing the future of tissue engineering. n Engineered bladder tissues, created with autologous cells seeded on collagen–PGA scaffolds and wrapped in omentum after implantation, are currently undergoing clinical trials in patients requiring bladder augmentation. n Today’s strategy for tissue engineering predominantly depends upon autologous cells from the diseased organ of the patient. In the case of extensive end-stage organ failure, a tissue biopsy may not yield enough normal cells for culture and ultimate transplantation. n Primary autologous human cells cannot always be expanded from a particular organ, as is the case with the pancreas. In these cases, pluripotent human ES cells can become a source of cells from which the desired tissue can be derived. n Nonseeded acellular matrices have functioned well clinically for onlay urethral repairs. However, when tubularized urethral repairs are performed, seeded tubularized collagen matrices are required. n The kidney is surely the most challenging organ in the urinary system to be reconstructed by tissue engineering methods because of its complex structure and function. To develop a bioartificial kidney, some researchers are pursuing the replacement of isolated kidney function parameters with the use of extracorporeal cellular units, while others are aiming at the replacement of total renal function by tissue-engineered bioartificial structures. n As in renal replacement strategies, the complexity of liver tissue engineering has driven some effort toward liver replacement therapies with BAL replacement, but so far, human studies have been equivocal. New advances in decellularization technology are moving researchers closer to viable in vitro constructs. n  Cellularized, and possibly acellular, segments of collagen scaffold material offer a possible source of tissue for patch repairs of the tracheobronchial tree. n

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Studies into gastrointestinal tissue engineering have demonstrated that SIS, AlloDerm®, and other unseeded scaffolds can develop structural features of the normal intestine when implanted in the bowel, and may provide an off-the-shelf option for replacement patches for intestinal tissue. Additional studies have shown that cells obtained from intestines, seeded on a scaffold, and formed into a tubular shape developed tissues similar to those in the normal small bowel and have been successfully anastomosed to the native bowel in animal studies. More recent studies have demonstrated promising advances in enteric neuronal function, including internal anal sphincter constructs. n The most effective therapy for diabetes, since the discovery of insulin by Banting and Best in 1921, has been insulin replacement therapy, but this therapy does not match insulin levels appropriately to the blood glucose concentration. The ultimate goal, therefore, is to develop a permanent endogenous replacement for the lack of insulin, and many tissue engineering strategies have been directed toward this goal. n

CLASSICAL EXPERIMENT After several reports in the past about successful culture of animal and human urothelial and bladder smooth muscle cells, it was only in the 1990s that Southgate et al. (1994) characterized in detail cultured human urothelial cells and were able to induce urothelial stratification in vitro. They summarized their research as follows: The purpose of the work was to establish urothelium as an in vitro model for the study of proliferation, stratification, and differentiation in “complex” epithelia. Normal human urothelial cells were cultured in a serum-free medium. The effects of epidermal growth factor (EGF), cholera toxin (CT), extracellular calcium, and 13-cisretinoic acid on cell growth, morphology, phenotype, and cytodifferentiation were studied using phasecontrast microscopy and indirect immunofluorescence. Stratification-related changes were additionally analyzed by transmission electron microscopy. Under optimized conditions, long-term cultures were successful in 44 (74.5%) out of 59 specimens. Bacterial infection was the most common cause of failure (nine cases). Primary urothelial cells required an initial plating density of 1 × 104 cells/cm2 for survival; passaged cells survived much lower plating densities (2.53 × 102 cells/cm2).

CT significantly improved cell attachment, but neither CT nor EGF was essential for growth. By contrast, cells failed to proliferate without bovine pituitary extract. In media containing bovine pituitary extract, CT, and EGF, cultures had a mean population doubling time of 14.7 6 1.8 h, maintained a nonstratified phenotype, and expressed the cytokeratin (CK) profile of basal/ intermediate urothelium: CK7, CK8, CK17, CK18, and CK19, with variable expression of CK13. CK20 was not expressed in vitro. CK14 and CK16 were also expressed, suggestive of squamous metaplasia in culture, which could be inhibited with 13-cisretinoic acid. Increasing extracellular calcium from 0.09 to 0.9–4.0 mM slowed down cell proliferation, induced stratification and desmosome formation, and increased expression of E-cadherin. High calcium, EGF, CT, and retinoic acid did not induce markers of late/terminal urothelial cytodifferentiation. The above-mentioned study describes a simplified technique for the isolation and long-term culture of human urothelial cells. Urothelial cells in vitro are capable of rapid proliferation and can be induced to form integrated stratifying cell layers in high calcium medium.

19.8 Summary

STATE-OF-THE-ART EXPERIMENT In this study, transplantable urinary bladder neoorgans were reproducibly created in vitro from urothelial and smooth muscle cells grown in culture from canine native bladder biopsies and seeded onto preformed bladder-shaped polymers. The native bladders were subsequently excised from canine donors and replaced with the tissue-engineered neoorgans. In functional evaluations for up to 11 months, the bladder neoorgans demonstrated a normal capacity to retain urine, normal elastic properties, and histological architecture. This study demonstrated, for the first time, that successful reconstitution of an autonomous hollow organ is possible using tissue engineering methods (Oberpenning et al., 1999). Patients with end-stage bladder disease can be treated with cystoplasty using gastrointestinal segments. The presence of such segments in the urinary tract has been associated with many complications. The authors explored an alternative approach using autologous engineered bladder tissues for reconstruction.

Seven patients with myelomeningocele, a birth defect in which the backbone and spinal canal do not close before birth leading to loss of bladder or bowel control, aged 4–19 years with high-pressure, or poorly compliant bladders, were identified as candidates for bladder repair. A bladder biopsy was obtained

(a)

(b)

from each patient. Urothelial and muscle cells were grown in culture, and seeded on a biodegradable bladder-shaped scaffold made of collagen, or a composite of collagen and polyglycolic acid. About 7 weeks after the biopsy, the autologous engineered bladder constructs were used for reconstruction and implanted either with, or without an omental wrap. Serial urodynamics, cystograms, ultrasounds, bladder biopsies, and serum analyses were done. Follow-up range was 22–61 months with a mean of 46 months. Postoperatively, the mean bladder leak point pressure decrease at capacity, the volume, and compliance increase were the greatest in the composite engineered bladders with an omental wrap. Bowel function returned promptly after surgery. No metabolic consequences were noted, urinary calculi did not form, mucus production was normal, and renal function was preserved. The engineered bladder biopsies showed an adequate structural architecture and phenotype. This study showed that engineered bladder tissues, created with autologous cells seeded on collagen–polyglycolic acid scaffolds, and wrapped in omentum after implantation, can be used in patients who need cystoplasty (Atala et al., 2006) (Figures 19.11 and 19.12).

(c)

FIGURE 19.11 Radiographic cystograms 11 months after subtotal cystectomy. (a) Subtotal cystectomy without reconstruction (group A), (b) polymer-only implant (group B), and (c) tissue-engineered neoorgan (group C). Continued...

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STATE-OF-THE-ART EXPERIMENT Continued (a)

(b)

(c)

(d)

(e)

(f)

(g)

FIGURE 19.12 Histological and immunochemical analyses of implants 6 months after surgery. Hematoxylin and eosin histologic results are shown for (a) normal canine bladder (group A), (b) the bladder dome of the cell-free polymer reconstructed bladder (group B), and (c) the tissueengineered neoorgan (group C). Immunocytochemical staining of tissue-engineered neoorgan revealed; (d) positive staining of the epithelial cell layers with pancytokeratin AE1/AE3 antibodies; (e) urothelial differentiation-related membrane proteins, which constitute the apical plaques of the asymmetric unit membrane in normal urothelium; (f) anti-α smooth muscle actin staining of phenotypically normal smooth muscle; and (g) positive staining with S-100 antibodies binding to neural tissue and nerve sheaths.

References

REFERENCES Atala, A., et al., 1999. A novel inert collagen matrix for hypospadias repair. J. Urol. 162, 1148–1151. Atala, A., et al., 2006. Tissue-engineered autologous bladders for patients needing cystoplasty. ­Lancet 367, 1241–1246. Baccarani, U., et al., 2005. Human hepatocyte transplantation for acute liver failure: state of the art and analysis of cell sources. Transpl. Proc. 37, 2702–2704. Bader, A., Macchiarini, P., 2010. Moving towards in situ tracheal regeneration: the bionic tissue engineered transplantation approach. J. Cell. Mol. Med. 14, 1877–1889. Badylak, S.F., et al., 1989. Small intestinal submucosa as a large diameter vascular graft in the dog. J. Surg. Res. 47, 74–80. Baptista, P.M., et al., 2011. The use of whole organ decellularization for the generation of a vascularized liver organoid. Hepatology 53, 604–617. Barshes, N.R., et al., 2005. Support for the acutely failing liver: a comprehensive review of historic and contemporary strategies. J. Am. Coll. Surg. 201, 458–476. Bhargava, S., et al., 2008. Tissue-engineered buccal mucosa urethroplasty-clinical outcomes. Eur. Urol. 53, 1263–1269. Binnington, H.B., et al., 1973. A technique to increase jejunal mucosa surface area. J. Pediatr. Surg. 8, 765–769. Bliss, M., 1982. Banting’s, Best’s, and Collip’s accounts of the discovery of insulin. Bull. Hist. Med. 56, 554–568. Blomeier, H., Zhang, X., Rives, C., Brissova, M., Hughes, E., Baker, M., et al., 2006. Polymer s­ caffolds as synthetic microenvironments for extrahepatic islet transplantation. T ­ ransplantation 82 (4), 452–459. Brittberg, M., et al., 1994. Treatment of deep cartilage defects in the knee with autologous chondrocyte transplantation. N. Engl. J. Med. 331, 889–895. Buitinga, M., Truckenmüller, R., Engelse, M.A., Moroni, L., Hoopen Ten, H.W.M., van B ­ litterswijk, C.A., et al., 2013. Microwell scaffolds for the extrahepatic transplantation of islets of ­Langerhans. PLoS ONE 8 (5), e64772. Chandra, V., et al., 2009. Generation of pancreatic hormone-expressing islet-like cell aggregates from murine adipose tissue-derived stem cells. Stem Cells 27, 1941–1953. Chen, F., et al., 1999. Acellular collagen matrix as a possible “off the shelf” biomaterial for urethral repair. Urology 54, 407–410. Chen, M.K., Badylak, S.F., 2001. Small bowel tissue engineering using small intestinal submucosa as a scaffold. J. Surg. Res. 99, 352–358. Chen, M.K., Beierle, E.A., 2004. Animal models for intestinal tissue engineering. Biomaterials 25, 1675–1681. Cheung, A.T., et al., 2000. Glucose-dependent insulin release from genetically engineered K cells. Science 290, 1959–1962. Coraux, C., et al., 2005. Embryonic stem cells generate airway epithelial tissue. Am. J. Respir. Cell Mol. Biol. 32, 87–92. Danielsson, C., et al., 2006. Polyesterurethane foam scaffold for smooth muscle cell tissue engineering. Biomaterials 27, 1410–1415. De La Fuente, S.G., et al., 2003. Evaluation of porcine-derived small intestine submucosa as a biodegradable graft for gastrointestinal healing. J. Gastrointest. Surg. 7, 96–101.

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