# Straightforward Preparation of Polyurethane-Based Biomaterials with Dual Antibacterial and Anti-Adhesive Properties for Long-Lasting Catheter Applications

Baptiste Caron, Marc Maresca, Amélie Leroux, Marie Lemesle, Jean Louis Coussegal, Yohann Guillaneuf, Catherine Lefay

► To cite this version:

Baptiste Caron, Marc Maresca, Amélie Leroux, Marie Lemesle, Jean Louis Coussegal, et al., Straightforward Preparation of Polyurethane-Based Biomaterials with Dual Antibacterial and Anti-Adhesive Properties for Long-Lasting Catheter Applications. Advanced Materials Interfaces, 2025, 12 (24), pp.e00714. 10.1002/admi.202500714. hal-05440316

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# Straightforward Preparation of Polyurethane-Based Biomaterials with Dual Antibacterial and Anti-Adhesive Properties for Long-Lasting Catheter Applications

Baptiste Caron, Marc Maresca, Amélie Leroux, Marie Lemesle, Jean Louis Coussegal, Yohann Guillaneuf, and Catherine Lefay*

Catheter-associated infections are a major concern in hospitals, leading to both life-threatening for the patients and a high cost for society. The development of a straightforward and industrial route to make an antibacterial catheter is thus worthwhile. This study demonstrates that the use of 2 wt.% of an antibacterial Mel-quaternized poly(butyl methacrylate-block-N,N-dimethylaminoethyl methacrylate) P(BMA-b-DMAEMA) copolymer in combination with 2 wt.% of an antiadhesive poly(butyl methacrylate-block-poly(ethylene glycol) methyl ether methacrylate) P(BMA-b-PEGxMA) copolymer (with x the molecular weight of the PEG) as additives during the extrusion of the polyurethane matrix is an efficient method to produce antibacterial and antiadhesive PU materials without loss of activity after exposure to biologic media. The addition in the formulation of the antiadhesive copolymers enables protecting the surface from passivation and then to keep the contact possible between the bacteria and the antibacterial material. The antibacterial activity of the materials against E. coli and S. aureus is then preserved even after exposure to albumin, plasma, intralipids, or gastric acids. The prepared biomaterials also present no toxicity and are able to limit E. coli biofilm formation. Based on these results, this methodology can be realistically envisioned to elaborate long-lasting venous or enteral catheters.

## 1. Introduction

In hospitals and the medical field in general, bacterial contamination of medical devices and especially catheters ( urinary catheters or central venous catheter (CVC), probes, etc.) accounts for a significant proportion of nosocomial illnesses. Among the 35 million of whole catheters used annually in the USA, it is estimated that 3–8% of CVCs and 10–30% of urinary catheters are infected, leading to various complications for the patient, which in the most serious cases can be lethal. Overall, it is estimated that the yearly cost of catheter-associated infections ranges from $296 million to $2.3 billion.[1–3]

A further difficulty with healthcare-associated illnesses is the increasing resistance of pathogenic microbes to antimicrobial treatments.[4] Resistant strains are more common in device-associated infections than surgical site infections, according to a review of strains recovered from HAIs in the US between 2011 and 2014.[5] One method of preventing these infections is to use

antibacterial biomaterials whose antibacterial activity derives from the release of active compounds or molecules (release-killing) or which kill bacteria on contact (contact-killing).[6–12] The release-killing approach, for example of antibiotics such as rifampicin or silver nanoparticles,[13,14] has certainly been the most widely developed to date, but has the disadvantage of losing effectiveness over time (materials with no activity after complete release of the active compound) and potentially inducing resistance phenomena (particularly in the case of antibiotics with a slow release of sub-active dose) and toxicity. In contrast, the contact killing approach, which kills bacteria on contact by targeting the bacterial membrane, reduces the risk of bacterial resistance and, by not releasing active molecules into the environment, does not generate toxicity and/or pollution risks.[6–9] Most contact-active materials consist of quaternary ammonium compounds as the biocidal agent, which have shown efficacy in eliminating bacteria while being non-cytotoxic to human cells.[15] However, the major drawback of these materials is that their cationic surface is generally passivated during use

B. Caron, Y. Guillaneuf, C. Lefay
Aix-Marseille Université, CNRS
Institut de Chimie Radicalaire UMR 7273

Marseille 13397, France
E-mail: catherine.lefay@univ-amu.fr

B. Caron, A. Leroux, M. Lemesle, J. L. Coussegal
VYGON

Ecouen 95440, France

M. Maresca
Aix-Marseille Université, CNRS

Centrale Marseille, iSm2
Marseille 13397, France

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The ORCID identification number(s) for the author(s) of this article can be found under https://doi.org/10.1002/admi.202500714

© 2025 The Author(s). Advanced Materials Interfaces published by Wiley-VCH GmbH. This is an open access article under the terms of the Creative Commons Attribution License, which permits use, distribution and reproduction in any medium, provided the original work is properly cited.

DOI: 10.1002/admi.202500714
<figure><img src="image_4.png" /><figcaption>**Figure 1.** a) Schematic representation of a contact killing surface bearing a cationic polymer that could be passivated via protein deposition. The use of a hydrophilic polymer could avoid this passivation by inducing a hydrophilic barrier. b) Different methods to combine hydrophilic and cationic polymer to obtain antifouling and antibacterial properties.</figcaption></figure>

by the deposition of molecules (bacterial debris, proteins, etc.), which reduces or even annihilates their antibacterial activity, particularly in the presence of blood, urine, or other biological media (Figure 1a). It is commonly observed that when a medical device is inserted, proteins and other host molecules quickly coat its surface, creating a conditioning film.[16–19] This film forms a layer between the bacteria and the material, preventing them from coming into contact. As a result, the material can no longer kill bacteria on contact. Whereas urinary catheters normally get coated with proteins and electrolytes from the patient's urine, vascular catheters typically have a conditioning film made up of platelets and plasma proteins such as albumin, fibrinogen, and fibronectin.[16–19]

To alleviate this problem, biomaterials have been developed that combine antibacterial activity by contact killing with anti-adhesive properties, with the aim of combining these two activities within a material without one interfering with the other.[20–22] There have been several reports of surfaces that could alternate between killing and repelling in response to external stimuli such as heat, pH, light, and so on.[23] The practical use of these surfaces on indwelling devices is debatable since these external stimuli are challenging to administer within the human body. Furthermore, it is typically challenging to scale up the manufacturing of these surfaces due to multistep processes.

Surface modification might be achieved by combining the bactericidal and anti-adhesive components via the use of one

or two different macromolecules (Figure 1b).[22] For example, a copolymer of poly(N,N-(dimethylamino)ethyl methacrylate-co-2-methacryloyloxyethylphosphorylcholine) (P(DMAEMA-co-MPC)) was grafted from the polydimethylsiloxane (PDMS) substrate using surface-initiated reversible-addition fragmentation chain transfer polymerization.[24] The surface was then further quaternized to give it dual functions to effectively kill and repel both Gram-positive and Gram-negative bacteria.[24] In that case, a statistical copolymer was used, but block copolymers with either a first cationic block and then a hydrophilic anti-adhesive or the contrary were also tested (Figure 1b).[25,26]

To elaborate on both antibacterial and antiadhesive surfaces, Yang[27] and coworkers focused on the grafting of hydrophilic and anti-adhesive poly(N-hydroxyethyl acrylamide) (polyHEAA) synthesized by surface-initiated atom transfer radical polymerization (SI-ATRP), and cationic antibacterial poly((trimethylamino)ethyl methacrylate chloride) (polyMETAC) synthesized by surface-initiated photoiniferter-mediated polymerization (SI-PIMP). With a ratio of 50% between polyHEAA and polyMETAC chains, the grafted surfaces showed good antibacterial activity against *E. coli* and *S. aureus* after 72 h of incubation, without the surface being passivated by the deposition of bacteria (dead or alive) during at least 3 days.[27]

The biomedical catheters are commonly made of either silicones or mainly in soft aliphatic or aromatic polyether-based-polyurethane (PU), such as Pellethane or Tecoflex.[28,29] The
nature of the matrix and the geometrical form of catheters (hollow tubes instead of flat surface) requires to develop specific procedure to prepare anti-infective materials.[30,31] In these cases, the active coating is generally physically adsorbed (coating) or attached via a (poly)dopamine functionality or silane coupling, followed by harsh pretreatment. For example, Zhang et al.[32] functionalized the PU surface by plasma treatment and silane-based initiators to perform surface-initiated atom transfer radical polymerization of copolymers consisting of a bactericidal lower layer of Poly(2-(dimethyl decyl ammonium)ethyl methacrylate) (PQDMAEMA) and an antifouling upper layer based on poly(ethylene glycol) (PEG) integrated into the copolymer via a Schiff base structure. Since bacterial metabolism creates an acidic environment, the PEG layer preventing protein uptake is cleaved after a certain period of exposure to bacteria, allowing the quaternary ammonium layer to kill the bacteria. Nevertheless, although this methodology shows good results in terms of antibacterial activity against *S. aureus* and antifouling properties against proteins (Bovine Serum Albumin, BSA), it seems clear that once the PEG layer has been cleaved, the PU surface can no longer be protected against fouling. A similar approach was proposed by Xu and coworkers[33] with the same grafting strategy but with a lower layer of poly(3-[dimethyl-[2-(2-methylprop-2-enoyloxy)ethyl]azaniumyl]propane-1-sulfonate) (PDMAPS) extended by a poly(methacrylic acid) upper layer post-functionalized with an antimicrobial peptide as pendant groups. Another solution to allow efficient grafting is to prepare a prefunctionalized PU (PU with uretdione moiety[34] or PU with allyl moiety[35,36]) that could be post-functionalized by dipping[34] or thiol-ene reaction[35] to give an antibacterial/antiadhesive covalent coating.

However, whatever the approach used, all these methods are based on modifying the surface of materials by coating or grafting. These methods have the disadvantage of being multi-step, time-consuming, and difficult to envisage on a large scale and on tube geometries, limiting their use for a commercialized medical device.[10] In addition, the preparation of a pre-functionalized PU that could allow efficient grafting is required to register such new materials and then change the composition of the catheter bulk material by the manufacturer.

Bulk modification by polymer blending (extrusion) is a rather less explored route to impart antibacterial properties to common plastics, and only a few works have reported the introduction of polycationic polymers as additive. Xiu et al.[37] for example used this methodology to elaborate antibacterial acrylonitrile-butadiene-styrene (ABS) materials by coextrusion with a poly-methacrylic phosphonium salt but 10 wt.% was necessary to get only 90% antibacterial efficiency (Red Log = 1) against both *E. coli* and *S. aureus*. Xu et al.[38] developed a cationic additive bearing an allyl functionality (imidazole ionic liquid) that could be blended with commercially available PU and that could partially be grafted in the extruder due to the generation of a few radicals in the PU melt. The obtained materials with 3 wt.% of additive presented good antibacterial properties but with a non-negligible release of the cationic active species when soaked in water for 2 days.[38]

Our group developed a few years ago a simple and effective methodology to produce a wide library of antibacterial materials (polystyrene (PS), poly(methyl methacrylate) (PMMA), polyethylene terephthalate glycol (PETG), poly(ethylene) (PE), PU, etc.) by simple dispersion/co-extrusion of 2 wt.% of a specifically designed antibacterial copolymer within a polymer matrix (Figure 2a,b).[39–41] This methodology is promising since it could be extended to many materials, whatever the shape and size of the object and do not require changing the industrial process.

We showed in these previous studies that the introduction of 2 wt.% of such an additive did not modify the mechanical properties of the matrixes and that there were no detected leachables.[40] This approach has proved effective in providing materials active on *E. coli* and *S. aureus* under in vitro conditions, i.e., following the ISO 22196 procedure. However, such antibacterial materials were not tested to prepare medical devices and specially catheters. In view of the literature, it seems foreseeable that these materials will exhibit passivation phenomena when tested under conditions close to in vivo, i.e., in the presence of proteins or biological media. The aim of this work is first to investigate the antibacterial efficiency of our approach in the presence of proteins or biological media and then in case of the passivation of the materials to develop polyurethane products that are antibacterial by contact-killing and antiadhesive, using a simple process that can easily be scaled up to industrial scale. To achieve this, we will use the technology we have already developed, i.e., the dispersion of antibacterial copolymers within a matrix, but this time associating two diblock copolymers, the cationic antibacterial poly(butyl methacrylate-*block*-N,N-dimethylaminoethyl methacrylate) P(BMA-*b*-DMAEMA) copolymer quaternized with MeI and a second amphiphilic poly(butyl methacrylate-*block* - poly(ethylene glycol) methyl ether methacrylate) P(BMA-*b*-PEGₓMA) copolymer with x = 360, 500 or 950 Da for adding anti-adhesive properties (Figure 2b).

While the use of antibacterial copolymers based on methacrylic monomers containing a quaternary ammonium is well known, and PEG-based copolymers are already well reported for conferring anti-adhesive properties to a material, their conjugated association within a material by simple extrusion/mixing is, to our knowledge, a new process. The challenge is to obtain biomaterials that are both antibacterial and anti-adhesive, which could be used to produce medical devices that can be implanted in the body, i.e., in the presence of biological fluids, for long periods of time (long-lasting materials). Such properties are particularly sought-after in the case of venous, urinary, or enteral catheters.

To this end, we will first quantify the passivation rate of antibacterial PUs (obtained by dispersing 2 wt.% of the antibacterial copolymer in the presence of albumin, plasma and serum), then show that the use of PEGₓMA-based diblock amphiphilic copolymers is a preferred solution for limiting fouling of the material surface and thus maintaining good antibacterial activity against *S. aureus*, *E. coli*, methicillin-resistant *S. aureus* (MRSA), *A. baumannii* and *P. aeruginosa* after exposure to biological media (serum, plasma, albumin, gastric medium or intralipids).
<figure><img src="image_5.png" /></figure>

<figure><img src="image_6.png" /><figcaption>a) 3 steps :</figcaption></figure>

<figure><img src="image_7.png" /></figure>

<figure><img src="image_8.png" /><figcaption>**Figure 2.** a) Elaboration of antibacterial copolymer. b) Preparation of antibacterial PU materials by dispersion of an antibacterial copolymer used as an additive (2 wt.%). c) Elaboration of antibacterial AND antiadhesive PU materials by dispersion of two copolymers used as additives (2 wt.% of antibacterial AB copolymer and 2 wt.% of antiadhesive AA copolymer).</figcaption></figure>
<figure><img src="image_9.png" /><figcaption>**Figure 3.** Antibacterial activity of films of PU with 2 wt.% of AB copolymer against *E. coli* b) and e) and *S. aureus* a) and d) with various pretreatments: A) without any pretreatment (blank) and after pre-treatment with B) plasma, C) serum, and D) albumin. The Reference sample corresponds to PU film without AB copolymer. Results correspond to means ± S.D. (n = 4). Statistical significance was obtained using ANOVA (\*\*p < 0.01 and \*\*\*\*p < 0.001).</figcaption></figure>

## 2. Results and Discussion

### 2.1. Evaluation of the Adhesion of Biomolecules on Antibacterial PU Films

The antibacterial P(BMA-b-DMAEMA) copolymer quaternized with MeI was prepared by Nitroxide-Mediated Polymerization (NMP) as already reported.[39,40] The ratio of BMA to DMAEMA is set to 1:2, and the $M_n$ of the block copolymer called here AB for antibacterial is close to 20 000 g mol⁻¹ to avoid any leaching after being inserted into the matrix.[40] 2 wt.% of AB has been mixed and then dispersed by extrusion in a polyurethane matrix of Pellethane containing 20 wt.% of barium sulfate for imparting radio-opacity of the medical device. The final material was then pressed to obtain antibacterial PU films of 2 cm in diameter.

The antibacterial activity of the films is caused by the presence of cationic charges on the material surface, which kill bacteria upon contact.[40] It is thus plausible that these cationic charges may interact with other components of body fluids, such as proteins. A deposit of bacteria and bacteria debris could also pass

ate the material surface. To evaluate the rate of this eventual passivation phenomena, antibacterial PU films (with 2 wt.% of AB) have been exposed for 1 h at 37 °C to plasma, serum, or albumin before ISO 22196 antibacterial tests against *E. coli* and *S. aureus*.[41] As suspected, **Figure 3** clearly shows that the pre-exposure to biological media inhibited the antibacterial activity of the film against *E. coli* in the presence of Bovin Serum Albumin that is negative charged at neutral pH. The percentage of killing decreased from 100% to 67.6% ± 4.1%. This percentage of killing could be converted to a logarithm of bacterial reduction (Red Log) values, and in that case, a decrease from 4.3 to 0.5 was observed. Using human serum, the decrease of efficiency is lower, i.e., from 100 to 92.8% ± 2.6% (4.3–1.3 for the Red Log). It is interesting to note, however, that the reduction in antibacterial activity is lowest in the case of plasma, which is the medium that most closely resembles blood. The various interactions that could occur in such a complex medium already hindered partially the proteins to the cationic surface. Even if the percentage of killing is not statistically affected, the impact in the Red Log is more pronounced (4.3–1.9) and the materials could not be longer
<figure><img src="image_10.png" /></figure>

<table><thead><tr><th>b)</th><th>PBMA</th><th>-Z</th><th>PEG chain length (g/mol) – (n)</th><th>MAPEG conv (¹H NMR)</th><th>Theoretical M<sub>n</sub> of the block copolymer</th><th>M<sub>n</sub> of the diblock copolymer (SEC/THF-TFA)</th><th>Đ</th></tr></thead><tbody><tr><td></td><td></td><td>2a (AA<sub>360</sub>)</td><td>-H</td><td>360 – (8)</td><td>62</td><td>52,000</td><td>29,000</td><td>1.3</td></tr><tr><td></td><td></td><td>2b (AA<sub>500</sub>)</td><td>-CH<sub>3</sub></td><td>500 – (12)</td><td>60</td><td>69,000</td><td>60,000</td><td>1.8</td></tr><tr><td></td><td></td><td>2c (AA<sub>950</sub>)</td><td>-CH<sub>3</sub></td><td>950 – (22)</td><td>56</td><td>123,000</td><td>46,000<sup>a</sup></td><td>1.5</td></tr></tbody></table>

<caption>**Figure 4.** a) Synthesis by RAFT polymerization of antiadhesive poly(butyl methacrylate-*block*-PEG$_x$ methacrylate) P(BMA-*b*-PEG$_x$MA) diblock copolymers with *x* = 360, 500, or 950 Da. b) Characteristics of the anti-adhesive (AA) copolymers.</caption>

described as antibacterial (antibacterial threshold of 3 according to the ISO 22 196). A similar inhibition of the antibacterial activity against *S. aureus* was observed, with in particular a decrease for the Red Log from 4.3 to 0.8 and 2.3 after exposure to albumin and plasma, respectively.

## 2.2. Elaboration of Antibacterial and Antiadhesive PU Films

As stated in the introduction, there are many possibilities to combine anti-adhesive with antimicrobial properties on a specific material. To keep our approach that is to find a scalable and straightforward method to impart such properties, we thus investigated to disperse two different copolymers into the matrix via melt processing. We choose to keep the first AB one since we have already important data concerning its efficiency, safety, and non-leachability. We then prepared a new amphiphilic copolymer with a similar design, but we replaced the cationic DMAEMA block with a (polyethylene glycol) methacrylate to impart anti-adhesive properties.[42] More precisely, three poly(butyl methacrylate-block-PEG$_x$ methacrylate) P(BMA-*b*-PEG$_x$MA) diblock copolymers have been prepared by RAFT (Figure 4a) with three different PEG$_x$MA monomers having different PEG size (x corresponding to the molecular weight of the PEG, here 360, 500, or 950 Da).

Using 4-cyano-4-(phenylcarbonothioylthio)pentanoic acid (CDSPA) as RAFT agent, the first block of poly(butyl methacrylate) is obtained readily with a conversion of 80%, and properties close to those we obtained for the antibacterial block ($M_{n}$ = 9000 g mol$^{-1}$ and $Đ = 1.13$ determined by SEC in THF). A library of amphiphilic P(BMA-*b*-PEG$_x$MA) diblock copolymers (AA for anti-adhesive) was then prepared using this macroRAFT agent. The characteristics of the AA copolymers are gathered in Figure 4b. P(BMA-*b*-PEG$_{360}$MA) and P(BMA-*b*-PEG$_{500}$MA) diblock copolymers were analyzed by SEC in THF using a calibration based on Poly(methyl methacrylate) standards. The discrepancy of the theoretical $M_{n}$ and the experimental one is related to the difference in hydrodynamic volume behavior with the standards.[43] For the P(BMA-*b*-PEG$_{950}$MA) diblock copolymer, the SEC analysis was performed in DMF using a calibration based on Poly(methyl methacrylate) standards. Poly(PEG$_{950}$MA) is known to form comb-shaped structures with a very high grafting density.[44] The resulting material can therefore assume various conformations, such as starlike or brush-like polymers, depending on the relative length or ratio of the backbone to the side chain. Matyjaszewski and coworkers[44] thus reported that using a SEC analysis with a conventional calibration led to an experimental $M_{n}$ more than half of the theoretical one, and MALS analysis is necessary to obtain the correct value. Since the ¹H NMR analyses of the copolymer are coherent with the desired

<footnote>a SEC in DMF</footnote>
materials and the analyses of the diblock copolymers are also in good agreement with a controlled polymerization, we then used these copolymers for the preparation of anti-adhesive materials. Antibacterial and anti-adhesive PU films have been prepared like previously described by dispersion, extrusion and injection molding. Briefly, 2–4 w% of the AB and/or the AA copolymer(s) were blended with BaSO₄-containing Pellethane in an internal mixer, and the obtained masterbatch was extruded then pressed to prepare antibacterial and/or antiadhesive PU films.

According to our previous study, the characterization of the antibacterial surface is very difficult. AFM, Tof-SIMS, and FTIR analyses showed no detectable changes between the reference films and the film containing the AB copolymer due to a too low an amount of copolymer located on the surface.[40] Since we inserted a second copolymer into the matrix, we performed ATR analyses of the different film surfaces. The results are presented in Supplementary Information. Similar to the previous study,[40] the ATR analyses do not show detectable signals that could be attributed to the two copolymers. This is not surprising since the polyurethane (Pellethane) is a polyether-based material, similar to the PEG chains, and thus no specific signal could be assigned to the copolymer.

## 2.3. Protein Adhesion Measurements on Antiadhesive and Antibacterial PU Films

The hydrophilicity of the PU films with and without additives has been first evaluated by surface tension measurements. The pristine PU sample presented a value of 87°, that is in good agreement with reported values.[33,45] When only the antibacterial AB copolymer was added, the contact angle slightly decreased to 83.5°, which is due to the introduction of a low amount of polar cationic moieties on the surface of the film. When only the amphiphilic copolymers AA were dispersed in the film, the introduction of PEG chains on the surface decreased more strongly the contact angle between (77° and 55°), the value depending on both the PEG size and the amount of added AA copolymers (Figure 5a). These values are higher than the ones usually found when PEG or poly(PEGMA) was grafted directly onto the surface,[46,47] suggesting that the copolymer does not cover all the surface but is randomly distributed with a lower density.

We thus also determined the amount of non-specific protein adhesion using BSA as model protein and the BCA kit for quantitative measurements. The results are given in Figure 5b and showed that whatever the AA copolymer we showed a protein adhesion reduction between 22 and 39%, the main parameter being the size of the pendant PEG group and then the amount of AA added into the matrix. The lowering of the water contact angle as well as the reduction of protein adhesion let us envision to combine as expected such AA copolymers to the AB copolymer into PU matrix.

To confirm the results obtained with anti-adhesive films, BSA-FITC adsorption assays were conducted on PU films containing antibacterial and/or anti-adhesive copolymers. Analysis by epifluorescence microscopy (Figure 5c) and image analyses confirm first that Pellethane is prone to BSA adsorption and that the introduction of cationic groups (2 wt.% AB) strongly increases the adhesion, explaining the passivation of the surface since the bac-

teria cannot interact anymore with the surface. Second, the dispersion of only 2 wt.% of the anti-adhesive P(BMA-β-PEG₃₀₀MA) AA₃₀₀ diblock copolymer led to a near adhesion on the surface as expected with the previous spectroscopic BCA results. Lastly, when both copolymers were added, we could observe that the BSA adhesion decreased and gave values close or lower than the reference Pellethane, which could suggest an improved antibacterial efficiency even with the various pre-treatments. The same trend was observed with the AA₅₀₀ and AA₉₅₀ antiadhesive copolymers.

## 2.4. Mechanical Properties of Antiadhesive and Antibacterial PU Films

The incorporation of antiadhesive copolymer in the material aims to decrease the surface passivation by organic molecules, but the goal is to do it without modifying the mechanical properties of the pristine PU. We already showed that the introduction of 2 wt.% of AB copolymer do not modify the mechanical properties of the materials. In this study, we increase the amount of additive by adding to 2 wt.% of AB copolymer a second AA copolymer with a low Tg hydrophilic block that could have an impact of the mechanical properties. Tensile bars were prepared by injection molding, and tensile tests showed that whatever the AA copolymer, 2 wt.% did not significantly change the Young Modulus and the stress at break value of the material, proving that mechanical properties of the PU were preserved. One can also notice that a commercial sample (Ag zeolite) containing releasing silver ion nanoparticles presents greater variations with the reference PU compared to our materials (Figure 6).

## 2.5. Antibacterial Activity of the Double-Action PU Films After Exposure to Different Biologic Media

The antibacterial activity against *E. coli* and *S. aureus* was then assessed on PU films containing 2 wt.% of AB and films with 2 wt.% of AB and different amounts of AA copolymer (2–4 wt.% AA₃₀₀, AA₅₀₀, or AA₉₅₀) after different pre-treatments to evaluate the efficiency of this approach to avoid surface passivation. For the first tests, we focused on Human Serum Albumin and plasma since we previously showed that BSA is the worst case for the passivation of the materials and plasma mimics in vivo conditions. The ISO 22196 antibacterial tests after pre-treatments clearly show an important increase in the antibacterial efficiency as soon as the anti-adhesive copolymer is added. (Figure 7) Indeed, we observed that if targeting a reduction of Log(CFU) by 3, 2 wt.% of AA₅₀₀ or AA₉₅₀ is enough to avoid the passivation of the antibacterial material when exposed to plasma or albumin. These results are very encouraging to envision using this methodology to elaborate antibacterial PU for venous catheters, for example.

Once these two model strains tested, we also extended the testing to other clinically relevant strains that are methicillin-resistant *S. aureus* MRSA,[48] *P. Aeruginosa*[48] and *A. Baumanii*.[49] The tests were performed on the film containing 2 wt.% of AB and 2 wt.% of AA₅₀₀. The results are presented in Figure 7e,f. For all the strains a 1 h plasma exposure was performed before testing, and we observed that in all cases, more
<figure><img src="image_11.png" /></figure>

<table><tr><th></th><th>Wt%</th><th>AA<sub>360</sub></th><th>AA<sub>500</sub></th><th>AA<sub>950</sub></th></tr><tr><th>Contact angle (°)</th><td>2</td><td>77 ± 3</td><td>66 ± 3</td><td>62 ± 2</td></tr><tr><th></th><td>4</td><td>60 ± 3</td><td>58 ± 3</td><td>55 ± 3</td></tr></table>

<figure><img src="image_12.png" /></figure>

<table><tr><th></th><th>wt%</th><th>AA<sub>360</sub></th><th>AA<sub>500</sub></th><th>AA<sub>950</sub></th></tr><tr><th>BSA adhesion reduction (%)</th><td>2</td><td>22 ± 8</td><td>27 ± 3</td><td>32 ± 4</td></tr><tr><th></th><td>4</td><td>25 ± 4</td><td>35 ± 3</td><td>39 ± 3</td></tr></table>

<figure><img src="image_13.png" /></figure>

<table><caption>**Figure 5.** a) Images of contact angle measurements on PU films containing 2 or 4 wt.% of anti-adhesive copolymers (AA<sub>360</sub>, AA<sub>500</sub>, and AA<sub>950</sub>). b) Reduction of protein (BSA) adhesion (%) for PU films prepared by 2 or 4 wt.% of anti-adhesive copolymers (AA<sub>360</sub>, AA<sub>500</sub> and AA<sub>950</sub>). c) Determination of surface protein coverage by epifluorescence microscopy images of BSA-FITC adsorption assays on a) Pellethane, b) Pellethane + 2 wt.% AA<sub>360</sub>, c) Pellethane + 2 wt.% AB and d) Pellethane + 2 wt.% AB + 2 wt.% AA<sub>360</sub>.</caption><thead><tr><th>Sample</th><th>%Surface Protein Coverage</th></tr></thead><tbody><tr><td>a) Pellethane® Reference</td><td>12.1 %</td></tr><tr><td>b) Anti-adhesive Pellethane (2 w% AA<sub>360</sub>)</td><td>1.5 %</td></tr><tr><td>c) Antibacterial Pellethane® (2 w% AB)</td><td>34.6 %</td></tr><tr><td>d) Antibacterial and anti-adhesive Pellethane® (2 w% AB + 2w% AA<sub>360</sub>)</td><td>9.8 %</td></tr></tbody></table>
<figure><img src="image_14.png" /><figcaption>**Figure 6.** a) Tensile curves for the various PU samples. b) Young's modulus and c) stress at break values. Results correspond to means ± S.D. (n = 5). Statistical significance was obtained using ANOVA (\*p < 0.05; \*\*p < 0.01 and \*\*\*p < 0.005).</figcaption></figure>

than 99.9% of the bacteria were eliminated that corresponds to a red Log above 2.5. Except for MRSA, the Red Logs are always higher than 3, that corresponds to more than 99.99% of eliminated bacteria, confirming the antibacterial efficiency of the material whatever the bacterial strains.

To enlarge the application of these antibacterial and anti-adhesive PU materials, we studied their antibacterial activity after exposure to gastric media and intralipids to mimic in vivo conditions for enteral tubes application. As previously, if targeting a Red Log of 3, 2 wt.% of AA₅₀₀ or AA₉₅₀ is enough to get PU efficient against both *E. coli* and *S. aureus*, even if pre-exposed to gastric media or intralipids. The AA₃₆₀ copolymer was not evaluated at this stage as considered less efficient based on the previous results with plasma and albumin pre-treatments (Figure 8).

The antibacterial mechanism of methacrylate-based copolymers has been explored in the literature and is believed to primarily involve membrane destabilization, resulting in bacterial membrane lysis. We have already examined the mechanism of action of the antibacterial copolymer incorporated into a polymer matrix using a propidium iodide test.[40] The experiments conducted with various additive matrices (polylactide (PLA), high-density polyethylene (HDPE), low-density polyethylene (LDPE), and PETG) demonstrate a temporal increase in fluorescence values, thereby confirming that the antibacterial activity is attributed to the permeabilization of the bacterial membrane.

The comparison of the enhancement in fluorescence intensity between the positive control and the additivated matrices indicates that the permeabilization process of the copolymer integrated into the matrix is rapid (≈90% permeabilization in under 1 h), in contrast to the efficacy of leachable biocides, which typically necessitates several hours to be effective.[40]

To ensure that the mechanism is not perturbated with the introduction of a second amphiphilic copolymer, the membraneolytic activity of the films was investigated using ATP release assay after incubation with the Gram + *S. aureus*. Data showed in Figure 9 confirmed that films containing the two copolymers are able to cause strong membrane damages to *S. aureus*, even after a short period of contact (50% of permeabilization after 10 min) with maximal effect after 60 min of exposure.

## 2.6. Performance of the Antibacterial Material

The previous tests showed that such kind of materials have a good antibacterial property, but for interfaces intended for biomedical or implant applications, long-term immersion and repeated challenge tests are necessary to assess the performance of the material. To do these tests, we selected the material containing 2 wt.% of AB and 2 wt.% of AA₅₀₀ as the best compromise between efficiency and amount/cost of the additive.

We first focused on stability upon immersion into plasma and thus stored the material for different period of time from 1 day to 8 days. The results are presented in Figure 10a. As suspected, the bacterial efficiency of the material that was stored for 8 days was slightly lower than the material that was only exposed to 1 h to plasma, but the difference is not statistically different.

We then turned our attention to the repeated challenge test and thus performed the ISO 22196 test with a 1 h plasma exposure before test against *E. coli* as the most representative medium for a CVC application. With this protocol, we could observe on Figure 10b a small and still non statistically different decrease of the Red Log after the second test that could be due to a residual attachment of some bacterial debris or proteins.
<figure><img src="image_15.png" /><figcaption>c)</figcaption></figure>

<figure><img src="image_16.png" /><figcaption>b)</figcaption></figure>

<figure><img src="image_17.png" /><figcaption>d)</figcaption></figure>

<figure><img src="image_18.png" /><figcaption>e)</figcaption></figure>

<figure><img src="image_19.png" /><figcaption>f)</figcaption></figure>

<figure><img src="image_20.png" /></figure>

**Figure 7.** ISO 22196 tests of antibacterial (2 wt.% AB) and anti-adhesive films (0–4 wt.% of AA₃₆₀, AA₅₀₀, or AA₉₅₀) after plasma pretreatment against a) *E. coli* and b) *S. aureus* or albumin pre-treatments against c) *E. coli* and d) *S. aureus*. e) and f) Comparison of antibacterial activity (% of killing versus a pellethane film used as reference and Red Log CFU) of the antibacterial and antiadhesive film containing 2 wt.% of AB copolymer and 2 wt.% of AA₅₀₀ copolymer versus different Gram + and Gram – bacteria. Results correspond to means ± S.D. (n = 5). Statistical significance was obtained using ANOVA (*p* < 0.05; ** *p* < 0.01 and **** *p* < 0.001). RedLog is expressed in CFU.
<figure><img src="image_21.png" /><figcaption>**Figure 8.** ISO 22196 antibacterial tests on antibacterial and anti-adhesive PU films after gastric acid mimics pre-treatment against a) *E.coli* and b) *S.aureus* or intralipids pre-treatments against c) *E.coli* and d) *S.aureus*. Results correspond to means ± S.D. (n = 3). Statistical significance was obtained using ANOVA (\*p < 0.05).</figcaption></figure>

## 2.7. Inhibition of the Biofilm Formation

One of the limitations of medical devices such as catheters is the
formation of bacterial biofilm on their surface, which represents

<figure><img src="image_22.png" /><figcaption>**Figure 9.** *S. aureus* membrane permeabilization assay (ATP detection) with films of PU (reference, blank), PU+2%AA<sub>500</sub> (blue), PU+2%AB (red), PU+2%AB%+2%AA<sub>500</sub> (green). CTAB is used as a positive control.</figcaption></figure>

the main source of infection. It therefore seems important that
materials intended to remain in a biological environment (or in
the human body) for several days or weeks should limit, and if
possible, avoid biofilm formation. To address this issue, we com-
pared the ability of PU containing or not 2 wt.% AB and/or AA<sub>500</sub>
copolymer to prevent *E. coli* biofilm formation in static mode or
under orbital agitation, and after 24 or 48 h of incubation time
(Figure 11). Whatever the incubation time and the material com-
position, OD values measured with orbital shaking are always
lower than those measured in static mode. This is logically ex-
plained by the fact that orbital movement makes it more difficult
for bacteria to adhere to the surface. After 48 h of incubation time,
PU containing one or two kinds of copolymer (AB and/or AA<sub>500</sub>)
similarly decreased moderately the *E. coli* biofilm formation, via
either antibacterial or antifouling mechanisms. After 48 h, the
PU containing the combination of 2 wt.% of AB antibacterial
copolymer and 2 wt.% of AA<sub>500</sub> anti-adhesive copolymer drasti-
cally reduced the biofilm formation by 95% compared to a PU
without additives. The results using the combination of the AB
and AA<sub>500</sub> copolymers confirmed that the two mechanisms are
<figure><img src="image_23.png" /><figcaption>**Figure 10.** a) ISO 22196 antibacterial tests against *E. coli* on antibacterial and anti-adhesive PU films (PU+2%AB+2%AA₅₀₀) after 1, 24, or 8 days of immersion in plasma at 37 °C. b) ISO 22196 antibacterial tests on antibacterial and anti-adhesive PU films (PU+2%AB+2%AA₅₀₀) after 2 successive exposures to *E. coli* (1 h immersion in plasma before each test and samples immersed 1 h in ethanol between the tests). Results correspond to means ± S.D. (n = 3). Statistical significance was obtained using ANOVA (¹p < 0.05).</figcaption></figure>

needed to obtain an efficient reduction of *E. coli* biofilm formation on PU materials.

**2.8. Cytotoxicity and Hemocompatibility of Antibacterial and antiadhesive PU Films**

The innocuity of PU with 2 wt.% of AB antibacterial copolymer and/or 2 wt.% of AA₅₀₀ anti-adhesive copolymer was next evaluated using various cell types. In relation to their potential final use as catheters, the toxicity of the materials was tested using human vascular endothelial cells (HUVEC cells), human keratinocytes (HaCaT cells), and mouse fibroblasts (L-929) (Figure 12a). Whatever the tested cell, all the copolymer-based PU materials present a good biocompatibility (cell viability > 93%) with the lowest

values measured for the PU containing only the antibacterial AB copolymer. We already tested such copolymer in others organic matrices[40] and get always a good biocompatibility, that is also confirmed in PU. Even if all the materials presented no cytotoxicity, the biocompatibility is, as expected, improved by adding the antiadhesive AA₅₀₀ copolymer to the AB copolymer (cell viability increases to 96–99%). The same trend was observed by Agarwal et al.,[34] who showed that the viability of L929 fibroblasts on PU grafted with polyguanidinium (PU-PHMG) (85%) increased on surfaces additionally modified with dextran (PU-PHMG-Dex), returning to the value of the reference PU (95%).

To complete the previous cytotoxicity study, the rate of hemolysis of the antibacterial and anti-adhesive PU films was measured to validate the biocompatibility of the materials (Figure 12b).
<figure><img src="image_24.png" /><figcaption>**Figure 11.** Corrected Optical Density (ΔOD) measurements of *E. coli* biofilm formation on antibacterial, anti-adhesive or antibacterial and anti-adhesive PU surfaces after 24 h a) or 48 h b) incubation time in static mode or under orbital agitation. Results correspond to averages and standard deviations obtained on a series of 10 measurements per sample (n = 10). Statistical significance was obtained using ANOVA by comparing each sample with the others (***p < 0.005).</figcaption></figure>

Whatever the antiadhesive copolymer (AA<sub>500</sub> or AA<sub>950</sub>), 2 wt.% did not induce cytotoxicity against red blood cell (rate of hemolysis <1.2 %), confirming the possible use of these materials in contact with blood like for veinous catheters.

**3. Conclusion**

To prevent bacterial infection, catheters are usually treated either by coatings, impregnation, or by incorporation of additives like chlorhexidine, silver ions, antibiotics, or heparin. To have a long-lasting effect and avoid the toxicity of the released molecules, contact-killing materials were developed. Constructing polymer brushes and hydrogel layers on catheters is a common strategy to prevent pathogen colonization that has been widely studied in the literature. Nevertheless, even if several methods exist to coat/graft an open surface, it is very difficult to practically add in an industrial process a coating of the inside and outside surface of meter-long and narrow medically relevant tubes. We showed previously that the development of PU-based antibacterial materials by a dispersion and co-extrusion process using 2 wt.% of an antibacterial MeI-quaternized P(BMA-*b*-DMAEMA) copolymer as an additive has proved to provide effective antibacterial materials against *E. coli* and *S. aureus*, killing bacteria by a "contact-killing" mechanism. However, the antibacterial activity of the material is inhibited after exposure to biological media such as plasma or albumin, with the Red Log vs *E. coli* and *S. aureus* dropping from 4.3 to a value well below 3. This loss of activity can be explained by contamination of the material surface by debris or biological molecules. The covering of the surface then inhibits the material's antibacterial activity, as there is no longer any possible contact between the bacteria and the surface. To maintain antibacterial activity, even after pre-exposure to biological media such as plasma, albumin, intralipids, or gastric acids, different percentages (2 or 4%) of AA anti-adhesive copolymers of the P(MBMA-*b*-PEG<sub>x</sub>MA) type with x = 360, 500, and 950 Da were added to the material formulation to produce antibacterial and anti-adhesive PU.

After verifying that the addition of 2 wt.% of AA copolymer in addition to 2 wt.% of AB copolymers did not alter the mechanical properties of the PU materials, biological studies showed that 2 wt.% of AA<sub>500</sub> copolymer in addition to 2 wt.% of antibacterial AB copolymer was a formulation of choice for developing antibacterial and antiadhesive PUs, active against *E. coli*, *S. aureus*, *MRSA*, *A. baumannii* and *P. aeruginosa* even after pre-exposure to biological media and in particular plasma that mimics blood. These materials proved to be stable in biological conditions since they maintain their antibacterial activity even after 2 successive exposures to *E. coli* preceded by immersion in plasma or when stored in plasma for 8 days. These materials also revealed no toxicity to red blood cells, L929 fibroblasts, HaCaT, and HUVEC cells. In addition, the antibacterial and antiadhesive PUs were shown to prevent the formation of *E. coli*-derived biofilm. These results demonstrate that antibacterial and anti-adhesive PUs prepared by dispersion and co-extrusion of 2 wt.% of AB and 2 wt.% of AA<sub>500</sub> copolymers are a method of choice that could be optimistically envisioned to elaborate biomaterials suitable for long-lasting venous or enteral catheters. As part of this work, we have been able to produce kilograms of both copolymers and thus to produce prototypes of antibacterial and anti-adhesive catheters (Figure 13). The next step will be to study the activity of these materials in vivo.

To conclude, the methodology described in this paper offers numerous advantages and innovations compared to approaches already reported in the literature. Unlike grafting and coating methods, the dispersion and extrusion of additives into a polymer matrix is a technique that can easily be used on a large scale, as many industrial processes already use extrusion during the shaping stage. Furthermore, the use of copolymers rather than small molecules ensures the stability and nontoxicity of the material. This method for developing materials that are both antibacterial and anti-adhesive offers the possibility of using additives of a different chemical nature from that of the matrix. Methacrylic additives have been shown to be effectively incorporated into PU. It therefore seems realistic to use these same additives in other matrices (PE, polypropylene, or silicone). Finally, this approach has
<figure><img src="image_25.png" /></figure>

<table><thead><tr><th>Sample</th><th>% HUVEC viability</th><th>% HaCAT viability</th><th>% Fibroblast viability</th></tr></thead><tbody><tr><td>Reference</td><td>100.0± 7.0</td><td>100.0± 5.4</td><td>100.0± 9.8</td></tr><tr><td>AB (2 wt%)</td><td>92.9± 13.4</td><td>98.3± 8.3</td><td>96.6± 9.9</td></tr><tr><td>AA500 (2 wt%)</td><td>98.7± 12.9</td><td>102.6± 3.6</td><td>103.6± 12.9</td></tr><tr><td>AB (2 wt%) + AA500 (2 wt%)</td><td>96.7± 6.2</td><td>99.5± 11.5</td><td>96.0± 11.9</td></tr><tr><td>TX100</td><td>2.8± 0.4</td><td>1.6± 0.2</td><td>1.2± 0.2</td></tr></tbody></table>

b)

<figure><img src="image_26.png" /></figure>

<table><caption>**Figure 12.** a) Cell viability (in %) of i) HUVEC; ii) HaCAT and iii) L929 fibroblast after exposure to PU with no additive (reference) or PU containing 2 wt.% of AB and/or AA<sub>500</sub> copolymer. b) Hemolytic effect of the different antibacterial and antifouling compounds after a 1 h contact at 37 °C with a 8 vol. red blood cell suspension. Results were compared with a Triton X100 treated PU as a positive reference. Results correspond to means ± S.D. (n = 3). Results were compared with a Triton X100-treated PU as positive reference. Results correspond to means ± S.D. (n = 3).</caption><thead><tr><th rowspan="2"></th><th colspan="2">PU films</th><th rowspan="2">wt%</th><th rowspan="2">% Hemolysis</th></tr><tr><th>wt%</th><th>wt%</th></tr></thead><tbody><tr><td>PU</td><td>-</td><td>-</td><td>-</td><td>0.22 ± 0.19</td></tr><tr><td>PU + AB</td><td>2</td><td>-</td><td>-</td><td>0.56 ± 0.70</td></tr><tr><td>PU + AA<sub>500</sub></td><td>-</td><td>2</td><td>-</td><td>0.45 ± 1.08</td></tr><tr><td>PU + AA<sub>950</sub></td><td>-</td><td>-</td><td>2</td><td>0.89 ± 0.39</td></tr><tr><td>PU + AB +</td><td>2</td><td>2</td><td>-</td><td>1.12 ± 0.78</td></tr><tr><td>PU + AB +</td><td>2</td><td>-</td><td>2</td><td>0.79 ± 0.78</td></tr><tr><td>PU + TX100</td><td>-</td><td>-</td><td>-</td><td>100.22 ± 3.06</td></tr></tbody></table>
<figure><img src="image_27.png" /><figcaption>**Figure 13.** Picture of a) reference of pellethane-based catheter and b) pellethane-based catheter containing 2 wt.% of AB copolymer and 2 wt.% of AA₅₀₀ copolymer.</figcaption></figure>

the advantage of being able to provide materials that are active in a sustainable and stable manner under biological conditions close to in vivo (plasma) and in the case of strains encountered in clinical settings on catheters.

## 4. Experimental Section

**Materials:** Iodomethane (Mel, >99%, Acros), n-butyl methacrylate (BMA, 99%), 2-(dimethylamino)ethyl methacrylate (DMAEMA, 99%), acrylonitrile (ACN, ≥99%, contains 35–45 ppm monomethyl ether and hydroquinone as inhibitors), poly(ethylene glycol) methacrylate PEG₃₅₀MA (average M~n~ 360 Da, contains 500–800 ppm MEHQ as inhibitor), poly(ethylene glycol) methyl ether methacrylate PEG₅₀₀MA (average M~n~ 500 Da, contains 200 ppm BHT as inhibitor, 100 ppm MEHQ as inhibitor), poly(ethylene glycol) methyl ether methacrylate PEG₉₅₀MA (average M~n~ 950 Da, contains 300 ppm BHT as inhibitor, 100 ppm MEHQ as inhibitor) and 4-cyano-4-(phenylcarbonothioylthio)pentanoic acid (CDSPA, 97% (HPLC)) were purchased by Sigma-Aldrich. Azo-bisisobutyronitrile (AIBN, >98%) was purchased by TCI. BlocBuilder (>99%), an alkoxyamine based on the nitroxide SG1 (N-tert-butyl-N-[diethylphosphono (2,2-dimethylpropyl)]nitrooxide) and the 1-carboxy-1-methylethyl alkyl moiety, and SG1 were provided by Arkema.

Methanol (MeOH, laboratory reagent ≥99.6%), diethyl ether (contains BHT as inhibitor, puriss. p.a., ACS reagent, reag. ISO, reag. Ph. Eur., ≥99.8%), 1,4-dioxane (anhydrous 99.8%), and 2-methyltetrahydrofuran (MeTHF, BioRenewable, anhydrous, ≥99%, inhibitor-free) were purchased by Sigma-Aldrich. Toluene (≥99.8%, HPLC Grade) was purchased by Fischer Chemicals. Tetrahydrofuran (THF, pure-stabilized with BHT) and n-pentane (Pure) are provided from Carlo Erba. Pellethane 2363 90 AE + 20% BaSO₄ was sent by the company Vygon, but the virgin polymer provided from Lubrizol and Blanc Fixe Xr-Hn Barium Sulfate was originally purchased by Venator.

¹H spectra were obtained on a Bruker Avance DPX 400 MHz spectrometer. Tetramethylsilane (TMS) was used as an internal standard. The ¹H chemical shifts were referenced to the solvent peak for CDCl₃ (δ = 7.26 ppm). SEC analyses were performed using an EcoSEC system from

TOSOH equipped with a differential refractometer detector. THF was used as an eluent with 0.25 vol% toluene as a flow marker at a flow rate of 0.3 mL min⁻¹ after filtration using Alltech PTFE membranes with a porosity of 0.2 µm. The column oven was kept at 40 °C, and the injection volume was 20 µL. One ResiPore Precolumn (50 mm, 4.6 mm) and two ResiPore columns (250 and 4.9 mm) from Polymer Laboratories were used in series. The system was calibrated using PMMA standards from Agilent in the range 602–520000 g mol⁻¹. The SEC in DMF was performed using an EcoSEC instrument (Tosoh Bioscience) equipped with a solvent stocker, a dual flow refractive index detector, and a UV detector. The stationary phase was composed of 2 PSS Gram columns (300 mm × 7.8 mm) (Agilent) and a PSS Gram precolumn (50 mm × 7.8 mm). The whole system was thermostated at 50 °C. Mobile phase was filtered DMF (0.2 µm Nylon membrane filter) containing 0.1 M of NaNO₃ at 0.7 mL min⁻¹ for the sample part and 0.35 M l min⁻¹ for the reference one. Samples were prepared at a concentration of 0.25 wt.% in the mobile phase containing 0.25 vol.% of toluene, as a flow marker. Samples were filtered on 0.2 µm Nylon syringe filter before injection at a volume of 50 µL. Polymethyl methacrylate (PMMA) equivalent number-average and weight-average molar masses (M~n~ and M~w~) and dispersities D were calculated by means of a conventional calibration curve established by using PMMA standards from 0.602 to 520 kg mol⁻¹ (Agilent).

**Copolymer Synthesis and Characterization—Synthesis of Amphiphilic Cationic Antibacterial Copolymer:** Polymerization of the antibacterial copolymer (AB copolymer) follows the protocol described in previous work and presented in Figure 2.[³⁹,⁴⁰]

**Copolymer Synthesis and Characterization—Synthesis of Antiadhesive PEG-Based Copolymers:** The synthesis of the P(BMA-b-PEG₅₀₀MA) diblock copolymer (noted AA₅₀₀) is presented here as the typical two-step procedure used for all diblock copolymers. BMA (hydrophobic monomer, 11 g; 76.85 mmol), CDSPA (201 mg; 0.498 mmol), AIBN (32.62 mg; 0.199 mmol), and toluene (11 mL) were introduced in a 50 mL round-bottom glass flask fitted with a magnetic stir bar. After 30 min of nitrogen bubbling at room temperature, the media were heated at 70 °C under a nitrogen atmosphere. Samples were taken periodically to monitor conversion and the molecular weight. The reaction was stopped after 5 h. The conversion was determined by ¹H NMR using the peak of the –CH₂– close to the ester function of the monomer (δ = 4.11 ppm) and the peak of the
polymer (δ = 3.91 ppm). The PBMA was recovered by precipitation into a cold MeOH/H₂O (4/1 in volume) mixture and analyzed by SEC/THF to obtain the molecular weight (M~n~ = 9000 g mol⁻¹) and dispersity (D = 1.13) values.

After 24 h drying under vacuum, the PBMA was used to initiate the polymerization of the hydrophilic monomer (PEG~500~MA) and obtain the corresponding P(BMA-b-PEG~500~MA) diblock copolymer (also noted AA~500~). PBMA (0.5 g; 0.06 mmol), PEG~500~MA (5.8 g; 11.60 mmol), MeTHE (30 mL), and AIBN (5.26 mg; 0.03 mmol) were introduced in a 100 mL three-necked round-bottom glass flask fitted with a magnetic stir bar. The solution was deoxygenated by nitrogen bubbling for 30 min at room temperature before heating at 70 °C under a nitrogen atmosphere. The reaction was stopped when the conversion was close to 60% (6 h). The final P(BMA-b-PEG~500~MA) or AA~500~ diblock copolymer was recovered by precipitation in cold pentane, then dried under vacuum before analysis by ¹H NMR spectroscopy and SEC to obtain the composition and the number-average molecular weight, respectively.

**Elaboration of Antibacterial and Antiadhesive PU Films:** The preblending step was performed with a HAAKE Rheomix OS internal mixer driven by HAAKE PolyLab OS Rheodrive 7 controller. PUR pellets were dried at 90 °C under vacuum for 4 h before extrusion. Once dry, 50 g of pellets were manually loaded into the internal mixer with antibacterial/ anti-adhesive/antibacterial and anti-adhesive additives in the desired proportions (between 1 and 4 wt.%). The mixing was performed at 200 °C at 80 rpm for 3 min before recovering an antibacterial/antiadhesive/antibacterial and antiadhesive PUR masterbatch that was cooled with liquid nitrogen before being crushed in a FRITSCH Pulverisette 19 universal knife mill to obtain mixture chips. The extrusion step was performed using a Thermo Scientific HAAKE MiniLab II mini-extruder. The pellets were dried at 90 °C under vacuum for 4 h before extrusion. Once dry, 5 g of masterbatch chips were manually loaded into the extruder. Extrusion was performed at 210 °C at 80 rpm.

For mechanical testing, tensile specimens (type 5A) were prepared by injection molding using a Thermo Scientific HAAKE Minijet II injector. Small pellets of extruded material were collected in the nozzle heated to 200 °C of the injector and then injected with a pressure of 500 bar in a test specimen mold (tensile bar ISO527-2-5A) heated at 60 °C. Bioassays were performed on circular-shaped films of 3 cm diameter and 100 µm thickness. Upon exiting the extruder, the recovered rushes were cooled under air and then cut into small pellets, and finally shaped using a Specac hydraulic press equipped with heating platens and a thin film kit. The pellets were pressed between two aluminum sheets heated at 170 °C and under 2 tons of pressure for 10 s.

**Tensile Tests:** The tensile bars were analyzed using a Testometric TM M250-2.5CT bench coupled with the « winTestTM Analysis ». Tests were carried out in accordance with ASTM D412, which describes a tensile test protocol adapted to rubber and elastomeric materials. This method consists of carrying out tensile tests with a drawing speed of 500 mm min⁻¹ at ambient temperature and humidity. For each formulation, the Young's modulus and peak stress values were extracted by taking the statistical mean of 5 tensile bars.

**ISO 22196 Antibacterial Tests:** The bactericidal activity of the PU materials was evaluated as previously described.[41] The bacteria used in the study were either Gram-positive (*Staphylococcus aureus*, ATCC CRM-6538P, Methicillin-resistant *Staphylococcus aureus*, ATCC BAA 1717) or Gram-negative (*Escherichia coli*, ATCC 8739, *Pseudomonas aeruginosa*, CIP 107 398, and *Acinetobacter Baumanii* CIP 110 431) and were routinely grown on lysogeny broth (LB) agar plates. The bactericidal activity of the different matrices was evaluated according to an adapted ISO 22 196 procedure. Briefly, reference and antibacterial films of 3 cm in diameter were sterilized (UV irradiation for 15 min). These samples were placed in a six-well plate before being overlaid with a 112 µL drop of a bacterial suspension containing 2 × 10⁶ bacteria.mL⁻¹ in sterile water with 0.2 wt.% NB or PBS. This drop was then flattened by a piece of Stomacher homogenization bag cut into a 2 cm square and UV sterilized beforehand, to increase contact between the bacteria and the surface being tested. Films were then incubated at 37 °C for 24 h before adding 5 mL of TSB to dilute the bacteria and stop the contact killing. This new mixture was added to a further 5 mL

of TSB (1:10 dilution), followed by further dilutions to 1:100; 1:1000, and 1:10000 in sterile water. 10 µL of each dilution was placed in a quarter of a petri dish on an agar gel and then gently spread over the entire area. After one night at 37 °C, the number of colonies formed could be counted and the number of living bacteria determined. For each sample, the test was carried out in quintuplicate.

**Bacterial Membrane Permeabilization Assay:** The effect of the matrices on bacterial membrane integrity was evaluated through measurement of the release of intracellular ATP as previously described.[50] Briefly, pieces of matrices (squares of 4 × 4 mm) were sterilely prepared from extruded films and were placed at the bottom of the wells of sterile polypropylene 96-well microplates (Greiner BioOne). Log-phase bacterial suspension of *E. coli* or *S. aureus* was prepared as explained previously. Bacteria were diluted at a density of 10⁷ bacteria per ml in the same media used for the contact kill assay (0.2 wt.% NB). Twenty microliters of the bacterial suspension were added onto the surface of the matrices. Bacteria were then incubated at 37 °C for 10 or 60 min. At the end of the incubation period, 100 µL of ATP detection kit based on luciferase (Yelen, France) was added into each well, and the total volume (120 µL) was transferred into 96-well white plates for luminescence (Greiner BioOne). After 5 min of incubation, the luminescence of the wells was measured using a microplate reader (Biotek, Synergy Mx). Lysis buffer provided in the ATP detection kit was used as a positive control, giving 100% bacterial membrane permeabilization. Experiments were conducted in triplicate (n = 3).

**Biological Pre-Treatments:** Biological pre-treatments are carried out prior to ISO 22196 antibacterial testing to investigate the effect of surface fouling by biological media on the antibacterial properties of materials. In 6-well plates, samples (discs with D = 3 cm, thickness = 100 µm) were immersed in the tested biological liquid (2 mL) for 1 h at 37 °C. The biological liquid was removed, and surfaces were washed twice with PBS (2 mL) for 5 min under orbital shaking (400 rpm) and once with sterile water (2 mL) for 5 min under orbital shaking (400 rpm). Samples were then dried for 1 h at 37 °C before being sterilized under UV light for 15 min. Tested biological liquids were 0.1 g L⁻¹ albumin solution, human serum, or human plasma as blood-related media, and intralipids or mimic-gastric acid solution as enteral-related media.

**Repeatability Assay Against *E. coli*:** Repeated bacterial exposure tests were performed following ISO 22196 antibacterial tests against *Escherichia coli*, ATCC 8739 which were repeated twice on the same PU samples. A human plasma pre-treatment procedure was carried out prior to each ISO 22196 assays. Between the two ISO 22196 tests, the materials tested were washed twice with 70% ethanol by simple immersion for 1 h to eliminate any bacteria still alive on the film. For each sample, the test was carried out in quintuplicate.

**μ-BCA Protein Dosing Method:** Surfaces (Discs of D = 3 cm, thickness = 100 µm) were placed in wells of 6-well plates and washed with ethanol. These surfaces were stabilized in PBS (5 mL) for 1 h under orbital shaking. Then PBS was removed, and samples were immersed in a 1 mg mL⁻¹ BSA solution (3 mL) for 2 h. BSA solution is removed, and surfaces were washed once with PBS (5 mL) for 5 min under orbital shaking (400 rpm) and once with sterile water (2 mL) for 5 min under orbital shaking (400 rpm) to remove non-adsorbed BSA. To each sample was added a 1 wt.% SDS solution in water (2 mL) then they were placed in an ultrasonic bath for 25 min to desorb adhered BSA. Surfaces were removed and supernatant was conserved as analyte solutions. 100 µL of each solution is sampled in wells from 96-well plates. These solutions were completed with a reagent solution from μ-BCA assay kit (100 µL). Standard solutions with different concentrations of BSA were sampled (100 µL) in wells from 96-well plates and completed with reactive solution from μ-BCA assay kit (100 µL). Used 96-well plates were incubated for 2 h at 37 °C. During this time, CuSO₄ contained in the reactive solution with react: Cu(II) will be reduced as Cu(I) by BSA enzymes, and the solution's color will change from cyan-green to purple-pink. Finally, 96-well plates were analyzed spectrometrically to measure their absorbance at 562 nm. The quantity of adsorbed BSA is obtained using a calibration curve, which gives the mass of BSA adsorbed on the initial surfaces. Results were compared with reference materials without additives to know the reduction of adhered proteins (%).
**Prevention of Biofilm Formation:** The antibiofilm activity of the materials was investigated as previously described with minor modifications.[51] Liquid cultures of *E. coli* (ATCC 8739) in Lysogeny Broth (LB) were prepared from colonies grown on LB agar plates. Tubes were agitated at 200 rpm overnight at 37 °C. The next day, overnight bacterial culture was diluted 1/100 in LB and incubated at 37 °C under agitation at 200 rpm until bacteria reached log phase growth (i.e., OD 600 nm ≈0.5). Bacteria were then diluted in LB to reach a density ≈10⁸ bacteria per mL. The 4 mL of this bacterial culture were deposited on top of sterile films placed into 6-wells plates. Films incubated with LB alone were used as controls. After 24 or 48 h incubation at 37 °C under static or orbital agitation (at 200 rpm) conditions, planktonic bacteria were removed by aspiration, and the films were washed 5 times by immersion of the films in beakers containing distilled water. Films were then placed into new 6-wells plates and were stained with 2 mL of crystal violet stain at 0.1% for 15 min at room temperature. After staining, films were washed 5 times by immersion of the films in beakers filled with distilled water and were then dried 1 h at 37 °C. The crystal violet stain was then solubilized in 2 mL of 30% acetic acid (v/v) for 15 min under agitation at 200 rpm. One hundred microliters were finally transferred into 96-well plates, and the amount of solubilized crystal violet was measured at 570 nm. The OD 570 nm values of the films exposed to bacteria were corrected by the OD 570 nm values obtained with films exposed to LB alone.

**Hemolysis Assays:** The hemolytic activity of the solid antimicrobial materials was also determined. Briefly, 10 µL of human erythrocyte suspension (8% in PBS) was then added per well of the sterile 96-well microplate containing the solid antimicrobial materials. After 60 min of incubation at room temperature, 90 µL of PBS was added, and the microplate was centrifuged at 800 g for 5 min. The 50 µL were then carefully collected and transferred to a new 96-well microplate, and OD was measured at 405 nm. Hemolysis caused by copolymers was expressed as the percentage of hemolysis with Triton-X100 being used as a positive control.

**Cytotoxicity Assay:** The toxicity of the materials was evaluated as previously described40 using human vascular endothelial cells (HUVEC cells obtained from ECACC, Sigma-Aldrich, Lyon, France), human skin keratinocytes (HaCaT cells obtained from Creative Bioarray, Shirley, NY 11 967, USA), and mouse fibroblasts (L-929 cells (CCL-1) obtained from ATCC, Molsheim Cedex, France). Cells were routinely maintained on 75 cm2 flasks in a CO2 incubator at 37 °C in Dulbecco's modified essential medium (DMEM) supplemented with 10% fetal bovine serum (FBS), 1% Lglutamine, and 1% antibiotics (all from Thermo Fisher Scientific, Illkirch-Graffenstaden, France) for HaCaT and L-929 cells, or in endothelial cells specific medium (Sigma-Aldrich, Lyon, France) for HUVEC cells. To test the toxicity of the materials, cells were trypsinised, counted under microscope using Malassez cell, and seeded at 30000 cells per wells in 96-wells plates (100 µL per well). The next day, pieces of matrices (squares of 4 x 4 mm) were sterilely prepared from extruded films and were sterilely added to wells containing cells. Wells without matrices added and wells treated with Triton X-100 at 0.1% were used as untreated/negative and positive toxicity controls, respectively. After 48 h incubation, resazurin was added to each well at a final concentration of 30 µg mL⁻¹ and cells were further incubated 1 h at 37 °C. The fluorescence intensity of the wells was then measured using a microplate reader (SynerMix, Biotek, Ex 530 nm/Em 590 nm). The fluorescence values were normalized by the untreated controls and expressed as the percentage of cell viability.

**Supporting Information**

Supporting Information is available from the Wiley Online Library or from the author.

**Acknowledgements**

The authors thank VYGON, the Centre National de la Recherche Scientifique (CNRS), The SATT Sud-Est and Aix-Marseille Université (AMU) for

Financial support. VYGON is acknowledged for the PhD funding of Baptiste Caron. The authors thank Dr. Nicolas Vidal and the Yelen company for the ATP detection kit based on luciferase. The authors thank Isabelle Martin, Pr. Christine Bressy and Dr. Olivier Soppera for ATR analyses.

**Conflict of Interest**

B.C., M.M., A.L., M.L., J.L.C., Y.G., and C.L. are named inventors on patent applications (French Patent Application FR2414002 (2024)) filed by VYGON, the CNRS, Centrale Méditerranée, and Aix-Marseille Université on the anti-adhesive and antibacterial materials described in the current study.

**Data Availability Statement**

The data that support the findings of this study are available from the corresponding author upon reasonable request.

**Keywords**

additive, antiadhesive, antibacterial, catheter, diblock copolymer, extrusion, polyurethane

Received: August 4, 2025

Revised: October 20, 2025

Published online: November 11, 2025

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