Controlled Co-delivery of Anti-inflammatory Drugs from Bilayer Polymer Films Coating a Meniscus Implant

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Blanco, Gustavo Lou, Alba Pensado-López, Aldo Ummarino, and 3 more This is a preprint; it has not been peer reviewed by a journal. https://doi.org/ 10.21203/rs.3.rs-6869700/v1 This work is licensed under a CC BY 4.0 License Status: Published Journal Publication published 25 Aug, 2025 Read the published version in Drug Delivery and Translational Research → Version 1 posted 5 You are reading this latest preprint version Abstract Knee osteoarthritis (OA), a degenerative joint disease, is increasingly prevalent worldwide and often results from a meniscal deterioration that leads to meniscus removal. Replacing the damaged meniscus with a non-biodegradable prosthesis offers an innovative solution to prevent OA progression, particularly in older patients. However, the long-term use of anti-inflammatory drugs for pain relief and prosthesis integration can cause severe off-target side effects. The objective of this work was to design and develop drug-loaded bilayer polymer films to be used as coatings for a meniscus polycarbonate urethane (PCU). The developed bilayer polymer films enabled a sustained release of two anti-inflammatory drugs - dexamethasone (DEX) and celecoxib (CLX) - with distinct release kinetics (1-4 weeks for DEX and 6-9 months for CLX). This release profile was defined to modulate post-surgical and chronic inflammation within the knee joint, respectively. Two bilayer prototypes showed consistent biodegradation, drug release, drug loading, and reproducibility. Furthermore, the systems were sterile, biocompatible, and maintained the anti-inflammatory efficacy of the released drugs, effectively reducing pro-inflammatory cytokine secretion from human primary macrophages. Knee osteoarthritis meniscus implant controlled drug release polymer anti-inflammatory drug inflammation Figures Figure 1 Figure 2 Figure 3 Figure 4 Figure 5 Figure 6 Figure 7 Figure 8 Introduction Osteoarthritis (OA) is a multifactorial, degenerative, and chronic joint disease characterized by progressive cartilage degradation, bone remodeling, and an imbalance between synthesis and degradation processes, leading to pain, inflammation, stiffness, and loss of function [ 1 ]. It is the most common form of arthritis and a major cause of disability, affecting over 650 million individuals worldwide in 2020 [ 2 ]. Knee OA is particularly prevalent, with its increasing incidence attributed to longer life expectancy and rising obesity rates, the latter being a significant risk factor [ 3 ]. The meniscus plays a crucial role in knee function by providing stabilization, load distribution, and shock absorption [ 4 – 6 ]. However, meniscal injuries are common and often fail to heal due to the avascular nature of the inner meniscus, necessitating surgical intervention [ 7 , 8 ]. While partial meniscectomy is frequently performed to alleviate pain, it tends to accelerate joint degeneration and significantly increases the risk of OA development [ 9 ]. The use of meniscus implants or scaffolds that promote regeneration is a promising alternative to avoid the onset of OA after a meniscectomy [ 10 ]. To address this, meniscus implants and scaffolds have been developed to restore joint biomechanics and delay disease progression. Among them, the non-biodegradable NUsurface® prosthesis (Active Implants, Israel) serves as an artificial meniscus substitute after total meniscectomy. Approved in several European countries and Israel, it is currently undergoing FDA clinical trials [ 11 , 12 ]. Despite its advantages, implantation of the NUsurface® prosthesis could induce a foreign body reaction (FBR), requiring anti-inflammatory treatment to mitigate the patient’s pain, post-surgical inflammation and improve implant integration [ 13 , 14 ]. Current anti-inflammatory drugs have important limitations; for example systemic administration of COX-2 inhibitors, such as celecoxib (CLX), can cause gastrointestinal, renal, and cardiovascular side effects [ 15 , 16 ], whereas intra-articular (IA) corticosteroid injections, such as dexamethasone (DEX), provide short-term relief but pose risks of infection and patient discomfort [ 17 ]. To overcome these challenges, localized and sustained drug delivery strategies are needed to enhance therapeutic efficacy while minimizing systemic and local adverse effects [ 18 ]. Polymer coatings have been widely explored to improve implant biocompatibility, facilitate tissue integration, and enable controlled drug release. An optimized coating can provide site-specific drug release, shielding drugs from enzymatic degradation while reducing systemic toxicity [ 19 , 20 ]. CLX, a potential disease-modifying osteoarthritis drug (DMOAD), has shown protective effects in OA models by reducing inflammation and cartilage degradation [ 21 – 25 ], whereas a low-dose IA administration of DEX has shown chondroprotective potential when delivered IA, and its potential consideration as a DMOAD has been suggested. Therefore, the hypothesis of this study has been that a polymer bylayer coating enabling the sustained release of two drugs widely used in OA treatment, CLX for long-term inflammation control and short-term release of DEX for early-phase inflammation management, may improve implant performance and clinical outcomes [ 26 , 27 ]. Within this context, the current challenge in the field of meniscus-related degenerative pathologies is to design in situ therapies enabling controlled drug release with the goal of achieving a prolonged response and a reduction of systemic side effects [ 28 ]. In this study, we developed a bilayer polymer coating for the NUsurface® prosthesis intended to release these two anti-inflammatory drugs with distinct, controlled release kinetics. Using square-shaped polycarbonate urethane (PCU) implants (0.7 × 0.7 × 0.3 cm), we screened various polymer formulations to optimize controlled DEX release for short-term inflammation control (1–4 weeks) and CLX release for long-term pain and inflammation management (6–9 months). The coatings were characterized for their drug release kinetics, degradation profiles and biocompatibility. To assess their anti-inflammatory efficacy, we conducted in vitro studies in primary human monocyte-derived macrophages (HMDMs). Our findings provide a foundation for the development of an implantable, bioactive, drug-releasing meniscus prosthesis designed to enhance post-surgical outcomes and reduce OA progression. Materials and Methods 2.1. Materials Dexamethasone (DEX) and celecoxib (CLX) were supplied by Acofarma (Madrid, Spain) and Sigma-Aldrich (Missouri, USA), respectively. Poly (lactic-co-glycolic acid) (PLGA), of two different lactic acid:glycolic acid (LA:GA) ratios (PLGA 50:50, and PLGA 85:15), poly(caprolactone) (PCL) of low (LMW-PCL) and high molecular weight (HMW-PCL), poly(L-lactide) (PLLA) of low (LMW-PLLA) and high molecular weight (HMW-PLLA), and poly(lactic acid)-poly(ethylene glycol) di-block copolymer (PLA-PEG) were purchased from Evonik Industries (Darmstadt, Germany). LMW-PLGA LA:GA ratio 50:50 was obtained from PolySciTech, a division of Akina (Indiana, USA). Polyethylene glycol (PEG) of low (PEG 400) and high molecular weight (PEG 1450) were purchased from BASF (Ludwigshafen am Rhein, Germany). Sodium azide (NaN 3 ) was purchased from Sigma-Aldrich (Missouri, USA). Acetone was obtained from Fisher chemicals (New Hampshire, USA). Dichloromethane (DCM and acetonitrile (ACN) were distributed by Scharlau (Barcelona, Spain). Phosphate saline buffer (PBS), Tween®80, trifluoracetic acid (TFA) and methanol (MeOH) were supplied by Scientific (Nottingham, England), Merck (Darmstadt, Germany), Sigma-Aldrich (Missouri, USA) and VWR Chemicals (Pennsylvania, USA), respectively. The square-shaped implants made of polycarbonate urethane (PCU) were kindly donated by Active Implants (Israel). 2.2. Drug solubility in release buffer An excess of the drugs (DEX and CLX) was incubated in agitation (700 rpm) in 1 mL of PBS supplemented with Tween®80 concentrations ranging from 0 to 1% (w/v) for 24 hours. Then, samples were centrifuged at 10,000 rpm for 20 minutes, the supernatant was diluted in MeOH:H 2 O 65:35 (v/v), and the solutions were quantified by ultraperformance liquid chromatography (UPLC) with a TUV detector at 239 nm with a column Kinetex® 1.7 µm C18 100 Å, LC Column 50 x 2.1 mm acquired from Phenomenex (Torrance, CA, USA), maintaining the samples at 20°C in a Waters Acquity H-Class UPLC system (Waters, Milford, USA). 2.3. Production of drug-releasing polymer films To prepare the polymer-drug solutions, appropriate amounts of each drug (DEX or CLX) were dissolved in acetone and DCM, respectively [ 29 – 31 ]. In DEX-releasing polymers, PLGA (PLGA 50:50; intrinsic viscosity (IV) = 0.32–0.44 dL/g and PLGA 85:15; IV = 1.3–1.7 dL/g), LMW-PLGA (PLGA 50:50; MW = 8 kDa), and PLA-PEG (MW = 5.3 kDa) were dissolved to their final concentrations using acetone solutions of DEX of determined concentrations. Regarding CLX-loaded polymers, the polymers PCL, with intrinsic viscosities of 0.39 dL/g (~ 32 KDa; “LMW-PCL”) and 0.9 dL/g (~ 73 KDa; “HMW-PCL”) and PLLA, with IV of 1.0 dL/g (~ 74 KDa; “LMW-PLLA”) and 2.9 dL/g (~ 210 KDa; “HMW-PLLA”) as well as combinations of LMW-PCL (IV = 0.39 dL/g ~ 32 kDa) and HMW-PLLA (IV = 2.9 dL/g ~ 210 kDa) were dissolved to their final concentrations using DCM solutions of CLX of determined concentrations. The polymers used were selected based on their hydrophobicity and degradation rates. All these procedures were carried out under ambient conditions. These films were prepared by solvent casting. 70 µL of polymer-drug solutions were then cast on square-shaped meniscus implants (made of PCU) of 0.7 x 0.7 x 0.3 cm. The organic solvent was allowed to evaporate for 1 hour at room temperature. The resulting drug-loaded polymer films were then vacuum-dried for at least 24 hours (Supplementary Fig. 1A). 2.4. Production of bilayer drug-releasing polymer coatings To prepare the polymer-drug solutions, appropriate amounts of each drug (DEX or CLX) to achieve the desired concentrations were dissolved in acetone and DCM, respectively [ 29 – 31 ]. Polymers, LMW-PLGA (PLGA 50:50; MW = 6.9 kDa), and PLA-PEG (MW = 5.3 kDa) were dissolved at a concentration of 200 mg/mL using a solution of DEX of 5 mg/mL. Regarding CLX-releasing polymers, blends of PLLA (IV = 2.9 dL/g; MW ~ 210 kDa) and PCL (IV = 0.39 dL/g; MW ~ 32 kDa) were dissolved to their final concentrations using a solution of CLX of 30 mg/mL. The square-shaped PCU implants (0.7 × 0.7 × 0.3 cm), held in place by a needle, were immersed in the polymer-CLX solution and immediately withdrawn, allowing 10 min for drying. The cycle was repeated a total of 3 times. After an additional 3 hours of drying, 5 new immersion cycles were performed, in this case in the polymer-DEX solution and with a drying time of 15 minutes between cycles. Finally, the organic solvent was allowed to evaporate for 72 hours (Supplementary Fig. 1C). 2.5. Drug release evaluation Drug release studies were performed in agitation (450 rpm) at 37°C in PBS Tween®80 1% (w/v) to ensure sink conditions. The amount of drug (DEX and CLX) released was quantified by reverse phase ultra-performance liquid chromatography (UPLC) with a TUV detector at 239 nm using a column Kinetex® 1.7 µm C18 100 Å, LC Column 50 x 2.1 mm acquired from Phenomenex (Torrance, CA, USA), maintaining the samples at 20°C in a Waters Acquity H-Class UPLC system (Waters, Milford, USA) [ 32 ]. The mobile phase consisted of A: deionized H 2 O acidified with TFA 0.1% (v/v) and B: ACN acidified with TFA 0.1% (v/v) pumped with a flow rate of 0.1 mL/min. The injection volume was 5 µL, and the column oven temperature was set to 40°C. To control the UPLC/UV system as well as for data acquisition and processing, EMPOWER software was used. To quantify the amount of DEX and CLX, a calibration curve, ranging from 1 to 100 ppm was performed. The curves had a correlation coefficient (R 2 ) of 1 for DEX, and 0.9999 for CLX (n = 22). The validation procedure was carried out according to the ICH guidelines [ 33 , 34 ]. Limit of detection (LOD) and quantification (LOQ) were calculated directly from the calibration plots, as 3.3σ/S and 10σ/S, respectively, where σ is the standard deviation of intercept and S is the slope of the calibration plot [ 35 ]. The values were LOD = 1.73 ppm, and LOQ = 5.25 ppm for DEX, and LOD = 0.3 ppm, and LOQ = 0.92 ppm for CLX. 2.6. Drug loading assay PCU prostheses coated with drug-releasing polymers were immersed in 5 mL of a mixture of DCM/acetone 3/2 (v/v) for up to 24 hours to ensure polymer disolution. Samples of 200 µL were then transferred to microtubes containing 800 µL methanol and centrifuged at 10,000 rpm for 20 min. 200 µL of supernatant were further diluted with 800 µL MeOH:H 2 O 65:35 (v/v) and the concentration of drug was quantified by UPLC with a TUV detector at 239 nm using a column Kinetex® 1.7 µm C18 100 Å, LC Column 50 x 2.1 mm acquired from Phenomenex (Torrance, CA, USA), maintaining the samples at 20°C in a Waters Acquity H-Class UPLC system (Waters, Milford, USA) [ 32 ]. The mobile phase consisted of A: deionized H 2 O acidified with TFA 0.1% (v/v) and B: ACN acidified with TFA 0.1% (v/v) pumped with a flow rate of 0.1 mL/min. The gradient was from 25–60% of B in 6.5 min, and 6 min from 60–25% of B. The injection volume was 5 µL, and the column oven temperature was set to 40°C. To control the UPLC/UV system as well as for data acquisition and processing, EMPOWER software was used. To quantify the amount of DEX and CLX, a calibration curve, ranging from 1 to 100 ppm was used. The release buffer had no matrix effect on the quantification of the drugs. The curves had a correlation coefficient (R 2 ) of 1 for DEX, and 1 for CLX (n = 8). The values were LOD = 0.383 ppm, and LOQ = 1.162 ppm for DEX, and LOD = 0.187 ppm, and LOQ = 0.567 ppm for CLX. 2.7. Characterization of the drug-releasing polymer films 2.7.1. Differential Scanning Calorimetry (DSC) DSC measurements for polymers, drugs, and combinations of polymer and drug were recorded with a DSC Q1000 V9.9 (TA Instruments, New Castle, DE, USA) in standard aluminum sample pans. The samples were heated from 20°C to 400°C, depending on the sample analyzed, with a heating rate of 10°C/min in a nitrogen atmosphere. Data recording and processing of the first heating cycles were carried out with the software Advantage (TA Instruments, New Castle, DE, USA). 2.7.2. Powder X-ray diffraction (XRD) Diffraction measurements of crystalline powder were carried out by an Empyrean diffractometer of the PANAlytical brand. The X-rays were obtained from a sealed tube with Cu anode (λ(Kα1) = 1.5406 Å) and were collimated prior to incidence on the sample with optics including a W/Si bilayer mirror. The radiation emitted by the sample was collected with a "PIXcel3D" type solid state detector. The samples were mounted at ambient temperature on a flat base without signal (Si single crystal), to avoid the amorphous component different from that coming from the sample. Diffractograms were taken in an angular range of 2 to 40° with a step of 0.04 and a time per step of 8 s. To perform the mathematical adjustments of the obtained diffractograms, the program HighScore Plus: Version 3.0d was used. 2.8. Degradation of bilayer drug-releasing polymer coatings Drug-releasing polymer coated PCU implants were immersed in a volume of 10 mL of PBS (pH 7.4) supplemented with 1% (w/v) Tween®80. At defined time points, the coated squared-shaped PCU implants were removed from the release buffer for further analysis: 2.8.1. Field-emission scanning electron microscopy (FESEM) Drug-releasing polymer coatings were sputter coated with a layer of iridium and imaged in a Zeiss UltraPlus analytical FESEM with a beam voltage of 3 kV and a magnification ranging from 500 to 10,000X for the analysis of the surface of the coatings. Also, Zeiss EVO analytical FESEM with a beam voltage of 20 kV and magnification ranging from 60X to 5,000X was used for the measure of the coating thickness and analyzing side profile upon degradation after sagittal cut. 2.8.2. pH measurements pH was measured in the media where the polymer coated PCU implants were incubated over specific periods using a pH meter calibrated using standard buffer solutions of known pH values ranging from pH 2.0 to 10.0. 2.9. In vitro studies 2.9.1. Drug release evaluation Drug release studies to evaluate released drug activity in vitro were performed at 37°C in PBS Tween®80 0.05% (w/v) supplemented with penicillin/streptomycin 1% (v/v). The amount of drug (DEX and CLX) released was quantified by UPLC using the same method as described in section Drug loading assay (Section 2.6) . The release buffer collected and quantified at each time point was subsequently lyophilized. Afterwards, the powder was resuspended in sterilized miliQ water and evaluated in vitro . Stored release buffers were diluted to achieve a concentration of DEX and CLX (Table 1 ) as close as possible to their active anti-inflammatory concentration (10 µM for DEX and 25 µM for CLX), based on previous investigations [ 27 ]. A minimum dilution of 1:10 using RPMI was implemented for each sample, regardless of the concentration of drugs quantified in the media (Supplementary Fig. 2), to prevent possible toxicity by the release buffer (Tween ® 80 0.05% w/v). Table 1 Concentrations of DEX and CLX obtained from Prototype PLA-PEG and Prototype PLGA bilayer polymer coatings and total percentage (%) of drug released at each time point for the in vitro validation of their sterility and anti-inflammatory activity, and efficacy. CLX DEX Prototype Time point ppm µM % released ppm µM % released Dilution Prototype PLA-PEG 3 h 27.27 71.50 0.71 ± 0.22 193.63 493.38 39.08 ± 1.01 1:50 3 d 118.75 311.38 5.58 ± 0.97 97.58 248.65 89.16 ± 3.67 1:20 1 w 58.13 152.44 7.38 ± 1.01 22.95 58.49 93.70 ± 3.06 1:20 2 w 130.69 342.70 11.20 ± 1.16 17.41 44.36 96.63 ± 1.40 1:10 4 w 134.31 352.17 20.37 ± 1.58 1.81 4.60 ~ 100 1:15 Prototype PLGA 3 h 7.14 18.74 0.21 ± 0.02 16.40 41.80 6.28 ± 0.31 1:10 3 d 40.81 107.00 2.81 ± 0.32 45.82 116.74 44.76 ± 1.76 1:10 1 w 60.53 158.73 4.87 ± 0.31 48.45 123.46 63.51 ± 1.03 1:10 2 w 108.59 284.74 8.49 ± 1.87 25.69 63.46 72.61 ± 2.87 1:10 4 w 115.87 303.84 16.08 ± 3.44 10.49 26.75 82.77 ± 5.89 1:12 Abbreviations: CLX: Celecoxib. DEX: Dexamethasone. PLA-PEG: poly(lactic acid)-poly(ethylene glycol) di-block copolymer. PLGA: Poly(lactic-co-glycolic) acid. ppm: parts per million (µg/mL). h: hour. d: days. w: week. µM: Micromolar. Values represent the mean ± standard deviation (n=3). 2.9.2. In vitro control of endotoxin contamination The chromogenic LAL-test (LO50650U, Lonza) was used, following the manufacturer’s indications. Each of the samples used for in vitro studies showed endotoxin levels below 0,125 EU/ml (otherwise the samples were discarded) to prevent possible interferences in their immunotoxicity activity. 2.9.3. Monocyte isolation, differentiation, and treatment of human primary macrophages Monocytes were isolated from the whole blood of six healthy donors (buffy coats) following the Ficoll and Percoll separation method, as previously described [36], and subsequently counted with Trypan Blue staining 0.04% using a counting chamber (Kova international, BVS100H). In order to obtain human monocyte-derived macrophages (HMDMs), monocytes were resuspended in RPMI supplemented with penicillin/streptomycin and glutamine, subsequently counted, and 10 6 cells were plated in each well of 24-well plates to allow their attachment to the plate. After 30 minutes of incubation at 37 °C, media was removed, and a new media containing human recombinant M-colony stimulating factor (hrM-CSF, Peptrotech, 300-25) supplemented with fetal bovine serum (FBS), penicillin/streptomycin and glutamine were added for differentiation towards macrophages after 5 days. At day 5, macrophages were treated with appropriately diluted release buffers. Samples were added together with complete media for 24 hours. Non-treated HMDMs were used as the negative control, while HMDMs treated with LPS + IFNγ (100 ng/mL and 50 ng/mL, respectively) were used as the positive control (M1-like macrophages). All other samples were treated with free drugs (DEX, CLX, and DEX+CLX) at appropriate concentrations (10 mM for DEX and 25 mM for CLX) for comparison with equivalent concentrations of the drugs in release buffer from Prototype PLA-PEG or Prototype PLGA drug-releasing bilayer polymer coatings at determined times. 2.9.4. Evaluation of biocompatibility and toxicity To assess the biocompatibility of the drug-loaded polymer coatings using primary human macrophages, the AlamarBlue™ cell viability assay was performed. After exposure of macrophages to conditions described in Table 1 for 24 hours (relative day 6 of cell culture), which corresponds to the release buffer containing the drugs and products of polymer degradation at different time points, supernatants were collected for cytokine quantification, while 1 mL of AlamarBlue™ HS Cell Viability Reagent 10% (Invitrogen, A50101) was added to the remaining cells in each well of the 24-well plates, following the manufacturer’s protocol. Plates were incubated for 3 hours at 37 °C, protected from light. 100 µL from each well, in triplicate, were transferred to black 96-well plates, and fluorescence was measured at 560-590 nm using a Synergy H4 Microplate reader (BioTek). Non-treated cells were used as controls and considered as 100% cell viability. Cell viability was calculated according to the equation: % Cell viability = (Sample Fluorescence / Control Fluorescence) × 100. 2.9.5. Assessment of anti-inflammatory activity and efficacy by measurement of cytokine secretion After checking the absence of LPS contamination and non-toxic activity of the release buffer containing DEX and CLX at indicated time points, the anti-inflammatory activity of the drugs in the release buffer was compared with the free drugs at equivalent concentrations by ELISA. The anti-inflammatory activity was evaluated by quantifying the secretion of relevant inflammatory signals (TNF-α, CCL2, and PGE2) from primary human macrophages treated with stimulated with LPS + IFNγ alone, in combination with free drugs or exposed to the release buffer for 24 hours. ELISAs were performed using the commercial kit DuoSet® ELISA Development Systems (Bio-Techne) for TNF-α (DY210) and the Prostaglandin E2 ELISA kit (Cayman) for PGE2 (004CA514010-96), following the manufacturer’s instructions. ELISA assays for CCL2 were performed using a custom kit developed in the HUNIMED lab [36]. The amount of cytokine secretion (ng/ml) in macrophages treated with LPS + IFNγ (positive control) was normalized to 100% and, subsequently, values for the remaining samples were normalized and expressed as percentage (%) of control, according to the equation: % of Control = (ng/ml of Sample / ng/ml of Control) × 100. Results and discussion Our aim was to engineer a biodegradable bilayer polymer coating around a PCU meniscus prosthesis to release two anti-inflammatory drugs with controlled release kinetics. More precisely, dexamethasone (DEX) was intended to be released within 1-4 weeks to alleviate the acute inflammation caused mainly by the surgery. Meanwhile, celecoxib (CLX) was intended to be released within 6-9 months to relieve the chronic inflammation related to the prosthesis. Both drugs are expected to mediate the incorporation of the implant into the knee cavity, reducing potential adverse inflammatory reactions. Therefore, we developed and optimized the preparation of the bilayer coating, and we performed experiments to evaluate the activity of released drugs in vitro . For this, we followed a specific work plan consisting of: 1) The screening of different biodegradable and biocompatible polymers for the independent release of DEX and CLX with the desired release kinetics. 2) The development of drug-releasing bilayer polymer coatings results from the combination of the most promising prototypes selected for the independent release of the drugs. This development consisted of evaluating the capacity of the bilayer systems to release CLX and DEX in a required time frame, characterization of their physicochemical properties, and in vitro study of their degradation. 3) The in vitro evaluation of the sterility and biocompatibility of the drugs and polymer degradation products from the bilayer coated system, and the anti-inflammatory activity of these drugs in human primary macrophages. 3.1. Design and optimization of drug-loaded polymer films Polyesters were selected as the base polymers of the bilayer coating due to their controlled biodegradability, tunable release properties and favorable regulatory profile. The primary objective of this section was to evaluate the influence of the polymer type on the drug release profile while keeping other parameters—such as the preparation method and the number of polymer layers—constant. Although the ultimate goal was to achieve the controlled release of two different anti-inflammatory drugs from the same meniscus prosthesis with distinct release kinetics, the initial screening focused on evaluating the release of each drug independently within a single polymeric matrix. 3.1.1. CLX-loaded Polymer films Poly(caprolactone) (PCL) and poly(L-lactide) (PLLA) were chosen due to their slow degradation rate and, hence, long-term release capacity of CLX (6 to 9 months). This profile was expected to modulate a process of chronic inflammation [37,38]. The influence of molecular weight (MW) on degradation rate was explored by testing high molecular weight (HMW) (~210 kDa HMW-PLLA and ~73 kDa HMW-PCL) and low molecular weight (LMW) (~74 kDa LMW-PLLA and ~32 kDa LMW-PCL) variants of each polymer [39–43]. CLX-loaded films were fabricated by dissolving CLX and each polymer in dichloromethane (DCM) and casting the solution onto square-shaped PCU implants. After curing, the films and PCU implants were immersed in PBS with 1% (w/v) Tween®80, and CLX release was monitored via UPLC (Figure 1). The drug release study was halted when a plateau release phase was observed. The release studies revealed that CLX was consistently released faster from PCL films, irrespective of MW, as compared to from PLLA films. PLLA exhibited MW-dependent release kinetics, with HMW-PLLA providing more sustained release compared to LMW-PLLA [41]. The differences in release kinetics can be explained by the polymer chain length. Since none of these polymers is expected to degrade within a few weeks, the observed release behavior aligns with previous studies linking polymer chain length to drug diffusion [44]. The denser matrix created by HMW-PLLA restricted water penetration, slowing degradation and drug diffusion and, therefore, delaying drug release. As expected, due to the larger difference in MW, these effects are more pronounced in PLLA films (~140 kDa) than in PCL films (~41 kDa ). The plateau in release profiles is hypothesized to be a result of the interplay between drug diffusion and polymer degradation. Initially, the release is supposed to be diffusion-driven through water-filled pores, whereas the second phase corresponds to the release of the drug remaining trapped within the polymer matrix. Since polymer degradation is slow, this trapped drug remains unavailable, leading to incomplete release within the study timeframe. Thus, the release kinetics reflect both initial diffusion and delayed drug release associated with polymer degradation [45]. Although the release profile of HMW-PLLA might be considered aligned with the targeted release timeframe (6–9 months), the plateau observed after 20 weeks led to an incomplete drug release (30% of the total CLX released (Figure 1)). This incomplete release was attributed to limited polymer degradation, which could lead to prototype limitations such as drug recrystallization [45]. Although recrystallization was not observed or specifically analyzed in this study, prolonged entrapment of drug molecules within the polymer matrix under aqueous conditions is generally considered a risk factor for potential recrystallization in drug delivery systems [46]. To achieve intermediate CLX release kinetics, PLLA and PCL were blended at varying ratios (Figure 2). Increasing PCL content accelerated CLX release, a trend attributed to the formation of polymer-to-polymer interfaces that facilitated water penetration and diffusion of the drug near these interfaces [47–50]. Due to the limited miscibility of PLLA and PCL, PCL domains aggregate, lowering Tg locally and increasing chain mobility. This facilitates nucleation and improves PLLA’s mechanical properties, including flexibility and toughness [51]. Among tested formulations, both PLLA/PCL blends at 70/80 (w/w) and 80/70 (w/w) showed promising results with intermediate release profiles among the blends studied, releasing 40% of CLX after 3 months without plateauing. Given the variability in intra-articular CLX dosing across animal models (1-1.25 mg CLX/kg in mice [52], 0.46 mg CLX/kg/week over 5 weeks in rabbits [53], 0.0292 mg CLX/kg [54] or 1 mg/kg [55] in rats, and approximately 1 mg/kg in sheep [56]), it was assumed that higher doses of CLX could enhance therapeutic effects in both in vitro and in vivo models (e.g., reduced fibrous encapsulation, foreign body reactions, and modulation of signaling pathways). Hence, CLX concentration in selected formulations was increased 2.4-fold (from 12.5 mg/mL to 30 mg/mL), achieving a drug loading (DL) of 16.6%. This increase aimed to enhance the total amount of drug incorporated into the coating, bringing it closer to the intra-articular doses previously reported in the ovine model. The PLLA/PCL blend at an 80/70 (w/w) ratio was ultimately selected for its balanced release profile (Figure 2). These results led to two emerging candidates for sustained CLX release: HMW-PLLA with 20% DL, which provided a very slow release (<25% CLX released after 6 months), and PLLA/PCL 80/70 (w/w) with 16.67% DL, which achieved ~40% release in 3 months. Both formulations exhibited polymer-drug interactions based on thermal analysis (reductions in Tg, Tm, and Tcc) (Supplementary Figure 4A) and successfully incorporated CLX in a molecularly dispersed form, as confirmed by the absence of CLX crystalline peaks in XRD and the disappearance of its melting endotherm peak in DSC thermograms (Supplementary Figures 4A–B), both indicative of an amorphous drug state within the polymer matrix. However, only the PLLA/PCL 80/70 (w/w) blend met the required release profile and progressed to further studies. 3.1.2. DEX-loaded Polymer films To effectively address post-surgical acute inflammation, DEX was selected for its potent and pleiotropic anti-inflammatory activity and established clinical use. The goal was to achieve a rapid and controlled release over a 1 to 4-week period by incorporating DEX into an external polymer layer made of PLGA. A high molecular weight (MW) (24–38 kDa) was initially explored, but presented limitations due to its characteristic biphasic release profile (Supplementary Figure 3) —marked by an initial burst from surface-associated drug, followed by a less predictable sustained release phase driven by bulk polymer hydrolysis [57,58]. Its degradation rate is influenced by the lactic acid (LA) to glycolic acid (GA) ratio, with a 50:50 composition degrading the fastest due to enhanced hydrophilicity and reduced crystallinity [59]. Given the limitations of PLGA, particularly its biphasic release profile, an alternative amphiphilic matrix—poly(lactic acid)-poly(ethylene glycol) di-block copolymer (PLA-PEG)—was evaluated. Di-block copolymers like PLA-PEG can influence polymer-drug interactions differently than blends of individual polymers [60,61] due to the covalent linkage between the blocks, which creates a more defined and stable microphase separation compared to physical blends [62]. This structural organization can alter drug distribution and water accessibility, ultimately affecting drug release kinetics and solubility. For these reasons, di-block copolymers have been widely used in drug delivery systems [63]. As shown in Figure 3, PLA-PEG enabled a sustained and more complete release of DEX, for approximately one week at a polymer concentration of 200 mg/mL. This accelerated release was likely due to PEG-induced aqueous channel formation within the matrix. However, the slow degradation rate of the PLA block could prolong the integrity of the top layer, acting as a persistent barrier over the bottom CLX-releasing layer and thereby limiting CLX diffusion to the release medium over time. To improve control over the release profile and mitigate the limitations associated with PLGA’s biphasic behavior, the effect of PLGA molecular weight (MW) on DEX release was investigated. In addition to the high-MW PLGA used initially, low-MW (LMW) variants (8 kDa) were tested both as standalone matrices and in blends with PLA-PEG (Figure 3). Blends of PLA-PEG and LMW-PLGA 75/25 (w/w) showed release profiles similar to pure PLA-PEG, but a 50/50 (w/w) blend extended the sustained DEX release to two weeks (Figure 3). Increasing the relative amount of LMW-PLGA further enhanced this effect. Optimization studies with different PLA-PEG/LMW-PLGA ratios confirmed that higher LMW-PLGA content produced a sustained release within the desired 1–4-week window. Notably, pure LMW-PLGA exhibited a similar release profile, making it a practical alternative to PLA-PEG. Its selection not only met the therapeutic release target but also simplified the formulation by avoiding polymer blends. DSC and XRD analyses were performed to confirm the physical state of the incorporated drugs. For both DEX and CLX formulations, the disappearance of their characteristic melting peaks in DSC thermograms and the absence of crystalline reflections in XRD patterns confirmed that the drugs were molecularly dispersed within the polymer matrices (Supplementary Figures 4A and B, and 5A and B). For CLX-loaded systems, thermal events and crystallinity indicated good miscibility and a predominantly amorphous dispersion of the drug. Slight reductions in Tg and Tm—particularly in PLLA-based formulations—suggested plasticizing effects and increased polymer chain mobility [64–67]. These interactions are consistent with the initial diffusion-driven release observed in PLLA and PLLA/PCL blends. Notably, no direct correlation was found between polymer crystallinity and release kinetics, supporting the idea that other factors—such as porosity and polymer-drug interactions—played a more dominant role [68]. In the case of DEX, minor increases in crystallinity were observed in PEG-containing systems, suggesting weak drug-polymer interactions that may facilitate water uptake and diffusion. Conversely, PLGA and LMW-PLGA matrices remained largely amorphous, aligning with their more sustained release behavior. A full list of thermal parameters (Tg, Tm, Xc) is provided in Supplementary Tables 1 and 2. This comprehensive screening identified PLLA/PCL (80/70, w/w) with 16.6% DL as the most promising formulation for sustained CLX release, achieving higher drug incorporation while avoiding the plateau effect observed with pure PLLA. For DEX release, both PLA-PEG and LMW-PLGA (8 kDa, 50:50 LA:GA), each with a DL of 2.44%, achieved controlled release within the target 1–4 week window, with LMW-PLGA providing a slightly more sustained release profile and the advantage of a simpler formulation. 3.2. Engineering a bilayer drug-releasing polymer coating for the sequential release of DEX and CLX To enable the simultaneous release of CLX and DEX from the non-biodegradable PCU prosthesis (~0.7 × 0.7 × 0.3 cm), a bilayer polymeric coating was developed. The system consists of a first layer releasing CLX to be cast on top of the prosthesis, and a second layer releasing DEX, providing a shorter drug release. Polymer matrices previously identified as optimal candidates (Section 3.1.) were selected for testing in this configuration. While solvent casting was effective for coating one surface, dip coating was chosen for full-implant coverage due to its simplicity, versatility, and compatibility with multilayer deposition [18]. This section explores key aspects of bilayer development, including the impact of solvent interactions between layers and the selection of one optimized polymer matrix for each drug for final implementation. Given that the DEX layer is prepared using acetone, the potential effect of solvent exposure on the underlying CLX-loaded layer was evaluated. Since acetone can partially disrupt polymer-drug interactions, this investigation was crucial for ensuring stability and consistency in release kinetics. This was explored by DSC, XRD, and drug release evaluation— detailed in Supplementary Figures 6 and 7 . When acetone was applied to PLLA/PCL films, a slight acceleration in CLX release was noted, yet the overall release kinetics remained consistent. However, for pure PLLA, acetone significantly altered the release profile, accelerating CLX release from approximately 10% to over 60% by week 21 (Supplementary Figure 7). 3.2.1. Evaluation of drug release The release profiles of DEX from the two polymer matrices—PLA-PEG and LMW-PLGA— were evaluated while maintaining a constant CLX-releasing bottom layer composed of the PLLA/PCL blend. PCU implants were coated using a dip-coating approach, where CLX was first incorporated into the PLLA/PCL film using DCM, and DEX was then incorporated into either PLA-PEG or LMW-PLGA using acetone. The main distinction between these bilayer systems lies in the choice of the top-layer polymer, which influenced not only DEX release but also indirectly modulated CLX release kinetics. For clarity, these systems are referred to as Prototype PLA-PEG and Prototype PLGA (Figure 4A). Together, they provide valuable design insights for dual-drug delivery strategies targeting both acute and chronic phases of inflammation. Despite the identical CLX-releasing layer, the release profiles of CLX varied depending on the top-layer polymer (Figure 4B). When LMW-PLGA was used as the top layer, CLX release was faster, likely because the more rapid degradation of PLGA allowed earlier erosion or removal of the upper layer, facilitating earlier exposure of the CLX layer to the release medium. In contrast, the slower-degrading PLA-PEG formed a more persistent barrier, delaying access of the release medium to the CLX layer and thus slowing its diffusion. DEX release profiles were consistent with our prior results, where PLA-PEG exhibited an initial burst followed by a gradual tapering, while LMW-PLGA provided a more sustained and steady release. Both systems achieved complete DEX release within the first few weeks, effectively meeting the goal of addressing acute post-surgical inflammation. 3.2.2. In vitro degradation of PLA-PEG and PLGA bilayer coated implants Although both bilayer prototypes (PLA-PEG and PLGA) achieved the desired release kinetics for DEX and CLX, they exhibited distinct release profiles. Thus, using these prototypes, degradation studies were conducted to better understand the mechanisms behind their different release behaviors and to validate the hypothesis that PLA-PEG degrades more slowly than PLGA. Coated square implants were immersed in PBS supplemented with 1% (w/v) Tween®80, and periodically analyzed for trends in the change of pH and thickness, as well as surface morphology via FESEM. Over time, Prototype PLA-PEG showed a gradual decrease in thickness, with a pronounced reduction between 4 and 6 months, corresponding to clear structural degradation and coating disintegration (Figure 5A and Figure 6). By 12 months, the coating was visibly deteriorated, appearing uneven and incomplete, with clear signs of degradation and areas where the prosthesis surface was no longer covered. In contrast, Prototype PLGA initially exhibited a thickness increase—likely due to polymer swelling—followed by a steady decrease over the year (Figure 5A and Figure 7) [69,70]. The pH monitoring unveiled additional aspects of the degradation behavior not apparent from thickness measurements alone. While Prototype PLA-PEG maintained relatively stable pH values for the first four months, a noticeable decline occurred between months 4 and 6, consistent with the delayed onset of degradation observed in FESEM analysis and expected of PLA. In contrast, Prototype PLGA triggered a sharp pH drop by the second month, which then remained low for the rest of the study (Figure 5B). These trends reflect the faster breakdown of the PLGA matrix and further support the interpretation that the top-layer degradation rate directly influenced buffer access to the CLX-releasing layer. These findings help explain the differences observed in CLX release profiles. Although both prototypes had the same PLLA/PCL internal layer for CLX release, faster degradation and reduced tortuosity of the PLGA top layer likely allowed earlier buffer access to the underlying CLX matrix, accelerating drug release. In contrast, the slower degradation of PLA-PEG maintained a more effective diffusion barrier for longer. Thickness measurements were affected by sample cutting and imaging variability because of the angle of the image and should be interpreted as relative trends rather than absolute values (Supplementary Figure 8). In contrast, pH changes provided a more robust and reproducible indicator of degradation, though results may not fully reflect in vivo behavior, where pH buffering and enzymatic processes differ. Nonetheless, prior work suggests that PLGA degradation kinetics in vitro correlate with in vivo molecular weight loss [134]. 3.2.3. Evaluation of drug-releasing polymer coating reproducibility The reproducibility of the prototypes was assessed by quantifying the total amount of DEX and CLX per unit of area (µg/cm²) across multiple replicates of square-shaped PCU implants. The Prototype PLGA showed consistent DL for both DEX and CLX, indicating a robust and reproducible coating process with controlled dipping cycles and solvent evaporation. In contrast, Prototype PLA-PEG showed slight variability in drug incorporation between replicates, suggesting challenges in achieving uniform deposition. This may be attributed to uneven polymer film formation, variations in drug-polymer interactions, or greater sensitivity to coating conditions on the complex implant surface. These results indicate that Prototype PLGA offers greater reliability and process reproducibility , while Prototype PLA-PEG may require further optimization to ensure consistent DL (Supplementary Figure 9) . 3.3. In vitro assessment of sterility, biocompatibility, anti-inflammatory activity, and efficiency of the bilayer drug-releasing polymer-coated PCU implants Prior to evaluating the immunomodulatory potential of the bilayer-coated implants, a series of in vitro validations were conducted to ensure compatibility with biological assays. Drug release studies confirmed that both DEX and CLX were released in concentrations exceeding their respective effective thresholds (10 µM for DEX, 25 µM for CLX, based on literature [27]) at key time points—3 hours, 3 days, 1 week, 2 weeks, and 4 weeks (Table 1). These intervals were strategically selected to capture distinct drug activity phases: early DEX-driven effects, overlapping DEX and CLX activity, and later CLX-dominant responses. The release buffer also met sterility criteria, with endotoxin levels remaining below 0.125 EU/mL (Supplementary Figure 10), validating their suitability for macrophage-based assays. Moreover, no cytotoxic effects were observed in HMDMs exposed to release media from either prototype at any time point, indicating good biocompatibility of both the drugs released and the polymer degradation byproducts (Supplementary Figure 11). These findings are consistent with previous reports describing the non-toxic nature of polyester degradation products and further support the safety of the materials used [71]. With drug release, sterility, and cytocompatibility confirmed, the subsequent experiments focused on assessing the bioactivity of the released drugs by evaluating their capacity to modulate cytokine secretion in M1-like (classically activated with LPS + IFNg) pro-inflammatory macrophages. 3.3.1. In vitro evaluation of immunomodulatory activity The evaluation of cytokine secretion aimed to confirm that the drugs released from the bilayer polymer coatings retained their immunomodulatory activity. Specifically, it was sought to verify that DEX and CLX, when released at their respective time points, could suppress pro-inflammatory cytokine release from macrophages, events associated with acute and chronic inflammation. This assessment was essential to validate the therapeutic functionality of the coating system in an inflammatory environment. Cytokine secretion by M1-like macrophages was measured following exposure to release media collected from the implant incubations at defined time points. The three key pro-inflammatory cytokines selected were TNF-α (a central mediator of acute inflammation [72–74]), CCL2 (MCP-1) (regulates monocyte/macrophage recruitment and FBGC formation [75,76]), and PGE2 (a COX-2 pathway product involved in joint inflammation, vasodilation, and pain [77,78]). Both prototypes significantly suppressed TNF-α and CCL2 secretion across all time points (Figure 8A and B), demonstrating that the combination of DEX and CLX maintained its anti-inflammatory efficacy following their release from the coatings. Notably, TNF-α and CCL2 suppression was more closely linked to DEX activity, as expected during early time points. PGE2 secretion, in contrast, was more selectively reduced by CLX, consistent with its inhibition of COX-2 (Figure 8C) [79]. This effect was observed at all time points, regardless of the prototype, supporting the stability and retained bioactivity of CLX even after extended residence within the coating. Importantly, cytokine suppression levels were comparable to those obtained with free (non-encapsulated) drugs, validating the effectiveness of the delivery platform. These results confirm that both Prototype PLGA and Prototype PLA-PEG effectively preserved the bioactivity of the drugs during encapsulation, storage, and release. Conclusion Here we present a bilayer polymer coating developed to enable the sequential release of dexamethasone (DEX) and celecoxib (CLX), from a PCU meniscus prosthesis, with the final goal of mitigating both, acute and chronic inflammation observed after implantation. Through a systematic screening of polymer candidates, two bilayer prototypes—Prototype PLA-PEG and Prototype PLGA—were identified based on their ability to achieve the desired release kinetics: rapid DEX release over 1–4 weeks to prevent acute inflammation, and prolonged CLX release over 6–9 months for sustained immunomodulation. Polymer degradation studies revealed that Prototype PLGA, composed of a low molecular weight PLGA external layer, exhibited faster degradation, facilitating earlier exposure of the CLX-loaded PLLA/PCL bottom layer to the release medium. In contrast, Prototype PLA-PEG was degraded more slowly, increasing tortuosity and delaying CLX release. Although PLA-PEG allowed for higher DL, PLGA demonstrated more consistent and reproducible release profiles and better integration with the bilayer system. Both prototypes were endotoxin-free and biocompatible, and the released drugs maintained their anti-inflammatory activity, as demonstrated by the effective decrese in the secretion of pro-inflammatory cytokines by M1-like activated macrophages. Based on its favorable degradation kinetics, reproducibility, and compatibility with the bilayer architecture, Prototype PLGA was selected for further development. This next phase would involve optimizing the coating system and scaling up its application to actual meniscus prostheses for evaluation in a representative animal model, bringing the technology closer to clinical translation. Declarations Ethical statement Ethics approval and consent to participate: Studies using human-derived monocytes were conducted at Humanitas Research Institute (Milano, Italy) in accordance with Italian and European law. Monocytes were isolated from the buffy coats of blood donations. Buffy coats were obtained from anonymous healthy blood donors at Istituto Clinico Humanitas (ICH) (Milano, Italy). The ethical committee at ICH provided favorable confirmation for their use in basic scientific research activities. Their use does not involve any ethical problem regarding informed consent or the privacy of the donor. Consent for publication: Not applicable. Availability of data and materials: The datasets generated during and/or analysed during the current study are available from the corresponding author on reasonable request. Competing interests: María José Alonso is the founder and shareholder of LiberaBio. The remaining authors declare no conflict of interest. Funding: This project has received funding from the European Union's Horizon 2020 research and innovation programme under grant agreement No 814444 (MEFISTO). Alfonso F. Blanco has received research support through a predoctoral grant from Xunta de Galicia, Grant number ED481A 2022/066. Fernando Torres-Andón was also supported by a “Miguel Servet Grant” CP22/00106 by AES 2022, and Alba Pensado-López by a “Sara Borrell Grant” CD23/00173 by AES 2023, ISCIII, Spain. Authors' contributions: All authors contributed to the study conception and design. Material preparation, data collection, and analysis related to drug release studies, polymer selection, polymer coating evaluation, and optimization were performed by Alfonso F. Blanco and Gustavo Lou. In vitro study conceptualization was performed by Alfonso F. Blanco, Aldo Ummarino, Alba Pensado-López, and Fernando Torres-Andón. In vitro study preparation and data collection were performed by Aldo Ummarino and Alba Pensado-López. In vitro study analysis was done by Alfonso F. Blanco. The first draft of the manuscript was written by Alfonso F. Blanco, and all authors commented on previous versions of the manuscript. All authors read and approved the final manuscript. Acknowledgements: Figures were created with BioRender.com. Authors' information: Not applicable. Data Availability Statement: The datasets generated during and/or analysed during the current study are available from the corresponding author on reasonable request. References Horecka A, Hordyjewska A, Blicharski T, Kurzepa J. Osteoarthritis of the knee - biochemical aspect of applied therapies: a review. Bosn J Basic Med Sci 2022;22:488–98. https://doi.org/10.17305/BJBMS.2021.6489. Cui A, Li H, Wang D, Zhong J, Chen Y, Lu H. Global, regional prevalence, incidence and risk factors of knee osteoarthritis in population-based studies. EClinicalMedicine 2020;29–30:100587. https://doi.org/10.1016/j.eclinm.2020.100587. Heidari B. 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Supplementary Files Supplementary.docx floatimage1.jpeg Graphical abstract Cite Share Download PDF Status: Published Journal Publication published 25 Aug, 2025 Read the published version in Drug Delivery and Translational Research → Version 1 posted Editorial decision: Minor Revisions Needed 14 Jul, 2025 Reviewers agreed at journal 22 Jun, 2025 Reviewers invited by journal 16 Jun, 2025 Editor assigned by journal 13 Jun, 2025 First submitted to journal 12 Jun, 2025 You are reading this latest preprint version Research Square lets you share your work early, gain feedback from the community, and start making changes to your manuscript prior to peer review in a journal. As a division of Research Square Company, we’re committed to making research communication faster, fairer, and more useful. We do this by developing innovative software and high quality services for the global research community. Our growing team is made up of researchers and industry professionals working together to solve the most critical problems facing scientific publishing. Also discoverable on Platform About Our Team In Review Editorial Policies Advisory Board Help Center Resources Author Services Accessibility API Access RSS feed Manage Cookie Preferences © Research Square 2026 | ISSN 2693-5015 (online) Privacy Policy Terms of Service Do Not Sell My Personal Information {"props":{"pageProps":{"initialData":{"identity":"rs-6869700","acceptedTermsAndConditions":true,"allowDirectSubmit":false,"archivedVersions":[],"articleType":"Research Article","associatedPublications":[],"authors":[{"id":472179694,"identity":"771a76df-5d0c-42f0-b8e9-1c5a887465bc","order_by":0,"name":"Alfonso F. Blanco","email":"","orcid":"","institution":"University of Santiago de Compostela: Universidade de Santiago de Compostela","correspondingAuthor":false,"prefix":"","firstName":"Alfonso","middleName":"F.","lastName":"Blanco","suffix":""},{"id":472179695,"identity":"e6de2c7e-54b4-4734-bde2-ad63821449f8","order_by":1,"name":"Gustavo Lou","email":"","orcid":"","institution":"University of Santiago de Compostela: Universidade de Santiago de Compostela","correspondingAuthor":false,"prefix":"","firstName":"Gustavo","middleName":"","lastName":"Lou","suffix":""},{"id":472179696,"identity":"c3a8b23a-1867-46ba-bcf7-0fac8d3accab","order_by":2,"name":"Alba Pensado-López","email":"","orcid":"","institution":"HUNIMED: Humanitas University","correspondingAuthor":false,"prefix":"","firstName":"Alba","middleName":"","lastName":"Pensado-López","suffix":""},{"id":472179697,"identity":"7d125730-607f-4ad8-a916-c2adf955768d","order_by":3,"name":"Aldo Ummarino","email":"","orcid":"","institution":"HUNIMED: Humanitas University","correspondingAuthor":false,"prefix":"","firstName":"Aldo","middleName":"","lastName":"Ummarino","suffix":""},{"id":472179698,"identity":"f53e43c0-dc91-4e56-860c-8492f17d78b9","order_by":4,"name":"Fernando Torres-Andón","email":"","orcid":"","institution":"University of Santiago de Compostela: Universidade de Santiago de Compostela","correspondingAuthor":false,"prefix":"","firstName":"Fernando","middleName":"","lastName":"Torres-Andón","suffix":""},{"id":472179699,"identity":"6b98aecc-127e-43ca-ba0d-d15c9e79436e","order_by":5,"name":"José Crecente-Campo","email":"","orcid":"","institution":"University of Santiago de Compostela: Universidade de Santiago de Compostela","correspondingAuthor":false,"prefix":"","firstName":"José","middleName":"","lastName":"Crecente-Campo","suffix":""},{"id":472179700,"identity":"84a1098f-0bc6-4ddf-a603-1e3ce673d7dd","order_by":6,"name":"Maria J Alonso","email":"data:image/png;base64,iVBORw0KGgoAAAANSUhEUgAAAZAAAAAyAQMAAABI0h/eAAAABlBMVEX///8AAABVwtN+AAAACXBIWXMAAA7EAAAOxAGVKw4bAAAA6UlEQVRIiWNgGAWjYJCCAww8DAx8YGYFKVrYwMwzpFgF1sLYRoRK3fazDw/8kGGQZ2PvPfzh47zD9vwNzIc/4NNidibd4GAPD4NhG8+5NMmZ2w4nzjjAliaBV8uBNIYDPDxAJ0nkmDHzbjucYMDAY4bXYWbnnzEc/MPDYA/UYvyZd85hewMG/s/4HXYjjeEw0JZEoBYDad6Gw4wbgOGH32E3njEcluGRSG7jOWMmOeNYeuKMw2xm+LWcT2P++LbHxrafvcf4w4caa3v+9ubHeB0GBow9yMYyE1QPAj+IUjUKRsEoGAUjFQAAf51EFYQS/xAAAAAASUVORK5CYII=","orcid":"https://orcid.org/0000-0001-7187-9567","institution":"Universidade de Santiago de Compostela","correspondingAuthor":true,"prefix":"","firstName":"Maria","middleName":"J","lastName":"Alonso","suffix":""}],"badges":[],"createdAt":"2025-06-11 08:51:58","currentVersionCode":1,"declarations":"","doi":"10.21203/rs.3.rs-6869700/v1","doiUrl":"https://doi.org/10.21203/rs.3.rs-6869700/v1","draftVersion":[],"editorialEvents":[{"content":"https://doi.org/10.1007/s13346-025-01942-5","type":"published","date":"2025-08-25T15:57:13+00:00"}],"editorialNote":"","failedWorkflow":false,"files":[{"id":85367768,"identity":"7a93a0ab-be51-437f-b0af-551299e44108","added_by":"auto","created_at":"2025-06-25 07:04:54","extension":"png","order_by":1,"title":"Figure 1","display":"","copyAsset":false,"role":"figure","size":88796,"visible":true,"origin":"","legend":"\u003cp\u003eRelease kinetics of CLX expressed as the total percentage of drug released (%) from polymer films composed of polymers PCL and PLLA, both HMW and LMW, at a concentration of 50 mg/mL, with CLX loadings of 20%.\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003eAbbreviations:\u003c/strong\u003eCLX: Celecoxib. LMW: Low molecular weight. HMW: High molecular weight. DL: Drug loading. PLLA: poly(L-lactide) (PLLA). PCL: poly(caprolactone). Values represent the mean ± standard deviation (n=3).\u003c/p\u003e","description":"","filename":"1.png","url":"https://assets-eu.researchsquare.com/files/rs-6869700/v1/4dee7032a9b4a2f87587c70b.png"},{"id":85367767,"identity":"06e45aea-c3c0-4159-a4c7-51ac262689b7","added_by":"auto","created_at":"2025-06-25 07:04:54","extension":"png","order_by":2,"title":"Figure 2","display":"","copyAsset":false,"role":"figure","size":90611,"visible":true,"origin":"","legend":"\u003cp\u003eRelease kinetics of CLX expressed as the total percentage of drug released (%) from PLLA/PCL blends prepared at 150 mg/mL at different (w/w) ratios (ranging from 50/100 to 90/60) with a CLX loading of 16.6%.\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003eAbbreviations:\u003c/strong\u003e CLX: Celecoxib. PLLA: poly(L-lactide). PCL: poly(caprolactone). (w/w): weight-to-weight ratio. Values represent the mean ± standard deviation (n=3).\u003c/p\u003e","description":"","filename":"2.png","url":"https://assets-eu.researchsquare.com/files/rs-6869700/v1/9f8e2c468747727caa12bced.png"},{"id":85367765,"identity":"bcdec340-5733-450c-bd37-226a8cfc7c79","added_by":"auto","created_at":"2025-06-25 07:04:54","extension":"png","order_by":3,"title":"Figure 3","display":"","copyAsset":false,"role":"figure","size":105937,"visible":true,"origin":"","legend":"\u003cp\u003eRelease kinetics of DEX expressed as the total percentage of drug released (%) from pure PLA-PEG, pure LMW-PLGA, and PLA-PEG/PLGA blend films. LMW-PLGA was incorporated to the PLA-PEG matrix with different PLA-PEG/PLGA ratios ranging from 25/75 to 75/25 (w/w). Films were prepared at 200 mg/mL with constant drug loadings of 2.44%.\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003eAbbreviations:\u003c/strong\u003e DEX: Dexamethasone. PLA-PEG: poly(lactic acid)-poly(ethylene glycol) di-block co-polymer. PLGA: Poly(lactic-co-glycolic) acid. LMW: Low molecular weight. Values represent the mean ± standard deviation (n=3).\u003c/p\u003e","description":"","filename":"3.png","url":"https://assets-eu.researchsquare.com/files/rs-6869700/v1/611cd79c9662de0cc3fdeb4b.png"},{"id":85367772,"identity":"35847508-c8bd-4bcd-b187-d5ad4d1d712c","added_by":"auto","created_at":"2025-06-25 07:04:54","extension":"png","order_by":4,"title":"Figure 4","display":"","copyAsset":false,"role":"figure","size":830387,"visible":true,"origin":"","legend":"\u003cp\u003eSchematic representation of the two final prototypes of bilayer polymer coatings (A). Sequential release of DEX and CLX from bilayer polymer coatings composed of a first polymer coating of PLLA/PCL, prepared at 150 mg/mL at 80/70 (w/w) with CLX loading of 16.6%; and a second polymer coating of either PLA-PEG or LMW-PLGA, both prepared at 200 mg/mL with DEX loading of 2.44% (B).\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003eAbbreviations:\u003c/strong\u003e PCU: Polycarbonate urethane. CLX: Celecoxib. DEX: Dexamethasone. PLA-PEG: poly(lactic acid)-poly(ethylene glycol) di-block co-polymer. PLGA: Poly(lactic-co-glycolic) acid. PLLA: poly(L-lactide). PCL: poly(caprolactone). LMW: Low molecular weight. Values represent the mean ± standard deviation (n=3).\u003c/p\u003e","description":"","filename":"4.png","url":"https://assets-eu.researchsquare.com/files/rs-6869700/v1/d906038502b3517fa8f6a3a2.png"},{"id":85369505,"identity":"3b86d0e5-61c7-4611-8d61-dfd42a9a83a9","added_by":"auto","created_at":"2025-06-25 07:20:54","extension":"png","order_by":5,"title":"Figure 5","display":"","copyAsset":false,"role":"figure","size":263078,"visible":true,"origin":"","legend":"\u003cp\u003eChanges in thickness (A) and pH (B) of Prototype PLA-PEG and Prototype PLGA with time.\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003eAbbreviations:\u003c/strong\u003e PLA-PEG: poly(lactic acid)-poly(ethylene glycol) di-block co-polymer. PLLA: poly(L-lactide). PCL: poly(caprolactone). PLGA: Poly(lactic-co-glycolic) acid. µm: micrometers. Values represent the mean ± standard deviation (n≥1).\u003c/p\u003e","description":"","filename":"5.png","url":"https://assets-eu.researchsquare.com/files/rs-6869700/v1/6f460c26ae8a7932f4f41e93.png"},{"id":85367773,"identity":"d1a17ab7-1dd8-4b87-a9ba-fa102ac4c451","added_by":"auto","created_at":"2025-06-25 07:04:54","extension":"png","order_by":6,"title":"Figure 6","display":"","copyAsset":false,"role":"figure","size":1974608,"visible":true,"origin":"","legend":"\u003cp\u003eFESEM images of the Prototype PLA-PEG at different times of the biodegradation process. Images are divided into pairs per time point, a top image and a bottom image. The top images correspond with the sagittal cut of the bilayer polymer coating and the PCU implant and were obtained using Zeiss EVO analytical FESEM with a magnification of 100X. The thickness of the polymer coating was measured in these top images. The bottom images correspond with the view from above of the bilayer polymer coating sputter coated with iridium and were obtained using Zeiss UltraPlus analytical FESEM with a magnification of 1000X.\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003eAbbreviations:\u003c/strong\u003e FESEM: Field emission scanning electron microscopy. PLA-PEG: poly(lactic acid)-poly(ethylene glycol) di-block co-polymer. µm: micrometers. EHT: Electron high tension. WD: Working distance. Mag: Magnification.\u003c/p\u003e","description":"","filename":"6.png","url":"https://assets-eu.researchsquare.com/files/rs-6869700/v1/e0fd1376cbbbb552f8a340bd.png"},{"id":85367776,"identity":"4a7b0049-f79c-4aa8-a57e-9f420c8694d8","added_by":"auto","created_at":"2025-06-25 07:04:54","extension":"png","order_by":7,"title":"Figure 7","display":"","copyAsset":false,"role":"figure","size":1955616,"visible":true,"origin":"","legend":"\u003cp\u003eFESEM images of the Prototype PLA-PEG at different times of the biodegradation process. Images are divided into pairs per time point, a top image and a bottom image. The top images correspond with the sagittal cut of the bilayer polymer coating and the PCU implant and were obtained using Zeiss EVO analytical FESEM with a magnification of 100X. The thickness of the polymer coating was measured in these top images. The bottom images correspond with the view from above of the bilayer polymer coating sputter coated with iridium and were obtained using Zeiss UltraPlus analytical FESEM with a magnification of 1000X.\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003eAbbreviations:\u003c/strong\u003e FESEM: Field emission scanning electron microscopy. PLGA: Poly(lactic-co-glycolic) acid. µm: micrometers. EHT: Electron high tension. WD: Working distance. Mag: Magnification.\u003c/p\u003e","description":"","filename":"7.png","url":"https://assets-eu.researchsquare.com/files/rs-6869700/v1/942d233a446e3820c54b4f5c.png"},{"id":85368875,"identity":"4c491c9d-f87a-4654-b2b6-b7a9a7247603","added_by":"auto","created_at":"2025-06-25 07:12:54","extension":"png","order_by":8,"title":"Figure 8","display":"","copyAsset":false,"role":"figure","size":1237591,"visible":true,"origin":"","legend":"\u003cp\u003eTNFα (A), CCL2 (B), and PGE2 (C) secretion by human primary macrophages exposed to media where Prototype PLA-PEG and Prototype PLGA, both loaded with CLX and DEX, were incubated for specific periods of time. Analysis was carried out by ELISA.\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003eAbbreviations:\u003c/strong\u003e M0: Unactivated macrophages. Neg: Negative. Pos: Positive. Buf: Buffer. PLA-PEG: poly(lactic acid)-poly(ethylene glycol) di-block co-polymer. h: hour. d: days. w: weeks. PLGA: Poly(lactic-co-glycolic) acid. LPS: Lipopolysaccharide. IFN: Interferon. TNF: Tumor necrosis factor. CCL2: C-C motif chemokine ligand 2 or Monocyte chemoattractant protein-1 (MCP-1). PGE2: Prostaglandin E2. A significant comparison was performed using an ordinary one-way ANOVA followed by Tukey’s multiple comparison tests between LPS+IFNy and the rest of the groups. p-values \u0026lt; 0.05 were considered statistically significant (*). Also, (**) if \u003cem\u003ep\u003c/em\u003e-value \u0026lt; 0.01, (***) if \u003cem\u003ep\u003c/em\u003e-value \u0026lt; 0.001, (****) if \u003cem\u003ep\u003c/em\u003e-value \u0026lt; 0.0001. ns: not significative. Columns represent the mean ± standard deviation (n ≥ 5).\u003c/p\u003e","description":"","filename":"8.png","url":"https://assets-eu.researchsquare.com/files/rs-6869700/v1/b53d96d1351724c4233b8e1b.png"},{"id":90344851,"identity":"4e9a8bad-7ebc-4de0-a324-84a342b59fb9","added_by":"auto","created_at":"2025-09-01 16:06:16","extension":"pdf","order_by":0,"title":"","display":"","copyAsset":false,"role":"manuscript-pdf","size":8987399,"visible":true,"origin":"","legend":"","description":"","filename":"manuscript.pdf","url":"https://assets-eu.researchsquare.com/files/rs-6869700/v1/971dd82b-5932-4e1f-a55a-5acace2b290a.pdf"},{"id":85367771,"identity":"f7dedd3b-bfb7-43f5-8b8c-6d0c3b816d24","added_by":"auto","created_at":"2025-06-25 07:04:54","extension":"docx","order_by":1,"title":"","display":"","copyAsset":false,"role":"supplement","size":2211175,"visible":true,"origin":"","legend":"","description":"","filename":"Supplementary.docx","url":"https://assets-eu.researchsquare.com/files/rs-6869700/v1/f8857c51be6379943bfbf259.docx"},{"id":85370793,"identity":"8f6d71d2-e61c-430e-89c2-073f09e55bb2","added_by":"auto","created_at":"2025-06-25 07:28:54","extension":"jpeg","order_by":2,"title":"","display":"","copyAsset":false,"role":"supplement","size":228673,"visible":true,"origin":"","legend":"\u003cp\u003e\u003cstrong\u003eGraphical abstract\u003c/strong\u003e\u003c/p\u003e","description":"","filename":"floatimage1.jpeg","url":"https://assets-eu.researchsquare.com/files/rs-6869700/v1/304d4df4b9fe1664e66c9af9.jpeg"}],"financialInterests":"","formattedTitle":"Controlled Co-delivery of Anti-inflammatory Drugs from Bilayer Polymer Films Coating a Meniscus Implant","fulltext":[{"header":"Introduction","content":"\u003cp\u003eOsteoarthritis (OA) is a multifactorial, degenerative, and chronic joint disease characterized by progressive cartilage degradation, bone remodeling, and an imbalance between synthesis and degradation processes, leading to pain, inflammation, stiffness, and loss of function [\u003cspan citationid=\"CR1\" class=\"CitationRef\"\u003e1\u003c/span\u003e]. It is the most common form of arthritis and a major cause of disability, affecting over 650\u0026nbsp;million individuals worldwide in 2020 [\u003cspan citationid=\"CR2\" class=\"CitationRef\"\u003e2\u003c/span\u003e]. Knee OA is particularly prevalent, with its increasing incidence attributed to longer life expectancy and rising obesity rates, the latter being a significant risk factor [\u003cspan citationid=\"CR3\" class=\"CitationRef\"\u003e3\u003c/span\u003e].\u003c/p\u003e \u003cp\u003eThe meniscus plays a crucial role in knee function by providing stabilization, load distribution, and shock absorption [\u003cspan additionalcitationids=\"CR5\" citationid=\"CR4\" class=\"CitationRef\"\u003e4\u003c/span\u003e\u0026ndash;\u003cspan citationid=\"CR6\" class=\"CitationRef\"\u003e6\u003c/span\u003e]. However, meniscal injuries are common and often fail to heal due to the avascular nature of the inner meniscus, necessitating surgical intervention [\u003cspan citationid=\"CR7\" class=\"CitationRef\"\u003e7\u003c/span\u003e, \u003cspan citationid=\"CR8\" class=\"CitationRef\"\u003e8\u003c/span\u003e]. While partial meniscectomy is frequently performed to alleviate pain, it tends to accelerate joint degeneration and significantly increases the risk of OA development [\u003cspan citationid=\"CR9\" class=\"CitationRef\"\u003e9\u003c/span\u003e]. The use of meniscus implants or scaffolds that promote regeneration is a promising alternative to avoid the onset of OA after a meniscectomy [\u003cspan citationid=\"CR10\" class=\"CitationRef\"\u003e10\u003c/span\u003e]. To address this, meniscus implants and scaffolds have been developed to restore joint biomechanics and delay disease progression. Among them, the non-biodegradable NUsurface\u0026reg; prosthesis (Active Implants, Israel) serves as an artificial meniscus substitute after total meniscectomy. Approved in several European countries and Israel, it is currently undergoing FDA clinical trials [\u003cspan citationid=\"CR11\" class=\"CitationRef\"\u003e11\u003c/span\u003e, \u003cspan citationid=\"CR12\" class=\"CitationRef\"\u003e12\u003c/span\u003e]. Despite its advantages, implantation of the NUsurface\u0026reg; prosthesis could induce a foreign body reaction (FBR), requiring anti-inflammatory treatment to mitigate the patient\u0026rsquo;s pain, post-surgical inflammation and improve implant integration [\u003cspan citationid=\"CR13\" class=\"CitationRef\"\u003e13\u003c/span\u003e, \u003cspan citationid=\"CR14\" class=\"CitationRef\"\u003e14\u003c/span\u003e]. Current anti-inflammatory drugs have important limitations; for example systemic administration of COX-2 inhibitors, such as celecoxib (CLX), can cause gastrointestinal, renal, and cardiovascular side effects [\u003cspan citationid=\"CR15\" class=\"CitationRef\"\u003e15\u003c/span\u003e, \u003cspan citationid=\"CR16\" class=\"CitationRef\"\u003e16\u003c/span\u003e], whereas intra-articular (IA) corticosteroid injections, such as dexamethasone (DEX), provide short-term relief but pose risks of infection and patient discomfort [\u003cspan citationid=\"CR17\" class=\"CitationRef\"\u003e17\u003c/span\u003e]. To overcome these challenges, localized and sustained drug delivery strategies are needed to enhance therapeutic efficacy while minimizing systemic and local adverse effects [\u003cspan citationid=\"CR18\" class=\"CitationRef\"\u003e18\u003c/span\u003e].\u003c/p\u003e \u003cp\u003ePolymer coatings have been widely explored to improve implant biocompatibility, facilitate tissue integration, and enable controlled drug release. An optimized coating can provide site-specific drug release, shielding drugs from enzymatic degradation while reducing systemic toxicity [\u003cspan citationid=\"CR19\" class=\"CitationRef\"\u003e19\u003c/span\u003e, \u003cspan citationid=\"CR20\" class=\"CitationRef\"\u003e20\u003c/span\u003e]. CLX, a potential disease-modifying osteoarthritis drug (DMOAD), has shown protective effects in OA models by reducing inflammation and cartilage degradation [\u003cspan additionalcitationids=\"CR22 CR23 CR24\" citationid=\"CR21\" class=\"CitationRef\"\u003e21\u003c/span\u003e\u0026ndash;\u003cspan citationid=\"CR25\" class=\"CitationRef\"\u003e25\u003c/span\u003e], whereas a low-dose IA administration of DEX has shown chondroprotective potential when delivered IA, and its potential consideration as a DMOAD has been suggested. Therefore, the hypothesis of this study has been that a polymer bylayer coating enabling the sustained release of two drugs widely used in OA treatment, CLX for long-term inflammation control and short-term release of DEX for early-phase inflammation management, may improve implant performance and clinical outcomes [\u003cspan citationid=\"CR26\" class=\"CitationRef\"\u003e26\u003c/span\u003e, \u003cspan citationid=\"CR27\" class=\"CitationRef\"\u003e27\u003c/span\u003e]. Within this context, the current challenge in the field of meniscus-related degenerative pathologies is to design \u003cem\u003ein situ\u003c/em\u003e therapies enabling controlled drug release with the goal of achieving a prolonged response and a reduction of systemic side effects [\u003cspan citationid=\"CR28\" class=\"CitationRef\"\u003e28\u003c/span\u003e].\u003c/p\u003e \u003cp\u003eIn this study, we developed a bilayer polymer coating for the NUsurface\u0026reg; prosthesis intended to release these two anti-inflammatory drugs with distinct, controlled release kinetics. Using square-shaped polycarbonate urethane (PCU) implants (0.7 \u0026times; 0.7 \u0026times; 0.3 cm), we screened various polymer formulations to optimize controlled DEX release for short-term inflammation control (1\u0026ndash;4 weeks) and CLX release for long-term pain and inflammation management (6\u0026ndash;9 months). The coatings were characterized for their drug release kinetics, degradation profiles and biocompatibility. To assess their anti-inflammatory efficacy, we conducted \u003cem\u003ein vitro\u003c/em\u003e studies in primary human monocyte-derived macrophages (HMDMs). Our findings provide a foundation for the development of an implantable, bioactive, drug-releasing meniscus prosthesis designed to enhance post-surgical outcomes and reduce OA progression.\u003c/p\u003e"},{"header":"Materials and Methods","content":"\u003cdiv id=\"Sec3\" class=\"Section2\"\u003e \u003ch2\u003e2.1. Materials\u003c/h2\u003e \u003cp\u003eDexamethasone (DEX) and celecoxib (CLX) were supplied by Acofarma (Madrid, Spain) and Sigma-Aldrich (Missouri, USA), respectively. Poly (lactic-co-glycolic acid) (PLGA), of two different lactic acid:glycolic acid (LA:GA) ratios (PLGA 50:50, and PLGA 85:15), poly(caprolactone) (PCL) of low (LMW-PCL) and high molecular weight (HMW-PCL), poly(L-lactide) (PLLA) of low (LMW-PLLA) and high molecular weight (HMW-PLLA), and poly(lactic acid)-poly(ethylene glycol) di-block copolymer (PLA-PEG) were purchased from Evonik Industries (Darmstadt, Germany). LMW-PLGA LA:GA ratio 50:50 was obtained from PolySciTech, a division of Akina (Indiana, USA). Polyethylene glycol (PEG) of low (PEG 400) and high molecular weight (PEG 1450) were purchased from BASF (Ludwigshafen am Rhein, Germany). Sodium azide (NaN\u003csub\u003e3\u003c/sub\u003e) was purchased from Sigma-Aldrich (Missouri, USA). Acetone was obtained from Fisher chemicals (New Hampshire, USA). Dichloromethane (DCM and acetonitrile (ACN) were distributed by Scharlau (Barcelona, Spain). Phosphate saline buffer (PBS), Tween\u0026reg;80, trifluoracetic acid (TFA) and methanol (MeOH) were supplied by Scientific (Nottingham, England), Merck (Darmstadt, Germany), Sigma-Aldrich (Missouri, USA) and VWR Chemicals (Pennsylvania, USA), respectively. The square-shaped implants made of polycarbonate urethane (PCU) were kindly donated by Active Implants (Israel).\u003c/p\u003e \u003c/div\u003e \u003cdiv id=\"Sec4\" class=\"Section2\"\u003e \u003ch2\u003e2.2. Drug solubility in release buffer\u003c/h2\u003e \u003cp\u003eAn excess of the drugs (DEX and CLX) was incubated in agitation (700 rpm) in 1 mL of PBS supplemented with Tween\u0026reg;80 concentrations ranging from 0 to 1% (w/v) for 24 hours. Then, samples were centrifuged at 10,000 rpm for 20 minutes, the supernatant was diluted in MeOH:H\u003csub\u003e2\u003c/sub\u003eO 65:35 (v/v), and the solutions were quantified by ultraperformance liquid chromatography (UPLC) with a TUV detector at 239 nm with a column Kinetex\u0026reg; 1.7 \u0026micro;m C18 100 \u0026Aring;, LC Column 50 x 2.1 mm acquired from Phenomenex (Torrance, CA, USA), maintaining the samples at 20\u0026deg;C in a Waters Acquity H-Class UPLC system (Waters, Milford, USA).\u003c/p\u003e \u003c/div\u003e \u003cdiv id=\"Sec5\" class=\"Section2\"\u003e \u003ch2\u003e2.3. Production of drug-releasing polymer films\u003c/h2\u003e \u003cp\u003eTo prepare the polymer-drug solutions, appropriate amounts of each drug (DEX or CLX) were dissolved in acetone and DCM, respectively [\u003cspan additionalcitationids=\"CR30\" citationid=\"CR29\" class=\"CitationRef\"\u003e29\u003c/span\u003e\u0026ndash;\u003cspan citationid=\"CR31\" class=\"CitationRef\"\u003e31\u003c/span\u003e]. In DEX-releasing polymers, PLGA (PLGA 50:50; intrinsic viscosity (IV)\u0026thinsp;=\u0026thinsp;0.32\u0026ndash;0.44 dL/g and PLGA 85:15; IV\u0026thinsp;=\u0026thinsp;1.3\u0026ndash;1.7 dL/g), LMW-PLGA (PLGA 50:50; MW\u0026thinsp;=\u0026thinsp;8 kDa), and PLA-PEG (MW\u0026thinsp;=\u0026thinsp;5.3 kDa) were dissolved to their final concentrations using acetone solutions of DEX of determined concentrations. Regarding CLX-loaded polymers, the polymers PCL, with intrinsic viscosities of 0.39 dL/g (~\u0026thinsp;32 KDa; \u0026ldquo;LMW-PCL\u0026rdquo;) and 0.9 dL/g (~\u0026thinsp;73 KDa; \u0026ldquo;HMW-PCL\u0026rdquo;) and PLLA, with IV of 1.0 dL/g (~\u0026thinsp;74 KDa; \u0026ldquo;LMW-PLLA\u0026rdquo;) and 2.9 dL/g (~\u0026thinsp;210 KDa; \u0026ldquo;HMW-PLLA\u0026rdquo;) as well as combinations of LMW-PCL (IV\u0026thinsp;=\u0026thinsp;0.39 dL/g\u0026thinsp;~\u0026thinsp;32 kDa) and HMW-PLLA (IV\u0026thinsp;=\u0026thinsp;2.9 dL/g\u0026thinsp;~\u0026thinsp;210 kDa) were dissolved to their final concentrations using DCM solutions of CLX of determined concentrations. The polymers used were selected based on their hydrophobicity and degradation rates. All these procedures were carried out under ambient conditions. These films were prepared by solvent casting. 70 \u0026micro;L of polymer-drug solutions were then cast on square-shaped meniscus implants (made of PCU) of 0.7 x 0.7 x 0.3 cm. The organic solvent was allowed to evaporate for 1 hour at room temperature. The resulting drug-loaded polymer films were then vacuum-dried for at least 24 hours (Supplementary Fig.\u0026nbsp;1A).\u003c/p\u003e \u003c/div\u003e \u003cdiv id=\"Sec6\" class=\"Section2\"\u003e \u003ch2\u003e2.4. Production of bilayer drug-releasing polymer coatings\u003c/h2\u003e \u003cp\u003eTo prepare the polymer-drug solutions, appropriate amounts of each drug (DEX or CLX) to achieve the desired concentrations were dissolved in acetone and DCM, respectively [\u003cspan additionalcitationids=\"CR30\" citationid=\"CR29\" class=\"CitationRef\"\u003e29\u003c/span\u003e\u0026ndash;\u003cspan citationid=\"CR31\" class=\"CitationRef\"\u003e31\u003c/span\u003e]. Polymers, LMW-PLGA (PLGA 50:50; MW\u0026thinsp;=\u0026thinsp;6.9 kDa), and PLA-PEG (MW\u0026thinsp;=\u0026thinsp;5.3 kDa) were dissolved at a concentration of 200 mg/mL using a solution of DEX of 5 mg/mL. Regarding CLX-releasing polymers, blends of PLLA (IV\u0026thinsp;=\u0026thinsp;2.9 dL/g; MW\u0026thinsp;~\u0026thinsp;210 kDa) and PCL (IV\u0026thinsp;=\u0026thinsp;0.39 dL/g; MW\u0026thinsp;~\u0026thinsp;32 kDa) were dissolved to their final concentrations using a solution of CLX of 30 mg/mL. The square-shaped PCU implants (0.7 \u0026times; 0.7 \u0026times; 0.3 cm), held in place by a needle, were immersed in the polymer-CLX solution and immediately withdrawn, allowing 10 min for drying. The cycle was repeated a total of 3 times. After an additional 3 hours of drying, 5 new immersion cycles were performed, in this case in the polymer-DEX solution and with a drying time of 15 minutes between cycles. Finally, the organic solvent was allowed to evaporate for 72 hours (Supplementary Fig.\u0026nbsp;1C).\u003c/p\u003e \u003c/div\u003e \u003cdiv id=\"Sec7\" class=\"Section2\"\u003e \u003ch2\u003e2.5. Drug release evaluation\u003c/h2\u003e \u003cp\u003eDrug release studies were performed in agitation (450 rpm) at 37\u0026deg;C in PBS Tween\u0026reg;80 1% (w/v) to ensure sink conditions. The amount of drug (DEX and CLX) released was quantified by reverse phase ultra-performance liquid chromatography (UPLC) with a TUV detector at 239 nm using a column Kinetex\u0026reg; 1.7 \u0026micro;m C18 100 \u0026Aring;, LC Column 50 x 2.1 mm acquired from Phenomenex (Torrance, CA, USA), maintaining the samples at 20\u0026deg;C in a Waters Acquity H-Class UPLC system (Waters, Milford, USA) [\u003cspan citationid=\"CR32\" class=\"CitationRef\"\u003e32\u003c/span\u003e]. The mobile phase consisted of A: deionized H\u003csub\u003e2\u003c/sub\u003eO acidified with TFA 0.1% (v/v) and B: ACN acidified with TFA 0.1% (v/v) pumped with a flow rate of 0.1 mL/min. The injection volume was 5 \u0026micro;L, and the column oven temperature was set to 40\u0026deg;C. To control the UPLC/UV system as well as for data acquisition and processing, EMPOWER software was used. To quantify the amount of DEX and CLX, a calibration curve, ranging from 1 to 100 ppm was performed. The curves had a correlation coefficient (R\u003csup\u003e2\u003c/sup\u003e) of 1 for DEX, and 0.9999 for CLX (n\u0026thinsp;=\u0026thinsp;22). The validation procedure was carried out according to the ICH guidelines [\u003cspan citationid=\"CR33\" class=\"CitationRef\"\u003e33\u003c/span\u003e, \u003cspan citationid=\"CR34\" class=\"CitationRef\"\u003e34\u003c/span\u003e]. Limit of detection (LOD) and quantification (LOQ) were calculated directly from the calibration plots, as 3.3σ/S and 10σ/S, respectively, where σ is the standard deviation of intercept and S is the slope of the calibration plot [\u003cspan citationid=\"CR35\" class=\"CitationRef\"\u003e35\u003c/span\u003e]. The values were LOD\u0026thinsp;=\u0026thinsp;1.73 ppm, and LOQ\u0026thinsp;=\u0026thinsp;5.25 ppm for DEX, and LOD\u0026thinsp;=\u0026thinsp;0.3 ppm, and LOQ\u0026thinsp;=\u0026thinsp;0.92 ppm for CLX.\u003c/p\u003e \u003c/div\u003e \u003cdiv id=\"Sec8\" class=\"Section2\"\u003e \u003ch2\u003e2.6. Drug loading assay\u003c/h2\u003e \u003cp\u003ePCU prostheses coated with drug-releasing polymers were immersed in 5 mL of a mixture of DCM/acetone 3/2 (v/v) for up to 24 hours to ensure polymer disolution. Samples of 200 \u0026micro;L were then transferred to microtubes containing 800 \u0026micro;L methanol and centrifuged at 10,000 rpm for 20 min. 200 \u0026micro;L of supernatant were further diluted with 800 \u0026micro;L MeOH:H\u003csub\u003e2\u003c/sub\u003eO 65:35 (v/v) and the concentration of drug was quantified by UPLC with a TUV detector at 239 nm using a column Kinetex\u0026reg; 1.7 \u0026micro;m C18 100 \u0026Aring;, LC Column 50 x 2.1 mm acquired from Phenomenex (Torrance, CA, USA), maintaining the samples at 20\u0026deg;C in a Waters Acquity H-Class UPLC system (Waters, Milford, USA) [\u003cspan citationid=\"CR32\" class=\"CitationRef\"\u003e32\u003c/span\u003e]. The mobile phase consisted of A: deionized H\u003csub\u003e2\u003c/sub\u003eO acidified with TFA 0.1% (v/v) and B: ACN acidified with TFA 0.1% (v/v) pumped with a flow rate of 0.1 mL/min. The gradient was from 25\u0026ndash;60% of B in 6.5 min, and 6 min from 60\u0026ndash;25% of B. The injection volume was 5 \u0026micro;L, and the column oven temperature was set to 40\u0026deg;C. To control the UPLC/UV system as well as for data acquisition and processing, EMPOWER software was used. To quantify the amount of DEX and CLX, a calibration curve, ranging from 1 to 100 ppm was used. The release buffer had no matrix effect on the quantification of the drugs. The curves had a correlation coefficient (R\u003csup\u003e2\u003c/sup\u003e) of 1 for DEX, and 1 for CLX (n\u0026thinsp;=\u0026thinsp;8). The values were LOD\u0026thinsp;=\u0026thinsp;0.383 ppm, and LOQ\u0026thinsp;=\u0026thinsp;1.162 ppm for DEX, and LOD\u0026thinsp;=\u0026thinsp;0.187 ppm, and LOQ\u0026thinsp;=\u0026thinsp;0.567 ppm for CLX.\u003c/p\u003e \u003c/div\u003e \u003cdiv id=\"Sec9\" class=\"Section2\"\u003e \u003ch2\u003e2.7. Characterization of the drug-releasing polymer films\u003c/h2\u003e \u003cdiv id=\"Sec10\" class=\"Section3\"\u003e \u003ch2\u003e2.7.1. Differential Scanning Calorimetry (DSC)\u003c/h2\u003e \u003cp\u003eDSC measurements for polymers, drugs, and combinations of polymer and drug were recorded with a DSC Q1000 V9.9 (TA Instruments, New Castle, DE, USA) in standard aluminum sample pans. The samples were heated from 20\u0026deg;C to 400\u0026deg;C, depending on the sample analyzed, with a heating rate of 10\u0026deg;C/min in a nitrogen atmosphere. Data recording and processing of the first heating cycles were carried out with the software Advantage (TA Instruments, New Castle, DE, USA).\u003c/p\u003e \u003c/div\u003e \u003cdiv id=\"Sec11\" class=\"Section3\"\u003e \u003ch2\u003e2.7.2. Powder X-ray diffraction (XRD)\u003c/h2\u003e \u003cp\u003eDiffraction measurements of crystalline powder were carried out by an Empyrean diffractometer of the PANAlytical brand. The X-rays were obtained from a sealed tube with Cu anode (λ(Kα1)\u0026thinsp;=\u0026thinsp;1.5406 \u0026Aring;) and were collimated prior to incidence on the sample with optics including a W/Si bilayer mirror. The radiation emitted by the sample was collected with a \"PIXcel3D\" type solid state detector. The samples were mounted at ambient temperature on a flat base without signal (Si single crystal), to avoid the amorphous component different from that coming from the sample. Diffractograms were taken in an angular range of 2 to 40\u0026deg; with a step of 0.04 and a time per step of 8 s. To perform the mathematical adjustments of the obtained diffractograms, the program HighScore Plus: Version 3.0d was used.\u003c/p\u003e \u003c/div\u003e \u003c/div\u003e \u003cdiv id=\"Sec12\" class=\"Section2\"\u003e \u003ch2\u003e2.8. Degradation of bilayer drug-releasing polymer coatings\u003c/h2\u003e \u003cp\u003eDrug-releasing polymer coated PCU implants were immersed in a volume of 10 mL of PBS (pH 7.4) supplemented with 1% (w/v) Tween\u0026reg;80. At defined time points, the coated squared-shaped PCU implants were removed from the release buffer for further analysis:\u003c/p\u003e \u003cdiv id=\"Sec13\" class=\"Section3\"\u003e \u003ch2\u003e2.8.1. Field-emission scanning electron microscopy (FESEM)\u003c/h2\u003e \u003cp\u003eDrug-releasing polymer coatings were sputter coated with a layer of iridium and imaged in a Zeiss UltraPlus analytical FESEM with a beam voltage of 3 kV and a magnification ranging from 500 to 10,000X for the analysis of the surface of the coatings. Also, Zeiss EVO analytical FESEM with a beam voltage of 20 kV and magnification ranging from 60X to 5,000X was used for the measure of the coating thickness and analyzing side profile upon degradation after sagittal cut.\u003c/p\u003e \u003c/div\u003e \u003cdiv id=\"Sec14\" class=\"Section3\"\u003e \u003ch2\u003e2.8.2. pH measurements\u003c/h2\u003e \u003cp\u003epH was measured in the media where the polymer coated PCU implants were incubated over specific periods using a pH meter calibrated using standard buffer solutions of known pH values ranging from pH 2.0 to 10.0.\u003c/p\u003e \u003c/div\u003e \u003c/div\u003e \u003cdiv id=\"Sec15\" class=\"Section2\"\u003e \u003ch2\u003e2.9. \u003cem\u003eIn vitro\u003c/em\u003e studies\u003c/h2\u003e \u003cdiv id=\"Sec16\" class=\"Section3\"\u003e \u003ch2\u003e2.9.1. Drug release evaluation\u003c/h2\u003e \u003cp\u003eDrug release studies to evaluate released drug activity \u003cem\u003ein vitro\u003c/em\u003e were performed at 37\u0026deg;C in PBS Tween\u0026reg;80 0.05% (w/v) supplemented with penicillin/streptomycin 1% (v/v). The amount of drug (DEX and CLX) released was quantified by UPLC using the same method as described in section \u003cem\u003eDrug loading assay (Section 2.6)\u003c/em\u003e. The release buffer collected and quantified at each time point was subsequently lyophilized. Afterwards, the powder was resuspended in sterilized miliQ water and evaluated \u003cem\u003ein vitro\u003c/em\u003e.\u003c/p\u003e \u003cp\u003eStored release buffers were diluted to achieve a concentration of DEX and CLX (Table\u0026nbsp;\u003cspan refid=\"Tab1\" class=\"InternalRef\"\u003e1\u003c/span\u003e) as close as possible to their active anti-inflammatory concentration (10 \u0026micro;M for DEX and 25 \u0026micro;M for CLX), based on previous investigations [\u003cspan citationid=\"CR27\" class=\"CitationRef\"\u003e27\u003c/span\u003e]. A minimum dilution of 1:10 using RPMI was implemented for each sample, regardless of the concentration of drugs quantified in the media (Supplementary Fig.\u0026nbsp;2), to prevent possible toxicity by the release buffer (Tween\u003csup\u003e\u0026reg;\u003c/sup\u003e80 0.05% w/v).\u003c/p\u003e \u003cp\u003e \u003cdiv class=\"gridtable\"\u003e\u003ctable float=\"Yes\" id=\"Tab1\" border=\"1\"\u003e \u003ccaption language=\"En\"\u003e \u003cdiv class=\"CaptionNumber\"\u003eTable 1\u003c/div\u003e \u003cdiv class=\"CaptionContent\"\u003e \u003cp\u003eConcentrations of DEX and CLX obtained from Prototype PLA-PEG and Prototype PLGA bilayer polymer coatings and total percentage (%) of drug released at each time point for the \u003cem\u003ein vitro\u003c/em\u003e validation of their sterility and anti-inflammatory activity, and efficacy.\u003c/p\u003e \u003c/div\u003e \u003c/caption\u003e \u003ccolgroup cols=\"9\"\u003e \u003cdiv align=\"left\" class=\"colspec\" colname=\"c1\" colnum=\"1\"\u003e\u003c/div\u003e \u003cdiv align=\"left\" class=\"colspec\" colname=\"c2\" colnum=\"2\"\u003e\u003c/div\u003e \u003cdiv align=\"left\" class=\"colspec\" colname=\"c3\" colnum=\"3\"\u003e\u003c/div\u003e \u003cdiv align=\"left\" class=\"colspec\" colname=\"c4\" colnum=\"4\"\u003e\u003c/div\u003e \u003cdiv align=\"left\" class=\"colspec\" colname=\"c5\" colnum=\"5\"\u003e\u003c/div\u003e \u003cdiv align=\"left\" class=\"colspec\" colname=\"c6\" colnum=\"6\"\u003e\u003c/div\u003e \u003cdiv align=\"left\" class=\"colspec\" colname=\"c7\" colnum=\"7\"\u003e\u003c/div\u003e \u003cdiv align=\"left\" class=\"colspec\" colname=\"c8\" colnum=\"8\"\u003e\u003c/div\u003e \u003cdiv align=\"left\" class=\"colspec\" colname=\"c9\" colnum=\"9\"\u003e\u003c/div\u003e \u003cthead\u003e \u003ctr\u003e \u003cth align=\"left\" colname=\"c1\"\u003e\u0026nbsp;\u003c/th\u003e \u003cth align=\"left\" colname=\"c2\"\u003e\u0026nbsp;\u003c/th\u003e \u003cth align=\"left\" colspan=\"3\" nameend=\"c5\" namest=\"c3\"\u003e \u003cp\u003eCLX\u003c/p\u003e \u003c/th\u003e \u003cth align=\"left\" colspan=\"3\" nameend=\"c8\" namest=\"c6\"\u003e \u003cp\u003eDEX\u003c/p\u003e \u003c/th\u003e \u003cth align=\"left\" colname=\"c9\"\u003e\u0026nbsp;\u003c/th\u003e \u003c/tr\u003e \u003c/thead\u003e \u003ctbody\u003e \u003ctr\u003e \u003ctd align=\"left\" colname=\"c1\"\u003e \u003cp\u003ePrototype\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c2\"\u003e \u003cp\u003eTime point\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c3\"\u003e \u003cp\u003eppm\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c4\"\u003e \u003cp\u003e\u0026micro;M\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c5\"\u003e \u003cp\u003e% released\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c6\"\u003e \u003cp\u003eppm\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c7\"\u003e \u003cp\u003e\u0026micro;M\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c8\"\u003e \u003cp\u003e% released\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c9\"\u003e \u003cp\u003eDilution\u003c/p\u003e \u003c/td\u003e \u003c/tr\u003e \u003ctr\u003e \u003ctd align=\"left\" colname=\"c1\" morerows=\"4\" rowspan=\"5\"\u003e \u003cp\u003ePrototype PLA-PEG\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c2\"\u003e \u003cp\u003e3 h\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c3\"\u003e \u003cp\u003e27.27\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c4\"\u003e \u003cp\u003e71.50\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c5\"\u003e \u003cp\u003e0.71\u0026thinsp;\u0026plusmn;\u0026thinsp;0.22\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c6\"\u003e \u003cp\u003e193.63\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c7\"\u003e \u003cp\u003e493.38\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c8\"\u003e \u003cp\u003e39.08\u0026thinsp;\u0026plusmn;\u0026thinsp;1.01\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c9\"\u003e \u003cp\u003e1:50\u003c/p\u003e \u003c/td\u003e \u003c/tr\u003e \u003ctr\u003e \u003ctd align=\"left\" colname=\"c2\"\u003e \u003cp\u003e3 d\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c3\"\u003e \u003cp\u003e118.75\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c4\"\u003e \u003cp\u003e311.38\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c5\"\u003e \u003cp\u003e5.58\u0026thinsp;\u0026plusmn;\u0026thinsp;0.97\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c6\"\u003e \u003cp\u003e97.58\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c7\"\u003e \u003cp\u003e248.65\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c8\"\u003e \u003cp\u003e89.16\u0026thinsp;\u0026plusmn;\u0026thinsp;3.67\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c9\"\u003e \u003cp\u003e1:20\u003c/p\u003e \u003c/td\u003e \u003c/tr\u003e \u003ctr\u003e \u003ctd align=\"left\" colname=\"c2\"\u003e \u003cp\u003e1 w\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c3\"\u003e \u003cp\u003e58.13\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c4\"\u003e \u003cp\u003e152.44\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c5\"\u003e \u003cp\u003e7.38\u0026thinsp;\u0026plusmn;\u0026thinsp;1.01\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c6\"\u003e \u003cp\u003e22.95\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c7\"\u003e \u003cp\u003e58.49\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c8\"\u003e \u003cp\u003e93.70\u0026thinsp;\u0026plusmn;\u0026thinsp;3.06\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c9\"\u003e \u003cp\u003e1:20\u003c/p\u003e \u003c/td\u003e \u003c/tr\u003e \u003ctr\u003e \u003ctd align=\"left\" colname=\"c2\"\u003e \u003cp\u003e2 w\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c3\"\u003e \u003cp\u003e130.69\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c4\"\u003e \u003cp\u003e342.70\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c5\"\u003e \u003cp\u003e11.20\u0026thinsp;\u0026plusmn;\u0026thinsp;1.16\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c6\"\u003e \u003cp\u003e17.41\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c7\"\u003e \u003cp\u003e44.36\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c8\"\u003e \u003cp\u003e96.63\u0026thinsp;\u0026plusmn;\u0026thinsp;1.40\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c9\"\u003e \u003cp\u003e1:10\u003c/p\u003e \u003c/td\u003e \u003c/tr\u003e \u003ctr\u003e \u003ctd align=\"left\" colname=\"c2\"\u003e \u003cp\u003e4 w\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c3\"\u003e \u003cp\u003e134.31\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c4\"\u003e \u003cp\u003e352.17\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c5\"\u003e \u003cp\u003e20.37\u0026thinsp;\u0026plusmn;\u0026thinsp;1.58\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c6\"\u003e \u003cp\u003e1.81\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c7\"\u003e \u003cp\u003e4.60\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c8\"\u003e \u003cp\u003e~\u0026thinsp;100\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c9\"\u003e \u003cp\u003e1:15\u003c/p\u003e \u003c/td\u003e \u003c/tr\u003e \u003ctr\u003e \u003ctd align=\"left\" colname=\"c1\" morerows=\"4\" rowspan=\"5\"\u003e \u003cp\u003ePrototype PLGA\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c2\"\u003e \u003cp\u003e3 h\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c3\"\u003e \u003cp\u003e7.14\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c4\"\u003e \u003cp\u003e18.74\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c5\"\u003e \u003cp\u003e0.21\u0026thinsp;\u0026plusmn;\u0026thinsp;0.02\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c6\"\u003e \u003cp\u003e16.40\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c7\"\u003e \u003cp\u003e41.80\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c8\"\u003e \u003cp\u003e6.28\u0026thinsp;\u0026plusmn;\u0026thinsp;0.31\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c9\"\u003e \u003cp\u003e1:10\u003c/p\u003e \u003c/td\u003e \u003c/tr\u003e \u003ctr\u003e \u003ctd align=\"left\" colname=\"c2\"\u003e \u003cp\u003e3 d\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c3\"\u003e \u003cp\u003e40.81\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c4\"\u003e \u003cp\u003e107.00\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c5\"\u003e \u003cp\u003e2.81\u0026thinsp;\u0026plusmn;\u0026thinsp;0.32\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c6\"\u003e \u003cp\u003e45.82\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c7\"\u003e \u003cp\u003e116.74\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c8\"\u003e \u003cp\u003e44.76\u0026thinsp;\u0026plusmn;\u0026thinsp;1.76\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c9\"\u003e \u003cp\u003e1:10\u003c/p\u003e \u003c/td\u003e \u003c/tr\u003e \u003ctr\u003e \u003ctd align=\"left\" colname=\"c2\"\u003e \u003cp\u003e1 w\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c3\"\u003e \u003cp\u003e60.53\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c4\"\u003e \u003cp\u003e158.73\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c5\"\u003e \u003cp\u003e4.87\u0026thinsp;\u0026plusmn;\u0026thinsp;0.31\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c6\"\u003e \u003cp\u003e48.45\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c7\"\u003e \u003cp\u003e123.46\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c8\"\u003e \u003cp\u003e63.51\u0026thinsp;\u0026plusmn;\u0026thinsp;1.03\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c9\"\u003e \u003cp\u003e1:10\u003c/p\u003e \u003c/td\u003e \u003c/tr\u003e \u003ctr\u003e \u003ctd align=\"left\" colname=\"c2\"\u003e \u003cp\u003e2 w\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c3\"\u003e \u003cp\u003e108.59\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c4\"\u003e \u003cp\u003e284.74\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c5\"\u003e \u003cp\u003e8.49\u0026thinsp;\u0026plusmn;\u0026thinsp;1.87\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c6\"\u003e \u003cp\u003e25.69\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c7\"\u003e \u003cp\u003e63.46\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c8\"\u003e \u003cp\u003e72.61\u0026thinsp;\u0026plusmn;\u0026thinsp;2.87\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c9\"\u003e \u003cp\u003e1:10\u003c/p\u003e \u003c/td\u003e \u003c/tr\u003e \u003ctr\u003e \u003ctd align=\"left\" colname=\"c2\"\u003e \u003cp\u003e4 w\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c3\"\u003e \u003cp\u003e115.87\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c4\"\u003e \u003cp\u003e303.84\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c5\"\u003e \u003cp\u003e16.08\u0026thinsp;\u0026plusmn;\u0026thinsp;3.44\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c6\"\u003e \u003cp\u003e10.49\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c7\"\u003e \u003cp\u003e26.75\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c8\"\u003e \u003cp\u003e82.77\u0026thinsp;\u0026plusmn;\u0026thinsp;5.89\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c9\"\u003e \u003cp\u003e1:12\u003c/p\u003e \u003c/td\u003e \u003c/tr\u003e \u003c/tbody\u003e \u003c/colgroup\u003e \u003c/table\u003e\u003c/div\u003e \u003c/p\u003e \u003c/div\u003e \u003c/div\u003e\u003cp\u003e\u003cstrong\u003eAbbreviations:\u003c/strong\u003e CLX: Celecoxib. DEX: Dexamethasone. PLA-PEG: poly(lactic acid)-poly(ethylene glycol) di-block copolymer. PLGA: Poly(lactic-co-glycolic) acid. ppm: parts per million (\u0026micro;g/mL). h: hour. d: days. w: week. \u0026micro;M: Micromolar. Values represent the mean \u0026plusmn; standard deviation (n=3).\u003c/p\u003e\n\u003cp\u003e2.9.2. \u003cem\u003eIn vitro\u003c/em\u003e control of endotoxin contamination\u003c/p\u003e\n\u003cp\u003eThe chromogenic LAL-test (LO50650U, Lonza) was used, following the manufacturer\u0026rsquo;s indications. Each of the samples used for \u003cem\u003ein vitro\u003c/em\u003e studies showed endotoxin levels below 0,125 EU/ml (otherwise the samples were discarded) to prevent possible interferences in their immunotoxicity activity.\u003c/p\u003e\n\u003cp\u003e2.9.3.\u0026nbsp; \u0026nbsp;Monocyte isolation, differentiation, and treatment of human primary macrophages\u003c/p\u003e\n\u003cp\u003eMonocytes were isolated from the whole blood of six healthy donors (buffy coats) following the Ficoll and Percoll separation method, as previously described [36], and subsequently counted with Trypan Blue staining 0.04% using a counting chamber (Kova international, BVS100H). In order to obtain human monocyte-derived macrophages (HMDMs), monocytes were resuspended in RPMI supplemented with penicillin/streptomycin and glutamine, subsequently counted, and 10\u003csup\u003e6\u003c/sup\u003e cells were plated in each well of 24-well plates to allow their attachment to the plate. After 30 minutes of incubation at 37 \u0026deg;C, media was removed, and a new media containing human recombinant M-colony stimulating factor (hrM-CSF, Peptrotech, 300-25) supplemented with fetal bovine serum (FBS), penicillin/streptomycin and glutamine were added for differentiation towards macrophages after 5 days.\u0026nbsp;\u003c/p\u003e\n\u003cp\u003eAt day 5, macrophages were treated with appropriately diluted release buffers. Samples were added together with complete media for 24 hours. Non-treated HMDMs were used as the negative control, while HMDMs treated with LPS + IFN\u0026gamma; (100 ng/mL and 50 ng/mL, respectively) were used as the positive control (M1-like macrophages). All other samples were treated with free drugs (DEX, CLX, and DEX+CLX) at appropriate concentrations (10\u0026nbsp;mM for DEX and 25\u0026nbsp;mM for CLX) for comparison with equivalent concentrations of the drugs in release buffer from Prototype PLA-PEG or Prototype PLGA drug-releasing bilayer polymer coatings at determined times.\u0026nbsp;\u003c/p\u003e\n\u003cp\u003e2.9.4.\u0026nbsp; \u0026nbsp;Evaluation of biocompatibility and toxicity\u0026nbsp;\u003c/p\u003e\n\u003cp\u003eTo assess the biocompatibility of the drug-loaded polymer coatings using primary human macrophages, the AlamarBlue\u0026trade; cell viability assay was performed. After exposure of macrophages to conditions described in Table 1 for 24 hours (relative day 6 of cell culture), which corresponds to the release buffer containing the drugs and products of polymer degradation at different time points, supernatants were collected for cytokine quantification, while 1 mL of AlamarBlue\u0026trade; HS Cell Viability Reagent 10% (Invitrogen, A50101) was added to the remaining cells in each well of the 24-well plates, following the manufacturer\u0026rsquo;s protocol. Plates were incubated for 3 hours at 37 \u0026deg;C, protected from light. 100 \u0026micro;L from each well, in triplicate, were transferred to black 96-well plates, and fluorescence was measured at 560-590 nm using a Synergy H4 Microplate reader (BioTek).\u003c/p\u003e\n\u003cp\u003eNon-treated cells were used as controls and considered as 100% cell viability. Cell viability was calculated according to the equation: % Cell viability = (Sample Fluorescence / Control Fluorescence) \u0026times; 100.\u003c/p\u003e\n\u003cp\u003e2.9.5.\u0026nbsp; \u0026nbsp;Assessment of anti-inflammatory activity and efficacy by measurement of cytokine secretion\u0026nbsp;\u003c/p\u003e\n\u003cp\u003eAfter checking the absence of LPS contamination and non-toxic activity of the release buffer containing DEX and CLX at indicated time points, the anti-inflammatory activity of the drugs in the release buffer was compared with the free drugs at equivalent concentrations by ELISA. The anti-inflammatory activity was evaluated by quantifying the secretion of relevant inflammatory signals (TNF-\u0026alpha;, CCL2, and PGE2) from primary human macrophages treated with stimulated with LPS + IFN\u0026gamma; alone, in combination with free drugs or exposed to the release buffer for 24 hours.\u0026nbsp;\u003c/p\u003e\n\u003cp\u003eELISAs were performed using the commercial kit DuoSet\u0026reg; ELISA Development Systems (Bio-Techne) for TNF-\u0026alpha; (DY210) and the Prostaglandin E2 ELISA kit (Cayman) for PGE2 (004CA514010-96), following the manufacturer\u0026rsquo;s instructions. ELISA assays for CCL2 were performed using a custom kit developed in the HUNIMED lab [36].\u0026nbsp;\u003c/p\u003e\n\u003cp\u003eThe amount of cytokine secretion (ng/ml) in macrophages treated with LPS + IFN\u0026gamma; (positive control) was normalized to 100% and, subsequently, values for the remaining samples were normalized and expressed as percentage (%) of control, according to the equation: % of Control = (ng/ml of Sample / ng/ml of Control) \u0026times; 100.\u003c/p\u003e"},{"header":"Results and discussion","content":"\u003cp\u003eOur aim was to engineer a biodegradable bilayer polymer coating around a PCU meniscus prosthesis to release two anti-inflammatory drugs with controlled release kinetics. More precisely, dexamethasone (DEX) was intended to be released within 1-4 weeks to alleviate the acute inflammation caused mainly by the surgery. Meanwhile, celecoxib (CLX) was intended to be released within 6-9 months to relieve the chronic inflammation related to the prosthesis. Both drugs are expected to mediate the incorporation of the implant into the knee cavity, reducing potential adverse inflammatory reactions. Therefore, we developed and optimized the preparation of the bilayer coating, and we performed experiments to evaluate the activity of released drugs \u003cem\u003ein vitro\u003c/em\u003e. For this, we followed a specific work plan consisting of:\u003c/p\u003e\n\u003cp\u003e1) The screening of different biodegradable and biocompatible polymers for the independent release of DEX and CLX with the desired release kinetics.\u003c/p\u003e\n\u003cp\u003e2) The development of drug-releasing bilayer polymer coatings results from the combination of the most promising prototypes selected for the independent release of the drugs. This development consisted of evaluating the capacity of the bilayer systems to release CLX and DEX in a required time frame, characterization of their physicochemical properties, and \u003cem\u003ein vitro\u003c/em\u003e study of their degradation.\u003c/p\u003e\n\u003cp\u003e3) The \u003cem\u003ein vitro\u003c/em\u003e evaluation of the sterility and biocompatibility of the drugs and polymer degradation products from the bilayer coated system, and the anti-inflammatory activity of these drugs in human primary macrophages.\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003e3.1. Design and optimization of drug-loaded polymer films\u003c/strong\u003e\u003c/p\u003e\n\u003cp\u003ePolyesters were selected as the base polymers of the bilayer coating due to their controlled biodegradability, tunable release properties and favorable regulatory profile. The primary objective of this section was to evaluate the influence of the polymer type on the drug release profile while keeping other parameters\u0026mdash;such as the preparation method and the number of polymer layers\u0026mdash;constant. Although the ultimate goal was to achieve the controlled release of two different anti-inflammatory drugs from the same meniscus prosthesis with distinct release kinetics, the initial screening focused on evaluating the release of each drug independently within a single polymeric matrix.\u003c/p\u003e\n\u003cp\u003e3.1.1. CLX-loaded Polymer films\u003c/p\u003e\n\u003cp\u003ePoly(caprolactone) (PCL) and poly(L-lactide) (PLLA) were chosen due to their slow degradation rate and, hence, long-term release capacity of CLX (6 to 9 months). This profile was expected to modulate a process of chronic inflammation [37,38]. The influence of molecular weight (MW) on degradation rate was explored by testing high molecular weight (HMW) (~210 kDa HMW-PLLA and ~73 kDa HMW-PCL) and low molecular weight (LMW) (~74 kDa LMW-PLLA and ~32 kDa LMW-PCL) variants of each polymer [39\u0026ndash;43]. CLX-loaded films were fabricated by dissolving CLX and each polymer in dichloromethane (DCM) and casting the solution onto square-shaped PCU implants. After curing, the films and PCU implants were immersed in PBS with 1% (w/v) Tween\u0026reg;80, and CLX release was monitored via UPLC (Figure 1). The drug release study was halted when a plateau release phase was observed. \u003c/p\u003e\n\u003cp\u003eThe release studies revealed that CLX was consistently released faster from PCL films, irrespective of MW, as compared to from PLLA films. PLLA exhibited MW-dependent release kinetics, with HMW-PLLA providing more sustained release compared to LMW-PLLA [41]. The differences in release kinetics can be explained by the polymer chain length. Since none of these polymers is expected to degrade within a few weeks, the observed release behavior aligns with previous studies linking polymer chain length to drug diffusion [44].\u003c/p\u003e\n\u003cp\u003eThe denser matrix created by HMW-PLLA restricted water penetration, slowing degradation and drug diffusion and, therefore, delaying drug release. As expected, due to the larger difference in MW, these effects are more pronounced in PLLA films (~140 kDa) than in PCL films (~41 kDa ). The plateau in release profiles is hypothesized to be a result of the interplay between drug diffusion and polymer degradation. Initially, the release is supposed to be diffusion-driven through water-filled pores, whereas the second phase corresponds to the release of the drug remaining trapped within the polymer matrix. Since polymer degradation is slow, this trapped drug remains unavailable, leading to incomplete release within the study timeframe. Thus, the release kinetics reflect both initial diffusion and delayed drug release associated with polymer degradation [45]. \u003c/p\u003e\n\u003cp\u003eAlthough the release profile of HMW-PLLA might be considered aligned with the targeted release timeframe (6\u0026ndash;9 months), the plateau observed after 20 weeks led to an incomplete drug release (30% of the total CLX released (Figure 1)). This incomplete release was attributed to limited polymer degradation, which could lead to prototype limitations such as drug recrystallization [45]. Although recrystallization was not observed or specifically analyzed in this study, prolonged entrapment of drug molecules within the polymer matrix under aqueous conditions is generally considered a risk factor for potential recrystallization in drug delivery systems [46].\u003c/p\u003e\n\u003cp\u003eTo achieve intermediate CLX release kinetics, PLLA and PCL were blended at varying ratios (Figure 2). Increasing PCL content accelerated CLX release, a trend attributed to the formation of polymer-to-polymer interfaces that facilitated water penetration and diffusion of the drug near these interfaces [47\u0026ndash;50]. Due to the limited miscibility of PLLA and PCL, PCL domains aggregate, lowering Tg locally and increasing chain mobility. This facilitates nucleation and improves PLLA\u0026rsquo;s mechanical properties, including flexibility and toughness [51]. Among tested formulations, both PLLA/PCL blends at 70/80 (w/w) and 80/70 (w/w) showed promising results with intermediate release profiles among the blends studied, releasing 40% of CLX after 3 months without plateauing.\u003c/p\u003e\n\u003cp\u003eGiven the variability in intra-articular CLX dosing across animal models (1-1.25 mg CLX/kg in mice [52], 0.46 mg CLX/kg/week over 5 weeks in rabbits [53], 0.0292 mg CLX/kg [54] or 1 mg/kg [55] in rats, and approximately 1 mg/kg in sheep [56]), it was assumed that higher doses of CLX could enhance therapeutic effects in both \u003cem\u003ein vitro\u003c/em\u003e and \u003cem\u003ein vivo\u003c/em\u003e models (e.g., reduced fibrous encapsulation, foreign body reactions, and modulation of signaling pathways). Hence, CLX concentration in selected formulations was increased 2.4-fold (from 12.5 mg/mL to 30 mg/mL), achieving a drug loading (DL) of 16.6%. This increase aimed to enhance the total amount of drug incorporated into the coating, bringing it closer to the intra-articular doses previously reported in the ovine model. The PLLA/PCL blend at an 80/70 (w/w) ratio was ultimately selected for its balanced release profile (Figure 2).\u003c/p\u003e\n\u003cp\u003eThese results led to two emerging candidates for sustained CLX release: HMW-PLLA with 20% DL, which provided a very slow release (\u0026lt;25% CLX released after 6 months), and PLLA/PCL 80/70 (w/w) with 16.67% DL, which achieved ~40% release in 3 months. Both formulations exhibited polymer-drug interactions based on thermal analysis (reductions in Tg, Tm, and Tcc) (Supplementary Figure 4A) and successfully incorporated CLX in a molecularly dispersed form, as confirmed by the absence of CLX crystalline peaks in XRD and the disappearance of its melting endotherm peak in DSC thermograms (Supplementary Figures 4A\u0026ndash;B), both indicative of an amorphous drug state within the polymer matrix. However, only the PLLA/PCL 80/70 (w/w) blend met the required release profile and progressed to further studies.\u003c/p\u003e\n\u003cp\u003e3.1.2. DEX-loaded Polymer films\u003c/p\u003e\n\u003cp\u003eTo effectively address post-surgical acute inflammation, DEX was selected for its potent and pleiotropic anti-inflammatory activity and established clinical use. The goal was to achieve a rapid and controlled release over a 1 to 4-week period by incorporating DEX into an external polymer layer made of PLGA. A high molecular weight (MW) (24\u0026ndash;38 kDa) was initially explored, but presented limitations due to its characteristic biphasic release profile (Supplementary Figure 3) \u0026mdash;marked by an initial burst from surface-associated drug, followed by a less predictable sustained release phase driven by bulk polymer hydrolysis [57,58]. Its degradation rate is influenced by the lactic acid (LA) to glycolic acid (GA) ratio, with a 50:50 composition degrading the fastest due to enhanced hydrophilicity and reduced crystallinity [59]. \u003c/p\u003e\n\u003cp\u003eGiven the limitations of PLGA, particularly its biphasic release profile, an alternative amphiphilic matrix\u0026mdash;poly(lactic acid)-poly(ethylene glycol) di-block copolymer (PLA-PEG)\u0026mdash;was evaluated. Di-block copolymers like PLA-PEG can influence polymer-drug interactions differently than blends of individual polymers [60,61] due to the covalent linkage between the blocks, which creates a more defined and stable microphase separation compared to physical blends [62]. This structural organization can alter drug distribution and water accessibility, ultimately affecting drug release kinetics and solubility. For these reasons, di-block copolymers have been widely used in drug delivery systems [63]. As shown in Figure 3, PLA-PEG enabled a sustained and more complete release of DEX, for approximately one week at a polymer concentration of 200 mg/mL. This accelerated release was likely due to PEG-induced aqueous channel formation within the matrix. However, the slow degradation rate of the PLA block could prolong the integrity of the top layer, acting as a persistent barrier over the bottom CLX-releasing layer and thereby limiting CLX diffusion to the release medium over time. To improve control over the release profile and mitigate the limitations associated with PLGA\u0026rsquo;s biphasic behavior, the effect of PLGA molecular weight (MW) on DEX release was investigated. In addition to the high-MW PLGA used initially, low-MW (LMW) variants (8 kDa) were tested both as standalone matrices and in blends with PLA-PEG (Figure 3).\u003c/p\u003e\n\u003cp\u003eBlends of PLA-PEG and LMW-PLGA 75/25 (w/w) showed release profiles similar to pure PLA-PEG, but a 50/50 (w/w) blend extended the sustained DEX release to two weeks (Figure 3). Increasing the relative amount of LMW-PLGA further enhanced this effect. Optimization studies with different PLA-PEG/LMW-PLGA ratios confirmed that higher LMW-PLGA content produced a sustained release within the desired 1\u0026ndash;4-week window. Notably, pure LMW-PLGA exhibited a similar release profile, making it a practical alternative to PLA-PEG. Its selection not only met the therapeutic release target but also simplified the formulation by avoiding polymer blends.\u003c/p\u003e\n\u003cp\u003eDSC and XRD analyses were performed to confirm the physical state of the incorporated drugs. For both DEX and CLX formulations, the disappearance of their characteristic melting peaks in DSC thermograms and the absence of crystalline reflections in XRD patterns confirmed that the drugs were molecularly dispersed within the polymer matrices (Supplementary Figures 4A and B, and 5A and B). For CLX-loaded systems, thermal events and crystallinity indicated good miscibility and a predominantly amorphous dispersion of the drug. Slight reductions in Tg and Tm\u0026mdash;particularly in PLLA-based formulations\u0026mdash;suggested plasticizing effects and increased polymer chain mobility [64\u0026ndash;67]. These interactions are consistent with the initial diffusion-driven release observed in PLLA and PLLA/PCL blends. Notably, no direct correlation was found between polymer crystallinity and release kinetics, supporting the idea that other factors\u0026mdash;such as porosity and polymer-drug interactions\u0026mdash;played a more dominant role [68]. In the case of DEX, minor increases in crystallinity were observed in PEG-containing systems, suggesting weak drug-polymer interactions that may facilitate water uptake and diffusion. Conversely, PLGA and LMW-PLGA matrices remained largely amorphous, aligning with their more sustained release behavior. A full list of thermal parameters (Tg, Tm, Xc) is provided in Supplementary Tables 1 and 2.\u003c/p\u003e\n\u003cp\u003eThis comprehensive screening identified PLLA/PCL (80/70, w/w) with 16.6% DL as the most promising formulation for sustained CLX release, achieving higher drug incorporation while avoiding the plateau effect observed with pure PLLA. For DEX release, both PLA-PEG and LMW-PLGA (8 kDa, 50:50 LA:GA), each with a DL of 2.44%, achieved controlled release within the target 1\u0026ndash;4 week window, with LMW-PLGA providing a slightly more sustained release profile and the advantage of a simpler formulation.\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003e3.2. \u003c/strong\u003e \u003cstrong\u003eEngineering a bilayer drug-releasing polymer coating for the sequential release of DEX and CLX\u003c/strong\u003e\u003c/p\u003e\n\u003cp\u003eTo enable the simultaneous release of CLX and DEX from the non-biodegradable PCU prosthesis (~0.7 \u0026times; 0.7 \u0026times; 0.3 cm), a bilayer polymeric coating was developed. The system consists of a first layer releasing CLX to be cast on top of the prosthesis, and a second layer releasing DEX, providing a shorter drug release. Polymer matrices previously identified as optimal candidates (Section 3.1.) were selected for testing in this configuration. While solvent casting was effective for coating one surface, dip coating was chosen for full-implant coverage due to its simplicity, versatility, and compatibility with multilayer deposition [18]. This section explores key aspects of bilayer development, including the impact of solvent interactions between layers and the selection of one optimized polymer matrix for each drug for final implementation. \u003c/p\u003e\n\u003cp\u003eGiven that the DEX layer is prepared using acetone, the potential effect of solvent exposure on the underlying CLX-loaded layer was evaluated. Since acetone can partially disrupt polymer-drug interactions, this investigation was crucial for ensuring stability and consistency in release kinetics. This was explored by DSC, XRD, and drug release evaluation\u0026mdash; detailed in \u003cstrong\u003eSupplementary Figures 6 and 7\u003c/strong\u003e. When acetone was applied to PLLA/PCL films, a slight acceleration in CLX release was noted, yet the overall release kinetics remained consistent. However, for pure PLLA, acetone significantly altered the release profile, accelerating CLX release from approximately 10% to over 60% by week 21 (Supplementary Figure 7).\u003c/p\u003e\n\u003cp\u003e3.2.1. Evaluation of drug release\u003c/p\u003e\n\u003cp\u003eThe release profiles of DEX from the two polymer matrices\u0026mdash;PLA-PEG and LMW-PLGA\u0026mdash; were evaluated while maintaining a constant CLX-releasing bottom layer composed of the PLLA/PCL blend. PCU implants were coated using a dip-coating approach, where CLX was first incorporated into the PLLA/PCL film using DCM, and DEX was then incorporated into either PLA-PEG or LMW-PLGA using acetone. The main distinction between these bilayer systems lies in the choice of the top-layer polymer, which influenced not only DEX release but also indirectly modulated CLX release kinetics. For clarity, these systems are referred to as Prototype PLA-PEG and Prototype PLGA (Figure 4A). Together, they provide valuable design insights for dual-drug delivery strategies targeting both acute and chronic phases of inflammation.\u003c/p\u003e\n\u003cp\u003eDespite the identical CLX-releasing layer, the release profiles of CLX varied depending on the top-layer polymer (Figure 4B). When LMW-PLGA was used as the top layer, CLX release was faster, likely because the more rapid degradation of PLGA allowed earlier erosion or removal of the upper layer, facilitating earlier exposure of the CLX layer to the release medium. In contrast, the slower-degrading PLA-PEG formed a more persistent barrier, delaying access of the release medium to the CLX layer and thus slowing its diffusion.\u003c/p\u003e\n\u003cp\u003eDEX release profiles were consistent with our prior results, where PLA-PEG exhibited an initial burst followed by a gradual tapering, while LMW-PLGA provided a more sustained and steady release. Both systems achieved complete DEX release within the first few weeks, effectively meeting the goal of addressing acute post-surgical inflammation.\u003c/p\u003e\n\u003cp\u003e3.2.2. \u003cem\u003eIn vitro\u003c/em\u003e degradation of PLA-PEG and PLGA bilayer coated implants\u003c/p\u003e\n\u003cp\u003eAlthough both bilayer prototypes (PLA-PEG and PLGA) achieved the desired release kinetics for DEX and CLX, they exhibited distinct release profiles. Thus, using these prototypes, degradation studies were conducted to better understand the mechanisms behind their different release behaviors and to validate the hypothesis that PLA-PEG degrades more slowly than PLGA. Coated square implants were immersed in PBS supplemented with 1% (w/v) Tween\u0026reg;80, and periodically analyzed for trends in the change of pH and thickness, as well as surface morphology via FESEM. \u003c/p\u003e\n\u003cp\u003eOver time, Prototype PLA-PEG showed a gradual decrease in thickness, with a pronounced reduction between 4 and 6 months, corresponding to clear structural degradation and coating disintegration (Figure 5A and Figure 6). By 12 months, the coating was visibly deteriorated, appearing uneven and incomplete, with clear signs of degradation and areas where the prosthesis surface was no longer covered. In contrast, Prototype PLGA initially exhibited a thickness increase\u0026mdash;likely due to polymer swelling\u0026mdash;followed by a steady decrease over the year (Figure 5A and Figure 7) [69,70]. The pH monitoring unveiled additional aspects of the degradation behavior not apparent from thickness measurements alone. While Prototype PLA-PEG maintained relatively stable pH values for the first four months, a noticeable decline occurred between months 4 and 6, consistent with the delayed onset of degradation observed in FESEM analysis and expected of PLA. In contrast, Prototype PLGA triggered a sharp pH drop by the second month, which then remained low for the rest of the study (Figure 5B). These trends reflect the faster breakdown of the PLGA matrix and further support the interpretation that the top-layer degradation rate directly influenced buffer access to the CLX-releasing layer.\u003c/p\u003e\n\u003cp\u003eThese findings help explain the differences observed in CLX release profiles. Although both prototypes had the same PLLA/PCL internal layer for CLX release, faster degradation and reduced tortuosity of the PLGA top layer likely allowed earlier buffer access to the underlying CLX matrix, accelerating drug release. In contrast, the slower degradation of PLA-PEG maintained a more effective diffusion barrier for longer. Thickness measurements were affected by sample cutting and imaging variability because of the angle of the image and should be interpreted as relative trends rather than absolute values (Supplementary Figure 8). In contrast, pH changes provided a more robust and reproducible indicator of degradation, though results may not fully reflect \u003cem\u003ein vivo\u003c/em\u003e behavior, where pH buffering and enzymatic processes differ. Nonetheless, prior work suggests that PLGA degradation kinetics \u003cem\u003ein vitro\u003c/em\u003e correlate with \u003cem\u003ein vivo\u003c/em\u003e molecular weight loss [134].\u003c/p\u003e\n\u003cp\u003e3.2.3. Evaluation of drug-releasing polymer coating reproducibility\u003c/p\u003e\n\u003cp\u003eThe reproducibility of the prototypes was assessed by quantifying the total amount of DEX and CLX per unit of area (\u0026micro;g/cm\u0026sup2;) across multiple replicates of square-shaped PCU implants. The \u003cstrong\u003ePrototype PLGA\u003c/strong\u003e showed consistent DL for both DEX and CLX, indicating a robust and reproducible coating process with controlled dipping cycles and solvent evaporation. In contrast, \u003cstrong\u003ePrototype PLA-PEG\u003c/strong\u003e showed slight variability in drug incorporation between replicates, suggesting challenges in achieving uniform deposition. This may be attributed to uneven polymer film formation, variations in drug-polymer interactions, or greater sensitivity to coating conditions on the complex implant surface. These results indicate that \u003cstrong\u003ePrototype PLGA offers greater reliability and process reproducibility\u003c/strong\u003e, while \u003cstrong\u003ePrototype PLA-PEG may require further optimization to ensure consistent DL \u003c/strong\u003e(Supplementary Figure 9)\u003cstrong\u003e.\u003c/strong\u003e\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003e3.3. \u003cem\u003e In vitro\u003c/em\u003e assessment of sterility, biocompatibility, anti-inflammatory activity, and efficiency of the bilayer drug-releasing polymer-coated PCU implants \u003c/strong\u003e\u003c/p\u003e\n\u003cp\u003ePrior to evaluating the immunomodulatory potential of the bilayer-coated implants, a series of \u003cem\u003ein vitro\u003c/em\u003e validations were conducted to ensure compatibility with biological assays. Drug release studies confirmed that both DEX and CLX were released in concentrations exceeding their respective effective thresholds (10 \u0026micro;M for DEX, 25 \u0026micro;M for CLX, based on literature [27]) at key time points\u0026mdash;3 hours, 3 days, 1 week, 2 weeks, and 4 weeks (Table 1). These intervals were strategically selected to capture distinct drug activity phases: early DEX-driven effects, overlapping DEX and CLX activity, and later CLX-dominant responses.\u003c/p\u003e\n\u003cp\u003eThe release buffer also met sterility criteria, with endotoxin levels remaining below 0.125 EU/mL (Supplementary Figure 10), validating their suitability for macrophage-based assays. Moreover, no cytotoxic effects were observed in HMDMs exposed to release media from either prototype at any time point, indicating good biocompatibility of both the drugs released and the polymer degradation byproducts (Supplementary Figure 11). These findings are consistent with previous reports describing the non-toxic nature of polyester degradation products and further support the safety of the materials used [71]. With drug release, sterility, and cytocompatibility confirmed, the subsequent experiments focused on assessing the bioactivity of the released drugs by evaluating their capacity to modulate cytokine secretion in M1-like (classically activated with LPS + IFNg) pro-inflammatory macrophages.\u003c/p\u003e\n\u003cp\u003e3.3.1. \u003cem\u003eIn vitro\u003c/em\u003e evaluation of immunomodulatory activity\u003c/p\u003e\n\u003cp\u003eThe evaluation of cytokine secretion aimed to confirm that the drugs released from the bilayer polymer coatings retained their immunomodulatory activity. Specifically, it was sought to verify that DEX and CLX, when released at their respective time points, could suppress pro-inflammatory cytokine release from macrophages, events associated with acute and chronic inflammation. This assessment was essential to validate the therapeutic functionality of the coating system in an inflammatory environment. Cytokine secretion by M1-like macrophages was measured following exposure to release media collected from the implant incubations at defined time points. The three key pro-inflammatory cytokines selected were TNF-\u0026alpha; (a central mediator of acute inflammation [72\u0026ndash;74]), CCL2 (MCP-1) (regulates monocyte/macrophage recruitment and FBGC formation [75,76]), and PGE2 (a COX-2 pathway product involved in joint inflammation, vasodilation, and pain [77,78]).\u003c/p\u003e\n\u003cp\u003eBoth prototypes significantly suppressed TNF-\u0026alpha; and CCL2 secretion across all time points (Figure 8A and B), demonstrating that the combination of DEX and CLX maintained its anti-inflammatory efficacy following their release from the coatings. Notably, TNF-\u0026alpha; and CCL2 suppression was more closely linked to DEX activity, as expected during early time points. PGE2 secretion, in contrast, was more selectively reduced by CLX, consistent with its inhibition of COX-2 (Figure 8C) [79]. This effect was observed at all time points, regardless of the prototype, supporting the stability and retained bioactivity of CLX even after extended residence within the coating. Importantly, cytokine suppression levels were comparable to those obtained with free (non-encapsulated) drugs, validating the effectiveness of the delivery platform. These results confirm that both Prototype PLGA and Prototype PLA-PEG effectively preserved the bioactivity of the drugs during encapsulation, storage, and release.\u003c/p\u003e"},{"header":"Conclusion","content":"\u003cp\u003eHere we present a bilayer polymer coating developed to enable the sequential release of dexamethasone (DEX) and celecoxib (CLX), from a PCU meniscus prosthesis, with the final goal of mitigating both, acute and chronic inflammation observed after implantation. \u0026nbsp;Through a systematic screening of polymer candidates, two bilayer prototypes\u0026mdash;Prototype PLA-PEG and Prototype PLGA\u0026mdash;were identified based on their ability to achieve the desired release kinetics: rapid DEX release over 1\u0026ndash;4 weeks to prevent acute inflammation, and prolonged CLX release over 6\u0026ndash;9 months for sustained immunomodulation. Polymer degradation studies revealed that Prototype PLGA, composed of a low molecular weight PLGA external layer, exhibited faster degradation, facilitating earlier exposure of the CLX-loaded PLLA/PCL bottom layer to the release medium. In contrast, Prototype PLA-PEG was degraded more slowly, increasing tortuosity and delaying CLX release. Although PLA-PEG allowed for higher DL, PLGA demonstrated more consistent and reproducible release profiles and better integration with the bilayer system. Both prototypes were endotoxin-free and biocompatible, and the released drugs maintained their anti-inflammatory activity, as demonstrated by the effective decrese in the secretion of pro-inflammatory cytokines by M1-like activated macrophages.\u003c/p\u003e\n\u003cp\u003e\u0026nbsp;Based on its favorable degradation kinetics, reproducibility, and compatibility with the bilayer architecture, Prototype PLGA was selected for further development. This next phase would involve optimizing the coating system and scaling up its application to actual meniscus prostheses for evaluation in a representative animal model, bringing the technology closer to clinical translation.\u003c/p\u003e"},{"header":"Declarations","content":"\u003cp\u003e\u003cstrong\u003eEthical statement\u003c/strong\u003e\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003eEthics approval and consent to participate:\u003c/strong\u003e Studies using human-derived monocytes were conducted at Humanitas Research Institute (Milano, Italy) in accordance with Italian and European law. Monocytes were isolated from the buffy coats of blood donations. Buffy coats were obtained from anonymous healthy blood donors at Istituto Clinico Humanitas (ICH) (Milano, Italy). The ethical committee at ICH provided favorable confirmation for their use in basic scientific research activities. Their use does not involve any ethical problem regarding informed consent or the privacy of the donor.\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003eConsent for publication:\u003c/strong\u003e Not applicable.\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003eAvailability of data and materials:\u003c/strong\u003e The datasets generated during and/or analysed during the current study are available from the corresponding author on reasonable request.\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003eCompeting interests:\u003c/strong\u003e Mar\u0026iacute;a Jos\u0026eacute; Alonso is the founder and shareholder of LiberaBio. The remaining authors declare no conflict of interest.\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003eFunding:\u003c/strong\u003e This project has received funding from the European Union\u0026apos;s Horizon 2020 research and innovation programme under grant agreement No 814444 (MEFISTO). Alfonso F. Blanco has received research support through a predoctoral grant from Xunta de Galicia, Grant number ED481A 2022/066. Fernando Torres-And\u0026oacute;n was also supported by a \u0026ldquo;Miguel Servet Grant\u0026rdquo; CP22/00106 by AES 2022, and Alba Pensado-L\u0026oacute;pez by a \u0026ldquo;Sara Borrell Grant\u0026rdquo; CD23/00173 by AES 2023, ISCIII, Spain.\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003eAuthors\u0026apos; contributions:\u003c/strong\u003e All authors contributed to the study conception and design. Material preparation, data collection, and analysis related to drug release studies, polymer selection, polymer coating evaluation, and optimization were performed by Alfonso F. Blanco and Gustavo Lou. In vitro study conceptualization was performed by Alfonso F. Blanco, Aldo Ummarino, Alba Pensado-L\u0026oacute;pez, and Fernando Torres-And\u0026oacute;n. In vitro study preparation and data collection were performed by Aldo Ummarino and Alba Pensado-L\u0026oacute;pez. In vitro study analysis was done by Alfonso F. Blanco. The first draft of the manuscript was written by Alfonso F. Blanco, and all authors commented on previous versions of the manuscript. All authors read and approved the final manuscript.\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003eAcknowledgements:\u003c/strong\u003e Figures were created with BioRender.com.\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003eAuthors\u0026apos; information:\u003c/strong\u003e Not applicable.\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003eData Availability Statement:\u003c/strong\u003e The datasets generated during and/or analysed during the current study are available from the corresponding author on reasonable request.\u003c/p\u003e"},{"header":"References","content":"\u003col\u003e\n\u003cli\u003eHorecka A, Hordyjewska A, Blicharski T, Kurzepa J. Osteoarthritis of the knee - biochemical aspect of applied therapies: a review. 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Polymer (Guildf) 2020;191:122227. https://doi.org/10.1016/J.POLYMER.2020.122227.\u003c/li\u003e\n\u003c/ol\u003e"}],"fulltextSource":"","fullText":"","funders":[],"hasAdminPriorityOnWorkflow":false,"hasManuscriptDocX":true,"hasOptedInToPreprint":true,"hasPassedJournalQc":"","hasAnyPriority":false,"hideJournal":false,"highlight":"","institution":"","isAcceptedByJournal":true,"isAuthorSuppliedPdf":false,"isDeskRejected":"","isHiddenFromSearch":false,"isInQc":false,"isInWorkflow":false,"isPdf":false,"isPdfUpToDate":true,"isWithdrawnOrRetracted":false,"journal":{"display":true,"email":"[email protected]","identity":"drug-delivery-and-translational-research","isNatureJournal":false,"hasQc":true,"allowDirectSubmit":false,"externalIdentity":"ddtr","sideBox":"Learn more about [Drug Delivery and Translational Research](https://www.springer.com/journal/13346)","snPcode":"","submissionUrl":"https://www.editorialmanager.com/ddtr/default.aspx","title":"Drug Delivery and Translational Research","twitterHandle":"","acdcEnabled":true,"dfaEnabled":true,"editorialSystem":"em","reportingPortfolio":"Springer Hybrid","inReviewEnabled":true,"inReviewRevisionsEnabled":false},"keywords":"Knee osteoarthritis, meniscus implant, controlled drug release, polymer, anti-inflammatory drug, inflammation","lastPublishedDoi":"10.21203/rs.3.rs-6869700/v1","lastPublishedDoiUrl":"https://doi.org/10.21203/rs.3.rs-6869700/v1","license":{"name":"CC BY 4.0","url":"https://creativecommons.org/licenses/by/4.0/"},"manuscriptAbstract":"Knee osteoarthritis (OA), a degenerative joint disease, is increasingly prevalent worldwide and often results from a meniscal deterioration that leads to meniscus removal. Replacing the damaged meniscus with a non-biodegradable prosthesis offers an innovative solution to prevent OA progression, particularly in older patients. However, the long-term use of anti-inflammatory drugs for pain relief and prosthesis integration can cause severe off-target side effects. The objective of this work was to design and develop drug-loaded bilayer polymer films to be used as coatings for a meniscus polycarbonate urethane (PCU). The developed bilayer polymer films enabled a sustained release of two anti-inflammatory drugs - dexamethasone (DEX) and celecoxib (CLX) - with distinct release kinetics (1-4 weeks for DEX and 6-9 months for CLX). This release profile was defined to modulate post-surgical and chronic inflammation within the knee joint, respectively. Two bilayer prototypes showed consistent biodegradation, drug release, drug loading, and reproducibility. Furthermore, the systems were sterile, biocompatible, and maintained the anti-inflammatory efficacy of the released drugs, effectively reducing pro-inflammatory cytokine secretion from human primary macrophages.","manuscriptTitle":"Controlled Co-delivery of Anti-inflammatory Drugs from Bilayer Polymer Films Coating a Meniscus Implant","msid":"","msnumber":"","nonDraftVersions":[{"code":1,"date":"2025-06-25 07:04:49","doi":"10.21203/rs.3.rs-6869700/v1","editorialEvents":[{"type":"communityComments","content":0},{"type":"decision","content":"Minor Revisions Needed","date":"2025-07-14T15:11:52+00:00","index":"","fulltext":""},{"type":"reviewerAgreed","content":"","date":"2025-06-23T00:38:23+00:00","index":0,"fulltext":""},{"type":"reviewersInvited","content":"","date":"2025-06-16T20:18:33+00:00","index":"","fulltext":""},{"type":"editorAssigned","content":"","date":"2025-06-13T07:37:29+00:00","index":"","fulltext":""},{"type":"submitted","content":"Drug Delivery and Translational Research","date":"2025-06-12T04:57:04+00:00","index":"","fulltext":""}],"status":"published","journal":{"display":true,"email":"[email protected]","identity":"drug-delivery-and-translational-research","isNatureJournal":false,"hasQc":true,"allowDirectSubmit":false,"externalIdentity":"ddtr","sideBox":"Learn more about [Drug Delivery and Translational Research](https://www.springer.com/journal/13346)","snPcode":"","submissionUrl":"https://www.editorialmanager.com/ddtr/default.aspx","title":"Drug Delivery and Translational Research","twitterHandle":"","acdcEnabled":true,"dfaEnabled":true,"editorialSystem":"em","reportingPortfolio":"Springer Hybrid","inReviewEnabled":true,"inReviewRevisionsEnabled":false}}],"origin":"","ownerIdentity":"57a1957b-9deb-44cf-8686-1ddeff949616","owner":[],"postedDate":"June 25th, 2025","published":true,"recentEditorialEvents":[],"rejectedJournal":[],"revision":"","amendment":"","status":"published-in-journal","subjectAreas":[],"tags":[],"updatedAt":"2025-09-01T15:59:55+00:00","versionOfRecord":{"articleIdentity":"rs-6869700","link":"https://doi.org/10.1007/s13346-025-01942-5","journal":{"identity":"drug-delivery-and-translational-research","isVorOnly":false,"title":"Drug Delivery and Translational Research"},"publishedOn":"2025-08-25 15:57:13","publishedOnDateReadable":"August 25th, 2025"},"versionCreatedAt":"2025-06-25 07:04:49","video":"","vorDoi":"10.1007/s13346-025-01942-5","vorDoiUrl":"https://doi.org/10.1007/s13346-025-01942-5","workflowStages":[]},"version":"v1","identity":"rs-6869700","journalConfig":"researchsquare"},"__N_SSP":true},"page":"/article/[identity]/[[...version]]","query":{"redirect":"/article/rs-6869700","identity":"rs-6869700","version":["v1"]},"buildId":"XKTyCvWXoU3ODBz1xrDgd","isFallback":false,"isExperimentalCompile":false,"dynamicIds":[84888],"gssp":true,"scriptLoader":[]}