Patent Description:
Currently, there are about <NUM>,<NUM> patients per year in the U. who need replacement tracheal tissue. Causes for this include tracheal cancer, invasive infections of the trachea or bronchi, and trauma. There are no replacements currently available for trachea in humans. when a segment of trachea is resected, the only surgical option is to "pull together" the two ends of the trachea and sew them together, hoping that the anastomosis does not "pull apart" thereafter.

Currently in the U. , approximately <NUM>,<NUM> patients per year need an esophageal replacement. This is due primarily to esophageal cancer, though trauma and infection are causes a small number of cases of esophageal replacement. Currently, there is no available replacement for esophageal tissue. What is done currently to replace esophagus is one of two procedures. Either a segment of the stomach is loosened from its connections in the abdomen and brought up into the chest, to anastomose to the remnant esophagus; or, a segment of large bowel (i.e., colon) is resected from the patient and sewn in to replace the resected esophageal tissue. Both of these procedures have many complications and a viable esophageal replacement is certainly medically needed.

Every year in the U. , approximately <NUM>,<NUM> patients undergo cystectomy, and require a urinary conduit to drain urine outside the body [Heathcare Cost and Utilization Project, N. In almost all cases. bowel is harvested from the patient to form either a noncontinent urinary diversion, or a continent urinary diversion that is catheterized intermittently to drain urine through a continent stoma [<NPL>. Due to surgical simplicity and lower complication rates, creation of a noncontinent urinary conduit is the most common approach for draining urine following cystectomy. Most typically, a <NUM>-<NUM> length of ileum is harvested from the patient for use as the urinary conduit, and the remaining bowel is reanastomosed [<NPL>. One end of the harvested ileal segment is anastomosed to the patient's ureters, and the other end is then brought out to the skin to form a stoma through which urine can drain.

Though widely used. ileal conduits pose many problems that can lead to short-term and long-term complications [<NPL>. In the short term, patients may suffer from complications at the bowel harvest site, including anastomotic leaks and peritonitis. In addition, ileal urinary conduits may suffer from ischemia and necrosis, which can lead to perforation, anastomotic breakdown, and leakage of urine from the conduit. In the long term, many patients suffer from chronic hyperchloremic metabolic acidosis, due to resorption of urine electrolytes through the conduit wall. Since ileal conduits harbor bacteria, patients also commonly suffer from recurrent urinary tract infections and pyelonephritis, as bacteria from the conduit infect the more proximal urinary system. Hence, there is a significant medical need for an improved method for urinary diversion, that avoids many of the complications associated with the use of ileal conduits [ <NPL>]. <CIT> discloses decellularized tissue engineered constructs such as esophageal constructs, which are produced by seeding cells known to secrete extracellular matrix molecules such as collagen and elastin.

There is a continuing need in the art for replacements for these important conduits, as well as other tubular tissues in the body, such as ureters, urethras, intestine, etc..

Only anatomical conduits according to the claims are part of the invention. Anatomical conduits not falling within the scope of the claims are only present for illustrative purposes. The invention is an artificial esophagus for replacement of damaged tissue by implantation. The esophagus comprises substantially acellular extracellular matrix formed as a tube of greater than <NUM> diameter. The artificial esophagus has a suture retention of greater than <NUM> grams, wherein the extracellular matrix is produced and secreted by vascular smooth muscle cells grown in vitro on a tubular substrate.

The inventors have developed new surgical tools for repairing damaged anatomical conduits, including upper digestive tract. Because the conduits are formed by seeding cells on a tubular substrate, rather than as sheets which are subsequently wound to form a tube, or tubes that are layered, they are not subject to the risks of slippage or leakage can arise with a layered approach. Furthermore, no slippage or leakage occurs between the tissue and the tubular substrate (e.g., stent) because the tubular substrate or stent is integrated within the non-layered tissue. Moreover, there is no risk of unwinding of tissue layers. The conduits form a composite artificial tissue in which, in some embodiments, a stent is totally enveloped and encased within extracellular matrix that has been grown and secreted in situ. The extracellular matrix is a naturally occurring matrix that is produced by cells. The matrix is a naturally occurring extracellular matrix that is produced by living cells, because it is more readily remodeled by host cells after implantation, and because it is less likely to induce adverse host responses such as inflammation or calcification as compared to denatured, processed, or artificially cross-linked extracellular matrices. Because of its mode of manufacture, no gaps exist or form during and after implantation. The conduits are in effect composite tissues. The tissues may be a composite of cells and matrix (decellularized), polymer fragments, and an optional stenting material. The extracellular matrix may, for example, bridge stent struts and completely incorporate the stent material.

The tissues so formed have matrix (ECM) that fits "snugly" around the stent struts. The tissues adjoin, are connected with, abut, are next to, have a common boundary with, touch, are contiguous with, share a common border with the stent. They form a unitary composite tissue that is not subject to separation and deconvolution into constituent parts.

Cells used in the conduits can be allogeneic, autologous, syngeneic, or xenogeneic. Typically cells used in making the conduits are killed and/or removed prior to use. The killing and/or removal of cells diminishes the potential for adverse immune reactions. Killing and/or removal of cells leaves less than <NUM> %, less than <NUM> %, less than <NUM> %, less than <NUM>%, less than <NUM>%, or less than <NUM>% of the cells viable, as assessed by trypan blue staining, nucleotide incorporation, or protein synthesis. Remaining extracellular matrix is highly conserved among individuals, and among species, rendering it less likely to provoke an adverse immune reaction than live cells. Vascular smooth muscle cells are the type of cell that are used to make the extracellular matrix. These can be isolated from any vasculature of a human or other mammal, including from the aorta. Much of the secreted extracellular matrix comprises collagen. Collagen may comprises at least <NUM> %, at least <NUM> %, at least <NUM> %, at least <NUM> %, at least <NUM> %, at least <NUM> % of the extracellular matrix. Typically the extracellular matrix is grown until it achieves a thickness of at least <NUM> microns, at least <NUM> microns, at least <NUM> microns, at least <NUM> microns, at least <NUM> microns, at least <NUM> microns, at least <NUM> microns, or at least <NUM> microns. Diameter of the conduits may be controlled during manufacturing. They have an internal diameter of at least <NUM>, at least <NUM>, at least <NUM>, at least <NUM>, at least <NUM>, at least <NUM>, <NUM>, at least <NUM>, at least <NUM>.

Under some circumstances, it may be desirable to have live cells on or within the conduit. Such cells may be seeded upon the conduit and either grown in culture or grown in situ. The cells may be seeded in situ as well, by endogenous cells of the recipient which migrate and establish themselves on the artificial prosthesis. The cells may be derived from the patient or from another source. The cells may be useful for mimicking and recreating natural conditions in the host. Alternatively the cells may be used as in situ factories to produce a product that is desirable, such as a growth hormone, chemokine, blood factor, and the like. Suitable cells for seeding on an artificial esophagus include epithelial cells, endothelial cells, smooth muscle cells, and fibroblasts.

While the conduits are described as tubular, they may also contain one or more branches, so that the conduit is in the shape of, for example, a Y, X, T, or F. Conduits with such branches are considered tubular, as well. The conduits described may be implanted to replace, line, reinforce, or by-pass an existing physiological or implanted conduit.

Conduits which are made by growing cells on a tubular stent may have an additional advantage over conduits formed using rolled sheets inside and outside of a stent. The conduits made by growing cells may have the extracellular matrix rotationally fixed with respect to the stent. The inner and outer surfaces of extracellular matrix may additionally be rotationally fixed with respect to one another. Being so fixed, slippage and leakage is minimized. The two surfaces may be so fixed by, for example, interlacing the substrates upon which the cells secreting the extracellular matrix are grown. The cells and extracellular matrix may, for example, envelope or bridge the stent struts, and may completely incorporate the stent material.

Conduits that are grown by culturing cells on a substrate that encases a tubular stent, either on the inside or the outside or on both sides, have the advantage of being comprised of a single tissue that envelopes and encases the stent material. The resulting material is a true tissue-stent composite material. This configuration has many functional advantages over earlier systems that involve rolling a sheet of tissue around the inside or outside of a stent. For example, in situations where tissue is rolled around the stent, it can occur that the sheets of tissue do not fuse with each other, or do not fuse with the stent material. This lack of fusion of tissue sheets results in a construct in which pieces of tissue and stent material can slip relative to one another, resulting in a conduit which is structurally unstable. In contrast, by culturing cells on a scaffold which fully encases the stent material, the resulting conduit is comprised of a single piece of tissue-stent composite material, and contains no sheets of tissue which may move or slip relative to one another. Such conduits are then more highly suited to various applications wherein the conduit must be liquid-tight or air-tight. In addition, such conduits are more highly suited to serving as replacements for native tubular tissues, such as trachea, bronchus, intestine, esophagus, ureter, urinary conduit, or other tubular tissues which must function to contain liquid or air or both. In contrast, stents that are wrapped with sheets of exogenous tissue may be poorly suited for these applications, since the slippage of tissues and stent material can cause leakage of air or fluid. In addition, a single tissue-stent composite material displays superior handling properties for surgical implantation, in contrast to wrapped tissue sheets which can slide and become detatched from the stent material. In addition, a single tissue-stent composite material will withstand physiological stresses following implantation, such as pressurization, shear forces, fluid flow, and the like, while a stent encased in tissue sheets may delaminate and lose structural integrity upon exposure to physiological forces in the body.

Decellularization of the conduit may involve the killing and/or removal of cells from a scaffold or substrate. Any means known in the art may be used, including but not limited to the use of agitation and the use of detergents. The decellularization process must be balanced between the limits of being sufficiently harsh to kill or dislodge the cells and sufficiently gentle to maintain the extracellular matrix structure intact. Substantially acellular extracellular matrix remains after the decellularization process. The prosthesis contains less than <NUM> %, less than <NUM> %, less than <NUM> %, less than <NUM>%, less than <NUM>%, or less than <NUM>% of the cells viable, as assessed by trypan blue staining, BrdU nucleotide incorporation, TUNEL staining, or protein synthesis.

Prostheses or conduits may be stored prior to implantation in a recipient mammal. The storage may occur before or after decellularization has occurred. Storage may be at various temperatures, but typically will be at or below <NUM> deg C, <NUM> deg C, -<NUM> deg C, -<NUM> deg C, or -<NUM> deg C. Storage may be for at least hours, at least days, at least weeks, at least months or at least years. In addition, conduits may be stored at room temperature, at or below <NUM> deg C, <NUM> deg C, <NUM> deg C, <NUM> deg C, or <NUM> deg C. In general, it is not desirable to store the conduits at temperatures above <NUM> deg C.

Esophagus has several key functions in the body. First and most importantly, it must provide an air-tight and water-tight conduit that prevents the leakage of food into the surrounding mediastinum. Since all of the food and drink that we consume is contaminated with bacteria, it is essential that the esophagus retain all food material and prevent it from entering the chest/mediastinum, where it would cause infection with significant attendant morbidity/mortality. Additionally, it may impermeable to gas.

A second key function of the esophagus is to provide peristalsis, or rhythmic contractility, to force food from the upper esophagus into the stomach. To perform this function, the esophagus may be comprised mainly of intestinal smooth muscle that has rhythmic contractile capability. Other cells which may populate the artificial esophagus include such esophageal cells as epithelial cells, endothelial cells, smooth muscle cells, and fibroblasts. To maintain adequate tensile strength and prevent tearing, the esophagus also may have significant collagenous extracellular matrix. An esophagus will typically have a suture retention of greater than <NUM> grams, greater than <NUM> grams, greater than <NUM> grams, greater than <NUM> grams, greater than <NUM> grams, or greater than <NUM> grams. Its rupture strength may be greater than <NUM> MPa (<NUM> Hg), greater than <NUM> MPa (<NUM> Hg), greater than <NUM> MPa (<NUM> Hg), greater than <NUM> MPa (<NUM> Hg), greater than <NUM> MPa (<NUM> Hg), greater than <NUM> MPa (<NUM> Hg),or greater than <NUM> MPa (<NUM> Hg). Unlike the tracheal prosthesis, the esophagus may not require any "stenting function" to maintain patency.

Our tissue engineered esophagus consists of an engineered tissue that is made from vascular smooth muscle cells that are cultured on a degradable polymer scaffold made of PGA. After <NUM>-<NUM> weeks of culture, the engineered tissue is decellularized, to produce a substantially acellular engineered esophagus that can be stored on the shelf for months at a time.

Engineered esophagus may be made by culturing human cells on a substrate, that optionally is also encasing a stent. In these cases, the final conduit contains human extracellular matrix. However, engineered esophagus can be made using any mammalian or primate vascular smooth muscle cells, including human vascular smooth muscle cells. Such primates or mammals include, without limitation, pig, horse, donkey, cat, mouse, rat, cow, sheep, baboon, gibbon, and goat. Additionally, recipients of the prostheses can be without limitation mammals including, human, dog, pig, horse, donkey, cat, mouse, rat, cow, sheep, baboon, gibbon, and goat.

A poly(glycolic) acid (PGA) sheet is cut into <NUM> x <NUM> and <NUM> x <NUM> pieces. <NUM> x <NUM> PGA mesh is rolled into a tube and inserted inside a bare-metal stent (<NUM> x <NUM>). <NUM> x <NUM> PGA mesh is then sewn around the bare-metal stent using the absorbable PGA suture, sandwiching the stent between the two layers of PGA mesh. A crochet needle is applied carefully throughout the PGA/stent construct to interlace the two PGA layers together. A non-absorbable suture is sewn through both ends of PGA/stent construct to suspend the construct inside a specially designed bioreactor for trachea reconstruction. The construct is then dipped into <NUM> NaOH solution for <NUM> minutes to treat the surface of the PGA mesh followed by three rinsing in distilled water. The PGA/stent construct is then assembled in the bioreactor as shown in <FIG>.

Primary SMCs were isolated from dog aortas and expanded in T-<NUM> in <NUM>% fetal bovine serum (FBS) low glucose Dulbecco's Modified Eagle's Medium. <NUM> million SMC's of P2 and P3 were re-suspended in <NUM> of medium and seeded onto the PGA/Stent construct inside the bioreactor, as shown in <FIG>. The construct was cultured inside the bioreactor statically for <NUM> weeks in <NUM> of low glucose Dulbecco's Modified Eagle's Medium with <NUM>% FBS, basic fibroblast growth factor (<NUM> ng/ml), platelet derived growth factor (<NUM> ng/ ml), L-ascorbic acid, copper sulfate, HEPES, L-proline,L-alanine, L-glycine, and Penicillin G (<FIG>). Medium was changed <NUM> times per week and ascorbic acid was supplemented three times per week.

Engineered trachea (<NUM> in length) was first incubated in <NUM> CHAPS buffer (<NUM> CHAPS, <NUM> NaCl, and <NUM> EDTA in PBS) for <NUM> minutes at 37C° under high-speed agitation, followed by thorough sterile PBS rinsing. The engineered trachea was further treated with <NUM>, sodium dodecyl sulfate (SDS) buffer (<NUM> SDS, <NUM> NaCl, and <NUM> EDTA in PBS) for <NUM> minutes at 37C° with high-speed agitation. The engineered trachea then underwent <NUM> days of washing in PBS to completely remove the residual detergent. All decellularization steps were performed under sterile conditions. The decellularized engineered trachea was stored in sterile PBS containing penicillin 100U/mL and streptomycin <NUM>/mL at 4C°.

<NUM> x <NUM> PGA sheet is sewn into a cylindrical construct with absorbable PGA suture around a compliant silicone tubing (inner diameter = <NUM>) with a suture line that is axially aligned to the PGA cylindrical scaffold. Dacron cuffs are then sewn onto the ends of the PGA tubular construct, one on each end. The construct is dipped into <NUM> NaOH solution for <NUM> minutes to treat the surface of the PGA mesh followed by three subsequent wash in distilled water. The PGA scaffold and silicone tubing are assembled inside a bioreactor as shown in <FIG>.

<NUM> million dog SMCs of P2 and P3 were re-suspended in <NUM> of medium and seeded onto the PGA construct inside the bioreactor. The seeded construct was cultured inside the bioreactor connecting to a peristaltic pump, which creates cyclic radial strain of <NUM>% at <NUM>. The engineered esophagus was cultured in the pulsatile culture for <NUM> weeks and maintained with <NUM> of low glucose Dulbecco's Modified Eagle's Medium with <NUM>% FBS. basic fibroblast growth factor (<NUM> ng/ml), platelet derived growth factor (<NUM> ng/ ml), L-ascorbic acid, copper sulfate, HEPES, L-proline,L-alanine, L-glycine, and Penicillin G (<FIG>). Half of medium volume was changed <NUM> times per week and ascorbic acid was supplemented three times per week.

The engineered esophagus was cut into two <NUM>-length pieces and were first incubated in <NUM> CHAPS buffer (<NUM> CHAPS, <NUM> NaCl. and <NUM> EDTA in PBS) for either <NUM> or <NUM> minutes at 37C° under high-speed agitation, followed by thorough sterile PBS rinsing. The engineered esophagus pieces were further treated with <NUM> sodium dodecyl sulfate (SDS) buffer (<NUM> SDS, <NUM> NaCl, and <NUM> EDTA in PBS) for <NUM> or <NUM> minutes at 37C° with high-speed agitation. The engineered esophagus pieces were then washed with PBS for two days to completely remove the residual detergent. All decellularization steps were conducted under sterile conditions. The decellularized engineered esophagus pieces were stored in sterile PBS containing penicillin 100U/mL and streptomycin <NUM>/mL at 4C°.

Weights are hanged from a suture line threaded onto one side of engineered esophagus, <NUM> to <NUM> away from the edge. Weights are incrementally added to the suture until the suture is torn from the tissue. The total weight at which the tissue is torn is recorded in units of gram.

A <NUM>-mm diameter metal stent is encased with PGA scaffolding, sterilized, and seeded with human vascular smooth muscle cells. The stent-scaffold-cell structure is cultured within a bioreactor for a period of <NUM>-<NUM> weeks in the presence of a nutrient culture medium. <NUM> x <NUM><NUM> P2 human smooth muscle cells (SMCs) are seeded onto the scaffold constructs (polygycolic acid mesh wrapped around a <NUM>-mm diameter nitinol stent) with <NUM> diameter and <NUM> length. The tracheas were statically suspended on silicone tubing and cultured inside the bioreactor for <NUM> weeks. The bioreactor medium was composed of DMEM (high glucose), bFGF (5ng/ml), EGF (<NUM>. 5ng/ml), lactic acid (<NUM>/L), insulin (<NUM>. 13U/ml), Pen G 100U/ml, Proline/Glycine/Alanine solution, CuSO<NUM> (3ng/ml), and vitamin C (50ng/ml). TE tracheae were cultured in <NUM> of medium at all times and only half of the medium was replaced during every medium change. The bioreactor medium was changed <NUM>-<NUM> times per week and vitamin C was supplemented to the culture <NUM> times per week. Lactic acid was freshly added to the medium once a week. Tracheas were cultured in <NUM>% human serum for the first <NUM> weeks. From the <NUM>th week on, tracheas were grown with <NUM>% human serum.

After culture, the conduit is decellularized and stored under sterile conditions in phosphate buffered saline at <NUM> deg C. After several weeks of storage, the conduit is implanted into a nude rat recipient. The chest of a <NUM> nude rat was trimmed with a shaver. A <NUM> incision was made from the neck region with a pair of surgical scissor. Muscle and surrounding tissues were separated layer by layer until the trachea was exposed. A full circumferential segment of the trachea that constitutes two cartilaginous rings was removed. Due to release from tension, the gap expanded to approximately <NUM> (depending on the individual animals). <NUM> trachea was placed in between the gap and was anastomosed end-to-end to native trachea with at least <NUM> interrupted <NUM>-<NUM> Prolene sutures for each end. Finally, the muscle and surrounding tissues were sewn together with sutures layer by layer.

At explantation after either <NUM> or <NUM> weeks of implant, the engineered trachea becomes invested with host epithelial cells in the lumen of the airway. During the implant time, no animals were treated with antibiotics. The implanted trachea also becomes invested with other host cells including fibroblasts, and also becomes invested with host micro-vasculature both in the wall of the engineered trachea and in the lumen of the engineered trachea. The dense and rapid influx of microvasculature (seen as early as <NUM> weeks after implantation, by histological evaluation) aids in resistance to infection, since host leukocytes can easily gain access to the implanted tissue to fight any infecting organisms. Over the longer term, the engineered trachea may also become invested with cartilaginous cells of the native trachea, as well as smooth muscle cells that occupy the native tracheal wall and other airways. The implanted tracheas all resisted dilatation, rupture, and perforation, which could lead to device failure and to infection in the animal. In addition, the engineered implanted tracheas did not show any evidence of either immune rejection, or of bacterial or fungal infection, during the entirely of the implantation period. The implanted trachea may resist stenosis or scarring which limits air flow to the lungs.

Engineered, decellularized human tracheas were explanted from rat recipients at two week and six weeks. After two weeks of implantation, robust tissue formation was observed in the lumen of the engineered tracheas, with evidence of extensive micro vascularization. Also after two weeks, luminal tissue was immunostained and was strongly positive for cytokeratin-<NUM>, an epithelial marker. See <FIG> and <FIG>.

After <NUM> weeks of implantation, engineered tracheas displayed good incorporation into host tissues, with formation of fibrous tissue surrounding the implanted engineered trachea, some evidence of residual decellularized human matrix, as well as ingrowth of luminal tissue. There was also some evidence of host cell infiltration into the previously acellular matrix of the tracheal implant. The implanted, engineered trachea was physically intact, without evidence of distension, perforation, or anastomotic breakdown. No evidence of excessive leukocyte infiltration or infection was observed in explanted specimens.

These results overall show that producing engineered, decellularized tracheas is feasible. Engineered tracheas can be sutured into recipient airways and can conduct air and allow the recipient to survive for long time periods. Engineered tracheas do not exhibit evidence of infection after implantation, and rapidly become invested with host cells and tissues and microvasculature after only a few weeks. Cells and tissue that infiltrates the engineered tracheas is highly vascular, and also contains cells that are native to the respiratory system (pulmonary epithelium). Engineered tracheas remain mechanically robust and do not suffer from mechanical failures such as perforation, dilatation, rupture, or anastomotic breakdown.

Based on methods pioneered by Dahl, Niklason, and colleagues [<NUM>-<NUM>], we have developed methods to grow tubular engineered tissues from banked human smooth muscle cells (SMC) that are seeded onto a biodegradable scaffold and cultured in bioreactors. No cells are harvested from the recipient for this process. After <NUM> weeks of culture, the engineered tissues are comprised of SMC and the extracellular matrix they have produced, which is primarily type I collagen. These tissues are then decellularized, creating an acellular tubular tissue that has excellent mechanical characteristics (rupture strengths > <NUM>,<NUM> Hg) [<NUM>]. We have tested these tubular engineered tissues as arteriovenous grafts in a baboon model, and they have shown excellent function, biocompatibility, zero mechanical failures, and zero infection.

We believe that the acellular engineered tissues will mitigate many of the complications that are associated with ileal conduits. Because our tissues are non-living and repopulate gradually with host cells, conduit ischemia and the associated mechanical failures will be extremely unlikely. Because our tissues do not actively absorb electrolytes, they should not cause a metabolic acidosis. Because our conduits do not foster the growth of commensal bacteria, they should not trigger recurrent urinary tract infections. And because they are available off-the-shelf, complications due to bowel resection will be avoided. Our acellular tubular engineered tissues have many favorable properties that may make them superior to segments of small intestine for urinary diversion. Since our urinary conduit is pre-manufactured using banked cells and can be stored on the shelf, there is no need to resect a segment of intestine from the patient - surgery on the bowel is completely avoided. Since our conduit is non-living, there is essentially no risk of tissue ischemia after implantation. Rather, host cells gradually migrate into the acellular matrix, with formation of commensurate microvasculature. Since our conduit does not actively absorb its luminal contents, the risk of hyperchloremic metabolic acidosis is substantially reduced. And, since our conduit does not harbor intestinal flora, the risks of recurrent urinary tract infections should be markedly reduced. Hence, essentially all of the common complications associated with use of an ileal conduit could be reduced or obviated by our acellular engineered tissues [<NUM>].

Approximately <NUM>,<NUM> cystectomies are performed annually in the US, with bladder cancer being the leading indication. In patients with T1 disease refractory to conservative measures and in patients with T2 tumors, surgical removal of the bladder with possible resection of associated pelvic organs remains the contemporary standard of care [<NUM>]. Other, less common reasons for cystectomy include neurogenic bladder (when it threatens renal function), severe radiation injury to the bladder, and intractable incontinence as well as chronic pelvic pain syndromes in females. All currently available surgical options for construction of urinary diversions involve the use of a segment of small or large intestine (<FIG>). Though it is possible to build more complex, continent reservoirs, the majority of patients in North America undergoing cystectomy are reconstructed using the ileal conduit technique [<NUM>].

Shabsigh [<NUM>] reported that within <NUM> days of surgery, gastrointestinal complications occurred most commonly (<NUM>%), followed by infections (<NUM>%), wound related complications (<NUM>%). cardiac (<NUM>%), and genitourinary complications (<NUM>%). Electrolyte abnormalities, particularly metabolic acidosis, occur in <NUM>% of patients, though often of unknown clinical significance. Severe electrolyte disturbances occur in <NUM>% of patients with an ileal conduit [<NUM>, <NUM>]. Osteomalacia can result from chronic acidosis with consequent release of calcium from bones. Acute pyelonephritis occurs in <NUM>-<NUM>% of patients with colon and ileal conduits, and <NUM>% of patients with ileal conduits die of sepsis [<NUM>]. Cancer occurs in ileal conduits - anaplastic carcinoma and adenomatous polyps have been described. The reported rate for cancer in ileal conduits varies from <NUM>-<NUM>% of all patients, though cancers can take decades to develop [<NUM>]. Early bowel complications typically consist of anastomotic leaks, enteric fistulas, bowel obstruction, and prolonged ileus [<NUM>]. Bowel obstruction has been reported in as many as <NUM>-<NUM>% of patients, with the majority responding well to conservative treatment while approximately <NUM>% require surgery. Bowel anastomotic leak is a potentially devastating complication reported in <NUM>-<NUM>% of patients, which can lead to abscess formation, peritonitis, and sepsis [<NUM>].

A urinary conduit is cultured using human smooth muscle cells that are cultured on a tube of PGA mesh scaffold in a bioreactor as described. After a culture period of <NUM>-<NUM> weeks, the resulting tubular tissue is decellularized, and then stored in phosphate buffered saline at <NUM> deg C for a period of several months. Thereafter, a cynomolgus monkey (which is an old-world primate that is phylogenetically close to humans and is therefore unlikely to reject the human engineered tissue) is prepared for implantation of the urinary conduit. After induction of anesthesia, a laparotomy is performed and the ureters of both kidneys are isolated and excised from the bladder wall, which is oversewn. The ureters are anastomosed to the urinary conduit, the other end of which is anastomosed to the abdominal wall to allow urine to flow from the ureters, through the conduit, and outside the animal's body. After completion of the implantation, the abdomen is closed and the animal is recovered from anesthesia. Thereafter, the urinary conduit is seen to conduct urine to the outside of the body to a collecting bag. There is no evidence of leakage of urine into the abdomen or from the anastomoses with the abdominal wall or the ureters.

Claim 1:
An artificial esophagus for replacement of damaged tissue by implantation, said esophagus comprising:
substantially acellular extracellular matrix formed as a tube of greater than <NUM> diameter,
wherein the artificial esophagus has a suture retention of greater than <NUM> grams, wherein the extracellular matrix is produced and secreted by vascular smooth muscle cells grown in vitro on a tubular substrate.