**1. Introduction**

Polyurethane (PU) is one of the most bio- and blood-compatible materials currently used for fabrication of various medical devices, e.g. blood bags, vascular/ureteral catheters and artificial heart. Originally, PU was conceived with other copolymers, aiming at similar goals, i.e. enough versatility to successfully meet biomedical devices constraints, such as biocompatibility, resistance to sterilization, physical features invariance over time and infection resistance during indwelling. PU in particular, has also other key properties particularly suited to biomedical industry, including strength, versatility and low cost. The surface of PU can be chemically functionalized with organic and biologically active molecules, resulting in improved durability, compliance, acceptance and tolerance in the human body during implantation. These features additionally strengthen PU as an appealing candidate for biomedical applications.

Normally, several polymers such as natural rubber, polyethylene, polyvinylchloride, fluoropolymers, hydrogels and silicon are used in biomedical applications. Despite the widespread use of these materials, PU still covers a relevant and dominant role, thanks to its high blood and tissue biocompatibility for improving the quality of patient's life. It conjugates a good stability over long implantation times, excellent physico-mechanical and surface tuning properties via anchoring of molecules. PU has simply a unique mix of features, highly required for almost any medical device.

PU is also widely used in cardiovascular applications, in particular for the preparation of venous and intravenous catheters and balloons for angioplasty and angiography. It has also been successfully used for tissue replacement and augmentation in breast implants, facial reconstruction and body joints. Artificial organs based on PU such as heart, kidneys and lungs have already been developed. Thanks to its above described unique features, PU has recently been proposed for new promising application fields, e.g. controlled drug delivery devices. A deep review on the use of PU in medicine and medical devices is also available (Zdrahala and Zdrahala 1999).

© 2012 Cauda and Cauda, licensee InTech. This is an open access chapter distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited. © 2012 Cauda and Cauda, licensee InTech. This is a paper distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

Given its large use, in this chapter we will focus on the use of PU in urological applications, in particular as ureteral catheter, or stent, in endourology routines.

Polyurethane in Urological Practice 125

these encrustations can obstruct the device and impair the urinary flow, compromising with

Temporary prevention from encrustation includes stent replacement at regular intervals, modification of the type or size of catheters, washing the catheter and bladder with acidic, antiseptic or saline solutions (Arshad, Shah and Abbasi 2006). Antibiotics are still orally administered whenever stent is replaced or inserted for preventing infections (Reid 2001).

However all these approaches are mostly ineffective. For this reason stent surface modifications have been proposed to prevent bacterial and inorganic molecule adhesion. Various strategies have been conceived, using silver-coated surfaces (Leung *et al.* 1992; Multanen *et al.* 2000), surface modification towards hydrophobicity (Jansen *et al.* 1993) or functional groups creation with intrinsic antimicrobial activity. Heparin is a good candidate for solving these problems. In previous *in vitro* and *in vivo* studies, heparinization of medical devices showed a reduction of microbial colonization (Appelgren *et al.* 1996; Cauda *et al.* 2008; Ruggieri *et al.* 1987). Heparin is a highly sulfated, anionic polysaccharide known for its anti-coagulant and anti-thrombogenic properties (Piper 1946). It has a strong negative electrical charge able to prevent cells adhesion since bacterial' cell membrane surface is also negatively charged. Coating stents with heparin can be practical and low-cost. Over the last three decades several studies, especially in vascular medicine, were indeed reported, showing that this approach is effective (Appelgren, Ransjo, Bindslev et al. 1996; Hildebrandt

Another successful approach preventing biofilm formation comprises the use of Diamond-Like Carbon (DLC) coatings on the ureteral stent. DLC is a thermodynamically meta-stable state of carbon where diamond-like (sp3-hybridized) and graphite-like (sp2-hybridized) bonding coexist with a large fraction of sp3 bonds. Coatings can be prepared by miscellaneous deposition methods, e.g. ion deposition, sputtering, pulsed laser deposition and plasma-enhanced chemical vapor deposition, using accelerated hydrocarbon ions as film forming particles (Grill 1999). In general, they are characterized by high mechanical hardness and chemical inertness. Depending on the deposition conditions, the properties of DLC films can be adjusted depending on the applications, e.g. allowing enhancement of the wear and corrosion resistance of precision cutting and machining tools. These films are already used as protective coatings on magnetic hard disks and optical glasses. They showed excellent tribological and mechanical properties, corrosion resistance, biocompatibility, and hemocompatibility (Anne Thomson *et al.* 1991; Roy and Lee 2007; Voevodin and Donley 1996). Recent studies have been focused on their ability to decrease the formation of crystalline bacterial biofilm as well as stent related side effects and

The evaluation of the *in vivo* efficacy of both heparin and plasma deposited DLC-coated ureteral stent have been reported. Recent works show the superiority of both coated stents for preventing biofilm adhesion and encrustation compared to uncoated PU catheters

patient care and leading to kidney infections, sepsis and shock (Warren et al. 1994).

*et al.* 1999; Ruggieri, Hanno and Levin 1987).

discomfort (Laube *et al.* 2006; Laube *et al.* 2007).

(Cauda *et al.* 2009).

In urology, catheterization is defined as the insertion of tubes (stents), such as urinary catheters, into the patient's bladder through the urethra. A stent is usually a tube with ending coils at both sides (pig-tails), or with lateral holes to further improve urine drainage (JJ or double-J). Urethral stents are usually in latex, silicone or polyurethane, allowing patients' urine to drain freely from the bladder. Stents can also be used to inject liquids for treatment or bladder conditions diagnosis. A patient is typically catheterized in the case of acute or chronic urinary retention, orthopedic procedures that may limit movement, benign prostatic hyperplasia, incontinence, and effects of various surgical interventions involving bladder and prostrate.

In endourology, urine is drained by indwelling a catheter between the kidney and the patient's bladder, hence inserting the stent into the ureter. Nowadays, ureteral stenting has become a common procedure for safe urine drainage. It is effective for managing several diseases, such as ureteral obstruction by stones or clots, benign or malignant ureteral obstruction, or post-surgical treatments, i.e. ureteroscopy and ureteral surgery. Ureteral stents almost painlessly keep the ureteral lumen open, ensuring that urine flows while maintaining the correct renal function. Stents also promote ureteral healing and prevent strictures formation.

From a general urological viewpoint, these devices must be easily maneuverable, affordable, and radiopaque for a correct positioning under fluoroscopic guidance. Hence, ureteral stents are mainly fabricated in PU or silicone since some patients can be allergic or sensitive to latex after long-term use. Radiopacity is ensured by adding metallic salts, e.g. based on barium. According to a critical study on materials (Mardis et al. 1993), the use of strong materials, e.g. PU, permits the reduction of the stent wall thickness, enlarging the inner lumen and the size or number of lateral holes and increasing urine flow. Silicone stents are weaker than PU, and need to be fabricated with smaller inner diameters, compromising urine flow and increasing the lateral compressibility.

Despite the numerous advantages enabled by PU, some complications and challenges remain. Indeed, PU is not perfectly biocompatible, in the sense that it somehow affects the epithelial cells of the ureter (urothelium): urothelial ulceration and erosion may occur. Other complications are related to stent migration or fracture, erosion, development of ureteroarterial fistula, fever, infection, voiding symptoms including dysuria and hematuria (Arshad et al. 2006). Among these, the stent encrustation represents one of the most serious complications resulting from the use of PU double-J stents: a stent can be encrusted by inorganic salts flowing with urine, and bacterial colonies can grow on the surface. These infections are very common among the general population, sometimes leading to death. During infections, the bacteria grow in the internal lumen of the stent, forming the so-called "biofilm", normally an aggregation of bacteria with their extracellular products and several inorganic salts (Costerton 2007). This matrix covers the cells, leading to a reduced susceptibility to prophylactic antibiotics (Tenke et al. 2004). Moreover, the development of these encrustations can obstruct the device and impair the urinary flow, compromising with patient care and leading to kidney infections, sepsis and shock (Warren et al. 1994).

124 Polyurethane

bladder and prostrate.

strictures formation.

urine flow and increasing the lateral compressibility.

Given its large use, in this chapter we will focus on the use of PU in urological applications,

In urology, catheterization is defined as the insertion of tubes (stents), such as urinary catheters, into the patient's bladder through the urethra. A stent is usually a tube with ending coils at both sides (pig-tails), or with lateral holes to further improve urine drainage (JJ or double-J). Urethral stents are usually in latex, silicone or polyurethane, allowing patients' urine to drain freely from the bladder. Stents can also be used to inject liquids for treatment or bladder conditions diagnosis. A patient is typically catheterized in the case of acute or chronic urinary retention, orthopedic procedures that may limit movement, benign prostatic hyperplasia, incontinence, and effects of various surgical interventions involving

In endourology, urine is drained by indwelling a catheter between the kidney and the patient's bladder, hence inserting the stent into the ureter. Nowadays, ureteral stenting has become a common procedure for safe urine drainage. It is effective for managing several diseases, such as ureteral obstruction by stones or clots, benign or malignant ureteral obstruction, or post-surgical treatments, i.e. ureteroscopy and ureteral surgery. Ureteral stents almost painlessly keep the ureteral lumen open, ensuring that urine flows while maintaining the correct renal function. Stents also promote ureteral healing and prevent

From a general urological viewpoint, these devices must be easily maneuverable, affordable, and radiopaque for a correct positioning under fluoroscopic guidance. Hence, ureteral stents are mainly fabricated in PU or silicone since some patients can be allergic or sensitive to latex after long-term use. Radiopacity is ensured by adding metallic salts, e.g. based on barium. According to a critical study on materials (Mardis et al. 1993), the use of strong materials, e.g. PU, permits the reduction of the stent wall thickness, enlarging the inner lumen and the size or number of lateral holes and increasing urine flow. Silicone stents are weaker than PU, and need to be fabricated with smaller inner diameters, compromising

Despite the numerous advantages enabled by PU, some complications and challenges remain. Indeed, PU is not perfectly biocompatible, in the sense that it somehow affects the epithelial cells of the ureter (urothelium): urothelial ulceration and erosion may occur. Other complications are related to stent migration or fracture, erosion, development of ureteroarterial fistula, fever, infection, voiding symptoms including dysuria and hematuria (Arshad et al. 2006). Among these, the stent encrustation represents one of the most serious complications resulting from the use of PU double-J stents: a stent can be encrusted by inorganic salts flowing with urine, and bacterial colonies can grow on the surface. These infections are very common among the general population, sometimes leading to death. During infections, the bacteria grow in the internal lumen of the stent, forming the so-called "biofilm", normally an aggregation of bacteria with their extracellular products and several inorganic salts (Costerton 2007). This matrix covers the cells, leading to a reduced susceptibility to prophylactic antibiotics (Tenke et al. 2004). Moreover, the development of

in particular as ureteral catheter, or stent, in endourology routines.

Temporary prevention from encrustation includes stent replacement at regular intervals, modification of the type or size of catheters, washing the catheter and bladder with acidic, antiseptic or saline solutions (Arshad, Shah and Abbasi 2006). Antibiotics are still orally administered whenever stent is replaced or inserted for preventing infections (Reid 2001).

However all these approaches are mostly ineffective. For this reason stent surface modifications have been proposed to prevent bacterial and inorganic molecule adhesion. Various strategies have been conceived, using silver-coated surfaces (Leung *et al.* 1992; Multanen *et al.* 2000), surface modification towards hydrophobicity (Jansen *et al.* 1993) or functional groups creation with intrinsic antimicrobial activity. Heparin is a good candidate for solving these problems. In previous *in vitro* and *in vivo* studies, heparinization of medical devices showed a reduction of microbial colonization (Appelgren *et al.* 1996; Cauda *et al.* 2008; Ruggieri *et al.* 1987). Heparin is a highly sulfated, anionic polysaccharide known for its anti-coagulant and anti-thrombogenic properties (Piper 1946). It has a strong negative electrical charge able to prevent cells adhesion since bacterial' cell membrane surface is also negatively charged. Coating stents with heparin can be practical and low-cost. Over the last three decades several studies, especially in vascular medicine, were indeed reported, showing that this approach is effective (Appelgren, Ransjo, Bindslev et al. 1996; Hildebrandt *et al.* 1999; Ruggieri, Hanno and Levin 1987).

Another successful approach preventing biofilm formation comprises the use of Diamond-Like Carbon (DLC) coatings on the ureteral stent. DLC is a thermodynamically meta-stable state of carbon where diamond-like (sp3-hybridized) and graphite-like (sp2-hybridized) bonding coexist with a large fraction of sp3 bonds. Coatings can be prepared by miscellaneous deposition methods, e.g. ion deposition, sputtering, pulsed laser deposition and plasma-enhanced chemical vapor deposition, using accelerated hydrocarbon ions as film forming particles (Grill 1999). In general, they are characterized by high mechanical hardness and chemical inertness. Depending on the deposition conditions, the properties of DLC films can be adjusted depending on the applications, e.g. allowing enhancement of the wear and corrosion resistance of precision cutting and machining tools. These films are already used as protective coatings on magnetic hard disks and optical glasses. They showed excellent tribological and mechanical properties, corrosion resistance, biocompatibility, and hemocompatibility (Anne Thomson *et al.* 1991; Roy and Lee 2007; Voevodin and Donley 1996). Recent studies have been focused on their ability to decrease the formation of crystalline bacterial biofilm as well as stent related side effects and discomfort (Laube *et al.* 2006; Laube *et al.* 2007).

The evaluation of the *in vivo* efficacy of both heparin and plasma deposited DLC-coated ureteral stent have been reported. Recent works show the superiority of both coated stents for preventing biofilm adhesion and encrustation compared to uncoated PU catheters (Cauda *et al.* 2009).

Despite these recent advances, oral administration of antibiotics (e.g. ciproflaxin) cannot be avoided. Papers in this respect report that bacteria proliferation has been addressed with a local release of antibiotics or antiseptics (Cormio *et al.* 2001; John *et al.* 2007; Leung *et al.* 2001; Raad *et al.* 1997). Drug-Eluting Stent (DES) has the advantage of maximizing the local tissue levels of therapeutic agents while minimizing systemic toxicity. The problem has been faced with antibiotics incorporation (Gorman and Woolfson 2002) using novel biomimetic and bioactive silicones. For example, with a gentamicin-releasing urethral catheter, encrustation inhibition in a rabbit model has been shown in the short term (Cho *et al.* 2001). Other authors (Cadieux *et al.* 2006) reported on a triclosan-loaded ureteral stent implanted in rabbit bladders with bacterial infection. The study showed a significant decrease of urinary tract infection rate. However, a very high control on the delivery kinetics has been not achieved yet.

Polyurethane in Urological Practice 127

**indwelling** 

like biofilm formation, inorganic encrustation extent, stiffness, brittleness or failure of the PU material depending on the indwelling time will be examined. The surface chemistry and morphology characterization of the indwelled stents will also be evaluated according to the surface coating and biofilm formation. The characterization techniques used in this work included Field Emission Scanning Electron Microscopy (FESEM), Energy Dispersive

We review here the results collected from 2006 to 2010, concerning the characterization studies on uncoated, heparin- and DLC-coated PU double-J ureteral stents (all provided by Cook Ireland LTD) after indwelling. We enrolled 59 patients (from 45 to 75 years old). 49 patients showed unilateral ureteral obstruction, thus requiring ureteral stenting in the affected ureter. The patients with unilateral obstruction received an heparin-coated, DLC-

10 patients suffered from bilateral obstruction, therefore the coated and uncoated PU stents were indwelled at the same time in both ureters, respectively. Each patient with bilateral obstruction received randomly both the coated stent (with heparin or with a diamond-like carbon coating) in one ureter, and the uncoated one, thus the pure PU stent, in the other ureter. Stents indwelling was also studied for different periods of time. The stent types and

**stents Coating Indwelling time Type of** 

14 Heparin 1-3 months Unilateral 9 Heparin > 3 months Unilateral

19 Nonea 1-3 months Unilateral 3 None > 3 months Unilateral

4 DLC 1 month Unilateral

5b None 1 month Bilateral 5b Heparin 1 month Bilateral

5b None 1 month Bilateral 5b DLC 1 month Bilateral

b These stents were implanted bilaterally, thus inserting both the uncoated and the coated stents into both ureters

**Table 1.** Stents types and indwelling periods of the PU, heparin-coated and diamond-like carbon

**2.1. Experimental part: Patients, stents and characterization methods** 

Spectroscopy (EDS), and Infrared (IR) spectroscopy.

coated or an uncoated PU stent in the ureter to be treated.

indwelling times are reported in Table 1.

a None indicates that the stent surface is of pure PU.

respectively of the same patient.

coated PU stents.

**Number of** 

Another important problem with stents is that they require an additional cytoscopic procedure for their removal. Stents removal may be uncomfortable for the patient, in particular when these are encrusted, therefore requiring hospitalization and anaesthesia. Considering this big concern, the development of time bio-degradable stent materials has become a major keypoint. In principle, bio-adsorbable stents shall be designed to maintain their integrity for a given period of time, and undergo a dissolution process followed by spontaneous expulsion by the patient at the same time.

So far, the bio-degradable stents degrade very fast (typically 48h) or leave fragments removable by lithotripsy and ureteroscopy (Lingeman *et al.* 2003). Some tests on pigs with degradable poly-L-lactic-L-glycolic acid (PLGA) devices showed fragments embedded in the ureteral walls within cystic sacs, normally leading to fibrosis, inflammation and a large foreign body giant cell reaction (Olweny *et al.* 2002). These results show that further enhancement on their degradable characteristics is highly required. It is straightforward that the future of ureteral stents will head towards this direction, i.e. to chemically controlled biodegradation, combined to a concomitant release of biologically active molecules on target sites and with minimally invasive surgical procedures. Moreover, highly engineered stents can possibly work as a "scaffolds" for tissue regeneration via cell attachment and proliferation, finally controlling the local inflammation and healing (Zdrahala and Zdrahala 1999).

This chapter focuses in details on the use of PU catheters for endourological applications, providing new insights on the *in vivo* performances of PU stent. Given to our past expertise, we report on the influence of the PU surface treatments in preventing the encrustation and the formation of bacterial biofilm when implanted into the ureter.
