Reservoir-Type Subcutaneous Implantable DevicesContaining Porous Rate Controlling Membranes forSustained Delivery of Risperidone

RESEARCH ARTICLE www.advhealthmat.de Reservoir-Type Subcutaneous Implantable Devices Containing Porous Rate Controlling Membranes for Sustained Delivery of Risperidone Linlin Li, Andi Dian Permana,* Juan Domínguez-Robles,* Muh Nur Amir, Habibie Habibie, Qonita Kurnia Anjani, Li Zhao, Natalia Moreno-Castellanos, Ryan F Donnelly, and Eneko Larrañeta* Implantable drug delivery systems are crucial for achieving sustained delivery of active compounds to specific sites or systemic circulation. In this study, a novel reservoir-type implant combining a biodegradable rate-controlling membrane with a drug-containing core prepared using direct compression techniques is developed. The membrane is composed of poly(caprolactone) (PCL), and risperidone (RIS) served as the model drug. Characterization of both membranes and direct compressed pellets includes hardness testing, optical coherence tomography, mercury intrusion porosimetry, and surface morphology observation. In vitro release studies of RIS reveal that higher drug loading in the pellets extended-release duration up to 70 days when incorporated into membranes with four layers. Increasing the number of membrane layers slows the release rate further, ranging from 70 to 170 days depending on membrane thickness. Biocompatibility studies demonstrate that these implantable devices are non-toxic and biocompatible with cells in vitro. In vivo studies conduct in male Wistar rats demonstrate sustained release of RIS, with plasma levels showing a significant increase post-implantation at a relatively constant rate for up to 49 days. These results indicate that the developed implants have the potential to provide long-acting drug delivery to the systemic circulation. L. Li, Q. K. Anjani, L. Zhao, R. F Donnelly, E. Larrañeta School of Pharmacy Queen’s University Belfast Lisburn Road 97, Belfast BT9 7BL, UK E-mail: [email protected] The ORCID identification number(s) for the author(s) of this article can be found under https://doi.org/10.1002/adhm.202403689 © 2025 The Author(s). Advanced Healthcare Materials published by Wiley-VCH GmbH. This is an open access article under the terms of the Creative Commons Attribution License, which permits use, distribution and reproduction in any medium, provided the original work is properly cited. DOI: 10.1002/adhm.202403689 1. Introduction Implantable drug delivery systems (IDDS) are typically medical devices designed to deliver drugs to specific sites of action or systemic circulation in a sustained way.[1] Deansby and Parkes first applied this idea of IDDS in practice in 1938 when they implanted compressed pellets of crystalline estrone subcutaneously to investigate their influence in animal models.[2]IDDS, as a long-acting delivery strategy, offers many advantages. It allows targeted and localized drug delivery and can also bypass first-pass metabolism and degradation in the stomach compared to oral delivery systems. This leads to faster drug release and lower concentrations of drug required to achieve therapeutic effect. Therefore, it might potentially reduce side effects and improve patient compliance.[3]Although surgical procedures are required to initiate therapy, the long-acting delivery ranging from several weeks to months, or even a year, may outweigh this disadvantage.[3] It is difficult to give a classification of IDDS because of several exceptions and hybrids which may belong to multiple categories. Biodegradable implants have more advantages than other A. D. Permana, M. N. Amir, H. Habibie Faculty of Pharmacy Hasanuddin University Makassar 90245, Indonesia E-mail: [email protected] J. Domínguez-Robles Department of Pharmacy and Pharmaceutical Technology Faculty of Pharmacy University of Seville Seville 41012, Spain E-mail: [email protected] N. Moreno-Castellanos CICTA Department of Basic Sciences Medicine School Health Faculty Universidad Industrial de Santander Cra 27 calle 9, Bucaramanga 680002, Colombia Adv. Healthcare Mater. 2025, 2403689 2403689 (1 of 16) © 2025 The Author(s). Advanced Healthcare Materials published by Wiley-VCH GmbH www.advancedsciencenews.com www.advhealthmat.de categories because the polymers used are biodegradable and biocompatible, which guarantees safe absorption or excretion from the body. Moreover, the surgical procedure to remove the implant is not required, which can also potentially improve patient compliance. Most biodegradable implantable devices are made from a mixture of polymer and drug. After implantation, the drug will be released with the degradation of the polymer.[4]The release rate can be affected by different factors, such as the degradation rate of different polymers in vivo, the drug loading, and the solubility and permeability of the drug.[4]By changing these factors, the implant can be designed with different predetermined release rates. The polymers used to fabricate biodegradable implantable devices are mostly thermoplastic aliphatic polyesters such as poly(lactic acid) (PLA), poly(glycolic acid) (PGA), poly(lactic-coglycolic acid) (PLGA), and poly(caprolactone) (PCL). Ester, amide, and anhydride bonds are labile bonds that can be hydrolyzed or degraded by enzymes contained in these polymers. These polymers have been extensively researched due to their advantageous traits including biodegradability, biocompatibility, and mechanical strength.[5] PCL has been one of the most commonly used biodegradable polymers in IDDS because of its biocompatibility, biodegradability, lack of toxicity, and comparatively inexpensive price.[6] It has become an FDA-approved polymer that can be used in the fabrication of medical devices.[7]For example, it has been used in a long-term contraceptive device Capronor containing levonorgestrel.[8]The degradation time of PCL is longer than other polymers such as PLA, PGA, and PLGA, ranging from several months to years. The degraded products can either be metabolized in the tricarboxylic acid cycle or excreted through the kidney.[9] A potential use of biodegradable IDDS is the treatment of chronic mental illness as a sustained pharmacological treatment is needed. There are many injectable formulations that provide sustained drug delivery for schizophrenia treatment.[10] However, the use of solid implants could be used to extend the duration of the treatment. To date, a subcutaneous implant named DLP-114 used to treat schizophrenia is in phase II clinical trials.[11]This IDDS is developed by Delpor (Delpor Inc., San Francisco, CA, USA) using ProzorTM technology.[12]Asmall cylindrical reservoir with membranes at both ends is used as a carrier to load risperidone (RIS), an antipsychotic drug, and some excipients, which can maintain the acidic environment and then improve the solubility of the drug.[12]The Phase I trials have shown that it can maintain constant plasma levels with little fluctuation for several months and distribute the medicine in a zeroorder manner.[11]Meanwhile, a subcutaneous RIS implant was developed and applied in adult patients with schizophrenia to assess its pharmacokinetics. This RIS implant achieved therapeutic levels within around 2 days after implantation and released for 6 months at a constant rate. Phase II and III studies are currently being conducted to assess the safety and efficacy of RIS implants.[13] In this study a simple method to develop reservoir-type implants combining a biodegradable rate-controlling membrane and a drug-containing core prepared by direct compression of pharma excipients and RIS. Direct compression is the method using compression to produce solid dosage forms from the blend of ingredients, which has advantages such as low cost and simple operation.[14]The aim of this work was to design and develop RIS implantable devices for use with PCL. PCL films and directly compressed RIS pellets were fabricated separately and then incorporated. Following these steps, a series of characterizations, in vitro release studies, and in vitro biocompatibility studies were performed to select the most appropriate formulation, which can provide sustained drug release and then be applied to the animal study. 2. Results and Discussion 2.1. Fabrication and Characterization of Directly Compressed RIS Pellets Direct compression is a manufacturing process that produces solid dosage forms from a mixture of active ingredients and excipients.[14]Due to the advantages of low cost, shorter processing time, and simple operation, as well as the absence of the need for heat and/or solvents, it is preferred as the fabrication method for tablets or pellets.[15]Besides, it is the method commonly used in the pharmaceutical industry, thus scaling it up will be simple.[16] In this case, a small punch and die system (3 mm) was used to prepare the directly compressed pellets using a laboratory hydraulic press and a force of 1 ton applied for 20 s. These conditions are like the ones described previously for directly compressed tablet manufacturing. As detailed in Table 1, formulation R100 could not be formed, as the 100% RIS composition of the formulation adhered to the die and could not be removed. For well-formed and consistently directly compressed pellet preparation, the use of suitable excipients is vital. In this case, poly(ethyelene glycol), with a molecular weight of 8000 Da (PEG 8000), and Hydroxypropyl 𝛽-cyclodextrin (HP-𝛽CD) were used. PEG 8000, plays several vital roles in pharmaceutical formulations including solubilizing agent, matrix former, lubricant, plasticizer, and vehicle for drug delivery, due to its biocompatibility, non-toxicity, and regulatory acceptance.[17]HP𝛽CD, a cyclic oligosaccharide derivative, is also widely used as an excipient in pharmaceutical formulations. Some of the functions of HP-𝛽CD as an excipient include a solubilizing agent, stabilizer, permeation enhancer, and compatibility agent.[18]As shown in Table 1, five types of pellets have been fabricated, and most of the resultant pellets were robust under visual inspection. In this case, formulations R40P60, R50P50, R60P40, and R60C40 were selected to evaluate the influence of drug loading and the type of excipient in the drug release kinetics. Directly compressed formulations have been evaluated before to prepare implants for sustained drug release.[19]In these studies, the excipients selected to prepare the directly compressed pellets were NaCl and poly(vinyl pyrrolidone). In these studies, the method proposed was similar to the one developed here. Drugs were combined with excipients and subsequently compressed. Alternatively, other authors tried a combined approach that requires the combination of the drug with a thermoplastic, such as poly(lactide-co-glycolide), via compression/thermal processing.[20]The direct compression method used in the present work is scalable and does not require the use of temperature that could lead to the degradation of the drug cargo during manufacturing. Adv. Healthcare Mater. 2025, 2403689 2403689 (2 of 16) © 2025 The Author(s). Advanced Healthcare Materials published by Wiley-VCH GmbH 21922659, 0, Downloaded from https://advanced.onlinelibrary.wiley.com/doi/10.1002/adhm.202403689 by Readcube (Labtiva Inc.), Wiley Online Library on [12/06/2025]. See the Terms and Conditions (https://onlinelibrary.wiley.com/terms-and-conditions) on Wiley Online Library for rules of use; OA articles are governed by the applicable Creative Commons License www.advancedsciencenews.com www.advhealthmat.de Table 1. Summary of appearance of directly compressed pellets containing RIS. Formulation Morphology Observation R40P60 •Well formed •Robust R50P50 •Well formed •Robust R60P40 •Well formed •Robust R75P25 •Formed with a rough surface •Part of the formulation sticked to the die, made it hard to remove intact R100 N/A •Did not form well •the formulation sticked to the die, made it hard to remove intact R60C40 •Well formed •Robust Adv. Healthcare Mater. 2025, 2403689 2403689 (3 of 16) © 2025 The Author(s). Advanced Healthcare Materials published by Wiley-VCH GmbH 21922659, 0, Downloaded from https://advanced.onlinelibrary.wiley.com/doi/10.1002/adhm.202403689 by Readcube (Labtiva Inc.), Wiley Online Library on [12/06/2025]. See the Terms and Conditions (https://onlinelibrary.wiley.com/terms-and-conditions) on Wiley Online Library for rules of use; OA articles are governed by the applicable Creative Commons License www.advancedsciencenews.com www.advhealthmat.de Table 2. Response values for pellets containing RIS. (means ±s.d., n=5). Formulation Hardness [N] R40P60 38.2 ±9.31 R50P50 30.6 ±6.35 R60P40 28.8 ±8.67 R60C40 30.8 ±6.42 The hardness of the resulting pellets was evaluated. The purpose of testing the hardness of pellets is to evaluate the maximum force needed to fracture a pellet. Hardness plays an important role in further assembly processes and in transit. According to the results listed in Table 2., all pellets with different compositions performed well, with no significant difference in hardness (p>0.05) observed. These hardness results are in line with previously reported directly compressed pellets prepared with pharmaceutical excipients such as microcrystalline cellulose.[21]Accordingly, the obtained pellets could be handled appropriately. Higher hardness values will make the pellet disaggregation extremely slow impacting the amount of drug releases from the implants.[22]This is an important parameter as there is a balance between release duration and the amount of drug release to achieve therapeutic outcomes. 2.2. Fabrication of PCL Films PCL-based membranes were prepared by coating a rotary metal rod using PCL solutions in dichloromethane (DCM) as described in the Experimental Section. Several cycles of casting were required to achieve suitable membrane thickness. To determine a suitable composition and concentration of PCL film, three different concentrations of PCL solution in DCM were used to fabricate the PCL films. As shown in Figure 1a, films fabricated with a 20% PCL solution were formed. It is worth noting that due to the thin thickness of the film, it is not sufficiently hard to resist the force needed to remove the film from the rod, resulting in a rough surface. This may potentially affect the integrity of the films. Figure 1b shows the films fabricated with a 30% PCL solution. A smooth surface and uniform thickness can be observed. Films made from a 40% PCL solution are shown in Figure 1c. The higher concentration led to a higher viscosity in this case, therefore they did not spread evenly across the surface of the metal rod, resulting in a rough and nonuniform surface of the films. The 30% PCL formulation was chosen to fabricate films with different thicknesses of coatings. The physical appearance of the films is shown in Figure 2All the films were formed well, with proper thickness and smooth and uniform surface. Similar methods have been used before to obtain tubular membranes/scaffolds. These methods normally use electrospinning or 3D-printing on the surface of a rod or a cylindrical structure to obtain tubular objects.[23]An alternative to this is the use of hot-melt extrusion but this technique normally requires the use of large amounts of polymer and therefore it is not ideal for the preparation of small prototypes.[24] Optical coherence tomography (OCT) is an imaging technology that evaluates biological tissue. However, it can be used to visualize other samples including pharmaceutical products or medical devices.[25]In this case, OCT is used to observe the cross-section of films and assembled implantable devices. Figure 2a,d,g,j shows the images of films with different layers of coatings, in which it can be seen that every film is uniform, and no gap can be observed. Different thicknesses can also be observed, as shown in Figure 3a; the thickness increased with the increase of layers of coatings in a linear way. The thickness of the membranes for 4, 5, 6, and 7 layers was: 120 ±9, 211 ±6, 300.67 ±8, and 412 ±10 μm respectively. The coating process was consistent, with a significant difference (p<0.05) observed in these four films with different layers of coatings. These results indicate that the method proposed can be easily adapted to modify membrane thickness and that the overall thickness of the membrane will be simple to predict as there is a linear relationship between the number of layers and the membrane thickness. The thickness of these membranes is in line with previously reported films used as rate-controlling membranes in implantable devices ranging between 70 and 400 μm.[19a,24,26] Figure 1. Physical appearance of films with different concentrations of PCL. a) 20% PCL in DCM b) 30% PCL in DCM c) 40% PCL in DCM. Adv. Healthcare Mater. 2025, 2403689 2403689 (4 of 16) © 2025 The Author(s). Advanced Healthcare Materials published by Wiley-VCH GmbH 21922659, 0, Downloaded from https://advanced.onlinelibrary.wiley.com/doi/10.1002/adhm.202403689 by Readcube (Labtiva Inc.), Wiley Online Library on [12/06/2025]. See the Terms and Conditions (https://onlinelibrary.wiley.com/terms-and-conditions) on Wiley Online Library for rules of use; OA articles are governed by the applicable Creative Commons License www.advancedsciencenews.com www.advhealthmat.de Figure 2. Physical appearance and OCT images of cross-section of 30% PCL films with different thicknesses. a–c) 4 coatings d–f) 5 coatings g–i) 6 coatings j–l) 7 coatings. The scale bar of OCT images is 1 mm. Figure 3. Thickness of membranes with different number of layers (a) (means ±s.d., n=5). Pore size distribution curves of membrane samples (b). Correlation between total surface area (TSA) and membrane thickness (c). Adv. Healthcare Mater. 2025, 2403689 2403689 (5 of 16) © 2025 The Author(s). Advanced Healthcare Materials published by Wiley-VCH GmbH 21922659, 0, Downloaded from https://advanced.onlinelibrary.wiley.com/doi/10.1002/adhm.202403689 by Readcube (Labtiva Inc.), Wiley Online Library on [12/06/2025]. See the Terms and Conditions (https://onlinelibrary.wiley.com/terms-and-conditions) on Wiley Online Library for rules of use; OA articles are governed by the applicable Creative Commons License www.advancedsciencenews.com www.advhealthmat.de Table 3. Textural properties of the polymeric membranes obtained by MIP. Sample TSA [m2g−1]TPV[cm 3g−1] 4 coatings 20.281 0.089 5 coatings 21.195 0.108 6 coatings 21.374 0.089 7 coatings 23.846 0.138 The porosity of the membranes was evaluated using MIP (Figure 3b). All the analyzed samples had a similar monomodal pore size distribution with a peak value centered around 1 μm. However, the addition of more layers in the 7 coatings membranes sample produced the most uniform shape with an important increase of the pore size distribution intensity, which indicates a larger number of pores with same size of around 1 μm. Therefore, the thickness increase of the 7 coatings membrane samples resulted in a larger number of pores of the same size (around 1 μm) than the rest of the membrane samples. The obtained pore distribution is in line with previously reported PCLbased membranes showing pores ranging between 1 and 2 μm. The outcomes shown in Table 3evidenced an increase of the membrane total surface area (TSA) associated to the general increase of pores population after increasing the number of coatings. These results also confirmed the general trend by which most of the membranes had higher values of the total pore volume (TPV) when the number of coatings was increased as can be seen in Figure 3c. Finally, the higher pore in volume is directly correlated with a higher porosity percentage, as reported by Stewartetal. [27]It is important to mention that this is a more consistent method for preparing rate-controlling membranes for implantable devices. Previous studies reported the use of this type of membrane, but they were wrapped around a drug-containing core and sealed.[28]This procedure is more time-consuming and requires extensive membrane sealing in the edges to avoid drug leakages. This is minimized if membranes are prepared in a tubular shape. Figure 4shows SEM images of the membranes with different coatings. The membranes exhibit pore sizes ranging between 2 and 8 μm in diameter. These results are consistent with the pore sizes obtained using MIP (Figure 3b). It is important to note that for some membranes, multiple small pores slightly distort the pore distribution. The pore size distribution is represented as the frequency of the pores. These small pores constitute a small portion of the total pore area. The results from MIP indicate that these small pores do not contribute to the differential intrusion (Figure 3b). Additionally, the frequency of pores ranging between 2and8μm is similar across all samples, regardless of the number of coatings. 2.3. Preparation of RIS Implantable Devices and In Vitro Drug Release Experiments The mini RIS implantable device was assembled from a RIS pellet and a PCL tube. The melting point of PCL is between 50 and 60 °C,[29]so the sealing process can be easily performed using heated pliers. By adjusting the number of pellets loaded into the PCL tube, the drug loading of this implantable device can be adjusted. Figure 5apresents a completed RIS implantable device. The OCT cross-section image of an implantable device is shown in Figure 5b. In this case, it shows the RIS pellet was right filled in the film. The outer part is clearer than the inner part due to the penetration of the laser. The in vitro release studies were carried out to further understand the release behavior of RIS in implantable devices. All groups were performed until all the drug was completely released from the implantable devices or until the study was stopped. Initially, in vitro, release studies were conducted with implantable devices containing different compositions of drug and excipient under the same thickness of films. The aim of this was to determine the influence of drug composition. The first step was to study the release of RIS from the uncoated pellets. As presented in Figure 6, these curves are all similar and release all the drugs in 20 days. When evaluating similarity and difference factors, if the value of F1is between 0 and 15, and the value of F2is between 50 and 100, then the two release profiles are considered similar.[30]As listed in Table 4, the results suggest that the release from R50P50 and R60P40 can be considered equivalent. RIS released from R40P60 and R60P40 cannot be considered equivalent according to F1and F2results. Finally, the results obtained for the comparison of R40P60 and R50P50 curves showed some discrepancies between F1and F2factors but considering how close the calculated values were to the threshold limit it can be assumed that the curves are equivalent. The comparison of F1and F2factors between R40P60 in 4 coatings and R50P50 in 4 coatings, R40P60 in 4 coatings and R60P40 in 4 coatings, and R50P50 in 4 coatings and R60P40 in 4 coatings can also validate above conclusion. Furthermore, by comparing the F1and F2factors between R40P60 and R40P60 in 4 coatings, R50P50 and R50P50 in 4 coatings, and R60P40 and R60P40 in 4 coatings, it can be concluded that they were different in each comparison, which proved the impact of PCL films on slowing release rate. When the pellets were coated with the tubular rate-controlling membrane, the RIS release rate was slowed down significantly as presented in Figure 6. It is important to note that these results suggest that the use of PCL-based membranes can sustain the release of RIS for up to 70 days. The release duration is proportional to the drug loading. Lower drug loadings resulted in a shorter release duration (≈40 days). When comparing the results of the RIS release from the pellets with the release from the coated implants F1and F2factors indicate that the release curves are different. The differences between the coated implants are more obvious and it can be established that the higher the drug loading the slower is the release. The comparison of the curves using F1and F2confirms these results. This may be because PEG is a hydrophilic polymer that could assist in the solubilization of hydrophobic drugs. Similar results have been observed previously for other hydrophobic drugs such as olanzapine, where PEG increased the solubility of the drug.[28b] Drug release from reservoir systems is influenced by various parameters and typically follows a diffusion-controlled mechanism.[31]For the drug to be released, water must permeate the membrane, partially dissolving the drug within it. The dissolved drug then diffuses through the membrane and is released, creating a concentration gradient between the implant core and its external environment.[31]Also, it is important to mention that the concentration gradient and, the release rate, will be proporAdv. Healthcare Mater. 2025, 2403689 2403689 (6 of 16) © 2025 The Author(s). Advanced Healthcare Materials published by Wiley-VCH GmbH 21922659, 0, Downloaded from https://advanced.onlinelibrary.wiley.com/doi/10.1002/adhm.202403689 by Readcube (Labtiva Inc.), Wiley Online Library on [12/06/2025]. See the Terms and Conditions (https://onlinelibrary.wiley.com/terms-and-conditions) on Wiley Online Library for rules of use; OA articles are governed by the applicable Creative Commons License www.advancedsciencenews.com www.advhealthmat.de Figure 4. SEM images of the membranes and pore size distribution obtained from SEM images. Scale bar: 100 μm. Adv. Healthcare Mater. 2025, 2403689 2403689 (7 of 16) © 2025 The Author(s). Advanced Healthcare Materials published by Wiley-VCH GmbH 21922659, 0, Downloaded from https://advanced.onlinelibrary.wiley.com/doi/10.1002/adhm.202403689 by Readcube (Labtiva Inc.), Wiley Online Library on [12/06/2025]. See the Terms and Conditions (https://onlinelibrary.wiley.com/terms-and-conditions) on Wiley Online Library for rules of use; OA articles are governed by the applicable Creative Commons License www.advancedsciencenews.com www.advhealthmat.de Figure 5. a) Image of an accomplished RIS implantable device. b) Representative OCT image of the cross-section of the implantable device with an R60P40 formulated pellet in 4 coatings film. The scale bar of OCT images is 1 mm. tional to drug solubility inside the implant core.[31]If the concentration gradient remains constant, the system exhibits zeroorder release. To sustain this gradient, the core must contain a sufficient quantity of drug to maintain a saturated drug solution. This includes both dissolved drug and undissolved drug reserves. As drug molecules are released, the undissolved drug replenishes the dissolved portion, preserving saturation conditions within the device. However, as the drug load diminishes over time, the system eventually fails to maintain saturation, causing the release rate to slow. This depletion explains the deviation from linearity observed in Figure 6during the later stages of release. Another in vitro release study was performed with implantable devices using directly compressed pellets with the same formulation but coated with different layers of film to evaluate the impact of film thickness. As shown in Figures 7aand 8a, the release of both RIS-PEG pellets and RIS-HP𝛽CD pellets occurred rapidly, in 22 and 24 days, respectively. They cannot be compared with RIS pellets because pure RIS could not form a directly compressed pellet on its own. In Figure 7., all these devices showed sustained release of drugs, and no burst release was observed initially. The results indicated that the release behavior was significantly influenced by the thickness of the films. With an increase in film thickness, the release rate became slower, ranging from 70 to 170 days depending on the thickness of the film. The release rate correlates with membrane thickness in a linear way as shown in Figure 6d. These implantable devices exhibited a linear drug release rate during their release time. Thicker membranes will lead to a longer diffusion path for the dissolved drug. This leads to slower release kinetics as demonstrated previously.[26a,c,32] Figure 6. In vitro cumulative release profile of PEG-RIS pellets. Release studies were carried out in triplicate (n=3) and all results were expressed as means ±s.d. Figure 8shows the cumulative release profile of implantable devices made from R60C40 pellets and different layers of films. All four groups showed faster release in the first 10 days, after which the release rate became slower, exhibiting a linear drug release rate until all the drugs were released from the implantable devices. The time ranged from 100 to 180 days depending on the thickness of the film. The results suggested that the use of HP𝛽CD can be used to extend RIS release. However, as mentioned earlier, these curves showed a biphasic release pattern. The initial release rate is faster due to the presence of HP-𝛽CD in the pellet. Cyclodextrins and their derivatives have been extensively Table 4. Difference and similarity factors calculated of each release profile of implantable devices. Curve 1 Curve 2 F1 F2 R40P60 R50P50 16 53 R40P60 R60P40 26 43 R50P50 R60P40 11 62 R40P60 R40P60 in 4 coatings 142 18 R50P50 R50P50 in 4 coatings 151 19 R60P40 R60P40 in 4 coatings 63 20 R40P60 in 4 coatings R50P50 in 4 coatings 17 51 R40P60 in 4 coatings R60P40 in 4 coatings 26 42 R50P50 in 4 coatings R60P40 in 4 coatings 9 60 R60P40 in 4 coatings R60P40 in 5 coatings 23 38 R60P40 in 4 coatings R60P40 in 6 coatings 54 21 R60P40 in 4 coatings R60P40 in 7 coatings 58 20 R60P40 in 5 coatings R60P40 in 6 coatings 41 33 R60P40 in 5 coatings R60P40 in 7 coatings 47 30 R60P40 in 6 coatings R60P40 in 7 coatings 12 67 R60C40 in 4 coatings R60C40 in 5 coatings 36 37 R60C40 in 4 coatings R60C40 in 6 coatings 48 31 R60C40 in 4 coatings R60C40 in 7 coatings 48 31 R60C40 in 5 coatings R60C40 in 6 coatings 17 63 R60C40 in 5 coatings R60C40 in 7 coatings 19 60 R60C40 in 6 coatings R60C40 in 7 coatings 3 90 R60P40 in 4 coatings R60C40 in 4 coatings 29 32 R60P40 in 5 coatings R60C40 in 5 coatings 42 33 R60P40 in 6 coatings R60C40 in 6 coatings 19 60 R60P40 in 7 coatings R60C40 in 7 coatings 21 63 Adv. Healthcare Mater. 2025, 2403689 2403689 (8 of 16) © 2025 The Author(s). Advanced Healthcare Materials published by Wiley-VCH GmbH 21922659, 0, Downloaded from https://advanced.onlinelibrary.wiley.com/doi/10.1002/adhm.202403689 by Readcube (Labtiva Inc.), Wiley Online Library on [12/06/2025]. See the Terms and Conditions (https://onlinelibrary.wiley.com/terms-and-conditions) on Wiley Online Library for rules of use; OA articles are governed by the applicable Creative Commons License www.advancedsciencenews.com www.advhealthmat.de Figure 7. In vitro release profile of R60P40 pellet in different thicknesses of PCL tubes. a) Cumulative RIS released in percentage. b) Cumulative RIS released in milligrams. c) RIS released in mg per day. Release studies were carried out in triplicate (n=3) and all results were expressed as means ±s.d. d) Correlation between membrane thickness and RIS release rate. Figure 8. In vitro release profile of R60C40 pellet in different thicknesses of PCL tubes. a) Cumulative RIS released in percentage. b) Cumulative RIS released in milligrams. c) RIS released in mg per day. Release studies were carried out in triplicate (n=3) and all results were expressed as means ±s.d. Adv. Healthcare Mater. 2025, 2403689 2403689 (9 of 16) © 2025 The Author(s). Advanced Healthcare Materials published by Wiley-VCH GmbH 21922659, 0, Downloaded from https://advanced.onlinelibrary.wiley.com/doi/10.1002/adhm.202403689 by Readcube (Labtiva Inc.), Wiley Online Library on [12/06/2025]. See the Terms and Conditions (https://onlinelibrary.wiley.com/terms-and-conditions) on Wiley Online Library for rules of use; OA articles are governed by the applicable Creative Commons License www.advancedsciencenews.com www.advhealthmat.de [26] a) A. Korelidou, J. Domínguez-Robles, E. R. Magill, M. Eleftheriadou, V. A. Cornelius, R. F. Donnelly, A. Margariti, E. Larrañeta, Biomater. 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