1 Introduction

The term “stent” arose when in 1850s Charles Thomas Stent, a London dentist, developed a material used for dental impression, which was named “Stent's compound.”[1] During the First World War, this material was used to stabilize the skin grafts of soldiers,[2] and, since then, the word “stent” was extended to different medical specialties, including urology, to designate supporting devices. In 1949, when ureteral tubes were first described, they were made of polyethylene, a promising material due to its endurance and water-repellent nature.[3] Ever since, stent technology has undergone many developments regarding the bulk material, surface coatings, and design, mostly to tackle problems related to stent migration, infection, and encrustation. In the 1960s, silicone started to be explored as stent material due to its thermal resistance, which allows sterilization, and, consequently, decreases the potential for infections.[4] Advances on the stent surface, as hydrophobic and peptides coatings, were made to improve stent tolerability, and prevent encrustation and infection, by inhibiting bacterial adhesion.[5] Over the years, the ureteral stent design evolved from a straight tube to a tube with a distal bulb to prevent dislodgment, and, in the 1970s, to single J and double J stents.[6] Nowadays, the most common ureteral stents are 22–24 cm long flexible tubes made of polymeric materials, namely polyurethane or silicone, with strategic side-holes and double J ends.[6-8]

2 Clinical Need

In clinics, urological stenting is the standard procedure frequently done to restore compromised urological function caused by disorders of the urinary tract, such as kidney stones, strictures, and tumors.[7, 9] These devices are effective for the relief of upper urinary tract obstruction, prevention from stricture formation, support of ureteral healing, and management of urinary leakage.[10] Annually, over 1.5 million ureteral stents are used worldwide; however, it is estimated that more than 80% of patients suffer from a wide variety of stent-associated complications, whose prevalence is directly proportional to treatment duration.[11] Listing stent-related problems is key to identify current needs . This will guide the future directions for ureteral stent technology development and improve the overall stent function and patient care.

In long-term treatments, biofilm formation may occur, which may promote the onset of urinary tract infections (UTIs), encrustation and pain. In these cases, biofilm formation arises due to the deposition of a urinary conditioning film, composed by urinary proteins, ions and crystals at the stent's surface, which favors the interaction and adhesion of bacteria to the stent surface.[12, 13] Within biofilm environment, microorganisms are protected from the action of host defenses and antibiotics.[14] In a clinical study, it was found that the most common bacteria found on urinary stent surface biofilms are Escherichia coli, Enterococcus spp. and Staphylococcus species.[15] Bacterial colonization on stents can exceed 90% of patients, however only 20–45% of patients develop symptomatic urinary tract infection.[15-17] This pre-existing bacterial colonization can promote encrustation, a late complication that may affect the indwelling of the stent, and in some cases, might require surgery to facilitate stent removal.[18, 19] Reported encrustation development rates were 9.2%–26.8% before 6 weeks, 47.5%–56.9% between 6 and 12 weeks and 75.9%–76.3% thereafter.[10, 20] The most decisive factor for the appearance of encrustation is the stent dwell time, however different factors can affect stent encrustation and contribute to the heterogenous reported rates, including urine composition and pH, stent's material, surface topography, stent dwell time, and urine flow dynamics.[21] Besides bacterial infection and encrustation, different clinical disorders associated with stents were also reported, as physical distress, misplacement, stent fracture, and forgotten stent syndrome.[18] Moreover, around 58% of patients reported decreased work performance, and 32% expressed sexual dysfunction.[11] As a consequence of these hurdles, stent failure may occur, which leads to a significant negative impact on a patient's quality of life and represents an increase in healthcare economic burden.[11]

3 Advances in Ureteral Stent Technology

In order to reduce the abovementioned morbidity associated with ureteral stents and improve the performance of these devices, three key aspects need to be addressed, namely, constitutive material, coatings, and design of stents. In this section, these features will be reviewed and future directions for ureteral stent technology will be outlined.

3.1 Materials

To date, a gold standard material for ureteral stents in terms of mechanical performance, surface roughness, biocompatibility and cost-effectiveness fabrication has not yet been identified. Although there are biodegradable materials on the market, most stents are based on nondegradable polymeric and metallic materials, such as silicone, polyurethane, and nickel/titanium mixed alloys[8] (Figure 1).

Most common used materials for ureteral stents.

As previously stated, one of the main drawbacks associated to ureteral stent procedure is their long-term persistence causing patient discomfort and pain, and the need of a surgical procedure to remove the device. In this context, completely bioabsorbable stents may represent a solution, providing temporary urinary drainage and avoiding surgical removal.[22] In the last 20 years, many efforts were devoted to the development of completely biodegradable stents based on synthetic and natural polymers or metals.[5, 23-26] Among these, biodegradable shape memory polymers (SMPs) are an interesting class of materials for the ureteral stent field. Shape memory materials are defined as materials able to keep a temporary shape, as a consequence of the application of a stimulus, which may return afterwards to the original permanent shape. In 2009, Neffe and co-workers[27] reported a drug delivery ureteral stent based on biodegradable shape-memory polyurethane, where the device, due to its temporary shape, could be easily inserted, anchoring in situ after recovering to its permanent shape.

The biodegradable SMPs are associated with cutting-edge printing technology. 3D printing techniques were introduced in the 1980s and gained an ever-growing interest in several application fields, ranging from space to biomedical sciences. The 4D printing process was first introduced in 2014, by Skylar Tibbits,[28] to add time-dependent properties to 3D objects. This technique combines smart materials, such as the mentioned biodegradable SMPs, and additive manufacturing techniques for rapid fabrication using computer-aided design models.[28] Ly et al.[29] developed a protocol to obtain stents with a 1.75 mm filament starting from shape memory polyurethane mixed with carbon nanotubes which were processed by fused deposition modeling technique. The stimuli-responsive shape memory polyurethane was retained after 3D printing process, leading to a wider range of applications of SMPs and encouraging research in this direction. Among the additive manufacturing techniques, stereolithography allows to achieve high resolution substrate surface finishes, being limited by the narrow range of employable materials. Choong et al.[30] described the use of stimulus responsive tert-butyl acrylate (tBA)-co-di(ethylene glycol) diacrylate (DEGDA) network as shape memory resin for stereolithography technique. The stimulus-responsive mechanism arouses from the ultraviolet crosslinking of tBA monomers and DEGDA crosslinkers, using phenylbis (2,4,6-trimethylbenzoyl) phosphine oxide (BAPO), as photoinitiator. The shape memory performance of the printed objects was demonstrated, and optimal resin composition allowed to obtain full shape recovery. Ge et al.[31] synthesized different methacrylate polymers to develop a material platform with varied ranges of physicochemical and mechanical properties. Multiple SMPs, formed by blending different methacrylate constituents and forming strong interactions between them, were processed through projection microstereolithography in 3D structures, which exhibited interesting time-dependent properties. The first example of 4D stent-like structure was proposed by De Marco et al.,[32] by developing an indirect 3D printing approach, through the fabrication of a 3D sacrificial template exploiting the direct laser writing (DLW) technique. DLW enables the production of high-resolution 3D architectures with heights ranging from a few hundred nanometers up to several millimeters and layer thicknesses below 1 μm. A commercially available shape memory material (NOA63) was chosen to produce stents with two different complex structures. A positive photoresist was deposited on a silicon substrate, then the 3D shape was obtained through DLW equipment, followed by NOA63 casting and ultraviolet irradiation. Both stent-like structures showed a temperature-responsive behavior, being capable of recovering their original shape when immersed in water solution at 40 °C.

From the biomaterial engineering point of view, the 4D printing process represents an innovative tool able to overcome conventional manufacturing processes limitations. 4D printed stents were successfully produced for cardiovascular[33] and tracheal[34] applications. The processing of biodegradable SMP by 4D printing technique is a promising approach for low-cost mass production of ureteral stents, able to be easily applied, limiting patient's discomfort.

3.2 Coatings

The modification of the stent surface has been proposed to improve the stent performance and to add new functionalities. Among all the proposed strategies, here we will review the used materials for ureteral stent coatings, dividing them by their nature (organic/inorganic), highlighting the future perspective for each case. Table 1 summarizes the most relevant studies made on ureteral stent coatings and their major conclusions.

Table 1. Most relevant strategies for ureteral stent coatings Strategy Coating Material Major remarks Ref. Organic

Hydrogel

Hydrophilic

Poly(N,N-dimethylacrylamide) (PDMAA)

Poly(vinyl pyrollidone) (PVP)

3,4-Dihydroxyphenylalanine (DOPA)-conjugated polyethylene glycol (PEG)

Polyurethane

Silicone polyurethane

Higher lubricity and less friction. Significant reduction of the adherence of E. coli. Reduction of the side effects and complications on patients with hydrogel-coated stents.

Reduction of encrustation than the uncoated ones.

Reduction of attachment of uropathogens, in vitro. In vivo, using rabbit model, it was reported a reduction of 75% in the number of stent adherent E. coli.

[54, 55]a

[38]

[56]b

Hydrophobic Polytetrafluoroethylene (PTFE) Metal Avoid obstruction by urothelial hyperplasia, supporting its safety and applicability in ureteral stricture. [57]b), [58]a) Corethane Metal Prevention of obstruction in vivo. [59]b)

Biomolecule-based

“Living” coating

Immunoglobulins Polyurethane Reduction of E. coli adhesion. [42] Oxalate-degrading enzymes Silicone Reduction of encrustation without exhibiting toxicity. [43] Chondrocytes Polyglycolic acid mesh coated with 50:50 polylactic-co-glycolic acid Macroscopic examination of the engineered stents showed the presence of cartilaginous tissue. Biomechanical tests demonstrated that the cartilaginous cylinders were readily elastic and withstood high degrees of pressure. [44]b) Bladder epithelial cells poly (L-lactic acid) Promotion of cell proliferation, aiming for ureteral reconstruction. [45]b) Inorganic Amorphous carbon Diamond-like carbon Polyurethane Reduction of bacterial adhesion and prevention of struvite encrustation. In ten patients, DLC coating strongly limited the formation of an extended biofilm and showed a lower friction coefficient that further facilitated the placing and removal of the stent. [46]c), [60]a) Metals

Zinc oxide (ZnO) particles

Copper (Cu)-based coatings

Molybdenum disulfide (MoS2 ) and tungsten disulfide (WS2) nanostructured coatings, gold, SiO2, TiO2

Silicone/polyurethane Great antimicrobial activity and biocompatibility. Capability to limit encrustation. [48-50]b), [51] a)Clinical trial; b)In vivo study.

Among the organic materials used as coatings for ureteral stent surfaces, hydrophilic/hydrophobic polymers represent the majority. Nowadays, hydrophilic coatings are commercially available in the ureteral stent market, as AQ from Cook Medical, SL-6 from Applied Medical, HydroPlus from Boston Scientific, and heparin-based coating Endo-Sof Radiance, from Cook Medical. The working principle is based on the water trapping within the polymeric structure, which decreases the friction coefficient of the surface and prevents encrustation, by reducing adhesion phenomena at the biomaterial–tissue interface. [8][35] Nevertheless, various studies also pointed out that, due to the absorption of urinary solutes, hydrogel-coated stents could have the same[36] or even a higher risk[37] of becoming encrusted, comparing with uncoated ones made of the same substrate polymer. Moreover, other hydrophilic coatings, using polymers as the antifouling agents as poly(vinyl pyrollidone) (PVP) or 3,4-dihydroxyphenylalanine (DOPA) conjugated polyethylene glycol (PEG), are able to inhibit or destabilize biofilm formation, while also bestowing a beneficial lubricious effect.[38, 39] On the other hand, hydrophobic coatings, as polytetrafluoroethylene (PTFE) and corethane, have also been studied for ureteral stent coating application, and their effectiveness in preventing the luminal occlusion caused by urothelial hyperplasia has already been proven.[40, 41] Concerning organic approaches, the future research must involve more detailed studies regarding biomolecule-based coatings, which attempt to impart biomimetic and biocompatible properties to the stent surface. These strategies follow a similar rationale as the one made for the already approved heparin-based coating, however, apart from polysaccharides (heparin), other biomolecules can be applied. Among proteins, immunoglobulins[42] and oxalate-degrading enzymes[43] are interesting examples, being associated with less encrustation and reduced bacterial adhesion. For a quite innovative and distinct purpose, cells can also be used as “living” coating for ureteral stents.[44] The results demonstrated that this kind of design can promoted cell proliferation, suggesting that it could serve as alternative cell carrier for tissue engineered ureters.[45]

Among inorganic materials, the use of carbon-based materials as functional coatings for ureteral stent devices was also considered. Diamond-like carbon (DLC) coatings, especially due to their inert surface chemistry, were able to limit the formation of deposits and encrustations during long-time indwelling both in vitro and in vivo.[46, 47] As the UTIs in treated patients were significantly reduced, nowadays DLC coating is a commercial option in the ureteral stent market (Ureteral Stent Set–CarboSoft). In addition, different inorganic material-based solutions using metals are being proposed. Recently, zinc oxide (ZnO) gained considerable attention in the biomedical field due to its intrinsic antimicrobial activity and biocompatibility.[48] Copper (Cu)-based materials have also been investigated in the field of ureteral stent fabrication due to their antibacterial properties, and the prevention/limitation of encrustation, comparing with control samples.[49, 50] Other novel materials also recently considered include molybdenum disulfide (MoS2) and tungsten disulfide (WS2) nanostructured coatings.[51] Apart from the abovementioned strategies, many other materials, especially those belonging to the class of metals (as gold) and oxides (silicon dioxide (SiO2), titanium dioxide (TiO2)) are rising as potential candidates for the fabrication of inorganic coatings with improved functionalities.[52, 53] These materials show very interesting physical and chemical properties, which are tunable once externally activated.

Additionally, in the last decades, several advances have been obtained in urinary catheters (UC) seeking to reduce pathogen colonization. Although ureteral stents and UC are different devices, with distinct idiosyncrasies, the progress made on UC can be helpful for the development of new antimicrobial coatings for ureteral stents as these devices share a number of features: i) the fluid surrounding their surfaces is the same (urine), ii) the etiological agents that infect the devices are essentially the same (although some may be more deleterious than others, depending on the device), and iii) the range of shear forces caused by urine flow in a stent comprises the average shear value determined in UC (see Section 4). In this context, we will review some of the surface coatings that have been successfully tested in UC, and may, therefore, indicate the direction for ureteral stents advances. Table 2 describes the anti-biofilm strategies of different material coatings and their potential against several bacterial and fungal species. Surface coatings were grouped into four categories: i) release of antimicrobial agents, ii) contact-killing, iii) antiadhesive, and iv) biofilm architecture disruption.

Table 2. Coatings successfully tested for urinary catheters that are promising for ureteral stents Strategy Coating Material Microorganisms Major remarks Ref. Release of antimicrobial agents Metal (ions/nanoparticles) Silver

Silicone

Polyurethane

Latex

Gram-positive and Gram-negative bacteria

Inhibition of bacterial adhesion by 30%–99.9%; and biofilm formation by 1–6 Log, depending on the species.

In vivo studies demonstrated that silver-coated catheters reduced CAUTIs incidence by 27%–47%.

[63]a), [64] c), [65, 66]c), [67, 68]c), [69, 70] [71]c), [72]c), [73]c), [74, 75]c), [76]a), [77]b), [78]b) Noble metal alloy

Silicone

Latex

Reduction of CAUTIS incidence by 1.5%.

[79]a, [121]a, [122]a

Antimicrobial agents/disinfectants Antibiotic agents

Silicone

Polyurethane

B. subtillis

E. coli

E. faecalis

P. aeruginosa

P. mirabilis

MRSA

S. aureus

S. epidermidis

Delay on the onset of biofilm formation up 12 consecutive weeks;

Reduction of bacterial adhesion by 85–91%, with potent antimicrobial activity (83%–96%).

[85]b, [86]c,

, [70], [75]c, [87]c, [88]

Antifungal agents Silicone C. albicans Reduction of fungal adhesion on coated films by approximately 3 Log CFU. [89]c) Disinfectants

Silicone

Polyurethane

Latex

C. albicans

C. parapsilosis

E. coli

MRSA

P. aeruginosa

P. mirabilis

S. aureus

Reduction of microbial adhesion by 83%–99.8% and biofilm formation by 2 Log CFU.

Resistance to encrustation was increased up to 7 days.

[90-92]c), [82]c), [83]c), [84]c), [86]c) Contact-killing Antimicrobial peptides (AMP)