Reservoir-type intranasal implants for sustained release of risperidone: A potential alternative for long-term treatment of schizophrenia

Journal of Drug Delivery Science and Technology 99 (2024) 105973 Available online 15 July 2024 1773-2247/© 2024 The Authors. Published by Elsevier B.V. This is an open access article under the CC BY license (http://creativecommons.org/licenses/by/4.0/). Reservoir-type intranasal implants for sustained release of risperidone: A potential alternative for long-term treatment of schizophrenia Emilia Utomo a , Linlin Li a , Jiaqi Gao a , Qonita Kurnia Anjani a , Camila J. Picco a , Natalia Moreno-Castellanos b , Ryan F. Donnelly a , Juan Domínguez-Robles a,c,** , Eneko Larra˜ neta a,* a School of Pharmacy, Queen’s University Belfast, 97 Lisburn Road, Belfast, BT9 7BL, UK b CICTA, Department of Basic Sciences, Medicine School, Health Faculty, Universidad Industrial de Santander, Cra 27 calle 9, Bucaramanga, 680002, Colombia c Department of Pharmacy and Pharmaceutical Technology, Faculty of Pharmacy, University of Seville, 41012, Seville, Spain ARTICLE INFO Keywords: Intranasal Long-acting drug delivery Reservoir-type implant Risperidone Membrane ABSTRACT To overcome the challenges of the blood-brain barrier for drug delivery to the central nervous system (CNS), intranasal implants were developed to improve the management of CNS conditions, such as schizophrenia. In the present work, we developed and characterised a drug-containing implant consisting of two parts: a core layer made from risperidone (RIS) and water-soluble polymers, including poly(vinylpyrrolidone) (PVP) and poly (ethylene glycol) (PEG), and a coating layer made of poly(caprolactone) (PCL) membrane. The obtained implants, where the core layer contained 75 % w/w risperidone, were characterised using several techniques: scanning electron microscopy (SEM), thermogravimetric analysis (TGA), differential scanning calorimetry (DSC), and attenuated total reflectance-Fourier transform infrared (ATR-FTIR). Moreover, the in vitro release profile of RIS was studied, showing that the PCL membrane could extend the release of RIS from 2 days up to 100 days. The in vitro release profile of the PCL-coated implant exhibited a linear release over the first 10 days, followed by a slower release rate that reached another linear phase up to 40 days. Subsequently, the drug release rates progressively slowed down. Finally, the results of in vitro biocompatibility studies indicated that the intranasal implants were biocompatible and non-cytotoxic. These findings suggest that the implants prepared in this work have the potential to provide long-acting drug delivery for targeting the brain. 1. Introduction The presence of the blood-brain barrier limits the delivery of drugs into the central nervous system (CNS) [1,2]. Accordingly, the development of systems capable of bypassing this barrier to enhance drug administration into the brain has attracted attention over the past decades [3,4]. A potential alternative route of administration to maximise drug concentrations in the CNS is the intranasal route [5–10]. Several studies have described the efficacy of this route in enhancing brain drug levels. It is important to note that many conditions affecting the CNS are chronic and require regular pharmacological treatment. Normally, these drugs are administered via the oral route. Despite being the preferred route for drug administration, it presents challenges for the management of chronic conditions, as the patient needs to adhere to the treatment [11,12]. A potential alternative to solve this issue is long-acting drug delivery systems [13–17]. These systems are typically injected or implanted into the patient, providing sustained drug dosage after a single administration [13–18]. Long-acting drug delivery systems have been prepared in a wide range of shapes and materials including solid subcutaneous implants [14,19,20], injectable suspensions [21,22] or in situ forming implants/gels [23–26]. Their application is invasive, unlike the oral route, but their benefit in terms of adherence to treatment can outweigh this limitation [17,27]. A novel strategy that combines the advantages of long-acting drug delivery systems and the intranasal route of administration is drugeluting intranasal implants [28,29]. These systems can be implanted in the patient’s nasal cavity to provide sustained drug release. It is * Corresponding author. ** Corresponding author. School of Pharmacy, Queen’s University Belfast, 97 Lisburn Road, Belfast, BT9 7BL, UK. E-mail addresses: [email protected] (J. Domínguez-Robles), [email protected] (E. Larra˜ neta). Contents lists available at ScienceDirect Journal of Drug Delivery Science and Technology journal homepage: www.elsevier.com/locate/jddst https://doi.org/10.1016/j.jddst.2024.105973 Received 11 June 2024; Received in revised form 8 July 2024; Accepted 11 July 2024 Journal of Drug Delivery Science and Technology 99 (2024) 105973 2 important to mention that drug eluting intranasal implants have been described extensively for treatment of nasal polyps [30–32]. However, they are designed for localised drug delivery. Intranasal implants aimed for brain drug delivery are a novel approach. Such devices can be used to improve the management of conditions affecting the CNS, such as schizophrenia. Schizophrenia is a chronic condition that severely affects patients, distorting perception and their way of thinking [33]. The treatment of this condition relies heavily on pharmacological interventions. As mentioned earlier, adherence to treatment can be challenging in chronic conditions, and schizophrenia is no exception [34–36]. It has been reported that up to 75 % of patients suffering from this condition experience treatment adherence issues [34]. This can lead to relapse, decreasing the quality of life for patients while increasing healthcare-associated costs [36–38]. Additionally, drugs used to treat schizophrenia, such as risperidone (RIS) [39], have side effects [40–43] when are administered orally, affecting the quality of life of patients and potentially contributing to poor adherence [44,45]. Drug-eluting intranasal implants can be used for schizophrenia treatment as they provide sustained drug release, maximising the amount of drug administered into the brain while reducing systemic exposure and thereby reducing side effects. This work will describe the development of reservoir-type implantable intranasal implants for schizophrenia treatment. For this purpose, a drug-containing implant was prepared using water-soluble polymers and RIS. It was subsequently encapsulated inside a biocompatible ratecontrolling membrane made of poly(caprolactone) (PCL). The implants were extensively characterised, employing a comprehensive array of physicochemical techniques. Moreover, in vitro RIS release and biocompatibility studies were performed. 2. Materials and methods 2.1. Materials RIS was purchased from Enke Pharma-Tech Co., ltd (Cangzhou, China). Poly(ethylene glycol) (PEG) (Mw =600 Da) was purchased from Tokyo Chemical Industries (Tokyo, Japan). Poly(vinyl pyrrolidone) (PVP) (360 kDa) was purchased to Sigma-Aldrich (Steinheim, Germany). High (CAPA ®6506; Mw =50 kDa) and low (CAPA®2054; Mw =550 Da) molecular weight PCL were kindly donated by Ingevity (North Charleston, South Carolina, U.S.A). Phosphate saline buffered saline (PBS) (pH 7.4) tablets were obtained from VWR International (Leuven, Belgium. Dichloromethane (DCM), acetonitrile and methanol were purchased from Sigma-Aldrich (Dorset, UK) 2.2. Preparation of reservoir type implantable devices The reservoir-type implants consisted of two parts: the core implants prepared using hydrophilic polymers (PVP and PEG), and the coating layer made of PCL film. The preparation of this type of implant is illustrated in Fig. 1A. The core implants containing 75 % w/w of RIS were fabricated using a solvent casting method, utilising PVP (15 % w/ w) and PEG (10 % w/w) as the main polymers. Water was added in a 1:1 ratio to the solid content. The drug and polymers were mixed using a SpeedMixer™DAC 150.1 FVZ-K (Hauschild &Co. KG, Hamm, Germany) at 3000 rpm for 3 min. The mixture was cast using a syringe inside a silicone mould prepared as described previously [28]. Subsequently the mould containing the liquid formulation was centrifuged at 3500 rpm for 30 min to spread the formulation through the moulds. The implants were placed at 37 ◦C for 24 h to dry. The coating layer of PCL film was prepared using a combination of Fig. 1. Diagrams showing the protocols followed for: core implant preparation (A) and reservoir type assembly (B). Diagram showing the setup used for permeation experiments (C). E. Utomo et al. Journal of Drug Delivery Science and Technology 99 (2024) 105973 3 PCL 6506 and PCL 2504 in a 1:1 ratio. This ratio was selected based on previous experiments [46]. This mixture of PCL was subsequently dissolved in DCM to obtain a concentration of 20 % w/w. The PCL solution was then cast in a Petri dish and left at room temperature to allow DCM evaporation. Subsequently, the coating layer was cut based on the shape of the implant. The procedure used to prepare the reservoir-type implants can be seen in Fig. 1B. The core implant was placed between two coating layers, and a heated plier was used to seal the implant. 2.3. RIS loaded core implant characterisation A Hitachi TM3030 tabletop scanning electron microscope (SEM) (Hitachi, Tokyo, Japan) was utilized to examine the surface morphology of RIS-loaded implants. To assess interactions between RIS and the polymers, Fourier transform infrared (FTIR), differential scanning calorimetry (DSC) and thermogravimetric analysis (TGA) and were conducted. For DSC, samples weighing between 5 and 10 mg were heated in sealed aluminium pans from 25 to 200 ◦C at a rate of 10 ◦C/min under a nitrogen atmosphere using a DSC Q20 (TA Instruments, New Castle, USA). For TGA, small implant pieces of similar weight were heated to 500 ◦C at a rate of 10 ◦C/min under nitrogen flow using a TGA Q500 (TA Instruments, New Castle, USA). For DSC and TGA data analysis was performed with TA Universal Analysis software (TA Instruments, New Castle, DE, USA). On the other hand, FTIR analysis was performed using an Accutrac FT/IR- 4100 series spectrometer (Jasco, Essex, UK) equipped with a MIRacle™diamond attenuated total reflection (ATR) accessory was used. IR spectra were recorded between 600 and 4000 cm −1 , as an average of 32 scans and 4.0 cm −1 of resolution. The RIS content in the implants was determined by dissolving the implants and measuring the drug amount in the solution as described previously [28]. To quantify the amount of drug in solution High-performance liquid chromatography (HPLC) was used. The HPLC conditions are described in the following section. 2.4. HPLC method for RIS quantification RIS quantification was conducted using an HPLC system (Agilent Technologies, Stockport, UK). The column used was a Phenomenex® SphereClone™C18 ODS column (length: 150 mm length; ID: 4.60; Particle size: 5 μ m). The mobile phase consisted of an aqueous phase and an organic phase in a ratio of 15:85 % v/v. The aqueous phase contained sodium dihydrogen phosphate buffer (10 mM) and 1 % v/v of triethylamine (TEA). The pH of the aqueous phase was adjusted to 7.2 using orthophosphoric acid. The organic phase was a mixture of methanol and acetonitrile in a 75:25 % v/v ratio. The injection volume was set at 20 μ L with a flow rate of 1.2 mL/min, and detection was carried out using a UV detector set at 235 nm. All the HPLC experiments were carried out at room temperature. Chromatogram analysis was conducted using Agilent ChemStation®Software B.02.01. The method was previously validated according to ICH guidelines [28]. 2.5. In vitro RIS release The release rate of RIS from the implants was assessed by quantifying the drug amount released at specified time points. For this purpose, reservoir implants and the core section of the implants without membrane were tested. Each implant was placed in a flask filled with 250 mL of PBS (pH 6.5). Subsequently, the flasks were introduced in a shaking incubator at 37 ◦C, mimicking intranasal conditions [47], and agitated at 40 rpm. To prepare the release media the pH of a PBS (pH 7.4) solution was adjusted with phosphoric acid (85 %) to achieve the required pH value of 6.5. The release experiments were performed under sink conditions, with RIS solubility in PBS (pH 6.5) being 1.24 ±0.06 mg/mL. For the experiment of coated implants, at each time point, a 1 mL aliquot of the release medium was sampled and replaced with an equal volume of fresh PBS (pH 6.5) to maintain sink conditions. On the other hand, the release from the implants without membrane, the release was carried out over 48 h without replacing the release media as the implants were soluble and could not be retrieved from the media after hydration. Moreover, to maintain sink conditions, only a fragment of the implant was used in this case. The collected samples were quantified using the HPLC method described previously (section 2.4). If necessary, samples were diluted to ensure the concentration fell within the range of the linear calibration curve. 2.6. RIS permeation experiment In vitro permeation studies were conducted to evaluate the release and diffusion of RIS from the implants, simulating drug permeation from the nasal cavity to the brain as depicted in Fig. 1C. The experimental setup utilized side-by-side horizontal diffusion cells, each consisting of donor and receptor compartments with water jackets for temperature control. A cellulose membrane (Spectrum™Spectra/Por™2 RC Dialysis Cellulose Membrane Tubing, 12–14 kDa MWCO, Fisher Scientific, Vantaa, Finland), pre-soaked in PBS (pH 6.5) for 30 min, was secured between the compartments to prevent leakage and evaporation. Both compartments were filled with an equal volume of PBS (pH 6.5), 3 mL, and maintained at 37 ◦C under continuous stirring at 600 rpm. At specified intervals, the entire solution from the receptor compartment was collected and replaced with fresh PBS (pH 6.5) to avoid saturation. The collected samples were then analysed with the validated HPLC method described earlier. Dilutions with PBS (pH 6.5) were made when necessary to ensure sample concentrations fell within the calibration curve range. The experiment lasted for 7 days, with samples collected every 24 h. 2.7. In vitro cytotoxicity studies In this study, human dermal fibroblasts (HDF, ATCC PCS-201-010) were employed as model cells. HDF were cultured in Dulbecco’s Modified Eagle Medium (DMEM). This medium was supplemented with 10 % fetal bovine serum, penicillin-streptomycin (10,000 units/10 mg/mL), and 200 mM L-glutamine. The cells were maintained at 37 ◦C in a 5 % CO 2 atmosphere. Subculturing was done when the cells reached 80–90 % confluency. The test samples were the rate controlling membranes prepared using high and low molecular weight PCL. Films were cut into discs (diameter: 13 mm; thickness: 0.1 mm) before the cytotoxicity studies. Samples were sterilised through 30 min of UV-C irradiation. Cells were then seeded on top of the polymeric samples at a density of 2 ×10 4 cells/cm 2 using 12-well plates. Subsequently, samples were incubated for 7 days. Post incubation, the samples were subjected to various assays including cell viability, metabolic activity, lactate dehydrogenase (LDH) release, and cell proliferation tests. The cell viability test was conducted using a conventional MTT assay, adapted from a previous study [48]. In brief, the cell-seeded samples were rinsed using PBS to remove residual DMEM. Subsequently, MTT solution in DMEM was added to each well of a 96-well plate followed by an incubation step for 12 h at 37 ◦C. The incubation was carried out in presence of 5 % CO 2 to allow formazan crystal formation. The medium was subsequently removed, and the samples were washed using dimethyl sulfoxide (DMSO). The formazan crystals were dissolved in the DMSO. Finally, the absorbance of this solution was measured at 580 nm using a Synergy H1 microplate reader (Agilent Technologies, Santa Clara, USA). Untreated HDF cells served as the negative control, while 1 % Triton X-100 was used as the positive control. The relative metabolic activity of the cells on the biomaterials was assessed using an Alamar Blue assay. To start, the cell-seeded samples were washed using PBS to eliminate remaining rests of DMEM. Then, 15 μ L of Alamar Blue dye containing 100 μ L of phenol red-free DMEM was added to each well of a 96-well plate. Then, the plates were incubated at E. Utomo et al. Journal of Drug Delivery Science and Technology 99 (2024) 105973 4 37 ◦C with 5 % CO 2 for 4 h. Fluorescence was measured using a plate reader with excitation at 550 nm and emission at 590 nm. HDF cells seeded without test materials served as the control. The LDH assay was employed to quantify the extent of HDF cell membrane rupture. Samples were washed with PBS and the cells were detached from the samples. Then, the detached cells were re-suspended in DMEM. A 50 μ L aliquot of cell suspension was treated with 50 μ L of PBS containing a 2 % (v/v) of Triton X-100 for 30 min at 37 ◦C. Following this, 100 μ L of the lysed cell solution was taken from each well and then combined with 100 μ L of LDH substrate in a 96-well plate. Finally, the fluorescence of the samples was measured at 630 nm using an excitation wavelength of 490 nm. Cell proliferation was assessed using a DNA content assay kit. The DNA content in cells attached to the PCL-based films was determined using Quant-iT™PicoGreen®dsDNA Reagent Kit following the instructions provided by the manufacturer. In brief, samples were rinsed three times with PBS and then submerged in 1 mL of lysis buffer that contained 10 mM Tris (pH 8), 1 mM EDTA, and 0.2 % (v/v) of Triton X- 100. DNA was released by vortexing the samples for 10 s every 5 min over a 30-min period, while keeping them on ice. After thawing on ice and homogenizing for 10–15 min, the samples were combined with 100 μ L of the DNA-binding fluorescent dye solution. Subsequently, they were incubated for 30 min at 37 ◦C with 5 % CO 2 . Finally, fluorescence intensity of these solutions was measured at 520 nm using an excitation wavelength of 480. All the results obtained for the cytotoxicity studies were compared using an unpaired t-test. Values of p lower than 0.05 indicated a statistically significant difference between tested groups. 3. Results and discussion Implants were prepared using a “flower-like”design so they can be folded as showed in Fig. 2A. Once they are folded, they can be inserted through an endoscope and placed inside the nasal cavity. The dimensions of the implant have been adapted from a similar product designed for localised delivery of corticosteroids to prevent nasal polyp recurrence after sinus surgery [49]. Intranasal implants were prepared using a solvent casting method to create the core section of the implant containing high RIS content (75 % w/w). Fig. 2B shows an image of the core section of the implant. The core section of the implant is water-soluble; therefore, to slow down drug release from intranasal implants, a rate-controlling membrane coating should be added. For this work, a PCL-based membrane was selected, combining two different molecular weight PCLs as reported earlier [46]. This type of membrane presents a defined porosity and pore diameter of 20 % and 1–2 μ m, respectively. Due to this structure, these membranes have been used successfully to sustain the delivery of hydrophilic compounds for up to 200 days [46]. Accordingly, soluble RIS intranasal implants were coated with the aforementioned membrane, as can be seen in Fig. 2C. Prior to evaluating the performance of the intranasal implants, the drug-containing core implant was characterised. SEM micrographs showed that there were no obvious drug particles or crystals on the surface of the implants, suggesting that the drug is well mixed with the PVP/PEG matrix. However, it is important to note that RIS does not have high water solubility, and accordingly, the resulting slurry used to prepare the implants will contain pure RIS particles. This is evident from the colour of the implants, as they displayed a white colour due to the drug being in suspension within the structure. Nevertheless, due to the high-speed mixing process, the drug is expected to be distributed throughout the implant. Also, the presence of PEG will contribute to increase drug solubility within the polymer matrix as described previously [20,50]. To confirm homogeneity of drug within the implant the drug content in the implant was evaluated. It was found that the drug content was 74.3 ±0.9 %, aligning with the initial 75 % drug loading in the formulation used to cast the implant. The amount of RIS loaded per implant was measured to be 240 ±7.5 mg. The use of drug/polymer slurries for the preparation of solid objects have been reported before, especially for 3D-printing applications where the viscosity of the slurry allowed to add layers of material to form a 3D object. This approach has been used with aqueous based slurries to prepare oral dosage forms and parenteral devices [51–54]. On the other hand, slurries prepared using organic solvents have been described too for the preparation of oral dosage forms or subcutaneous implants [20, 55]. In these cases, the drug was homogeneously dispersed within the matrix, but drug crystals were found in the formulation. Fig. 2. Core implant dimensions diagram (A). Images of RIS-loaded core implant (B) and coated implant (C). SEM micrographs of RIS-loaded core implant (D). E. Utomo et al. Journal of Drug Delivery Science and Technology 99 (2024) 105973 5 DSC results confirmed that the drug loaded into the implant remained in its crystalline form, as the melting point of RIS can be clearly seen at around 170 ◦C (Fig. 3A). This is the only meaningful peak, as PVP and PEG did not show any obvious transition apart from the dehydration peak for PVP at around 100 ◦C. Interestingly, the peak obtained in the implant DSC was broader and less intense than that of the pure drug. Additionally, the melting point was slightly lower. These results indicate that the crystallinity of the drug was affected when combined with PVP and PEG. A fraction of the initial RIS loading will be soluble in water and will interact with PVP and PEG in solution. PVP is a well-known polymer that forms interactions with hydrophilic and hydrophobic drugs [56], so it is not surprising that the drug shows some interactions with the polymers. These results are similar to the ones obtained previously when combining RIS with biodegradable polymers, PCL and poly(lactic-co-glycolic acid) (PLGA) [28]. Also equivalent behaviour was found for other drugs, such as paracetamol or olanzapine, loaded in polymeric formulations prepared using aqueous slurries [51, 54]. The DSC results can be confirmed by TGA (Fig. 3B). Pure RIS started to degrade at around 240 ◦C. However, the drug loaded into the core implants started the degradation process at slightly higher temperatures of around 250 ◦C. Interestingly, the degradation of pure PEG starts at lower temperatures than the core implant. Accordingly, the combination of the drug with these two polymers suggests the onset of non-covalent interactions, leading to the stabilisation of these molecules. FTIR was used to characterise the core implant. Fig. 3C shows the FTIR spectra of RIS. The characteristic peaks for the molecule can be seen clearly: bands found around 3060 and 1640 cm −1 can be attributed to the stretching of the aromatic C–H bonds and the carbonyl group of the δ-lactam ring, respectively. Additionally, the FTIR of the pure drug shows aromatic C–C stretching at around 1600 and 1530 cm −1 . Weaker peaks around 1400 cm −1 , assigned to C–N and C–O angular deformations of the oxazole ring, were also detected. C–N stretching bands for the oxazole ring were observed around 1350 cm −1 , while the stretching of the tertiary amine of the piperidine ring was observed at around 1190 cm −1 . Finally, the aryl fluoride group of the RIS molecule can be observed at around 1130 cm −1 . On the other hand, the FTIR spectra of PVP show characteristic peaks at around 3450 cm −1 for O–H stretching, around 2950 cm −1 for C–H stretching, and around 1650 cm −1 for C – – O stretching. Also, the characteristic band for the C – – N of the pyridine ring can be observed at around 1500 cm −1 . The spectra of pure PEG show bands at around 2890, 1350, and 1100 cm −1 attributed to C–H stretching, C–H bending, and C–O stretching vibrations, respectively. The FTIR spectra of the core implant show the characteristic peaks of RIS, PVP, and PEG. During the formulation, no chemical bonds were formed; therefore, there were no new peaks in the FTIR spectra of the formulations. Thermal analysis results suggested some interactions between RIS and the polymers. This can be observed in Fig. 3D, which presents a magnified version of the FTIR spectra between 1700 and 1500 cm −1 . In this figure, it can be seen that the RIS carbonyl peak overlaps with the carbonyl peak for PVP, but there is a subtle peak displacement to higher wavenumbers. Similar behaviour can be observed for the aromatic C–H stretching band at around 1530 cm −1 . This behaviour has been reported for RIS complexes with cyclodextrins [57–59]. Also, the FTIR of the core implant presented a double peak at around 1100 cm −1 that would correspond to the overlapping of characteristics peaks of RIS and the C–O stretching vibration of the PEG. After characterising the core implant, the next step was to evaluate RIS release from the resulting implants. Fig. 4A shows that the release from the PVA/PEG/RIS core implant material is fast, with more than 50 % of the drug cargo released after 2 h and the release completed after 24 h. As anticipated, these implants require a coating to slow down drug release to provide sustained drug release over prolonged periods. Therefore, coated implants were tested. Coated implants showed slower drug release kinetics, providing RIS release for up to 100 days (Fig. 4B). These implants did not show an initial drug burst release, as the PCL- Fig. 3. DSC (A), TGA (B) and FTIR (C) curves for pure RIS, PVP, PEG and RIS-containing core implant. FTIR of pure RIS, PVP and core implant between 1700 and 1500 cm −1 (D). E. Utomo et al. Journal of Drug Delivery Science and Technology 99 (2024) 105973 6 based membrane controls drug release. PCL and poly(lactic acid)-based membranes have been used before to sustain drug release [46,60,61]. Semipermeable membranes are capable of providing zero-order release [62], as the drug concentration gradient depends on drug solubility within the implant. If a supersaturated system is achieved, the system can provide zero-order release as fluid permeates inside the membrane, subsequently dissolving the drug in the fluid to permeate through the membrane. If the implant contains a supersaturated drug suspension, a constant drug release can be achieved. In this case, a linear release can be observed over the first 10 days, while subsequently, the release rates slow down to reach another linear release phase up until 40 days. In the following days, the drug release rates slow down progressively. This can be explained by the presence of PVP and PEG in the implant, which contribute to increasing RIS solubility during the initial phases of the release [28]. Initially, RIS release rates are higher as the solubility of the drug inside the implant is higher when PVP/PEG are present. These two polymers have high water solubility and, unlike RIS, will dissolve faster and therefore will be released during the first days of the experiment. After the polymers are released, the release is proportional to RIS solubility on its own inside the device. The final change in the release rate is mainly due to the depletion of RIS inside the implant. As mentioned earlier, if the concentration of drug dissolved inside the implant decreases, this will lead to a decrease in the release rates. Fig. 4C shows the daily release rate, highlighting the three main stages of the release from the implants. It is also important to mention that the dose obtained for a single implant is therapeutically relevant, as the required RIS dose for most patients ranges between 1 and 2 mg per day orally. Considering the oral bioavailability of the drug, around 70 % [63], the required dose of RIS ranges between 700 μ g and 1.4 mg. The target site of action for RIS is the CNS and considering that intranasal drug delivery shows potential to achieve higher drug concentrations within the brain while reducing systemic drug exposure, the required doses could be even lower. This suggests that implant dimensions can be reduced to load less drug inside. However, the presented results indicate that the proposed intranasal implants can provide sustained drug release of therapeutically relevant doses in vitro. We have previously reported the use of intranasal implants for RIS delivery through monolithic implants [28,29]. These implants were prepared using biodegradable polymers, including PCL and PLGA. All these implants showed some initial burst release due to the diffusion of the drug loaded on the surface. Additionally, these systems provided faster drug release, showing release profiles of up to 28 days after in vitro testing. PLGA showed longer release profiles, but after 28 days, the release rate was extremely slow. Accordingly, coated implants manufactured in this work present advantages over the previously mentioned monolithic type implants. Intranasal implants should release the drug cargo into the nasal mucosa, and subsequently, the drug should permeate through the tissue to reach the brain. To simulate the permeation of RIS through nasal mucosa over several days, a cellulose membrane was used. The main reason for choosing this model is that it is not possible to perform the experiment over prolonged periods using biological tissue, as it will decompose and degrade, limiting the results obtained. Additionally, the use of cellulose membranes to mimic nasal mucosa has been previously described by other authors. Karasulu et al. compared the permeability of indomethacin across different synthetic membranes (including cellulose and nylon) and compared the results with drug permeation across excised sheep nasal mucosa [64]. The results suggested that cellulose membranes provided the closest permeation profiles to those obtained for the excised tissue. In a different study, Bartos et al. showed similar results [65]. Accordingly, cellulose membranes were used to evaluate RIS permeation in a side-by-side Franz cell setup. The results obtained can be seen in Fig. 4D. The permeation profile followed a linear pattern as it is proportional to the drug concentration in the donor compartment. The permeation is proportional to the release profile, as seen in Fig. 4D. However, permeation across the simulated nasal mucosa is slower. The main reason for this is that the volume of fluid present in the Franz cell setup is smaller, and therefore, the solution could become saturated more quickly. This will maintain a constant drug administration. Additionally, this suggests that the release duration might be even slower in a real scenario, as the amount of fluid present in the nasal mucosa will be limited. For in vivo studies to ensure that enough fluid permeates inside the implant to facilitate RIS dissolution and release, implants might need to be placed in a vial with fluid prior to the implantation process. In this case, fluid will permeate inside and start the release process. Therefore, Fig. 4. RIS release profiles from uncoated RIS-loaded PVP/PEG materials (A) and coated implants (B). RIS daily release rate as a function of time (C). RIS permeation/release profiles over a period of 7 days (D). (means +S.D., n ≥3). E. Utomo et al. Journal of Drug Delivery Science and Technology 99 (2024) 105973 7 once the implant is applied, it will already contain fluid, ensuring that it will start working immediately, avoiding lag periods. This is not a new strategy as it has been described before for other reservoir-type implantable devices [66]. These results align with the results obtained previously for the permeation of RIS released from monolithic implantable devices [28]. After evaluating the performance of the resulting implants, the final step was to assess the biocompatibility of the PCL-based membranes. PVA and PVP are FDA-approved polymers and they have been extensively used for intranasal formulations [67–70], and accordingly, the tests focused on the PCL-based membrane. It is well known that PCL is a biocompatible and biodegradable material [71], but a solvent casting method using DCM is required to prepare these membranes. Therefore, a cytotoxicity test was developed to ascertain the cytocompatibility of the resulting rate-controlling membranes. For this purpose, cell viability/- proliferation, metabolic activity, and LDH release were evaluated (Fig. 5). The results obtained for the metabolic activity and cell viability tests were similar. It can be seen that the cells treated with the PCL-based membrane showed a slight reduction in cell viability of around 25 % (p <0.0001). The obtained values are higher than 70 % of the control cell viability, and therefore it can be considered that they are not cytotoxic [72]. Interestingly, the materials induced the release of LDH, which is used to determine cellular toxicity (p <0.0001). LDH is released from damaged cells. The obtained results, approximately 25 % LDH release, are in line with the results obtained previously for high molecular weight PCL implants prepared using a similar method [28]. Alternatively, the metabolic activity obtained in the treated cells is significantly lower than the control (p =0.009). However, the results obtained in the cell viability and cell proliferation (p =0.069) indicate that the resulting materials are not cytotoxic. 4. Conclusion The intranasal implants containing RIS were successfully fabricated and characterised in this study. The implant consisted of two parts: a core layer prepared from RIS and water-soluble polymers (PVP and PEG), and a coating layer made of PCL film. RIS was homogenously distributed within the core implant. Moreover, thermal analysis and FTIR results suggested interactions between RIS and the two watersoluble polymers. Furthermore, the in vitro release profile of RIS showed that the coating layer made of PCL film could extend the release of RIS from 2 days up to 100 days. Notably, the performed reservoir-type implant exhibited a linear release over the first 10 days, followed by a slower release rate that reached another linear phase up to 40 days. Subsequently, the drug release rates progressively slowed down. Moreover, a single implant was able to provide sustained drug release of therapeutically relevant RIS dose in vitro. The cytotoxicity study of the implant showed that the intranasal implants were biocompatible and non-toxic to the tested model cells (HDF). Accordingly, the findings from this study suggest that the implants have the potential to provide longacting drug delivery for targeting the brain. Therefore, further studies, such as in vivo pharmacokinetic and biodistribution studies using suitable animal models, are required to evaluate the delivery of RIS from this intranasal implantable device to the brain. Additionally, studies related to safety, sterility, and stability should be conducted prior to clinical application. CRediT authorship contribution statement Emilia Utomo: Writing –original draft, Validation, Methodology, Investigation, Formal analysis, Data curation, Conceptualization. Jiaqi Gao: Investigation. Qonita Kurnia Anjani: Writing –review &editing, Writing –original draft. Camila J. Picco: Investigation. Natalia Moreno-Castellanos: Writing –review &editing, Methodology, Investigation, Data curation. Ryan F. Donnelly: Supervision, Resources, Funding acquisition. Juan Domínguez-Robles: Writing –review &editing, Writing –original draft, Supervision. Eneko Larra˜ neta: Writing –review &editing, Writing –original draft, Supervision, Resources, Funding acquisition, Conceptualization. Declaration of competing interest The authors declare that they have no known competing financial Fig. 5. Cell viability, metabolic activity, LDH assay, and cell proliferation tests for PCL-based membranes (means +S.D., n ≥3). E. Utomo et al. Journal of Drug Delivery Science and Technology 99 (2024) 105973 8 interests or personal relationships that could have appeared to influence the work reported in this paper. Data availability Data will be made available on request. Acknowledgments EU would like to acknowledge the Indonesian Endowment Fund for Education (Lembaga Pengelola Dana Pendidikan/LPDP). EL would like to acknowledge the Academy of Medical Sciences (SBF005\1011) and EPSRC (EP/X525625/1) for their financial support for this work. 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