Krane et al. [28] eluted halofungione on a urethral stent. This alkaloid acts as a selective inhibitor of collagen type I. the stent was inserted in in rat urethras with strictures created by electrocautery. The use of halofungione prevented the formation of new type I collagen in the rat urethra.

#### *2.3 DES and DCB Containing Anti-Cancer Drugs*

Stents releasing anti-proliferative agents has been used in great extent in interventional cardiology. The combination of limiting the cell proliferation and avoiding the formation of fbrosis made this drug category especially suitable for preventing coronary vessel restenosis. The advantages in treating restenosis led to the creation of ureteral and urethral stents and balloons.

#### **2.3.1 Paclitaxel**

Paclitaxel inhibits the mitosis by stabilizing the microtubules of the cell. It is the most commonly used antiproliferative drug in studies evaluating devices for urinary stenosis.

Aiming to investigate the ability of antiproliferative drugs to permeate membrane models, Barros et al. [29] used ureteral stents with doxorubicin and paclitaxel. The researchers concluded that both drugs were restricted in the ureter and only a fraction of the total drug passed all the layers. Kram et al. [30] used a DES containing paclitaxel in rat ureters to examine its inhibitory effect in hyperplastic proliferation. The rats underwent ureteroureterostomy followed by the insertion of a drug-eluting or a conventional stent. The DES was found signifcantly more effective in reducing the cell proliferation in the side of the anastomosis. In a similar manner, Liatsikos et al. [31] compared DES to conventional stents in terms of infammation and the tissue hyperplasia occurring in porcine ureter. The ureters with the indwelling DES showed increased patency in urography compared to the ureters with the conventional stent. Shin et al. came to the same conclusions after examining a custom made paclitaxel stent in canine urethra [32] (Fig. 2).

Biodegradable stents have been also evaluated in an experimental study. Wang et al. used a biodegradable paclitaxel DES in rabbit urethra and observed the absorption of the stent in 12 weeks. Moreover, the treated urethral mucosa showed no signs of fbrosis while the urethras of the control group showed signs of fbrosis [33].

In addition to DES, DCBs with paclitaxel have been investigated (Fig. 3). Barbalias et al. [3] studied how paclitaxel is distributed in layers of rabbit urethra. The study included the dilation of the posterior urethra and subsequent dilation with a paclitaxel-coated balloon. Histological analysis showed that paclitaxel penetrates the urethral layers and especially the urothelial barrier. A similar study by Liourdi et al. [34] proved that dilation with DCBs containing paclitaxel resulted in distribution of the drug in every layer of the ureter. While the forementioned studies proved that paclitaxel could be distributed in through the urothelium to the all the layers of the urethra and the ureter, these experiments were performed in animal models which did not have strictures.

ROBUST I is a multi-centered study prospective study examining the safety and effcacy of Optilume™ Drug Coated Balloon (DCB; Urotronic, Plymouth, MN) in patients with recurrent bulbar strictures. Visasoro et al. published the frst-year

**Fig. 2** DES vs bare metal stent effcacy comparison in a porcine model

**Fig. 3** Ureter optical coherence tomography before—after DCB

results results of 53 patients that have been treated in their center using the paclitaxelcoated balloon mentioning 70% anatomic success, which was defned as diameter of lumen equal or greater than 14 Fr [35]. The same research group published the 2 years results announcing 70% success in having at least 50% improved IPSS sore for 2 years after the surgery [36]. Flow rate and post-void residual urine volumes were also improved. There researchers did not encounter any severe adverse effects. The long-term results of the study are expected after a 5-year follow up has been completed. However, it is crucial to bear in mind that every patient in this study prior to use of the DCB balloon received either an uncoated balloon or direct visualization internal urethrotomy treatment until their urethral diameter was increased by 50%. Consequently, safe assumptions for the use of paclitaxel-balloons as monotherapy for the treatment of bulbar strictures cannot be made yet Fig. 4.

Apart from paclitaxel, the effcacy of other anti-proliferative agents such as zotarolimus and sirolimus have been evaluated. Kallidonis et al. [37] using DESs containing zotarolimus in pigs and rabbits showed that zotarolimus-eluting stents reduce infammation and tissue hyperplasia in comparison to the control conventional stents. Kim et al. used two types of sirolimus-eluting stents containing different concentrations of the agent in male rat models [38]. The researchers found that the use of sirolimus stents reduced the formation of granulation tissue compared to the conventional stents and that the use of stents with sirolimus concentration of 450 μg/cm2 resulted to less layers of epithelial growth and greater rate of apoptosis in comparison to the 90 μg/cm2 stent and the conventional stent.

Many pre-clinical studies testing drug eluting devices have shown impressive results. The ability to prevent fbrosis, tissue proliferation or bacterial adhesion while reducing the symptoms and improving the patient's quality of life underline the great potential of these devices. That said, achieving a stable delivery of the drug, unaffected by the urine fow could be a key factor to lead to signifcant results in the human trials. Assumptions for the effcacy of the devices in humans cannot be made safely due to the lack of large clinical studies.

**Fig. 4** Use of DCB in urethral strictures

#### **References**

**Open Access** This chapter is licensed under the terms of the Creative Commons Attribution 4.0 International License (http://creativecommons.org/licenses/by/4.0/), which permits use, sharing, adaptation, distribution and reproduction in any medium or format, as long as you give appropriate credit to the original author(s) and the source, provide a link to the Creative Commons license and indicate if changes were made.

The images or other third party material in this chapter are included in the chapter's Creative Commons license, unless indicated otherwise in a credit line to the material. If material is not included in the chapter's Creative Commons license and your intended use is not permitted by statutory regulation or exceeds the permitted use, you will need to obtain permission directly from the copyright holder.

# **Methods and Materials for Drug Eluting Urinary Stents Design and Fabrication**

**Irene Carmagnola, Giulia Giuntoli, and Gianluca Ciardelli**

## **1 Introduction**

After urinary stenting, patients often suffer from mid- and long-term complications, such as infections, bacterial colonization, encrustations, or stent obstruction which are related to the design, materials and surface properties of the stent.

Drug eluting stents (DES) is an advance technology that can reduce the morbidity associated with stenting, by locally releasing loaded drugs in a timecontrolled manner.

The frst DES were introduced in the earlies'00 for cardiovascular applications to address the problems of restenosis associated with bare metal stents after coronary angioplasty [1]. In urology, DES could potentially solve or reduce a variety of stent-related and time-dependent complications, such as infections and obstruction, which are often related to encrustation and bioflm formation and which can dramatically result in stent failure [2]. Moreover, they could also fnd application for the management of cancer therapies [3] (Fig. 1).

Common stents are made of "inert" materials to minimize the foreign body reaction. Nonetheless, these stents are affected by several clinical problems. For instance, encrustation is caused by the deposition of urine constituents (such as

PolitoBIOMed Lab, Politecnico di Torino, Turin, Italy e-mail<EMAIL_ADDRESS><EMAIL_ADDRESS>
G. Ciardelli

Department of Mechanical and Aerospace Engineering, Politecnico di Torino, Turin, Italy

PolitoBIOMed Lab, Politecnico di Torino, Turin, Italy

Department for Materials and Devices of the National Research Council, Institute for the Chemical and Physical Processes (CNR-IPCF), Pisa, Italy e-mail<EMAIL_ADDRESS>
I. Carmagnola (\*) · G. Giuntoli

Department of Mechanical and Aerospace Engineering, Politecnico di Torino, Turin, Italy

**Fig. 1** Scheme of the concept of anti-cancer drug eluting biodegradable ureteral stent as a potential drug delivery system

proteins, ions and minerals) over the stent's surface [4]. This phenomenon occurs due to the physico-chemical characteristics of patient's urine, such as urine pH, composition and fow dynamics, as well as stent's constitutive materials or surface characteristics. Encrustation not only compromises stent drainage potential but also favor bacterial colonization and bioflm formation [5].

Different strategies can be applied to feature the stent with drug delivery systems able to reduce the incidence of encrustation and bioflm development by increasing the drug effectiveness and limiting any side effect associated with systemic delivery.

In this chapter is frstly reported an overview of the materials and manufacturing methods for conventional urinary stents, then are discussed the engineered strategies for the design and fabrication of drug eluting stents. These strategies can be divided into two main categories which are discussed in more detail in the next paragraphs: (1) DES obtained by surface functionalization and coating techniques, or (2) direct manufacturing.

#### **2 Conventional Urinary Stents**

#### *2.1 Urinary Stent Materials*

The constitutive materials of urinary stents have great infuence on their performances and durability. Materials are selected based on several requirements such as biocompatibility, mechanical strength and fexibility, surface properties, ease processability and cost-effectiveness.

Non-biodegradable polymeric and metallic materials are the main constitutive materials used for urinary stenting. The frst ureteral stent, described by Herdman in 1949, was made of polyethylene due to its desirable properties such as fexibility, strength, hydrophobicity and bio-inertness [6]. However, it was soon abandoned due to the occurrence of encrustations and the high risk of fracturing. Today, the most common polymeric non-biodegradable materials include silicones, polyurethanes, and other proprietary materials [7]. Between them, silicones have shown better performances in terms of encrustation resistance and ions deposition hindering [8].

Compared to polymeric stents, metallic stents made of titanium, nickel–titanium alloys (e.g. nitinol), or stainless steel have superior mechanical properties, and are often chosen to treat severe conditions, such as malignant ureteric strictures, or when long indwelling times are required [9]. Metallic stents suffer from encrustation as the polymeric ones but have higher migration rates and production costs.

More recently, biodegradable and/or bioresorbable stents have been introduced as novel class of temporary stents which can be dissolved or absorbed in the body [10]. These stents have the advantage of decreasing patient discomfort, eliciting fewer complications—for instance they are less prone to encrustation-, and reducing the healthcare costs (e.g. secondary procedures for stent replacement or removal). Biodegradable polymers include poly(l-lactide) acid (PLLA), polyglycolide (PGA) and polycaprolactone (PCL), whereas biodegradable metals include magnesium and its alloys [11]. Although very promising, no biodegradable nor bioresorbable stents have yet the U.S. Food and Drug Administration (FDA) regulatory approval for urological applications.

#### *2.2 Conventional Stent Manufacturing*

Extrusion is the most common technique for manufacturing polymeric hollow tubes from thermoplastic materials. Plastic materials, in the form of pellets, granules, fakes or powders, are fed from a hopper into the barrel of the extruder machine; the molten polymer is forced by a rotating screw into a die, which confers to the product its fnal shape. To obtain a hollow section, a pin or a mandrel is placed within the die, while, to prevent the tube from collapsing, a positive or negative pressure can be applied to the internal lumen or to the outside diameter of the tube, respectively. For tubes with multiple lumens, more than one pin can be inserted into the die.

Coextrusion is typically used to produce a multilayered tube by extruding simultaneously two or more discrete layers of different materials using the same die. This technique can be exploited, for instance, to encapsulate non-medical materials between two medical-material layers.  In the former, blown air expand the polymeric tube against a mold, while the latter, similarly to extrusion, forces the polymer through a conical die.

Metal based stents are obtained by extrusion or die casting, followed by surface treatments such as chemical or plasma etching, plasma treatment micro-electro discharge machining, and laser cutting [12]. Laser cutting is a rapid prototyping technique that ensures a high-precision, burr-free cut, reliability and fexibility of the design.

#### **3 Drug Eluting Stents**

Drug eluting stent is an effective means for local drug delivery to the urinary tract. It can potentially solve a variety of upper urinary tract problems, such as stentrelated urinary tract infections and discomfort, ureteral stricture, and neoplastic diseases. There are many strategies for loading drugs, including: (1) hot melt extrusion; (2) soaking the polymers into drug solution (dipping); (3) CO2 impregnation; (4) nanofbers; (5) nanoparticles. However, the release of drug elutes on the surface of biostable stents is often unsustainable and uncontrollable.

Some previous researches had coated drugs to the surface of biostable stents, and the results were not satisfactory due to the uncontrolled drug release. Alternatively, drugs or active agents can be continuously released in a controlled manner from a drug-eluting biodegradable stents (BUS), when the stent degrades. BUS are potentially a powerful tool to contrast the most frequent adverse effects reported by patients experiencing that are pain and diffculties in urinary tract.

Several surface engineering strategies have been applied to minimize encrustation and bacterial adherence over the stent surface, through antimicrobial (bactericidal), anti-fouling (bacteriostatic), lubricating or drug eluting coatings [13].

Natural antibacterial coatings include glycosaminoglycans and heparin—natural components of the urine that were found to delay encrustation; hydrophilic coatings such as phosphoryl-choline or hydrogels demonstrated the ability to inhibit bioflm formation thanks to the hydrophilic environment that hinders proteins adsorption, thus bacterial adhesion; similarly, chitosan is used for its intrinsic antibacterial properties. Other coatings, like diamond-like carbon or polytetrafuoroethylene, diminish the surface friction, facilitating stent insertion and placement while lowering patient discomfort.

#### *3.1 Coating Strategies*

Stents surface can be directly coated with active agents by several methods, such as impregnation by dipping or supercritical fuid technology [3], crystallization, spraycoating, or layer-by-layer techniques. The frst are most traditional and used, instead layer-by-layer technique is quite innovative in the urinary stent felds.

Drugs immobilization by impregnation is a two-step procedure: frst, a conventional polymeric or metallic stent is fully immersed into a solution containing the drug; subsequently the solvent evaporates leaving the surface coated with the drug.

Antibacterial agents/drugs can also be directly spray-coated onto the surface of the stent [14]. For example, Nasongkla et al. deposited chlorhexidine-loaded nanospheres on the surface by high-pressure emulsifcation-solvent evaporation technique. In detail nanosphere suspension was perpendicularly sprayed to the spinning silicone tubes for 10 s by air pump spray gun with the fow rate of 0.2 mL/s. Then silicone tubes were drained under air fow. The drug release ability of the coated silicone tubes were tested in artifcial urine and the chlorhexidine was sustained release for 2 weeks. Additionally, nanoparticles based coating showed antibacterial activity against common bacteria causing urinary tract infections up to 15 days.

Similarly, direct crystallization allows to crystallize drug onto a substrate through temperature-dependent or microdrop spray, ensuring a slower release than amorphous drug layers due to lower dissolution rate and at the same time limiting the loaded drug amount [15]. However the direct crystallization method is characterized by burst drug release kinetics.

With solvent-based polymer spraying, a drug-eluting coating can be obtained with few steps, including [16]: drugs incorporation into the polymeric solution by mixing, drug/polymer formulation spraying onto the surface; solvent evaporation. Coating characteristics, such as thickness, can be easily tailored by controlling solvent evaporation kinetics or spaying parameters. DES with prolonged antibiotic action have been obtained by spray coating drug-loaded nanospheres solutions, obtaining a smooth and homogeneous coating after several spray cycles; a sustained drug release up to 30 days occurs after the initial burst release of the loaded drug [14].

Layer-by-layer technique was frstly proposed by Decher et al. in the beginning of 1990s and it based on the alternating exposure of a charged substrate to solutions of molecules with opposite charges. After each deposition step, the substrate is washed to avoid cross-contamination of the polyelectrolyte solutions and eliminate the excess polyelectrolytes. The LBL technique allows to obtain homogeneous coating with precise and tunable architecture. It is also applicable to substrates of any shape and dimension, it is environmentally friendly and all the deposition processes can be carried out in mild conditions with low-cost manufacturing. The coating features are easily tunable by adjusting the experimental parameters, such as pH, ionic strength and polyelectrolyte concentration [15–17]. Tzanov and coworkers proposed the layer-by-layer technique to modify the surface of silicone urinary catheters to achieve infection-preventive coatings, as summarized on Fig. 2. Aminocellulose nanospheres positively charged were combined with the hyaluronic acid (HA) polyanion to build a layer-by-layer construct on silicone surfaces. Silicone supports were previously functionalized by polymerizing 3-(aminopropyl) triethoxysilane, ensuring the deposition of the frst negatively charged HA layer, the multilayer coating was build-up employing a multi-vessel automated dip coater system to obtain 5, 10 and 100 bilayers. LbL coating antibioflm activity was tested against *P. aeruginosa*. The inhibition of bioflm formation was already

**Fig. 2** Scheme of the antibacterial surfaces fabrication exploiting LbL techniques on silicone based surface. (From [18])

demonstrated with 5 bilayers coating. The antibioflm effciency of 10 bilayer multilayer coating on a Foley catheter was additionally validated under dynamic conditions using a model of the catheterized bladder in which the bioflm was grown during 7 days.

Although the manufacturing process is quite simple, DES obtained by direct coating methods are short term solutions and do not ensure sustained drug delivery. In fact, due to the rapid drug elution times, 90% of the drug is released within 2 days via burst release; once the thin biodegradable coatings are fully degraded or the drug content completely eluted, DES can be compared to conventional stents. To delay the complete depletion of the drug, additional coatings can be used for drug protection, among which the most common are hydrogels. Solutions employing drug delivery systems overcome the limited effectiveness of delivery observed for DES obtained by direct coating methods and provide the stent with sustained and controlled drug release over few weeks. The approach of using polymer coating layers has several functions: (1) delay the complete depletion of the active agent; (2) operate as drug loading system; (3) control drug release kinetics; (4) act as biocompatible coating after drug depletion.

Stents with a hydrophilic hydrogel coatings are currently commercially available (e.g., Universa® Soft Ureteral Stent-COOK® Medical). Hydrogel based coatings are design to reduce surface friction upon deployment which facilitates the placement of the stent and lowers patient discomfort and to be exploited for the controlled release of drugs and other biologically active compounds. For instance, antiproliferative drugs can be spray-coated onto the surface of the stent and be further

covered with an additional hydrogel layer to protect them upon stent insertion and further prevent a burst release [19]. Similarly, hydrogel or other polymeric materials can also work as a carrier due to a direct or indirect incorporation of the targeted drug into the polymeric matrix. Indirect incorporation includes the use of nanospheres/nanoparticles in which the drug can be loaded [20].

Drug-loaded nanoparticles can be prepared by several encapsulation methods which are reviewed by Pinto Reis et al., and others [21, 22]. Most of these procedures are simple and have industrial scale-up applicability, besides high encapsulation effcacy and improved pharmacokinetics and pharmacodynamics.

#### *3.2 DES Direct Manufacturing*

Drug eluting stents may be fabricated in a single process using direct manufacturing techniques. Drug delivery is a typical application of multilayered extruded tubes obtained by coextrusion [23]. In detail reservoir implants contain a drug loaded core encased in a rate controlling sheath. Co-extrusion to produce a core-sheath or multiple layer system requires multiple single screw extruders: centering and process adjustments to control dimensions can be diffcult. The materials in the layers should stick together; generally chemical similarity predicts adhesion.

In the last decades additive manufacturing techniques have been emerged as alternative tool for the fabrication of drug-eluting implants combining the advantages of a targeted local drug therapy over longer periods of time with a manufacturing technique that easily allows modifcations of the implant shape to comply with the individual needs of each patient [24]. Research until now has been focused on several aspects of this topic such as 3D-printing with different materials or printing techniques to achieve implants with different shapes, mechanical properties or release profles.

#### *3.3 DES Bioabsorbable Urinary Stents*

Drug-eluting technologies can be successfully combined with completely biodegradable devices. A biodegradable stent would eliminate the need for the patient to undergo a stent removal procedure, the problems of chronic indwelling stents (encrustation, stone, formation, infection), as well as the complication feared by urologists knows as the forgotten stent.

Biodegradable urinary stents (BUS) include both bioabsorbable natural or synthetic polymers, or metals. Polymeric BUS are not yet in available for urological applications but represent an important area of investigation.

BUS present several benefts such as higher resistance to encrustation and elimination of stent removal/replacement procedures, with an overall reduction

of healthcare costs. The most used biodegradable polymer is PLA. Polylactic acid (PLA) belongs to the family of aliphatic polyesters commonly made from α-hydroxy acids. It is both biocompatible and biodegradable and it is one of the most promising polymers in various applications, including biomedical feld, due to its mechanical, thermoplastic and biological properties. PLA is a biodegradable polymer, which degrade in physiological environment by macromolecular scission into smaller fragments, and then into stable end-products. It is possible to produce PLA textile structures, by different techniques, such as extrusion of the polymer into mono- and multiflaments, which may be achieved by melt, dry, dry-wet-jet spinning or electrospinning.

Different techniques can be used to incorporate drugs/active agents on BUS. Freeze-casting has been proposed as a new fabrication method to obtain biodegradable porous stents with increased urine drainage and reduced risk of refux [25]. This simple technique involves few steps: (1) preparation of a stable suspension or polymeric solution (usually water-based); (2) injection into a mold; (3) solidifcation by freezing within the mold; (4) solvent removal by sublimation (freeze-drying). Porous biodegradable DES could be further obtained by dispersing the selected drug in the polymeric solution using impregnation techniques. A method for obtaining drug-eluting BUS is supercritical CO2-impregnation, reported by Barros et al. [26]. This technique has several advantages over conventional wateror solvent-based impregnation, since supercritical CO2 shows high diffusivity and low viscosity, resulting in a fast diffusion of the solute molecules into the polymer matrix (Fig. 3). Moreover, CO2 is chemically inert in a wide range of conditions, has low toxicity, environmental sustainability, besides being ready available and low-cost [27].

Nowadays, nanotechnologies are widely used in many felds, including biomedical applications. Among nanotechnologies, electrospinning (ES) is rapidly emerging as a simple, versatile, and cost-effective method for the fabrication of smooth non-woven fbers with controllable morphology and tunable porosity, from a charged polymer solution or melt. Electrospinning process involves the application of a high electric feld to produce micro and nanofbers. The features of the end fbers depend both on solution properties, such as polymer molecular weight, concentration and conductivity, and operating parameters, such as fow-rate, applied voltage and tip-collector distance.

ES has high encapsulation effciency for drug loading, controlled residence time, desirable delivery of encapsulated drug at a predictable rate, better stability, high surface contact area, degradability, and satisfactory softness and fexibility.

Chew et al. reported many studies about the fabrication of biodegradable eluting ureteral stent by using double-needle electrospinning [28, 29]. To address the main

**Fig. 3** Supercritical CO2-impregnation process of ketoprofen into a polymeric matrix. (From [26])

I. Carmagnola et al.

**Fig. 4** Schematic illustration of preparation and antitumor effect of EPI-loaded PCL/PLGA electrospun fbers

challenges of BUS, PLGA degrading within 8 weeks was selected as stent basic material in combination with polycaprolactone (PCL) which degrades over 6 months. By using the technique of double nozzle electrospinning, researchers were able to produce a ureteral stent with different mass ratios of PCL/PLGA that degrade gradually from proximal end to the distal end.

PCL/PLGA based nanofbers can be exploited to have a controlled drug release. Ding et al. developed a biodegradable drug-loaded ureteral scaffolds able to maintain long-term effective drug concentrations in the lesion sites [30]. Epirubicin delivery was assessed on different ration PCL/PLGA electrospun fbers. Emulsionelectrospinning technology was used to fabricate a core–sheath structured EPIloaded PCL/PLGA nanofber capsules, displaying a sustained EPI release and controlled degradation *in vitro* and *in vivo* (Fig. 4).

#### **4 Conclusions**

Drug eluting ureteral stents, and in general urinary stents, were introduced to performed a drug delivery aiming to obtain a local treatment as well as to overcomes the main issues related to urinary stenting implantation. Drugs and/or active agents can be directly loaded in the stent structure or can be introduced through a surface coating. Although very simple, DES obtained by direct coating methods are short term solutions; once the thin biodegradable coatings are fully degraded or the drug content completely eluted, DES can be compared to conventional stents.

Drug-eluting technologies can then be combined with biodegradable bioabsorbable stents in order to eliminate the need for stent removal procedure. However some disadvantages remain still unsolved. In the last decades innovative manufacturing approaches and methods, such as nanotechnologies and additive manufacturing techniques, provide to scientists new tools for the design and fabrication on smart and custom-made urinary stents, able to go towards perfectly to patient needs.

#### **References**

**Open Access** This chapter is licensed under the terms of the Creative Commons Attribution 4.0 International License (http://creativecommons.org/licenses/by/4.0/), which permits use, sharing, adaptation, distribution and reproduction in any medium or format, as long as you give appropriate credit to the original author(s) and the source, provide a link to the Creative Commons license and indicate if changes were made.

The images or other third party material in this chapter are included in the chapter's Creative Commons license, unless indicated otherwise in a credit line to the material. If material is not included in the chapter's Creative Commons license and your intended use is not permitted by statutory regulation or exceeds the permitted use, you will need to obtain permission directly from the copyright holder.

# **Preventing Bioflm Formation and Encrustation on Urinary Implants: (Bio)coatings and Tissue Engineering**

**Noor Buchholz, Petra de Graaf, Julia E. de la Cruz, Wolfgang Kram, Ilya Skovorodkin, Federico Soria, and Seppo Vainio**

# **1 Introduction**

Even though urinary stents and catheters have been commonly applied in medicine for several decades and still are constantly being modifed and optimized, their structure and performance still requires further improvement. A major drawback of urinary implants is the deposition of organic and non-organic substances

N. Buchholz (\*)

P. de Graaf

W. Kram

I. Skovorodkin · S. Vainio

e-mail<EMAIL_ADDRESS><EMAIL_ADDRESS>
F. Soria

N. Buchholz, P. de Graaf, Julia E. De la Cruz, W. Kram, I. Skovorodkin, F. Soria and S. Vainio contributed equally with all other contributors.

U-merge Scientifc Offce, London-Athens-Dubai, Athens, Greece

Department of Urology, University Medical Center Utrecht, Utrecht, The Netherlands e-mail<EMAIL_ADDRESS>
J. E. de la Cruz U-merge Scientifc Offce, London-Athens-Dubai, Athens, Greece

Jesus Uson Minimally Invasive Surgery Centre Foundation, Caceres, Spain e-mail<EMAIL_ADDRESS>
Department of Urology, University Medical Center Rostock, Rostock, Germany e-mail<EMAIL_ADDRESS>
Laboratory of Developmental Biology, Disease Networks Research Unit, Faculty of Biochemistry and Molecular Medicine, Kvantum Institute, Infotech Oulu, University of Oulu, Oulu, Finland

Foundation, Jesús Usón Minimally Invasive Surgery Center, Cáceres, Spain e-mail<EMAIL_ADDRESS>
on their surface leading to bioflm formation resulting in encrustations, blockages, and infections.

In this chapter, we will present some modern biological research approaches to this problem. Promising research lines are stent coatings with antibodies, enzymes and various bioactive compounds. We will also discuss the possibility for making urinary implants more "tissue friendly" by designing biomimetic surfaces. Finally, in accordance with the paradigm "repair or regrow" we will touch on tissue engineering approaches to replace artifcial urinary implants by those generated *in vitro* or *in vivo* from homologous tissue.

#### *1.1 Antibody Coating*

Despite signifcant differences between urine and blood as the extracellular environment, designing of urinary implants has always been inspired by research approaches in cardiovascular stent engineering.

Antibody stent coating is a technology that was successfully applied to improve clinical performance of cardiovascular stents and might be considered for further optimization of urinary stents.

Coating of vascular stents with certain types of antibodies (namely: CD34, CD133 and VEGFR2) enhances migration of endothelial progenitor cells (EPCs) circulating in the blood to the site of injury, thus accelerating the healing and regeneration of vascular tissue [1, 2]. At the same time, populating of the stent surface and surrounding tissue with EPCs prevents migration of infammatory and smooth muscle cells which can cause neo-intimal hyperplasia, and re-stenosis and thrombosis of damaged blood vessels [1–3]. Antibody-coated cardiovascular stents showed promising results in preclinical trials [3]. Other antibodies with anti-infammatory, anti-platelet and anti-proliferative effects also have shown their effciency in reducing stenosis and neo-intimal hyperplasia *in vitro* [4, 5].

Despite the success of antibody stent-coating in cardiology, regrettably to our knowledge no preclinical or even experimental researches in this feld have been presented for urinary implants to date. Hypothetically, antibodies of implant coatings could attract urinary and cell components that prevent bioflm formation and urothelial hyperplasia on one hand, and promote peri-implant tissue healing on the other hand in cases where such implants are used temporarily after injuries or operations to the urinary tract. As a prerequisite for such a research approach, efforts must be made to identify urinary and cellular components with the desired properties frst.

#### **2 Enzyme Coating**

Antimicrobial enzymes (AE) have been tested as part of urinary catheter coatings. AE are found in immune systems of living organisms where their task is to attack pathogenic microbes. Hydrolytic AE destroy the structure of their foes, whereas oxidative AE trigger the production of antimicrobial molecules inside them. Quorum quenching AE interfere with bacterial quorum sensing which leads to inhibition of cell aggregation and virulent compounds [6].

AE can be attached to surfaces of medical devices, either permanently or ready to be released. Integrated methods for a controlled release include chelation or metal binding, disulfde [7], physical and ionic bonds [8].

Permanent or irreversible binding has however the advantage of better stability and decreased leaching. This can be achieved by crosslinking with linker molecules, entrapment, microencapsulation and covalent bonding [6].

For instance, cellobiose dehydrogenase (CDH), using cello-oligosaccharides as electron donors to produce H2O2, inhibited different urinary microbes including MRSA in the presence of either cellobiose or extracellular polysaccharides (EPS). The latter are essential in bioflm formation. Therefore, CDH could act as an antimicrobial agent "on demand" whenever beginning bioflm formation is triggering the reaction [9].

*Oxalobacter formigenes* is part of our gut microbiome and has been linked to non-infectious stone formation. It can degrade oxalate with the help of oxalatedegrading enzymes. These enzymes have been attached to stent coatings resulting in a 53% reduction in encrustation [10, 11].

When compared with current antibiotics or other antimicrobial agents used as active parts of urinary implant coatings, AE have certain advantages. They are highly specifc targeting only a particular bacterium without disturbing the natural microbiome. Bacterial resistance to enzymes is very rare. Care must be taken not to overdose the enzymes though in order to prevent such resistance development.

Currently, AE are expensive to produce which puts them at a disadvantage to cheaper alternatives like silver and antibiotic coatings. And it has to be borne in mind that they are proteins. That means they can get denatured during i.e. sterilization, storage and transport [6].

#### **3 Biomimetic Stents**

Natural surfaces have been the envy of many researchers. However, they are diffcult to mimic and usually outperform their artifcial copies. If it comes to internal and external surfaces, researchers try to get as close to the properties of natural surfaces as possible. Natural surfaces are made to repel or let seep through whatever is physiologically required by their environment. Blood and blood vessels form such an environment. With cardiology leading the way for a long time in stent research, it is no surprise that the attempt to create biomimetic surfaces gets its push from cardiology as well.

Although stents have been used extensively in cardiology, they do have inherent problems such as infammatory responses, thrombosis, endothelial hyperproliferation, delayed re-endothelialization, and ultimately stenosis and thrombotic obstruction [12]. An endothelium-like stent could alleviate many of these problems. A metal (copper)–catechol–(amine) (MCA) coated stent has thus been shown to facilitate a fast regeneration of a functional endothelium [12].

Earlier, a phosphorylcholine coating was effective in resisting platelet adhesion and prolong plasma recalcifcation time signifcantly. Contact angle measurements showed that the surface rearranged to become more hydrophilic at the polymer/ water interface [13].

In blood vessels, biomimetic surfaces may create an enhanced anti-thrombogenic, anti-infammatory and anti-proliferative micro-environment. In addition, such coatings could be made drug eluting [14]. Biomimetic stents will improve biocompatibility in the future [12]. Whether lessons learned from cardiology can be applied to a urinary environment will be the subject of future research. In theory, biomimetic surfaces have the ability to resist bacterial adhesion and consecutive bioflm formation.

#### **4 Bioactive Nanocoating**

As soon as urinary implants are inserted into a recipient, different biological materials start to accumulate on their surfaces. These substances, especially extracellular matrix (ECM) proteins, attract cells which migrate from surrounding tissue and form a bioflm thus affecting shape, mechanical properties and functionality of the implant. Application of nanomaterials [15, 16] can mimic cell—ECM interaction and, depending on what is required, can either suppress or induce cell migration.

To date, optimization of the structural and functional performance of the implants was mostly achieved by surface coating with different bioactive molecular substances such as growth factors and immunosuppressive and anti-infammatory drugs.

Interaction between cells, tissues and implanted mechanical objects strongly depend on the surface to volume ratio and, consequently, varies signifcantly when the objects are represented at macro-, micro- or nano-scale size ranges. In particular, this concerns particles < 10–15 nm and irregular in size distribution which affects greatly such material properties as roughness, friction and surface dislocation [17].

Materials with a particular 3D nanostructure can be generated either by coating of the surface with nanoparticles or modifcation of the surface by chemical and physical treatments. Combining these methods is even more effective. A controllable release of nanoparticles to the surface of the implant creates an additional modality for surface coating which helps to reduce potential toxicity.

Different materials could be proposed for nanocoating of urinary implants with each of them possessing certain advantages and disadvantages as to their applicability and functionality. Polymer-based and lipid nanoparticles (i.e. liposomes) are usually biocompatible and can be easily biodegradable. They are suitable for transportation of biologically active compounds to the cells due to their similarity to natural nanoparticles—extracellular vesicles, which take part in cell-cell communication [18].

Carbon nanotubes and fullerenes likewise possess a high biocompatibility since carbon is one of the main components of living tissues. They are smaller in size and have a high surface-to-volume ratio. Since carbon nanoparticles can form complex three-dimensional structures, they are especially useful for targeted delivery of particular bioactive substances.

Metal-based nanoparticles produced usually from transition metals like zinc, gold, silver and copper, are described broadly in the scientifc literature. However, despite demonstrated good antimicrobial properties their antifungal and anti-viral effciencies have not been established.

Toxicity of nanocoating particles has not been a problem so far [19]. However, a deep understanding is required due to the wide range of nanomaterials proposed for medical application and the great variety of their physico-chemical properties and ensuing potential effects on biological tissues.

## **5 Tissue Engineering and Regenerative Medicine as Future Research Lines**

This research line has preliminarily looked at urinary stents in the lower urinary tract. Stents were available for the treatment of lower urinary tract obstruction from the late 1980s. Despite initial enthusiasm, further studies questioned their usefulness as a primary treatment for urethral stricture disease [20]. Fibrotic tissue ingrowth is the main complication in permanently implanted stents but it can occur even in temporarily indwelling stents.

An ideal urinary stent should promote tissue healing, reduce fbrosis and scar formation, and maintain the physiological functions of the lower urinary tract. In addition, it would be desirable if they would be biodegradable and could stimulate their gradual replacement by regenerated autologous tissue.

Even though this is yet wishful thinking, some progress has been made in the feld of vascular and coronary artery stenting where implantation of complex tissue composites rather than a simple stent is required. How can this be translated into urology, and more specifcally into the treatment of lower urinary tract stricture disease?

There are few experimental trials on the designing of urethral stents using tissue engineering approaches [21, 22]. To our knowledge though, these clinical and preclinical applications have not proceeded beyond phase two trials.

The main problem of tissue engineered grafts or stents is the lack of appropriate acceptance by the host leading to fbrosis or rejection. If the wound bed consists of intact and viable tissue, transplantation of epithelium only might be suffcient [23]. In cases of fbrosis or developmental defects, vascularized grafts should be generated, e.g. by populating scaffolds with endothelial cells of endothelial progenitors [24–26].

In contrast to vascular stents, where the cells required for tissue regeneration can migrate from the blood, in urinary stents the recruitment of the cells from urine is rather unlikely though not impossible. If the urethral graft is well vascularized, blood fow might deliver important cells and cell progenitors to the sites of tissue

regeneration. To make this approach clinically relevant, comprehensive studies on the role of infammatory cells in tissue regeneration (which must be induced) and fbrosis (which must be prevented) are required.

In addition to *in situ* approaches where biodegradable urinary stents could be gradually replaced by regenerated tissue, we must consider cases where the tissue damage caused by trauma or disease is beyond the regenerative potential of the organism. In these cases, reconstruction of parts of the urinary tract must be done completely in vitro, using synthetic or naturally derived scaffolds [27, 28], manufactured by 3D printing [29] and electrostatic spinning [30], and populated by different types of cells (including autologous cells and progenitors) [31, 32]. Application of endothelial cell-attracting extracellular matrix into a graft scaffold and pre-vascularization of a transplant on vasculature-rich tissues (such as omentum) [31], as well as engineering of perfusable blood vessels of the transplant in vitro [33, 34] could be used for vascularization of the grafts. Combination of a "cell sheet technology" with bioprinting of cellularized scaffolds has a great potential for future perspectives of engineering of urinary stents.

Another completely different approach based on a "let mother nature do the job" paradigm can also be considered for designing urinary stents. The fundamental discovery of iPS cells [35] and the not yet that well-known tissue engineering technology of "blastocyst complementation" [36, 37] cumulatively might open the possibility to grow patient-specifc organs (including different components of the lower urinary tract) in humanized animals. With this technique, certain gene modifcations in the host organism prevents formation of the target organ during embryogenesis. Patient-derived iPS cells delivered to the blastocysts of immunocompromised animals will rescue the developmental program and form the target organ which will consist of human cells. After maturation in the animal host, the complete tissue/ organ or its parts can be transplanted to the patient. As an alternative to the "blastocyst complementation" assay, based on the use of "humanized animals", it is also proposed to transplant specifc human progenitor cells to actual sites of organogenesis thus leading to formation of target organs that consist of patient-derived cells [38]. These futuristic approaches, however, have not been tested in preclinical trials and can only be considered as a proof of concept to date.

#### **6 Conclusions**

In this chapter we have summarised modern biological approaches to improve the structure, function and performance of urinary stents. Some have been already applied in urinary stent production whilst others have been tested in the feld of vascular stents, such as antibody or biomimetic coating. Bioengineering approaches aiming at the generation of complete analogs of damaged urinary tissue from autologous patient-derived cells represent a more futuristic outlook. Nevertheless, we hope that the rapid development of advanced multidisciplinary research platforms in modern biomedicine will make these approaches feasible in the near future.

#### **References**

**Open Access** This chapter is licensed under the terms of the Creative Commons Attribution 4.0 International License (http://creativecommons.org/licenses/by/4.0/), which permits use, sharing, adaptation, distribution and reproduction in any medium or format, as long as you give appropriate credit to the original author(s) and the source, provide a link to the Creative Commons license and indicate if changes were made.

The images or other third party material in this chapter are included in the chapter's Creative Commons license, unless indicated otherwise in a credit line to the material. If material is not included in the chapter's Creative Commons license and your intended use is not permitted by statutory regulation or exceeds the permitted use, you will need to obtain permission directly from the copyright holder.

# **Preventing Bioflm Formation and Encrustation on Urinary Implants: (Bio)molecular and Physical Research Approaches**

**Ali Abou-Hassan, Alexandre A. Barros, Noor Buchholz, Dario Carugo, Francesco Clavica, Filipe Mergulhao, and Shaokai Zheng**

# **1 Introduction**

Stents and catheters are used to facilitate urine drainage within the urinary system [1]. When such sterile implants are inserted into the urinary tract, ions, macromolecules and bacteria from urine, blood or underlying tissues accumulate on their surface. This often results in the formation of bioflm causing infections that can

A. Abou-Hassan (\*)

Sorbonne Université, CNRS UMR 8234, PHysico-chimie des Électrolytes et Nanosystèmes InterfaciauX, Paris, France e-mail<EMAIL_ADDRESS>
A. A. Barros 3B's Research Group, University of Minho, Barco Guimãraes, Portugal e-mail<EMAIL_ADDRESS>
N. Buchholz U-Merge Scientifc Offce, London-Athens-Dubai, Athens, Greece

D. Carugo Department of Pharmaceutics, UCL School of Pharmacy, University College London, London, UK e-mail<EMAIL_ADDRESS>
F. Clavica · S. Zheng ARTORG Center for Biomedical Engineering Research, Faculty of Medicine, University of Bern, Bern, Switzerland e-mail<EMAIL_ADDRESS><EMAIL_ADDRESS>
F. Mergulhao LEPABE, Faculty of Engineering, University of Porto, Porto, Portugal e-mail<EMAIL_ADDRESS>
A. Abou-Hassan, Barros, N. Buchholz, D. Carugo, F. Clavica, F. Mergulhao and S. Zheng contributed equally with all other contributors.

be responsible for discomfort and complications in patients [2]. Therefore, urinary tract implants may be considered as out-of-equilibrium systems where different phenomena acting at various times and length scales occur. This leads to their reshaping by deposition and encrustation of chemical and biological species on their surface and the formation of bacterial bioflms and mineral crystals. Due to the continuous nature of biological fuids, these phenomena are in perpetual dynamics and self-organization, which can complicate their study in the human body [3, 4]. Giving this complexity of the "system", a multidisciplinary input with different scientifc approaches is needed to better understand and fnd solutions to this problem. In this chapter, we outline research strategies addressing biocompatibility, the use of antisense molecules, non-pathogenic bacteria and bacteriophages, and physical methods to prevent or inhibit bioflm formation and encrustation.

#### **2 Biodegradable Metal Stents**

Biodegradable metals are very appealing for urinary stent applications since they combine enhanced radial strength with a prolonged but controlled degradation time. Therefore, biodegradable metallic urinary stents (BMUS) constitute a promising research strategy overcoming some of the current stent limitations. The potential of biodegradable metals for urological applications was explored frst by Lock et al., who investigated the effcacy of Magnesium–4%Yttrium (Mg–4Y), AZ31, and commercially pure Mg as antibacterial BMUS. They showed a signifcant decrease of *E. coli* viability in the presence of Mg alloys after 3 days compared with a commonly used commercially available polyurethane stent [5]. Zhang et al. explored the potential of the ZK60 Mg alloy and pure Mg for urinary applications in a rat model. ZK60 had a faster degradation rate than pure Mg and neither of the metals showed toxicity during the three weeks implantation time [6]. More recently, Tie et al. used a Mg alloy in a large animal model (*Guangxi Bama Minipig*) as a BMUS. The Mg alloy (ZJ31) presented a homogeneous degradation, excellent biocompatibility and antibacterial activity compared with stainless steel—a commonly used material for non-degradable metallic urinary stents [7].

Zinc (Zn) has a slower degradation time than Mg, with low tissue toxicity and good antibacterial activity. Champagne et al. compared pure Zn, Zn–0.5 Mg, Zn–1 Mg, Zn–0.5% aluminium (Zn–0.5Al), pure Mg and Mg–2Zn–1% manganese (Mg–2Zn1Mn), a commercially available Mg alloy. Zn-containing metals degraded more slowly, and more homogeneous corrosion was obtained for Zn–0.5Al [8].

Biodegradable metal stents in urology have only been explored by a few research groups to date, but these have shown good potential in terms of improved biocompatibility and antibacterial activity.

Mg has been studied the most, but Zn is another promising component. An alloy of both might combine the good biocompatibility and antibacterial properties of Mg with the slower degradation and increased homogeneity of Zn.

## **3 Molecular and Biological Approaches to Prevent Bioflm Formation and Encrustation**

#### *3.1 Antisense Molecules*

Pathological protein synthesis, either as under- or overproduction, is a crucial part in many disease processes. Halting these abnormal syntheses might open the disease to effective therapies otherwise not available [9]. Antisense technology (AT) has been investigated for malignant, infectious, infammatory and metabolic diseases [10]. It modulates protein synthesis by inhibiting gene expression through pairing an antisense nucleic acid sequence base with its complementary sense RNA strand. This stops translation into the target protein [9]. In addition, it can disturb other functions of RNA molecules such as splicing, folding, protein binding, microRNA activities, and RNA-mediated telomerase action [11]. AT is very target specifc. For researchers, it is highly interesting since more and more underlying molecular pathways are getting identifed for major diseases offering new opportunities to interfere in these. However, AT is not mature enough to overcome some inherent problems such as limited *in-vivo* stability, mode of application, and potential side effects [10]. AT may be used to address urinary implant contamination, infection, and bioflm formation (BF). Bioflms contain extracellular polysaccharides and nucleic acids. The presence of extracellular polysaccharides results in well-structured and strong bioflms [12], which in turn makes bacteria embedded within the bioflm up to 5000 times more antibiotic-resistant [13, 14]. Stopping the synthesis of extracellular polysaccharides can stop BF and/or weaken their bacteriaprotective structure. AT would represent an early intervention whilst bioflm is still forming [15]. Especially in chronic infections with BF targeted by AT in its early stages, antibiotics can remain effective and remove or stop further bioflm activity [16].

Recently, it has been shown that the common urinary bacterium *Enterococcus faecalis* gene (efaA) is crucial in BF. Anti-sense *efaA* peptide nucleic acids could decrease it [17]. Research on AT to manipulate BF has only just begun. First the the genetic aspects governing bacterial BF processes need to be better understood [18]. Nevertheless, since one major drawback of urinary stents is the BF, antisense technology may be a promising approach to tackle this inherent stent problem in the future.

#### *3.2 Non-Pathogenic Bacteria*

Non-pathogenic bacteria can be used to reduce bioflm formation by pathogenic organisms using various mechanisms such as displacement, exclusion, and competition. In the displacement strategy, non-pathogenic cells or their metabolites disrupt the structure of a pre-formed pathogenic bioflm. Alternatively, pathogen exclusion can occur by blocking adhesion sites, and competition for nutrients or growth factors can inhibit the development of pathogenic strains [19]. Additionally, nonpathogenic bacteria can also modulate the immune system affecting pathogenic cells.

Non-pathogenic bacteria are able to produce a range of compounds, including biosurfactants, bacteriocins and extracellular polymeric surfaces (EPS), that can be detrimental to the development of pathogenic organisms or affect their adhesion to a surface. It has been shown that the production of biosurfactants may interfere with the microbial adhesion of pathogens, including those that are found in the urinary tract [20]. Bacteriocins were also shown to be useful given their high potency, stability, and low toxicity [21–23]. EPS comprises a large group of high-molecularweight polymers produced by different metabolic pathways in various organisms with proven antibioflm properties [24]. The production of a vast array of molecules, including lactic acid, fatty acids, enzymes, and hydrogen peroxide, with the potential to control pathogenic bioflms, has also been identifed in non-pathogenic cells [19].

Compared to other coating strategies, the use of non-pathogenic cells to coat medical devices may be advantageous because the coating is alive. This allows for the self-renewal of the anti-pathogenic activity, whereas conventional coatings eventually become covered by biomass which may reduce their effectiveness [25].

A number of hurdles have to be overcome for the broad application of nonpathogenic bacteria to protect the surface of urological stent implants. For example, although it has been shown that a certain degree of protection can be obtained for short time periods, the stability and activity of the coating for longer periods of time must be carefully assessed. If the protective effect relies on the viability of the nonpathogenic bacteria (for instance, to produce interfering molecules), this can be an issue. Also, if translocation of the non-pathogenic bioflm occurs (for instance, due to shear forces caused by urine fow), the coating effcacy can be compromised.

#### *3.3 Bacteriophages*

Viruses that use bacteria as their hosts are called bacteriophages. Whilst duplicating in the bacteria, they disrupt the metabolism of their hosts in several ways. Lytic bacteriophages destroy the host cell membranes and cells. Lysogenic bacteriophages use the functioning bacterium to multiply whilst letting it live. Lytic bacteriophages can therefore function as antibacterial agents. They are readily available, selective as to their hosts, and non-toxic for surrounding tissue cells. As a consequence, they have been discussed as a coating constituent for medical implants [26]. In early experiments, bladder catheters were pre-treated with lytic *Staphylococcus epidermidis bacteriophage 456*. This led to a signifcant decrease in intraluminal bioflm formation [27].

As with antibiotics, bacteria can become resistant to bacteriophages. This may be overcome using a mixture of several lytic viruses [28–30]. Silicone bladder catheters coated with hydrogel and pre-treated with such a mixture were indeed effcient

against multi-bacterial bioflms. It was proposed that such coating could tackle multi-bacterial bioflms by adapting the viral mixture [31].

Whilst the available evidence stems from experiments on bladder catheters only, the use of bacteriophages on urinary stents seems intuitive and promising. More importantly, since we live in an era of increasing and complex global bacterial resistance to antimicrobials, bacteriophages might represent an alternative approach in the future.

## **4 Physical Strategies to Prevent Bioflm Formation and Encrustation**

#### *4.1 Electrical Charges*

The role of electrostatic charge is pivotal in bacterial adhesion. Most bacterial genera have a net negative charge as determined from quantifcation of their zetapotential. Therefore, two types of engagement can be derived as antibacterial strategies, namely, material as repellent or as contact-killing agent. The frst strategy implies that materials with high negative charge can be deployed as anti-bacterial stent material or coating to repel bacterial cells [32]. Heparin, having the highest negative charge density of known biological molecules [33], has been a popular candidate as stent coating material. However, its effcacy against bioflm formation has been controversial [33–35]. The second strategy relies on a positively charged surface and permeabilization of the bacterial cell membrane that leads to the leakage of intracellular material and eventual cell death. One approach involves grafting polyethyleneimine (PEI) [36] micro-brushes onto polyurethane stents followed by an alkylation process [37, 38]. The resulting micro-structure of PEI brushes with positive charges showed a reduction in both bioflm and encrustation development in *in vitro* and *in vivo* experiments [37]. Another choice of positively charged material is chitosan, which works as an antimicrobial against fungi, Gram-positive and Gram-negative bacteria through various modes of actions [39, 40]. A freezecasting process was proposed to make entire ureteral stents from chitosan [41]. The best material and engagement strategy are yet to be concluded from further investigations.

#### *4.2 Enhancing and Maintaining Ureteral Peristalsis*

In physiological conditions, ureteral peristalsis moves the urine from the renal pelvis to the bladder. Although the insertion of ureteral stents can initially increase ureteral peristalsis, indwelling stents eventually lead to its cessation [42]. The mechanisms leading to aperistalsis in stented ureters are still unclear. A few models

[43] and experimental studies reported the possibility of artifcially inducing ureteral peristalsis by: (1) electrical stimulation [44–46], (2) mechanical stimulation e.g. applying distension and/or (3) pharmacological treatment [45]. A possible strategy against encrustation and bioflm in stented ureters could be based on the 'fushing effect' of the ureteral peristalsis: if peristalsis can be preserved in stented ureters in the long term, the movements of the ureteral wall could eliminate encrustations and bacterial deposits from the stent surface. Haeberlin et al. demonstrated that peristalsis can be electrically induced in stented ureters. Catheters were inserted into *ex-vivo* ureters (the size of the catheters was comparable to conventional ureteral stents) and propagating contractions of the ureteral wall were observed after each electrical stimulation [46]. Since these experiments were only conducted in *ex-vivo* ureters in the short term (up to 3 h), *in-vivo* experiments are required to demonstrate the possibility of long-term preservation of the peristalsis by artifcial electrical stimulation.

#### *4.3 Ultrasound Waves*

Ultrasound comprises longitudinal pressure waves with a frequency > 20 kHz. It represents a clinically viable modality of delivering mechanical stimulation within the body to achieve both therapeutic and diagnostic outcomes. It has also been demonstrated that ultrasound exposure can cause detachment of bacterial bioflms from different surface types and can promote the transport of antibiotics into planktonic or bioflm-forming bacterial cells [47]. Surface acoustic waves (SAW) are a type of sound waves that are transmitted along a surface, and the resulting vibrations have been identifed as a factor reducing bacterial adhesion onto solid surfaces [48]. This approach has been adopted to counteract bacterial bioflm formation in bladder catheters, whereby an ultrasound transducer is coupled with the extracorporeal segment of the catheter. Upon activation, the transducer generates SAWs in the frequency range 100–300 kHz, resulting in surface oscillations with amplitudes of 0.2–2 nm that propagate over the catheter surface. In previous studies in a rabbit model, this method has been evaluated both *in vitro* and *in vivo* for inhibiting bacterial adhesion on Foley bladder catheters. It has been shown that SAW-activated catheters had a signifcantly lower bioflm load *in vitro,* and that this effect was greater when lower SAW intensities were employed (in the range 0.05–0.20 mW/ cm2 ). These fndings were confrmed *in vivo*, where the average number of days until the development of a urinary tract infection was extended to 7.3 ± 1.3 days in the SAW-catheter group, compared to 1.5 ± 0.6 days in a non-treated, control group [49].

A commercially available SAW-activated catheter (UroShieldTM) has been developed by NanoVibronix Inc. (USA). Zillich et al. investigated its effcacy and safety through a randomized, double blinded clinical study on 22 patients, in which catheters were deployed for an average of 9 days. Patients having the UroShieldTM catheter reported less pain and bladder spasm, and showed a marked reduction in bioflm formation [50]. More recently, a double blinded randomized controlled trial assessed 55 patients who had an indwelling urinary or suprapubic catheter for > 1-year, and had a treated urinary tract infection during 90 days prior to the commencement of the study. The large majority of patients having the SAW-activated UroShieldTM showed a signifcant reduction in bacterial load compared to the control group [51]. To the best of our knowledge, a similar approach has not yet been investigated for ureteric stents. Given the demonstrated effcacy of SAW on stent encrustation, this surely represents an interesting future research strategy. When developing such a strategy, there are some technical aspects to be considered. As we are dealing with fully intracorporeal devices, remote powering and control of the SAW activation, and a careful investigation into the propagation properties of the stent materials is needed. For the latter, geometrical features of a stent (e.g., presence of side holes) that may affect SAW propagation and the resulting surface displacement feld must be considered, too.

#### *4.4 Biosensors*

One major problem of urinary stents and catheters is blockage by encrustation. If blockage occurs in a bladder catheter, it causes painful retention of urine and can provoke severe urinary tract infection and urosepsis. Often, the blockage results from urine infection with urease producing organisms, predominantly *Proteus mirabilis*. Urease generates ammonium which leads to an elevation of urinary pH. This leads to the precipitation of struvite and apatite, which then form a crystalline bioflm encrusting and blocking the urinary catheter. Biosensors are sensors that would alert patients and carers early of an ongoing encrustation and impending resulting blockage. A survey of the current literature shows that such sensors are mainly visual. A pH sensor based on a silicone-based strip incorporating a pH indicator (bromothymol blue) was integrated into an indwelling urinary catheter [52]. A change from yellow to blue indicating impending blockage occurred 19 days before the actual blockage in early human trials. Catheters can also be designed to integrate a pH dependent luminescent material [53]. A lanthanide (Eu) pH-responsive probe that can be incorporated in a hydrogel catheter coating was described. Upon elevation of pH in the presence of urease, the luminescence turns off. However, the system was not tested neither *in vivo* nor *in vitro*. Another approach to provide early warning of encrustation and blockage is to associate a 'trigger' layer, usually EUDRAGIT®S 100, on a hydrogel layer encapsulating a pH reporter or antibacterial agent [54, 55]. Upon elevation of the urinary pH, the upper layer dissolves, triggering the release of a pH indicator such as carboxyfuorescein or bacteriophages. Both approaches were tested in an *in vitro* bladder model, which provided a 12 h advanced warning of blockage and a 13–26 h advanced warning of delayed catheter blockage. The above are early and simple examples of pH-indicating visual biosensors. Because they are visual, they will only work on catheters where an extracorporeal part remains visible. However, the idea of biosensors to indicate early stent

encrustation is an appealing one. Fully intracorporeal stents could be equipped with micro- or nano-technological wireless sensors for the same purpose.

#### **5 Conclusions**

In this chapter, we presented a brief but comprehensive overview of future research strategies in the prevention of urinary device encrustation with an emphasis on biodegradability, molecular, microbiological and physical research approaches. The large and strongly associated feld of stent coatings and tissue engineering is outlined elsewhere in this book.

There is still plenty of room for future investigations in the felds of material science, surface science, and biomedical engineering to improve and create the most effective urinary implants. In an era where material science, robotics and artifcial intelligence have undergone great progress, futuristic ideas may become a reality. These ideas include the creation of multifunctional programmable intelligent urinary implants (core and surface) capable to adapt to the complex biological and physiological environment through sensing or by algorithms from artifcial intelligence included in the implant. Urinary implants are at the crossroads of several scientifc disciplines, and progress will only be achieved if scientists and physicians collaborate using basic and applied scientifc approaches.

#### **References**

**Open Access** This chapter is licensed under the terms of the Creative Commons Attribution 4.0 International License (http://creativecommons.org/licenses/by/4.0/), which permits use, sharing, adaptation, distribution and reproduction in any medium or format, as long as you give appropriate credit to the original author(s) and the source, provide a link to the Creative Commons license and indicate if changes were made.

The images or other third party material in this chapter are included in the chapter's Creative Commons license, unless indicated otherwise in a credit line to the material. If material is not included in the chapter's Creative Commons license and your intended use is not permitted by statutory regulation or exceeds the permitted use, you will need to obtain permission directly from the copyright holder.

# **Ten Steps to Strategic Planning for the Urinary Stents of the Future**

**Federico Soria**

# **1 Introduction**

To summarise all the knowledge in the current book and to allow its use both at clinical practise and its application in patients, as well as in the improvement of urinary stents, the simplest way is to build a decalogue that provides a global vision of the requirements for the improvement of these medical devices.

# **2 Understanding the Side Effects and Complications Related to Urinary Stents**

An in-depth knowledge of the side effects, complications, their pathophysiology and, above all, the etiopathogenesis associated with urinary stents is essential on the way to reduce the effects on patients, as well as to improve urinary stents. This knowledge allows urologists to identify symptoms early, as well as to arrange therapeutic measures to alleviate these symptoms [1]. Mainly antimicrobials, alphablockers or antimuscarinics to reduce discomfort in the lower urinary tract and, of course, analgesics. Knowledge and research into the etiopathogenesis of each of the adverse effects allows researchers to focus their research [2]. The detection of the cause of each adverse effect allows the identifcation of whether it is caused by the stent design itself, by the biomaterial or by a weak coating; these three factors are responsible for the majority of adverse effects related to urinary stents. We

F. Soria (\*)

Foundation, Jesús Usón Minimally Invasive Surgery Center, Cáceres, Spain e-mail<EMAIL_ADDRESS>
differentiate between adverse effects, which we consider inherent to the urinary stents themselves, such as vesicoureteral refux, bioflm formation, and complications which, although related to the stents, are due to a malfunction of these medical devices. Among these complications, migrations, perforation, etc. are the most important. Therefore, the frst factor to take into account is always knowledge of the adverse effects and complications produced by stents [3].

#### **3 Proper Indication for the Use of Urinary Stents**

It is clear that the simplest way to reduce the harmful effects associated with urinary stents is to reduce their use. This is the frst choice in the face of the high percentage of associated complications. Since it is impossible to avoid their use due to their evident benefcial effects on patients, it would be necessary to determine in which type of patients their use outweighs the adverse effects. Unfortunately, this is currently the case with the use of metallic stents both at the ureteral or urethral area, with very high complication rates; their use is reduced to a very small number of patients and in many cases exclusively as a palliative treatment.

According to the current scientifc literature, the use of ureteral stents after endourological treatment of ureteral or renal lithiasis is approximately 80%. This is a very high percentage, which means that the population susceptible to stent-related problems is very high. Unfortunately, both European and American guidelines cannot defne with great scientifc evidence the indications for urinary stenting. Stenting is well indicated in complicated ureteroscopies, but the difference between a complicated URS and a non-complicated URS always remains the surgeon's decision. As a result, since there is no criteria for deciding when it is mandatory to place a stent and, above all, when it is not mandatory, the use of these devices is on the rise. Although it is true that a great advance in this aspect is that stenting times have been reduced in an attempt to reduce adverse effects [4]. These effects and complications have been shown to be signifcantly related to the stenting time, increasing adversely in prolonged stenting times, generally longer than 6 weeks [5].

So a decrease in their use or at least a shortening of the stenting time, without delays in the removal date, would be associated with a better quality of life for patients.

#### **4 Biomaterials**

Another of the cornerstones on which the improvement of current urinary stents is based is the research being carried out on biomaterials that allow their use in the urinary tract. As can be seen, this point is critical, as the weaknesses demonstrated by the polymers, metals or their alloys currently in use are one of the main reasons for encrustation, bacterial and even fungal contamination, and sometimes stent fracture. The development of new biomaterials with better characteristics and suitable for the urinary environment will reduce these side effects. Certainly, the future of biomaterials to overcome the limitations they present in the urinary environment depends fundamentally on three factors. The frst is to improve their mechanical properties in order to be effective in extrinsic strictures of malignant origin. Secondly, combining biomaterials in the same stent to combine the advantages of each, reducing their weaknesses. Finally, it is possible to coat the biomaterials so that they are not in contact with urine, so that only the coating is affected and the inner part keeps all its properties intact [6].

#### **5 Coatings**

This area of knowledge is probably where most resources are being allocated, as research into coatings that prevent or reduce bioflm formation is an issue that involves not only urinary stents, but virtually all implantable medical devices, catheters, prostheses, implants, catheters, etc. Thanks to coatings it is possible to isolate the rest of the biomaterials that make up the stent from the urine, as well as to combat the formation of bioflm that is associated with bacterial contamination as well as encrustation. Therefore, the search for new coatings is of great importance to improve the durability of urinary stents [7]. The aim of these coatings is to provide an "antibiotic free solution" to bioflm formation. To this end, a number of strategies have been developed, as described in the previous chapters. The use of agents with antimicrobial properties has been emphasised: metals such as Zn2+, Ag1+, CuO; superhydrophilic coatings, hydrogels [8]. Mainly in this section, AMPs, antimicrobial peptides, which are proteins with antimicrobial properties, stand out, especially CWR11, RK1 and RK2. Efforts are also being made to detect probiotics that compete against bioflm-forming bacteria and prevent their development. As well as agents with anti-adhesive properties that prevent bacterial adhesion to the stent surface by preventing the action of bacterial adhesins, thanks to bacteriocins [9].

#### **6 New Designs**

Another essential element in the improvement of stents and thus the decrease in the adverse effects associated with their use in the urinary tract is the development of new urinary stent designs. It is noteworthy that the design of pigtail ureteral stents has remained virtually unchanged over the last four decades, despite their obvious side effects. Many efforts have been made to reduce the effects related to the bladder pigtail, which is associated with dysuria and LUTS. Therefore modifcations of this pigtail, reducing its size, changing its conformation, have been presented and evaluated in patient trials. Despite the decrease in patient discomfort, they have not demonstrated clear scientifc evidence, and their use is currently not established in daily clinical practice [10]. To prevent vesicoureteral refux, stents with anti-refux systems have also been designed, which have not shown a clear improvement over conventional ureteral stents [11]. However, the development of intraureteral stents, or stents with a small bladder tip to facilitate their removal, has shown scientifc evidence regarding the improvement in the quality of life of patients, making them a very interesting option for the present and future for certain patients [12, 13]. Magnetic ureteral stents for removal without the need for cystoscopy have also shown less painful and faster removal [14].

A further design innovation that have proven to be very useful and that were unthinkable decades ago is the possibility of removing metallic, ureteral or urethral stents. This design improvement is extremely attractive and broadens the indications for these stents in the urinary tract [15]. As has been seen in recent years and is expected in the coming years, design variability will reduce the discomfort associated with current designs. The goal is to personalise stents for each patient. The availability of more stent designs will allow choice, which with current plastic stents is almost impossible at the moment.

#### **7 Biodegradable Stents (BUS)**

One of the premises for the development of urinary stents is that they should all be biodegradable. In order to achieve the requirements that defnes an ideal stent. It is diffcult to understand that in the twenty-frst century a surgical procedure is necessary to remove a stent. Avoiding cystoscopic removal of stents, avoiding anesthesia in pediatrics patients and avoiding the "forgotten stent" are short-term goals [16]. The development of BUS has expanded in recent years because the most important limitations in its development have been overcome. Firstly, the control of degradation, making this rate controllable thanks to the selection of polymers and copolymers, natural–synthetic or metallic, and above all the use of combinations of different biomaterials with different degradation rates [17]. On the other hand, the control of degradation fragments is a key limitation, since this is a major drawback, in particular when this type of stent is placed at the ureteral lumen level. BUSs must degrade gradually and fragment into small pieces smaller than 2 mm to ease their evacuation. A strategy that has been described for this type of stent and that is related to its design is the ability to degrade from distal to proximal, so that, despite degradation, the stent continues to perform its function as an internal scaffold [18, 19]. One of the most important current challenges is the preservation of the mechanical properties of the BUS, regardless of the nature of the biomaterials that comprise it. Therefore, a balance between the rate of degradation and the maintenance of the mechanical properties of the stents is necessary, which is of great importance in ureteral stents, but is completely mandatory for segmental stents at the urethral level, when used as an internal scaffold, cellularised or not, after treatment of urethral strictures.

#### **8 Drug Eluting Stents (DES)**

In an attempt to reduce lumen restenosis after vascular stenting, DESs were introduced. Neointimal hyperplasia resulted in in-stent reestenosis in 20–30% of cases after intervention with bare metallic stents. DES were developed not only to act as vascular scaffolds in the diseased coronary artery but also to reduce the relatively high rates of "in-stent reestenosis" and subsequent "target lesion revascularization" compared to its predecessor Bare Metallic stents. DESs have the potential of endoluminal release of pharmacological anti-proliferative substances and reduce the hyperplastic reaction by inhibiting the smooth muscle cell cycle and their proliferation. With the excellent background of vascular stents, applications in the urinary tract are a very encouraging feld of development. The idea is to take advantage of the stent to add such an innovative and promising feature as local drug delivery. In this regard, local release of anti-infammatory, analgesic, or even antiproliferative drugs to reduce urothelial hyperplasia related to urinary metallic stents, or chemotherapy in the upper urinary tract are some of the drugs that have been evaluated [20].

This delivery system would avoid systemic drugs side effects. Possibly reduce the total daily drug dose. As well as using drugs with a short half-life. An important factor is that this urinary delivery system avoids drug absorption, distribution and metabolism, as the urinary tract is a watertight system with low absorption capacity and the drugs are constantly eliminated through urination. A very important consideration is that with these delivery systems, there is the possibility of maintaining urine drug levels in the optimal therapeutic range [21, 22]. Compared to current bladder or pyeloureteral instillation systems, the improvement of patient satisfaction is to be expected.

#### **9 Urine and Infection**

The association between UTI and urinary stents, mainly ureteral stents, is one of the most common complications in patients. It should not be forgotten that the prevalence of bacterial colonisation of urinary stents is between 42 and 90%, leading to the development of bacteriuria and UTI [23]. One of the current problems, which needs urgent evaluation to allow for the improvement of stents, is related to the laboratory techniques used for the quantifcation or detection of urinary bacteria. Regarding the analysis of bioflm and bacteriuria associated with CDJs, despite a low sensitivity of 21–40% and a specifcity of 46–64%, culture is the method of choice for detecting bacterial colonisation of the stent and asymptomatic bacteriuria. The frst by direct culture of stent fragments and the latter by culture of urine samples [3, 23, 24]. As a result, there is currently no consensus on the ideal microbiological technique to make a fast and, above all, consensual determination that allows for the standardisation of clinical and experimental studies.

It is evident that the aim with current stents is that they remain in place for as long as necessary, since the rates of colonisation and bacteriuria increase considerably with time [3, 23–25]. With regard to bioflm formation, an incidence of 34–66% is found when the stent remains in place for less than 2 months, compared to 75–100% when the stent remains in place for more than 3 months [23, 25]. The incidence of asymptomatic bacteriuria ranges from 7 to 33% in patients with less than 1 month of stenting, 21–50% between 1 and 3 months and when the stent remains in place for more than 3 months, the incidence can reach up to 54% [24, 25]. With regard to the bacterial strains that make up the bioflm and those present in the urine, a large discrepancy has been demonstrated between stent cultures and urine cultures, showing that there is no direct correlation between the bacteria that colonise the stents and those that cause UTI [25].

Unfortunately, the source of colonisation is unlikely to be eradicated, as contamination is mainly at the time of insertion, through skin bacteria or the urethral microfora itself. This could justify the fact that in some series a double incidence of colonisation is observed in women compared to men, given the short length of their urethra and the risk of contamination; as well as the highly frequent presence of Gram+ bacteria in the stents, bacteria commonly present in the distal urethra and vaginal fora.

This susceptibility of urinary stents to contamination, being aware that urine has its own microbiome, must be taken into account in both clinical and research settings. The development of new biomaterials and coatings with antimicrobial properties is therefore one of the milestones for the development of safe and more effective stents. Especially since antibiotic prophylaxis has not shown clear scientifc evidence in reducing colonisation of urinary stents. Therefore, only with contamination prevention and strategies to reduce formation is it possible to make progress on this issue, as bacteria in a bioflm can usually survive the presence of antimicrobial agents at a concentration, 1000–1500 times higher than the concentration that kills planktonic cells of the same species.

#### **10 Drugs to Change the Composition of Urine**

One of the promising strategies that may reduce the side effects of urinary stents, mainly related to encrustation and possibly also bacterial contamination, is the possibility of changing the composition of the medium in which the stent is placed, which is the urine. The main efforts are being made to alter the composition by oral administration of compounds that modify the urinary pH. Modifcation of the urinary pH alone causes a very important change as it affects both microorganisms and the precipitation capacity of compounds that are dissolved in the urine and which, due to their supersaturation, can precipitate and cause incrustation on the surface of urinary stents. In addition to pH modifcation, it is possible to administer crystallisation inhibitors that signifcantly reduce the risk of lithiasis formation or incrustation [26].

This strategy has begun to show encouraging results in clinical studies evaluating potassium sodium hydrogen citrate, or l-methionine and phytin, reducing the occurrence of stent encrustation. In addition, the synergistic ability of many compounds may allow combinations of these compounds to achieve better results in this area. It is clear that urinary stent fouling is multifactorial, but within these causes the composition of the urine is the main factor that triggers this phenomenon, along with the composition of the stent [27]. The availability of this tool and the fact that it is so easy to apply, usually orally, and safe, suggests that this is an important way to reduce the adverse effects of stents. Not only adverse effects, but also future designs with biodegradable materials that can be modulated in this sense or to activate drug release in DES.

#### **11 Receptor-Based Stents and Tissue Engineering**

Another future strategy for the development of urinary stents is, as with DES, to make more proft from the device in the urinary tract. This attempt to expand the benefts of stents is aimed on the one side at obtaining data from the urinary tract, and on the other at allowing the stents to be bio-coated and to facilitate tissue engineering applications.

The development of stents with nano pressure sensors, which can provide information on intrapyelic or intravesical pressure, or with other sensors capable of stimulating ureteral peristalsis. The miniaturisation of this type of sensors allows them to be incorporated into the surface of the stent and send wireless information of great interest.

The possibility of coating stents to promote tissue regeneration, or proper healing, is one of the future hopes of research. In particular, their use would be extremely useful as a scaffold after the treatment of complicated stenosis, mainly at the urethral level. Biocovered stents could reduce fbrosis and thus the formation of stricture scars. Biodegradable biocovered stents would allow their function as an internal scaffold and cellular vehicle to be followed by their controlled disintegration [7].

#### **References**

<sup>1.</sup> Tomer N, Garden E, Small A, Palese M. Ureteral stent encrustation: epidemiology, pathophysiology, management and current technology. J Urol. 2021;205(1):68–77. https://doi.org/10.1097/ JU.0000000000001343.

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**6.3 Action, frequency, and variety: from bottom-up to top-down and the other way round**

"*Hoping less and acting mor*e" is sometimes a conclusion of philosophical reasoning on happiness [1]. Action is a condition of control over our life and adaptive behavior. The data reported in this chapter also indicate that freeing cognitive resources for action is important for happiness. The analyses mentioned in this chapter concerning brain default mode network suggest that self-oriented thought and rumination interfere with cognitive performance and positive affective experience. Affective regulation is costly. It removes resources that become unavailable for successful goaldirected behavior [137] and can undermine our perception of action opportunities [138]. Perceiving less opportunity for action could then be deleterious for effective action and for a global feeling of achievement and potential realization. Top-down pathways to happiness can be taken into account to prevent the consequences of nonrealistic expectations, broaden one's standpoint on life experience, and reappraise situations. The education of brain, cognition, and behavior could contribute to help individuals develop strategies namely supported by frontal brain areas, and aiming at i) organizing the conditions of positive experience (e.g., frequency and variety of positive events), and ii) focusing activity on action rather than on hopes.