14.1.1 The bladder: structure and function

The bladder is a complex organ whose primary function is to store variable volumes of urine for extended periods of time. By retaining urine at safe, physiological pressures, the bladder protects the kidneys from damage (Thomas, 1997). The remarkable capacity and compliance of the bladder are dependent on the structural, biomechanical and biological properties of the smooth muscle wall and the highly specialised urothelial lining, which provides both urinary barrier and mechanosensory functions (Birder et al., 2012). In common with all tissue engineering, the ability to deliver successful (i.e. safe and functional) engineering of partial or whole bladder organ constructs requires fit-for-purpose biomaterials and a comprehensive understanding of bladder structure, cell/tissue biology and physiology. Given that until recently the urinary bladder was considered to be a passive urine storage organ, it is unsurprising that past attempts to reconstruct it with unsuitable materials have resulted in failure.

The mammalian bladder is composed of four distinct layers, with an outer serosal layer surrounding the detrusor muscle compartment, which is made up of three loosely arranged layers of smooth muscle (Fig. 14.1). Concentrically within this, the lamina propria is a viscoelastic collagenous connective tissue supporting a variety of cellular structures, including blood vessels, sensory and motor neurons. A basal lamina separates the lamina propria from the urothelium. The urothelium itself is a transitional epithelium comprising three stratified zones: a single row of basal cells attached to the basement membrane, several layers of intermediate cells and a single, overlying row of superficial ‘umbrella’ cells that abuts onto the luminal space. The function of the urothelium as a urinary barrier occurs primarily at the level of the superficial cells, with the paracellular barrier maintained by intercellular tight junctions (Acharya et al., 2004; Varley et al., 2006) and the transcellular barrier provided by multiple thickened plaques of asymmetric unit membrane (AUM) embedded in the outer leaflet of the apical membrane (Hicks, 1965). The AUM plaques are constituted by the interactions of four uroplakin (‘urothelium-plaque’) proteins and are a unique feature of urothelium (Wu et al., 1990; Yu et al., 1994; Olsburgh et al., 2003). The AUM plaques are formed in the Golgi apparatus and transported to the apical membrane as fusiform vesicles (Tu et al., 2002), thereby providing a source of membrane for accommodating changes in surface area and maintaining a low pressure environment during bladder filling.

Such is the relationship between urothelial structure and function that loss of one component of the AUM can have devastating effects on urothelial structure and transcellular barrier properties (Hu et al., 2000, 2002). Thus, urothelial differentiation antigens not only provide objective markers of urothelial cytodifferentiation but, by virtue of their role in the urothelium, may also be regarded as surrogate markers of urinary barrier function. Unfortunately, expression of these markers is not invariably reported in bladder tissue-engineering reports, leading to discrepancy in the interpretation of some studies.

An important feature of the urothelium is its high regenerative capacity. Thus, although the urothelium is regarded as a ‘stable’ or quiescent tissue with an extremely slow rate of cell turnover, which may be as long as 200 days (Hicks, 1975), it is able to undergo rapid proliferation in response to acute injury (Peyton et al., 2012). Lavelle and colleagues performed a controlled study of selective urothelial damage in rats, which showed that recovery of transcellular and paracellular components of the urinary barrier occurred within 72 hours, with the intermediate cells undergoing rapid maturation to form differentiated umbrella cells (Lavelle et al., 2002). The excellent regenerative and differentiation capacity of urothelium is critical to maintaining the urine-proofing properties of the bladder and has positive implications for developing tissue-engineering strategies.

14.1.2 The clinical need for bladder reconstruction

With over 120 000 new cases reported in Europe (International Association for Cancer Research) and 150 000 in the USA (National Cancer Institute) in 2012, bladder cancer is the ninth most common cancer diagnosis worldwide. Those patients requiring cystectomy (mainly for muscle invasive bladder cancer) represent the largest group requiring surgical reconstruction of the lower urinary tract. Current approaches for urinary diversions following cystectomy commonly involve reconfiguring bowel in the form of orthotopic ileal neobladders, ileal conduit stomas or continent pouches (Studer et al., 2004; Stein and Skinner, 2006).

Benign disease processes or surgical interventions that render the bladder unstable, under high pressure, or lacking in capacity or compliance can result in a range of clinical problems ranging from mild to severe chronic urinary incontinence to irreversible kidney damage caused by raised upper tract pressures. Examples include neuropathic bladder (e.g. secondary to myelomeningocele, multiple sclerosis or spinal cord injury), severe detrusor instability and end-stage interstitial cystitis. There have been recent improvements in the medical management of these conditions using anticholinergics and especially, the intravesical injection of Botulinum Toxin-A (‘Botox’) or sacral neuromodulation/neurostimulation in selected cases (Fowler et al., 2012). However, a more permanent surgical augmentation remains the clinical need for those patients who develop a small-capacity, poorly compliant bladder, where intractable incontinence or pain destroys quality of life, or where serious kidney damage is imminent (Cain and Rink, 2010; Biers et al., 2012). As discussed in detail below (Section 14.3.1), although the surgical autoaugmentation of the bladder using bowel is considered the ‘gold standard’ treatment for end stage bladder disease, it is associated with significant clinical complications that are driving research to find alternative approaches.

14.2.1 Concepts and strategies

An ideal tissue-engineered urinary bladder would mimic the range of functions fulfilled by the normal healthy bladder. During filling and voiding, the bladder undergoes dramatic changes in volume and is exposed to considerable mechanical forces. Adequate compliance is critical to accomplish the low pressure storage of urine and protection of the kidneys in the upper urinary tract. The development of sensory self-voiding function is outside current objectives and in all current and proposed bladder reconstruction strategies, it is anticipated that voluntary emptying will be aided by clean intermittent self-catherisation (CISC), either via the urethra or via an ileal conduit stoma or a vesicotomy, such as described by Mitrofanoff (1980).

Novel approaches for bladder reconstruction can be categorised as biomaterials-based, cell-based or combined (i.e. tissue engineering) strategies. The former, involving implantation of a biomaterial, is a passive approach that relies on the regenerative capacity of the host for full integration, whereby the material becomes cellularised and is eventually resorbed and replaced. The alternative is to harvest and expand cells from an appropriate host tissue in vitro, prior to transplanting the cells back into the body, with or without a biomaterial ‘scaffold’ to provide structure. The theoretical advantage of this latter approach is that clinically useful numbers of autologous cells are generated by propagation in the controlled environment of the laboratory, prior to surgical reimplantation into the host as a nascent or functional tissue.

A major challenge with all approaches is that where the underlying pathology of the host is unresolved, biomaterial integration and/or sourcing of fully functional, autologous cells for ex vivo tissue-engineering approaches may prove impossible. In this context, it is noteworthy that most experimental tissue-engineering models use healthy animals and problems can emerge when promising approaches are transferred to a clinical ‘disease setting’ (Bhargava et al., 2008).

14.2.2 Biomaterials

An ideal biomaterial scaffold should provide both structural support and adequate access to cells and nutrients to enable cells to engraft, survive, interact and be maintained. A more ambitious aim is that biomaterials provide instructive or ‘niche’ environments to support specific tissue development and differentiation. It is well-established that biomechanical properties such as stiffness and bioactive features that modulate cell-matrix interactions may influence cell phenotype and tissue function (Li and Xie, 2005; Rizvi and Wong, 2005; Rohman et al., 2007; Baker et al., 2009; Engelhardt et al., 2010). For example, human mesenchymal stem cells may be directed along neuronal, muscle or bone lineages by varying the stiffness of the scaffold (Engler et al., 2006). However, it is important to realise that there is no blueprint and tissue structure is not preformed in nature, but is an emergent property of development.

Several categories of biomaterial scaffold have been described for soft tissue applications:

Scaffolds may be generated from extracted biological or synthetic polymers using a variety of processing techniques, including electrospinning (Baker et al., 2006), phase separation (Rowlands et al., 2007), gas foaming (Mooney et al., 1996), particulate leaching (McGlohorn et al., 2004; Baker et al., 2011), inkjet-printing (Roth et al., 2004) and chemical cross-linking (Park et al., 2002) (reviewed in Wiesmann and Lammers, 2009). These techniques have been used to create scaffolds of different shapes and porosity to facilitate cell engraftment. Scaffolds may be further functionalised by incorporation, surface adsorption or chemical attachment of growth and other bioactive factors, for example, to enhance angiogenesis and encourage vascularisation (Mikos et al., 1993; Wang et al., 2008; Rohman et al., 2009; Lee et al., 2010; reviewed by Chen et al., 2010; Kaully et al., 2009).

A particular advantage of animal-derived natural matrices is that following decellularisation, they retain tissue-specific architectures and extracellular matrix (ECM) proteins (Bolland et al., 2007), thus providing a wide range of biological and physical material properties specified by the nature of the originating tissue (reviewed by Gilbert et al., 2006; Davis et al., 2010). The high degree of conservation of matrix proteins between species (collagens, laminins and fibronectins) means that these matrices tend to be non-immunogenic and represent natural substrates for influencing cellular repopulation and tissue integration (Marcal et al., 2012).

14.3.1 Vascularised tissue grafts

The use of reconfigured vascularised or pedicled host tissue grafts to augment the bladder has a long history. Currently, the most commonly performed procedure for end-stage bladder disease involves replacing or augmenting the bladder with a vascularised segment of the patient’s own bowel. This procedure of enterocystoplasty involves isolating a full thickness segment of bowel on its vascular pedicle, detubularising it along the antimesenteric border and incorporating it into the bi-valved bladder as an augmentation cytoplasty or, after cystectomy, as an orthotopic neo-bladder or conduit (Greenwell et al., 2001; Beier-Holgersen et al., 1994).

Enterocystoplasty was first described in a canine model in 1888 and then in man a year later, but it was not until the mid-twentieth century that the technique became popular for the treatment of the contracted, tuberculous bladder (Tizzoni and Foggi, 1888; Von-Mikulicz, 1889; Couvelaire, 1950). Stomach (gastrocystoplasty), small intestine (ileocystoplasty) and large intestine (colocystoplasty) have all been used as the reconstructing segment, but in the UK, ileocystoplasty is the most commonly performed procedure (Thomas, 1997).

Despite many patients experiencing the benefits of improved continence, improved urodynamic parameters and greater control over voiding, enterocystoplasty carries with it the potential for a number of serious complications. These are mainly attributable to the fact that bowel mucosa is structurally and physiologically unsuited to exposure to urine and include both early complications associated with all major abdominal surgery and specific, longer-term complications of enterocystoplasty, including spontaneous perforation of the bladder, mucus production by the bowel epithelium, bladder stones, bacteriuria, metabolic disturbances and malignancy (reviewed by Thomas, 1997; Greenwell et al., 2001).

Given that the side effects of enterocystoplasty are related to the long-term interaction of urine with the bowel mucosa, the logical progression would be to remove the bowel epithelium to leave the raw muscle surface facing the lumen – so-called seromuscular enterocystoplasty. Experimentally, in rabbit, canine, porcine and bovine surgical models, this approach has resulted in graft fibrosis and shrinkage, attributable to severe inflammation secondary to urinary exposure and irritation or infection of the graft and to ischaemia or damage to the intestine during dissection (Motley et al., 1990; Salle et al., 1990; Aktug et al., 2001; Fraser et al., 2004; Hafez et al., 2005). This phenomenon is independent of which side of the bowel wall faced the lumen. Severe fibrosis was also observed when a non-seeded vascularised capsule-flap of abdominal wall or gracilis muscle was incorporated into the rat bladder (Schoeller et al., 2004). However, in a series of 129 human bladder augmentations using demucosalised intestine, Lima and colleagues showed that fibrosis and shrinkage was prevented by the use of a silicon balloon left in the bladder for 2 weeks post-augmentation (Lima et al., 2004). Pedicled omental flaps to repair or augment the bladder (omentocystoplasty) have been largely successful both clinically and in animal models, particularly when used to close defects associated with vesico-vaginal fistula (Kiricuta and Goldstein, 1972).

Although demucosalisation of the bowel prior to incorporation into the bladder has inevitably resulted in graft fibrosis and shrinkage, when urothelium has been allowed to cover the augmenting graft, shrinkage occurred minimally or not at all (Aktug et al., 2001; Schoeller et al., 2004; Hafez et al., 2005). Hafez and colleagues (2005) developed an aerosol transfer technique in a porcine model using urothelial and bladder smooth muscle cell suspensions in fibrin glue. Autologous urothelial cells with or without smooth muscle cells, isolated at hemicystectomy, were sprayed onto demucosalised colon and then incorporated into the remaining bladder. After 6 weeks, this led to the development of a stratified, multilayered uroplakin-positive urothelium atop of a bladder or colonic smooth muscle submucosa, respectively, and no inflammation was described. Although the procedure did not involve propagation of urothelial cells in culture, it feasibly could do, the point of interest being whether in vitro-generated cells would remain capable of developing into a morphologically differentiated urothelial tissue after transplanting back in vivo.

A cell-engineering adaptation of enterocystoplasty has been described in a pig model wherein in vitro-propagated autologous urothelial cell sheets were implanted onto a vascularised, de-epithelialised host smooth muscle segment used to augment the bladder (Fraser et al., 2004). The urothelium was transplanted from cell culture to the surgical site using a Vicryl™ mesh carrier. The advantage of this ‘composite cystoplasty’ strategy over a full tissue-engineered approach is that the in vitro component of the procedure is confined to propagation of a single, highly regenerative cell type, the urothelium, which is combined with a preformed, innervated and vascularised smooth muscle host tissue.

In the most recent surgical series (Turner et al., 2011), the technique of extra-luminal dissection described by Hafez and colleagues (Hafez et al., 2003) was adapted to produce the de-epithelialised segment of bowel, and this was combined with an in vitro-generated functionally differentiated autologous urothelium (Turner et al., 2008) at the time of the composite cystoplasty reconstruction (Plate IX, between pages 354 and 355). Seven pigs underwent successful bladder augmentation using this technique and when sacrificed at 3 months, the bladder augments were found to be viable with no evidence of fibrosis or contraction. When examined histologically, all the augmented segments were completely covered with urothelium. Importantly, there was no evidence of colonic mucosal or crypt regrowth and unlike the initial study (Fraser et al., 2004), only minimal inflammatory changes were observed (Turner et al., 2011). This approach appears promising and is at the point of translation to clinic.

14.3.2 Free tissue grafts

Shortly after the first colocystoplasty was described in 1912, attempts were made to incorporate free biological tissue into the bladder. First, Neuhof (1917) used a free fascial patch in dogs and since then split skin grafts, placenta, peritoneum and dural membrane have all been used as patches (Draper et al., 1952; Kelami et al., 1970; Hutschenreiter et al., 1978; Fishman et al., 1987). There have been mixed results reported, with complications often arising as a result of normal functioning of the donor tissue (such as hair growth on skin grafts), alongside more general problems such as graft contraction and stone formation.

Nevertheless, the appeal of using free biological tissue persisted. Stenzl et al. (2000) performed detrusor myectomy using free latissimus dorsi (LD) grafts in four dogs. This approach was based on Carpentier’s LD cardiac wrap for patients with severe cardiomyopathy, which was the first recorded case of substituting non-skeletal muscle with skeletal muscle (Carpentier and Chachques, 1985). The procedure has been transferred to the clinical setting and initial clinical experiences were reported in 24 patients with bladder acontractility who required clean intermittent catheterisation (Gakis et al., 2011). Although the first clinical results appear promising, this procedure is still considered experimental.

14.3.3 Acellular matrices

The decellularisation of an allogeneic or xenogeneic tissue can potentially provide a bio- and tissue-compatible polymeric scaffold or matrix for recellularisation by the recipient’s own cells. Decellularised matrices retain the tissue-specific architecture with potential for tissue-specific cell–matrix interaction and differentiation cues. However, they also carry the potential risk of contamination by xenogeneic factors and for immunological reaction in the case of incomplete decellularisation. Another potential problem is the inherent biological variability of the source.

The two most common preparations described for potential use in bladder reconstruction are porcine SIS (Zhang et al., 2000; Kropp et al., 2004) and bladder acellular matrix graft (BAMG) (Dahms et al., 1998). In their natural tissue states, these matrices are heavily populated with cells and hence must undergo extensive decellularisation to remove all potentially immunogenic material. Non-cross-linked tissue matrices have been described as slow release vehicles of naturally occurring growth factors because, once implanted, they slowly degrade, acting as a scaffold for new ECM proteins produced by the in-growing cells (Kim et al., 2000; Badylak, 2002).

14.3.4 Natural ECM

The ECM has been used extensively as a xenogeneic and allogeneic biomaterial for cells of many types, reflecting its natural evolution as a tissue scaffold. Collagen, the most abundant protein within the ECM and the major structural protein in the body, is largely responsible for the strength and conformability of natural materials. Collagen has been shown to encourage cell growth, have minimal immunogenicity and can be readily purified and moulded into the desired form, making it an ideal tool for tissue-engineering applications (Elbahnasy et al., 1998; Hattori et al., 2006). Purified collagen, however, when used for reconstruction in the urinary tract, has been shown to lose its tensile strength and to be susceptible to tearing during suturing (Elbahnasy et al., 1998).

To overcome these problems, collagen has been reinforced with synthetic materials (see below), and with natural tissues, including a pedicled omental flap (Hattori et al., 2006). The latter investigators employed a porcine in vivo model to demonstrate that collagen sponge became vascularised when combined with omentum for 7 days in vivo and that only when pre-conditioned in this way was the collagen sponge able to support passive engineering of the bladder. This approach has important implications for other natural or synthetic biomaterials, as it provides a strategy for in vivo-preintegration of a scaffold for subsequent use in passive tissue engineering.

14.3.5 Synthetic grafts

The obvious advantages of synthetic materials can be seen in the usage of standardised raw materials and processing reproducibility, resulting in lower production costs compared to biological matrices.

The incorporation of synthetic materials alone into the bladder has largely been met with failure, primarily as a result of biological and mechanical incompatibility. Plastics, polyvinyl sponge, polytetrafluoroethylene (Teflon™) and Japanese paper have all been used to reconstruct the bladder with variable results, but none has been pursued to the present day and the use of such materials is considered obsolete (Bohne and Urwiller, 1957; Kudish, 1957; Kelami et al., 1970; Fujita, 1978). Perhaps the most promising report was the experimental bladder reconstruction in rabbits with a 6.25 cm² poly(epsilon-benzyloxycarbonyl-L-lysine) membrane (Koiso et al., 1983). By 6 months it was reported that the resorbable membrane was completely replaced with normal urothelium and smooth muscle and there were no recorded complications. Despite such a positive study, no follow-on or clinical studies have ensued.

Bladder wall constructs comprising scaffolds seeded with urothelial and smooth muscle cells have been the most extensively researched strategy for bladder reconstruction, with a consensus that seeded constructs are superior over non-seeded scaffolds in terms of limiting graft shrinkage and loss of function (Yoo and Meng, 1998; Oberpenning et al., 1999; Atala et al., 2006; Zhang et al., 2006; Jayo et al., 2008; Tanaka et al., 2010). In a subtotal cystectomy canine model (about 80% removal of the bladder) with a follow-up of 2 years, a PLGA-based biodegradable synthetic polymer matrix seeded with autologous urothelial and smooth muscle cells was reported to result in tissue formation similar to the native bladder, including a three-layered detrusor muscle. Urodynamic studies revealed similar viscoelastic characteristics compared to a control group in which the native bladder was re-implanted and the dogs were able to void by increasing their abdominal tone. Moreover, the constructs were reported to grow during skeletal maturation of the young animals (Jayo et al., 2008).

Atala and colleagues were the first to demonstrate the feasibility of seeding cells onto a purely synthetic matrix for implantation in vivo (Oberpenning et al., 1999). PLGA is a well-characterised biomaterial with predictable biodegradability properties and is widely used as Vicryl™ sutures and meshes within the field of surgery. It is non-toxic and biocompatible with both normal human urothelial and bladder smooth muscle cells (Pariente et al., 2001, 2002; Scriven et al., 2001). These qualities make Vicryl™ an attractive candidate for combination with natural materials to form implantable constructs for bladder reconstruction. Oberpenning et al. (1999) used a polyglycolic acid (PGA) mesh, moulded into the shape of a bladder and coated with PLGA, and seeded the outer and inner surfaces of the biomaterial with autologous smooth muscle and urothelial cells, respectively. The constructs were then implanted in vivo onto a bladder base remaining after trigone-sparing cystectomy in dogs. Once coated with fibrin glue, the construct was wrapped with omentum and the animals were monitored for up to 11 months. There were no reported complications and mean capacity and compliance of the neobladders were similar to measurements pre-operatively. At 3 months, the polymer had degraded, leaving a vascularised, innervated tissue composed of organised smooth muscle bundles and a stratified urothelium, which was positive with antibodies against AUM. It is unlikely that the same degree of regeneration would have occurred without an omental wrap, such is its ability to induce neovascularisation.

Methods to improve cell attachment and proliferation on synthetic materials have also been explored. One solution is to coat the synthetic material with a biological substance or to use a surface modification procedure prior to seeding to encourage attachment. For example, in vitro, smooth muscle cells have been shown to attach and proliferate on a biodegradable polyesterurethane foam pre-treated with fetal bovine serum (Danielsson et al., 2006) and on plasma-coated, electrospun polystyrene (Baker et al., 2006). Alternatively, the material can be combined with one of the aforementioned natural materials to act as a biodegradable scaffold, giving strength and conformability to the structure. It is perhaps surprising, given the results described by Oberpenning et al. that when combined with PLGA, collagen hybrid matrices have shown mixed results in vitro. In one study, smooth muscle cells were able to proliferate and retain expression of differentiation markers on a gel-based construct, but not on a sponge, the opposite being the case for urothelial cells, which stratified on a sponge but not a gel, although unequivocal markers of urothelial differentiation were not shown (Nakanishi et al., 2003).

As biomaterial properties can have differential effects upon proliferation, migration and differentiation of different cell types, this must be taken into consideration when developing the ideal synthetic material. For instance, smooth muscle cells adopted a more natural organisation when grown on electrospun polystyrene scaffolds where fibres were aligned rather than showing a random distribution (Baker et al., 2006). Similarly, urothelial and smooth muscle cells showed improved growth properties on materials where the elastic modulus most closely matched that of the bladder (Rohman et al., 2007).

Meanwhile, Atala and colleagues (2006) have made the transition from canine model to clinical trials. Collagen-only and collagen–PGA hybrid scaffolds were seeded with autologous smooth muscle and urothelial cells and implanted into nine patients with severely neuropathic bladders. Three had a collagen-only implant, one had collagen-only implant with an omental wrap and three patients had the collagen–PGA hybrid scaffold with an omental wrap. Two patients were lost to follow-up and one patient with a collagen-only implant underwent conventional augmentation because of progressively rising intravesical pressures. For the remaining patients, although followed up annually for up to 5 years, not all results were available at each time point and only four had investigations in the fifth year. There were minimal or modest increases in capacity and compliance of the bladders, with the best outcome in patients receiving cell-seeded collagen-coated PGA scaffolds that were wrapped in omentum as a vascular bed. The new bladder tissue was described as having a normal structure, with smooth muscle and stratified urothelium; however, the differentiation status of the urothelium was not reported.

More recent developments in the field include the incorporation of polymers, such as PLGA nanoparticles into decellularised matrices, e.g. SIS or BAMG grafts (Mondalek et al., 2008, 2010; Geng et al., 2011; Roth et al., 2011). With the aim of improving the consistency and biocompatility of the biomaterial, the incorporation of hyaluronic acid PLGA nanoparticles into SIS was reported to enhance angiogenesis and smooth muscle cell regeneration in a canine partial cystectomy model (Mondalek et al., 2010; Roth et al., 2011). Another group has incorporated vascular endothelial growth factor (VEGF)-loaded PLGA nanoparticles into BAMG and showed good biocompatibility in vitro and in vivo (Geng et al., 2011), although no results have yet been reported for its use in bladder reconstruction.

14.4.1 Static conditioning

It is widely accepted that cells lose functional and differentiated characteristics when isolated from their host tissue and propagated in cell culture. For all active tissue-engineering strategies involving cell-seeded approaches, the critical question is whether this loss is reversible and thus whether (and how) cultured cells can be induced to differentiate and form functional tissue equivalents.

For example, normal human urothelial (NHU) cells grown in monoculture adopt a proliferative, highly regenerative, but non-differentiated state (Southgate et al., 1994). It has been shown possible to switch these in vitro-propagated cells to a stratified, differentiated and functional barrier urothelium following manipulation of the growth medium (Cross et al., 2005). The resultant urothelium had functional barrier properties, as assessed by a high transepithelial electrical resistance (TER) of > 3000 Ω cm² and low diffusive permeability to urea, water and dextran. This technique was adapted and applied to porcine urothelial cells in vitro (Turner et al., 2008) prior to application in a surgical model of composite cystoplasty (described in Section 14.3.1) (Turner et al., 2011). Progress has also been made in the identification of the molecular pathways involved in inducing urothelial differentiation, with the nuclear receptor, peroxisome proliferator activated receptor gamma (PPARγ), identified as a key regulator of urothelial differentiation (Varley et al., 2004, 2006, 2009). Notably, although a urothelial stem cell is purported to reside basally in situ (Gaisa et al., 2011), it has been shown that both basal and suprabasal-derived urothelial cells demonstrate equal proliferation and differentiation potential in vitro, thus negating the need to isolate specific progenitor populations for tissueengineering purposes (Wezel et al., 2014).

14.4.2 Biomechanical conditioning

Further improvement of the biomimetic properties of in vitro-generated bladder tissue may be achieved by simulating the physical environment of the bladder (Korossis et al., 2006). So-called ‘functional tissue engineering’ employs an external bioreactor to condition cells seeded onto a scaffold by controlling nutrition and providing appropriate mechanostimulation. Mechanosensitivity is a requirement in all cells and allows them to respond appropriately to physiological signals, as well as to insults such as physical stress and osmotic pressure gradients (Hamill and Martinac, 2001); for example, in vitro studies of myofibroblasts showed that proliferation and biosynthetic activity changed with the degree of mechanical stress (Grinnell, 1994). The bladder fills with urine passively and undergoes several fill–void cycles daily, and despite the large and often rapid changes in volume, the urinary barrier remains intact. The urothelium manages this by being exquisitely sensitive to stretch, with the superficial cells mobilising the AUM-containing fusiform vesicles to open onto the luminal membrane in response to filling, thus maintaining an appropriate surface area (Truschel et al., 2002). Replication of the fill–void cycle would seem to be an appropriate measure when generating a biomimetic tissue in vivo and may have significant consequences for tissue functionalisation.