A new hydrophilic biodegradable ureteral stent restrain encrustation both in vitro and in vivo

Abstract

Ureteral stents have been widely used as biomedical devices to treat some urological diseases for several decades. However, the encrustation complications hamper the long-time clinical use of the ureteral stents. In this work, a new type of biodegradable material for the ureteral stents, methoxypoly(ethylene glycol)-block-poly(L-lactide-ran-Ɛ-caprolactone) (mPEG-PLACL), is evaluated to overcome this problem. The results show that the hydrophilicity and degradation rate in artificial urine of mPEG-PLACL are both significantly increased. It is worth noting that the mPEG-PLACL shows a lower amount of encrustation after immersing the stents in the dynamic urinary extracorporeal circulation (DUEC) model for 7 days. In addition, 71% Ca and 92% Mg are inhibited in vivo by quantitative analysis. Pathological analysis exhibit that the mPEG-PLACL cause less diffuse mucosal hyperplasia after 7 weeks of implantation. All the results indicate that this new type of biodegradable material had an excellent potential for the ureteral stents in the future.

Keywords

Ureteral stents degradation hydrophilicity encrustation surface morphology

Introduction

The transport of urine from the renal pelvis to the bladder depends on the ureter. The human urinary system requires ureteral stents for adjuvant therapy when it develops lesions, such as urinary tract obstruction, ureteral injury or malignant tumors.¹ However, the complication of encrustation frequently occurs after the implantation of ureteral stent.² In the absence of infection, the predominant urinary stone types are calcium oxalate monohydrate and calcium oxalate dihydrate.^3,4 Even this calcium salt precipitate was found on the ureteral stents in patients with normal urinary calcium levels.⁵ When infection occurs, especially with urease-producing bacteria, urease could decompose urea, leading to an elevated pH of urine. In this case, the solubility of magnesium and calcium salts is reduced, forming calcium phosphate and magnesium ammonium phosphate.⁶ Encrustation in the lumen and the wall of ureteral stents could obstruct the flow of urine and eventually lead to the kidney infections, sepsis and shock.^4,7 Therefore, preventing the encrustation is an urgent problem in the field of ureteral stents research.

At present, there are several strategies to prevent encrustation. Studies have shown that increasing the wall shear stress (WSS, the tangential force per unit area exerted on the solid boundary by the fluid in motion.) of urine on the ureteral stent could reduce the bacterial attachment to the ureteral stent surface.^8–10 De Grazia et al. demonstrated that manufacturing several side holes on ureteral stent at the proximal and distal of the ureteral occlusion could improve WSS and prevent short-term bacterial adhesion.¹¹ This strategy needs to modify the ureteral stent according to the specific conditions of different patients,¹¹ which is highly targeted, including different patients and symptoms, and the economy is limited. The other strategy is to reduce the surface friction coefficient of the ureteral stents by increasing the hydrophilicity of the stents surface,^12,13 which could prevent encrustation deposition, enhance comfort and reduce urinary system damage.^14–16 Ko et al. investigated a new anti-encrustation coating composed of monomethoxypolyethylene glycols (mPEG) and 3,4-dihydroxyphenylalanine (DOPA₃), which could be applied on the surface of the ureter stents to resist the biofilm formation and encrustation in human urine.¹⁷ In addition, Cauda et al. had compared two different polyurethane stents that were able to prevent inorganic encrustation; the first stents had hydrophilic hydrogel coating (PU-HY) and the other had water-mediated swelling chitosan coating (CHI).¹⁸ The authors found that the PU-HY stents showed a lower level of encrustation than the CHI stents after immersing them in a dynamic model for 6 months in vitro. However, these improvements are limited to the surface of the ureteral stents. Moreover, non-degradable ureteral stents require extubation and may possibly be forgotten in human body, which can cause secondary harm to the patient.¹⁹

In the field of degradable biomaterials, many synthetic polymers, including polylactic (PLA), polycaprolactone (PCL), poly(glycolic acid-co-lactic acid) (PGLA), poly(L-lactide-co-ɛ-caprolactone) (PLACL) etc., which have been confirmed by the Food and Drug Administration (FDA) as biomaterials, have broad prospects in the field of tissue engineering.^20–22 PLA was widely researched as the ureteral stents for a long time.^23–26 Lumiaho et al. proved that the self-reinforced poly-L, D-lactide (SR-PLA96) as ureteral stents could reduce the vesicoureteral inflammatory response, due to the outstanding biocompatibility.²⁷ Nevertheless, the hydrophobicity, brittleness and slow degradation rate of PLA limit its clinical application in the urinary system.^28–30 At the same time, it has been reported that PCL was flexible, high elongation at break and similar processing temperature to PLA.³¹ In addition, the polyether structure of mPEG enabled itself to couple a large number of water molecules and reduce its friction coefficient, thereby reducing the adhesion of encrustation on the surface of the ureter stents.³² Therefore, the copolymerization of mPEG, CL and LA could improve the hydrophilicity and smoothness for the ureteral stents. The degradation speed of this copolymer also could be accelerated.

In this study, we were committed to evaluating a novel type of biodegradable polymer, namely methoxypoly(ethylene glycol) -block-poly(L-lactide-ran-Ɛ-caprolactone) (mPEG-PLACL) (Figure 1), aiming to design and develop a biodegradable, anti-encrustation and biocompatible ureteral stent. The ureteral stents were extensively characterized for the surface, degradation and anti-encrustation effects both under in vitro and in vivo conditions using various analytical techniques including water contact angle, electronic universal material testing machine, SEM, EDS and ICP-OES. This study will provide useful information concerning the feasibility of the hydrophilic polymers as commercially available urinary biomaterials.

Experimental

Materials

The methoxypoly(ethylene glycol)_x-block-poly(L-lactide-ran-Ɛ-caprolactone) (mPEG _x -PLACL, x = 3%, 5%, 8% w/w, the molar ratio of LA/CL was 75/25) was obtained from our laboratory. The reagents of artificial urine were listed in Tables 1 and 2.

Table 1.

Artificial urine for degradation.³⁵

Reagent	Mass concentration (g/L)
CaCl₂	0.6664
Na₂SO₄	2.4310
MgSO₄	0.3572
NH₄Cl	2.3215
KCl	6.0650
NaH₂PO₄·2H₂O	1.3300
Na₂HPO₄·12H₂O	2.8196
Na₃Cit·2H₂O	0.5840
NaCl	6.7725

Table 2.

Composition of artificial urine in the DUEC.³⁸

Reagent	Mass concentration(g/L)
KH₂PO₄	3.80
MgCl₂	1.80
Urea	8.00
CaCl₂	1.35
Albumin	1.00
Urease	0.20

Preparation the ureteral stents of mPEGx-PLACL and PLACL

Using a micro twin-screw extruder, the mPEG_x-PLACL and PLACL were extruded into tube with the processing temperature of 145 °C. The outer diameter and thickness of the ureteral stents were 2.0 mm and 0.4 mm respectively.^33,34 The ureteral stents were cut to 5.0 cm for the next tests.

Water contact angle (WCA)

In order to measure the hydrophilicity of the stents surface, optical water contact angle (JC2000D1, POWEREACH, Shanghai, China) measurement were carried out by the sessile drop technique. Within 100 s, the pictures which were water drops with a volume of 1.5 µL on the stent were acquired every 10 s with the integrated camera. Using the technical software fit the water drop curve, and then the contact angle value could be gotten at the liquid–solid interface. The measurements were repeated five times for each sample of PLACL, mPEG₃-PLACL, mPEG₅-PLACL and mPEG₈-PLACL, averaging the left and right sides of the recorded angle.

Degradation of mPEGx-PLACL and PLACL in vitro

The ureteral stents of mPEG₃-PLACL, mPEG₅-PLACL and mPEG₈-PLACL were separated into three groups. PLACL was used as the control. Each group had 30 samples which were soaked in artificial urine (pH=5.37) (Table 1) and degraded for 50 days in a constant temperature incubator at 37 °C. The ratio of artificial urine to sample was 50: 1 (v/w).³⁴

Every 10 days, the samples were taken out, washed with distilled water, dried under decompression for 72 hours.³⁶ Five tests were carried out to show the changes of the ureteral stents in degradation, including the weight loss ratio (n = 6), water absorption ratio (n = 6), tensile strength (n = 5), extension at break (n = 5) and morphology (n = 1).

Water absorption and weight loss ratio

The weight loss ratio and water absorption of the samples was calculated by the equations (1) and (2):

Weight loss (%) = \frac{W_{0} - W_{dry}}{W_{0}} \times 100 %

(1)

Water absorption (%) = \frac{W_{wet} - W_{dry}}{W_{0}} \times 100 %

(2)

Where W₀ was the original dry weight of the sample, W_dry was the dry weight of the sample and W_wet was the wet weight of the sample after degradation in different periods.

Tensile strength and extension at break

The tensile strength and extension at break of the samples were measured by Electronic Universal Material Testing Machine (SANS, CMT4503, Shenzhen, China). Each sample (50 mm × 2 mm) was fixed between fixtures to ensure a constant gauge distance (30 mm).The upper clamp was raised at a fixed rate (100 mm/min) until the fracture occurred across the centre of each sample.³⁷ All determinations were repeated five times. The tensile strength and extension at break were obtained from the stress-strain diagram.

Morphology

The morphology of the samples was investigated by scanning electron microscope (SEM, S-520, HITACHI, Japan) with 10 kV acceleration voltages. Surfaces were coated with a thin gold layer to improve the electric conduction.

Encrustation in vitro

The DUEC (Figure 2) was set up to investigate the resistance of mPEG_x-PLACL to encrustation in vitro. The DUEC was consisted of an artificial urine reservoir (a glass container of 1 L in volume), a micro peristaltic pump (ChuangRui, BW100, China) and transparent silicone tubes (Saint-Gobain, Tygon 3350, France) for connecting the entire system (Figure 2). The DUEC was placed in 37 °C and 5% v/v CO₂ incubator (Yiheng, BPN-80CH, China) to simulate physiological conditions. The flow rate of the artificial urine (Table 2) which was different with the artificial urine for degradation was maintained at 10 ml/min.³⁸ 1 liter artificial urine was stored in the reservoir. 400 ml of new artificial urine was replaced on a daily basis.³⁸ At the same time, any crystalline deposited in the tubes was also removed every day to prevent the blockage. Before setting up the DUEC, all the instruments used in the whole system need to be sterilized at 120 °C for 1 hour in autoclave (BKQ-B100II, BIOBASE, China). During the first day of the experiment, the changes of pH were investigated every hour for the first 10 and 24th hours. Five stents (30 mm × 2 mm) of each sample were kept in the DUEC for 7 days for the next experiment.

Encrustation in vivo

A total of 45 male Wistar albino rats weighing 180–200 g were used. Methoxypoly(ethylene glycol)_x-block-poly(L-lactide-ran-Ɛ-caprolactone) (x = 5% w/w) was selected as the most effective anti-encrustation modification group (n = 5) based on the degradation performance and anti-encrustation properties in vitro. The hydrophobic PLACL served as a negative control (n = 5). To investigate whether the ureteral stents could cause bladder inflammation, it was necessary that we need the sham operation group (n = 5) which was performed the same procedure without implanting any stent. The rats were anesthetized with sodium pentobarbital (30 mg/kg) before implanting operation. Then, a 2-3 mm long stent was implanted into the rat bladder and kept for 3, 5 and 7 weeks (Figure 3). At the end of each period, five rats in each group were sacrificed. The bladder and stents were removed for further characterizations and histopathological examinations.

Encrustation analysis

Following the specified immersion period, the samples were taken out, dried in vacuum oven at 37 °C for 48 hours for the next analysis. SEM, equipped with energy dispersive spectroscopy (EDS), was used to characterize the morphology and composition of the sample surface and the formation of encrustation. Adhered inorganic calcium and magnesium salt were quantitatively analyzed by inductively coupled plasma optical emission spectrometry (ICP-OES, 5110, Agilent, the US). The presence of urinary encrustation on the surface of the sample was expressed as the weight of calcium and magnesium present per unit weight of the sample.

Result and discussion

Water contact angle (WCA) analysis

Biomaterials with a certain degree of hydrophilicity can be used as the ureteral stents to prevent the formation of the encrustation. WCA is the most common method for evaluating the hydrophilic and hydrophobic properties of the surface. As shown in Figure 4, the WCA values of the four stents remained almost invariable during 100 s contacting. The WCA value of the control PLACL was 78.65°. However, with the amount of mPEG increased, the WCA values of mPEG₃-PLACL, mPEG₅-PLACL and mPEG₈-PLACL were significantly decreased to 67.16°, 64.52° and 49.4° after 100 s contacting, respectively. These results indicate that the stents of hydrophilicity is provided by the mPEG terminal segment.^39,40

Degradation in vitro

Water absorption

In the PLACL control group, the water absorptions were around 0.33%, 1.0%, 2.3%, 1.45% and 1.6% at day 10, 20, 30, 40 and 50, respectively (Figure 5(a)). But for the increasing proportion of the mPEG, the hydrophilic chain in mPEG_x-PLACL, the water absorptions were much higher than those of the control during the degradation process. Among them, the water absorption of mPEG₅-PLACL reached the peak level of 9.69% at 30 days, after which there was a slight decrease. The similar phenomenon has been reported by XingYi et al.⁴¹ The authors found that the water absorption of the composite films significantly improved by the hydrophilic part. MPEG₈-PLACL was completely degraded (losing the mechanical strength as a complete degradation) after 10 days soaking due to its high mPEG content.

Weight loss ratio

The weight loss ratios of biodegradable stents also increased during the whole degradation period. The weight loss ratios of the stents with mPEG, hydrophilic terminal segment, were higher than those of the PLACL group (Figure 5(b)). Although the content of mPEG in mPEG₃-PLACL was lower than that of mPEG₅-PLACL, the weight loss ratios were almost the same. This result was consistent with the WCA that the hydrophilicity of mPEG₃-PLACL and mPEG₅-PLACL was similar.^41,42 We also found that the weight loss rate of mPEG₈-PLACL was 3.5% at 10th day. Park J. H. using GPC determined the molecular weights of PLACL and mPEG-PLACL during degradation time. The authors found that the strong hydrophilicity of mPEG chain increased the degradation rate of copolymers.⁴³ Similarly, as the hydrophilicity of the stents increases, the weight loss rate of these ureteral stents has also increased while degrading.

Tensile strength

The ureteral stents need to be implanted into the human body, and the mechanical properties of the stents changed as degradation progresses. Figure 5(c) showed the curves of the tensile strength of the stents as a function of degradation time. At the end of degradation, the tensile strength of PLACL, mPEG₃-PLACL, and mPEG₅-PLACL were reduced by 71%, 75% and 76%, respectively. The tensile strength of mPEG₈-PLACL had reduced from 6.35 MPa to 1.98 MPa after 10 days of degradation. Due to increasing the hydrophilicity, the mPEG₈-PLACL showed a very rapid loss of mechanical properties.⁴⁴ From Figure 5(c), it can be observed that the both hydrophobic and hydrophilic stents decreased uniformly after 30 days degradation. This result indicates that the copolymers had almost become small molecular substances at that time,³⁰ which causes the degradation rate to decrease slowly.

Extension at break

The elongation at break gradually decreased as shown in Figure 5(d) with the progresses of the degradation. All the stents lost elasticity at 50th day, indicating that the degradation had reached the late stage. Moreover, the elongations at break of mPEG_x-PLACL were lower than PLACL throughout degradation process. This could be due to the proportion of mPEG_x-PLACL’s soft segment Ɛ-caprolactone decreased, compared to PLACL. The elasticity of all the stents has been decreasing with the progress of degradation. According to previous reports, water could penetrate into the amorphous regions composed of mainly Ɛ-caprolactone units easier than the hard domains.⁴⁵ As consequence, the hydrolysis rate of Ɛ-caprolactone, soft segment, was accelerated, which was shown by a decrease in elasticity.

Surface morphology

The surface morphology of the PLACL, mPEG₃-PLACL and mPEG₅-PLACL were also inspected by SEM after 0, 30 and 50 days degradation. Figure 6 showed that the three stents had very smooth surfaces during the initial stages of degradation. The smooth surfaces of the stents could even facilitate implantation for patient and minimize the mucosal injury which might reduce the risk of infection.^13,46 In addition, stents with a hydrogel coating have been shown to resist encrustation, biofilm development and infection better than that of non-surface modification.⁴⁷ The surface of the hydrophilic stents showed uniform cracks with the further degradation, while the surface of the PLACL showed uneven size and irregular holes at 30 and 50 days after degradation. It has been reported that cracks and pores on the surfaces of the hydrophilic stents could enhance its water absorption during the degradation, thereby causing the stents to swell and accelerate its degradation.⁴³ After 30 days of degradation, PLACL had the fewest voids, while mPEG₅-PLACL had the largest number of dense cracks. This result also explained that PLACL had higher elongation than mPEG₃-PLACL and mPEG₅-PLACL in Figure 5(d). When all the stents were degraded to 50 days, their changes cause these stents to lose mechanical properties.

Encrustation in vitro

Multiple infections often occurred in patients with primary diseases, especially urinary tract infections. It mainly caused by proteus mirabilis which were the prime urease positive pathogen in the ureteral system. The urease decomposed urea to produce ammonia and carbon dioxide so that the pH of the urine finally fixed at alkalinity.^48,49 Under this condition, calcium ion and magnesium ion would form insoluble precipitations, namely struvite (magnesium ammonium phosphate, NH₄·MgPO₄·2H₂O) and hydroxyapatite (calcium phosphate, Ca₁₀(PO₄)₆·H₂O), respectively.⁵⁰ In order to simulate the process of the formation of encrustation, we monitor the pH of artificial urine in DUEC within 24 hours. Figure 7 showed that the pH of fresh artificial urine was 5.53, but it was finally fixed to 8.1 after 24 hours as the urea was decomposed. This result revealed that the DUEC successfully established the conditions for encrustation formation.

Due to the high crystallization potential of artificial urine in DUEC, 7 days in vitro study is equivalent to several months of in vivo study.^51,52 It can be clearly seen that the surface of PLACL were covered with encrustation after 7 days immersing through Figure 8(a), and the surface of the stents cannot be easily distinguished. On the contrary, according to Figure 8(b) to (d), the mPEG₃-PLACL, mPEG₅-PLACL and mPEG₈-PLACL showed scattered and thin packed aggregates over its entire surface, which proved that the stents with better hydrophilicity and smoother surfaces had less inorganic salt deposition.⁵³

We further used EDS to analyze the main elemental composition of the encrustation. Based on the summary of the consequences in Figure 8(e) and morphological analysis of SEM, the different encrustation types in Figure 8(e) can be easily distinguished. The prismatic structure in Site 2 clearly represented the struvite crystal and its chemical formula was NH₄·MgPO₄·2H₂O.⁵⁴ The small spherical structures appearing at Site 3 represented typical characteristics of hydroxyapatite Ca₁₀(PO₄)₆(OH)₂.⁵⁵ After qualitative analysis by SEM and EDS, the amounts of deposited Ca and Mg were quantitatively determined using ICP-OES, listed in Table 3. The PLACL soaking in the DUEC for 7 days showed the highest amount of encrustation with 1.6634 mg/g Ca and 0.3498 mg/g Mg. The Ca deposition of mPEG₃-PLACL, mPEG₅-PLACL and mPEG₈-PLACL decreased 1.1169 mg/g, 1.2960 mg/g and 1.4086 mg/g, respectively. Furthermore, Mg deposition was much lower than Ca on the surfaces of all the stents. Table 3 also indicated that mPEG₃-PLACL, mPEG₅-PLACL and mPEG₈-PLACL inhibited 46%, 53% and 64% Mg deposition, respectively. This result showed that the smooth and hydrophilic surface played an important role in resisting the deposition of encrustation. Khandwekar et al.⁵⁶ reported that PU films modified with PVP-I complex were more hydrophilic and even more resistant to encrustation. In other studies, hydrophillic biomaterials had been proven to possess several properties that were beneficial for medical device applications, including lubricity, biocompatibility and resistance to infection and encrustation.^2,57,58 These reports were consistent with our results.

Table 3.

Amount of encrustation on different stents after 7 days soaking in DUEC.

Stent	Ca (mg/g)	Mg (mg/g)
PLACL	1.6634	0.3498
mPEG₃-PLACL	0.5465	0.1882
mPEG₅-PLACL	0.4574	0.1643
mPEG₈-PLACL	0.2548	0.1248

Encrustation in vivo

The results of experiments in vitro cannot fully prove the anti-encrustation performance of the ureteral stents in vivo. In this regard, comparisons between experiments in vivo and in vitro are necessary. Due to the fact that the human urine and the rat urine have very similar characteristics in terms of the composition and encrustation,^59,60 Wistar albinism rats were selected as experimental animals in vivo. According to the degradation performance and anti-encrustation properties in vitro, methoxypoly(ethylene glycol)_x-block-poly(L-lactide-ran-Ɛ-caprolactone) (x = 5% w/w) was chosen for the tests in vivo as the most promising ureteral stent. The PLACL served as control group. At the same time, we need the sham operation group to investigate that whether the implanting ureteral stents could cause bladder pathological changes. The surfaces morphology of the stents which removed at each time period was qualitatively evaluated by SEM (Figure 9). The PLACL stents showed severe encrustation and even blockage at each period, while mPEG₅-PLACL revealed almost no encrustation at each time point. Furthermore, we used ICP-OES to quantitatively characterize encrustation as a supplement to the observation of surface morphology (Table 4). Compared with the hydrophobic PLACL, mPEG₅-PLACL reduced the deposition of 71% Ca and 92% Mg at the 7th week, respectively. This result was consistent with the experiment in vitro and proved the feasibility of the application of the DUEC in the investigation of anti-encrustation.

Table 4.

The encrustation amounts on PLACL and mPEG₅-PLACL stents after 3, 5 and 7 weeks in vivo (Ca and Mg represent the deposition of hydroxyapatite and struvite on the ureteral stent, respectively).

	PLACL			mPEG₅-PLACL
	3 weeks	5 weeks	7 weeks	3 weeks	5 weeks	7 weeks
Ca (mg/g)	2.35	3.64	5.24	0.33	0.85	1.51
Mg (mg/g)	107.14	134.57	157.36	7.58	9.25	12.48

This implanting surgery did not cause any rat death after 7 weeks of surgery. Besides, all the rats were dissected and the main organs in the abdominal cavity showed no obvious abnormalities (Figure 10(a)). Haematoxylin eosin (H&E) staining microscopy photographs were given in the Figure 10(b). After 7 weeks implanting, the muscle layer of bladder was thick and the mucosal cells were enlarged. It was remarkable that the diffuse mucosal hyperplasia occurred in different degrees in each group. The rates of diffuse mucosal hyperplasia were 100% in the PLACL group, 40% in the mPEG₅-PLACL group, and 20% in the sham group. Previous reports showed that the struvite and hydroxyapatite would be formed in the urinary system due to bacteria adhesion on implants.⁶¹ There were several mechanisms to solve this problem by controlling the hydrophobicity, surface roughness, surface functional groups and electrostatic interactions.⁶² Studies have proven that immobilizing the biocompatible hydrophilic PEG or mPEG can be used to prevent bacterial adhesion, even the higher molecular weight PEG and longer brushes can resist bacterial adhesion more strongly.^63–65 Hezi-yamit et al. demonstrated that polymers with hydrophilic surfaces prevent the inflammatory activated monocytes from adhering to their surface.⁶⁶ Besides, According to previous reports, the hydrophilic material surface had lower protein and cell adhesion, but the biocompatibility was not significant affected.⁶⁷ These findings perfectly agree with the present results. This makes our development as a further ideal candidate for a urinary biomaterial.

The major limitation of our study is that we were unable to assess the possibility of calcium oxalate encrustations in patients without bacterial infection.

Conclusion

In this study, we successfully evaluate the surface, degradation and anti-encrustation both in vitro and in vivo of the ureteral stents. Firstly, the degradation rate of the ureteral stents is increased due to the enhancement of the surface hydrophilicity. In addition, experiments both in vitro and in vivo demonstrate that mPEG-PLACL can offer the function of anti-encrustation and the superior biocompatibility. All in all, the properties exhibit by this ureteral stents might be used as future degradable anti-encrustation urinary system medical devices.

Footnotes

Acknowledgments

Lulu Guo and Hongchen Xie from the College of Public Health of West China Hospital of Sichuan University provided assistance.

Declaration of Conflicting Interests

The author(s) declared no potential conflicts of interest with respect to the research, authorship, and/or publication of this article.

Funding

The author(s) disclosed receipt of the following financial support for the research, authorship, and/or publication of this article: This research was supported by the Department of Science and Technology of Sichuan Province (14zc2001) and the West Light Foundation of the Chinese Academy of Sciences (Z1633-27).

ORCID iD

Yu Zhang

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