# Disease-Adaptive Drug Delivery to the Inflamed Intestinal Mucosa Using Poly(Lactic-Co-Glycolic Acid)-cyclodextrin Hybrid Nanocarriers

J. Schreiner, F. E. B. Brettner, S. Steigert, A. Haessler, R. Mols, S. Gier, N.

T. L. Jung, S. Vogel-Kindgen, Susanne Florin Muschert, P. Augustijns, et al.

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## To cite this version:

J. Schreiner, F. E. B. Brettner, S. Steigert, A. Haessler, R. Mols, et al., Disease-Adaptive Drug Delivery to the Inflamed Intestinal Mucosa Using Poly(Lactic-Co-Glycolic Acid)-cyclodextrin Hybrid Nanocarriers. Advanced Therapeutics, 2025, Advanced Therapeutics, 8 (2), pp.2400368. 10.1002/adtp.202400368. hal-04973661

HAL Id: hal-04973661

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# Disease-Adaptive Drug Delivery to the Inflamed Intestinal Mucosa Using Poly(Lactic-Co-Glycolic Acid)-cyclodextrin Hybrid Nanocarriers

Jonas Schreiner, Felix E. B. Brettner, Sebastian Steigert, Annika Haessler, Raf Mols, Stefanie Gier, Nathalie Jung, Sarah Vogel-Kindgen, Susanne Muschert, Patrick Augustijns, and Maike Windbergs*

Fluctuating severity of symptoms is a common hallmark of many inflammatory disorders, including inflammatory bowel disease (IBD). Addressing the pH changes during active and resting phases in IBD-affected tissue, a disease-adaptive nanocarrier system is designed for oral administration, enabling pH-dependent local drug release. The hybrid carrier combines poly(lactic-co-glycolic acid) and an amphiphilic cyclodextrin derivative, with physicochemical properties and drug release kinetics controlled by adjusting polymer ratios. The systems exhibited baseline drug release at pH 5 with increased rates at pH 2, which is characteristic of actively inflamed IBD tissue. Assessing the impact of biomolecule adhesion, biocorona formation was studied using ex vivo human intestinal fluids. Corona composition highly depended on the patient's prandial state and the nanocarrier matrix, with proteins predominating in the fasted state and lipids in the fed state. Notably, differences in the attachment of proteins and free fatty acids are detected in the latter. Transport studies using human in vitro models of the inflamed intestine revealed mucosal accumulation, facilitating localized drug delivery and effectively reducing cytokine levels to basal concentrations. This hybrid system highlights the potential of disease-adaptive drug release for inflammatory disease treatment and underscores the impact of biocorona formation on therapeutic performance in the gastrointestinal tract.

## 1. Introduction

Inflammatory bowel diseases (IBD) are a group of long-term inflammatory disorders of the intestinal tract, affecting more than 7 million people worldwide and thus posing a high economic burden on healthcare systems.[1-3] The 2 main forms comprise Crohn's disease (CD) and ulcerative colitis (UC), characterized by chronic and recurrent inflammation of the whole gastrointestinal tract or the colon, respectively, and severe impairment of quality of life for patients.[4-7] During disease progression, active symptom flares alternate with so-called resting periods, which are specified by milder symptoms, and the affected tissue areas are inhomogeneously spread throughout the intestinal tract. Currently, there is no cure for IBD, and disease management relies on a range of surgical and medical interventions, with the primary goal to induce and maintain remission. Further, the existing treatment regimes need continuous adjustment. While a lower basal dosage is applied during the resting phases to

J. Schreiner, F. E. B. Brettner, A. Haessler, S. Gier, N. Jung, S. Vogel-Kindgen, M. Windbergs
Institute of Pharmaceutical Technology Goethe-University Frankfurt
60438 Frankfurt am Main, Germany
E-mail: windbergs@em.uni-frankfurt.de

S. Steigert, R. Mols, P. Augustijns
Drug Delivery and Disposition
KU Leuven
Leuven 3000, Belgium
S. Muschert
University of Lille
Inserm
CHU Lille
U1008, Lille F-59000, France

© The ORCID identification number(s) for the author(s) of this article can be found under https://doi.org/10.1002/adtp.202400368

© 2025 The Author(s). Advanced Therapeutics 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/adtp.202400368
prevent acute inflammation and protect the intestinal tissue from long-term risks like enteric cancer, the treatment must be adjusted to higher drug doses in relapsing phases.[8–10] To date, steroidal anti-inflammatory drugs, the gold standard for treating IBD, are currently either administered orally, leading to systemic adverse effects, or locally, where they primarily target distal colonic segments.[11,12] Consequently, innovative formulations are urgently needed for targeted, local, and especially disease-adaptive drug delivery to the intestine, focusing on affected tissue regions, limiting absorption in non-inflamed tissue, and increasing the drug concentration in affected tissue.

Although the intestine is often favored for systemic delivery based on its high drug absorption rates, these qualities drastically limit the potential for a local delivery approach of small molecules via the oral route, for which systemic absorption should be limited.[13–15] The immediate and effective transport of most drugs across the intestinal mucosa upon release from their carrier minimizes their residence time at the inflamed tissue sites and can potentially lead to severe side effects by entering the systemic circulation.[16–18] Different strategies can be employed to mitigate these effects, such as the use of prodrugs, which reduce unwanted side effects and improve targeted drug delivery to the afflicted tissue. Recently, targeted nanocarriers (NCs) have been introduced as valuable tools revolutionizing disease diagnosis, treatment, and theranostics in IBD. Such carriers can be designed to take advantage of pathophysiologically altered conditions to achieve disease-adaptive drug delivery. In this context, exploiting pH alterations in IBD-affected tissue, in particular, poses a promising therapeutic strategy. Due to an accelerated secretion of lactic acid by activated immune cells, which infiltrate the intestinal tissue during acute inflammation events and increase glucose consumption via glycolysis based on their enhanced oxygen and energy demand, the physiological pH is drastically acidified.[19,20,21] While pH measurements in the intestinal lumen of IBD patients show minimal differences compared to healthy individuals, these values are likely skewed due to the high buffering capacity of intestinal fluids and fail to reflect the lower pH at inflammatory hotspots.[20,21] These localized microenvironments are also sites for the secretion of acidic components, further contributing to a decrease in pH values.[22]

This acidification can be exploited using NCs made of pH-responsive biomaterials, which modulate their release profile based on environmental pH. These comprise ionizable polymers, such as polymethacrylates or acid-labile materials that are commonly used in the design of oral NCs.[23] Although this approach has been proven successful in targeting specific regions of the intestine, polymethacrylates are often not biodegradable. Furthermore, pH-responsive materials frequently show limited cellular uptake or have poor drug-loading capabilities.[24,25] To address these limitations and allow for spatially resolved drug delivery, novel cell-penetrating delivery systems incorporating biocompatible, acid-degradable components and materials that improve drug encapsulation are required.

This study introduces a disease-adaptive nanocarrier (NC) platform comprised of heptakis-6-O-hexanoyl-β-cyclodextrin (CDOC₆) and poly(lactic-co-glycolic acid) (PLGA) capable of spatially controlled drug release by leveraging the pathophysiologically altered and local pH fluctuations in IBD. The

system ensures a baseline release of the anti-inflammatory drug curcumin under physiological pH conditions while achieving higher drug concentrations in more acidic environments, corresponding to the resting and active phases of IBD, respectively.

## 2. Materials and Methods

**Materials:** Unless otherwise stated, all materials were purchased from Merck (Taufkirchen, Germany). Ultrapure water was obtained via reverse osmosis (Purelab flex 2, Veolia Water Technologies Deutschland GmbH, Celle, Germany). All primary antibodies were purchased from BIOMOL GmbH (Hamburg, Germany). Fluorescent dyes and secondary antibodies were obtained from Thermo Fisher Scientific (Waltham, MA, USA). Resomer RG 502H (PLGA) was purchased from Evonik Industries AG (Darmstadt, Germany).

**Nanocarrier (NC) Preparation:** Heptakis-6-O-hexanoyl-β-cyclodextrin (CDOC₆) was synthesized as previously described.[26] Poly(lactic-co-glycolic acid) (PLGA) and CDOC₆ were mixed in proportions of 100:0, 80:20, 50:50, 20:80, and 0:100 and dissolved in tetrahydrofuran (THF) for nanoprecipitation. Dissolved PLGA and CDOC₆ were dropped into the aqueous phase under constant stirring at room temperature (RT) with a flow rate of 5 mL h⁻¹, controlled by a syringe pump (Pump 11 Pico Elite Plus, Harvard apparatus, March-Hugstetten, Germany). THF was evaporated under reduced pressure after the nanoprecipitation was finished. The resulting NC formulations were referred to as P100:C0, P80:C20, P50:C50, P20:C80, P0:C100. Curcumin-loaded NCs were prepared by adding 10% (m/m) curcumin into the THF solution before nanoprecipitation. The theoretical loading capacity of curcumin was defined as the amount of curcumin related to the polymer mass. Polyethylene glycol-poly(lactic-co-glycolic acid) nanocarriers (PEG-PLGA NCs) were produced accordingly using PLGA and PEG-PLGA with a mass ratio of 3:1. All NCs were stored at 4 °C until further use.

**Scanning Electron Microscopy (SEM):** Single droplets of each NC type were applied onto aluminum foil and immediately frozen in liquid nitrogen for sample preparation.[26] Afterward, the solvent was evaporated under reduced pressure. A gold-palladium layer was applied to the samples (SC7620 mini Sputter Coater, Quorum Technologies, Lewes, UK) before imaging using a scanning electron microscope with an electron high tension of 30 kV (Zeiss EVO10 microscope, Carl Zeiss AG, Oberkochen, Germany).

**Physicochemical Characterization:** The intensity-dependent hydrodynamic diameter, polydispersity index (PDI), and ζ potential of the NCs were determined at 25 °C by dynamic light scattering (DLS) and electrophoretic light scattering (ELS) using a Zeta-sizer NanoZS (Malvern Panalytical, Malvern, UK). All measurements were performed in ultrapure water and were conducted for each NC batch to ensure consistent quality. Loading capacity (LC) and encapsulation efficiency (EE) of curcumin-loaded NCs were indirectly measured after centrifugation of the nanosuspension at 21500×g for 20 min at 4 °C and quantification of curcumin in the supernatant at 425 nm (Spark multimode microplate reader, Tecan, Männerdorf, Switzerland). LC and EE of the NCs were
evaluated for 3 independent batches, measured in triplicate and calculated according to Equations 1 and 2, respectively.

$$LC (\%) = \frac{\text{mass of drug in nanocarriers}}{\text{mass of nanocarriers}} \times 100 \quad (1)$$

$$EE (\%) = \frac{\text{mass of encapsulated drug into nanocarriers}}{\text{initial mass of used drug}} \times 100 \quad (2)$$

**Attenuated Total Reflection-Infrared Spectroscopy (ATR-IR):** ATR-IR spectra of dried NCs were recorded with an ALPHA II Platinum FT-IR spectrometer (Bruker, Billerica, MA, USA) via the OPUS-TOUCH software (Version 8.1). For each NC type, 24 spectra over a shift range of 3100-400 cm⁻¹ at a resolution of 4 cm⁻¹ were recorded, averaged, and normalized.

**Differential Scanning Calorimetry (DSC):** To record thermograms of the precipitated and dried samples, a DSC 3 (Mettler Toledo, Columbus, OH, USA), equipped with an Intra-Cooler TC100MT (Peter Huber Kältemaschinenbau AG, Offenburg, Germany) was used. A 5 K min⁻¹ heating rate and a 50 ml min⁻¹ nitrogen flow were applied for every measurement. All samples were analyzed in cold-welded aluminum sample pans. All thermograms (mW/g) were normalized to the largest peak with OriginPro 2018 (OriginLab Corporation, Northampton, MA, USA).

**Raman Spectroscopy:** Raman spectroscopy was performed using an Alpha300R+ confocal Raman microscope (Witec GmbH, Ulm, Germany) equipped with a 532 nm laser set to a power of 20 mW in front of the 50x objective (NA 0.8). Spectra were acquired from dried NCs with 0.5 s integration time and 10 accumulations. Cosmic ray reduction and spectral background removal (shape function, window size: 100 points) were conducted in the Project FOUR software (Witec GmbH, Ulm, Germany). Data was then exported to Matlab (Version 2022a, The Mathworks Inc., Natick, USA), where all further processing steps were performed using in-house written scripts. Spectra were trimmed to Raman shifts of 400–3140 cm⁻¹ and denoised using singular value decomposition. Next, spectra were z-scored, followed by principal component analysis.

**In Vitro Drug Release Testing:** Curcumin release profiles were obtained in Fed State Simulated Intestinal Fluid (FeSSIF) as well as acidified FeSSIF. FeSSIF was prepared according to standard-procedure in the literature.[27] Briefly, NaOH, glacial acetic acid, and NaCl were dissolved in purified water, and the pH was adjusted to 5.0 using NaOH. Afterward, sodium taurocholate and a lecithin solution in methylene chloride were added. Methylene chloride was evacuated under reduced pressure at 40 °C. The second release medium was prepared using FeSSIF, acidified with 1 N HCl to pH 2.0. Each NC suspension was mixed with FeSSIF or acidified FeSSIF, respectively. The mixtures were incubated at 37 °C and stirred at 300 rpm. Samples were drawn at predetermined time points (0, 0.5, 1, 3, 5, 8, and 24 h), and the sample volume was replaced with fresh, pre-warmed medium. After centrifugation of the sample at 21500×g for 20 min at 4 °C, the amount of released curcumin was evaluated via absorbance measurement at 425 nm (Spark multimode microplate reader, Tecan, Männerdorf, Switzerland). All release experiments were carried out with 3 different batches of curcumin-loaded NCs, and each sample was analyzed in triplicate.

**Investigation of Drug Release Mechanisms:** The mass transport mechanisms of curcumin from the hybrid nanoparticles were analyzed by fitting 2 distinctive mathematic equations, one considering diffusion and one considering surface erosion as the predominant drug release mechanism. The diffusion-based theoretical consideration proposed by Siepmann & Siepmann assumes Fickian diffusion via Equation 3.[28]

$$\frac{M_t}{M_\infty} = 1 - \frac{6}{\pi^2} \sum_{n=1}^{\infty} \frac{\exp\left[-\frac{Dn^2\pi^2t}{R^2}\right]}{n^2} \quad (3a)$$

$M_t$ and $M_\infty$ represented the cumulative amounts of drug released at time t and at infinite time (24 h), respectively. D was the diffusion coefficient of the drug within the system, and R stood for the sphere's radius. A constant carrier size, homogeneous drug distribution, and sink conditions were assumed throughout the entire observation period.

The Hopfenberg equation was fitted to verify if surface erosion is the underlying drug release mechanism with the hybrid nanoparticles. This theoretical assumption is valid for spheres, cylinders, and slabs.[29,30]

$$\frac{M_t}{M_\infty} = 1 - \left(1 - \frac{k^0 \times t}{c^0 \times a}\right)^n \quad (3b)$$

$c_0$ denotes the uniform initial drug concentration within the system; a is the radius of a cylinder or sphere or the half-thickness of a slab; n is a “shape factor” representing spherical ($n = 3$), cylindrical ($n = 2$) or slab geometry ($n = 1$). To evaluate the fit quality, the calculated over the experimentally obtained drug amounts released (cumulative release in %) have been plotted for each model assumption and sample analyzed, and the $R^2$ values of the linear trendlines have been calculated.

**Analysis of the Intestinal Biomolecular Corona:** The different components of the biomolecular hard corona were analyzed to quantify proteins, peptides, and lipids. The human intestinal fluids (HIF) used in this work were aspirated from the duodenum and jejunum of healthy volunteers in the fasted and fed state following a previously published protocol.[31] Approval for the study was granted by the Ethics Committee Research of UZ/KU Leuven in Belgium (S53791). All participants provided written informed consent before their inclusion in the study. For sample preparation, 500 µL of NC suspension were incubated with 100 µL of fetal calf serum (FCS, Lonza, Verviers, Belgium), digested Ensure Plus (Abbott Nutrition, Chicago, USA), or fasted or fed state human intestinal fluid for 1 h in an incubator shaker at 37 °C (175 rpm). After incubation, the suspension was centrifuged (Centrifuge 5804 R, Eppendorf SE, Hamburg, Germany) at 12500 xg (20 min, 4 °C), and the supernatant was sampled. The remaining supernatant was discarded, and the NCs were washed by re-dispersion in ultrapure water and centrifuged (12500 xg, 20 min, 4 °C). The resulting supernatant was discarded, and the NCs were dispersed again in ultrapure water. Negative controls lacking either NCs, intestinal fluids, or both were prepared accordingly. The protein amount was determined by quantifying tryptophan in the washed and -dispersed particulate samples previously described in the literature.[32] Briefly, 5 µL of the sample was added to a black 96-well-plate and thoroughly mixed with 200 µL urea buffer
solution (8 M urea, 0.1 M Tris buffer, pH 8.5). Successively, the tryptophan content was determined via fluorescence measurement at an emission wavelength of 350 nm with excitation at 295 nm (Infinite M200 Plate Reader, Tecan, Männerdorf, Switzerland). Lipids and lipid digestion products were extracted and quantified from supernatants as previously described.[33] Briefly, lipid extraction was performed by adding 1-octadecanol (internal standard) in chloroform-isooctane (1:1 v/v) and 1 M HCl to 100 µL of sample. After mixing, the chloroform-isooctane-mixture was added, and the lipids were extracted by shaking. After collection of the bottom layer obtained via centrifugation (20000× g, 15 min, 4 °C), the samples were diluted 5-fold in isooctane and, following injection of 1 µL, analyzed via HPLC coupled to a charged aerosol detector (CAD), as previously reported.[33] For quantification, calibration curves were established using a surrogate reference product per lipid class. Bile salt quantification in supernatants was performed as previously described.[31,33] Briefly, a Waters LC-MS/MS system, comprising a Xevo TSQ micro triple quadrupole mass spectrometer with an electrospray ionization (ESI) source, connected to an Acquity UPLC H-class system (Waters) was utilized to determine bile salt quantity. Separation involved a Kinetex XB-C18 column (2.6 µm, 100 Å, 50 × 2.1 mm) maintained at a temperature of 35 °C. Bile salts were separated via gradient elution, with the mobile phase consisting of acetonitrile, methanol, water, and 5 mM ammonium acetate buffer (pH adjusted to 3.5 with acetic acid) after injection of 5 µL of the sample. Samples were diluted in a methanol-water-mixture (1:1 v/v) containing 200 nM deuterated cholic acid (C-d4) and chenodeoxycholic acid (CDC-d4) as internal standards. Positive and negative ESI modes were employed for analysis. Total bile salt concentrations were calculated as the sum of individual bile salts. For all components of the biomolecular corona, the association factor was determined to unveil the amount of the respective biomolecule compared to the amount of polymer contained in the NC. The association factor was calculated by dividing the determined mass of the biomolecule by the mass of the matrix substance in the sample via Equation 4. Unless otherwise stated, all experiments were performed in triplicate.

$$ \text{Association factor [\%]} = \frac{\text{mass of bound biomolecules}}{\text{mass of matrix substance}} \quad (4) $$

**General Cell Culture:** The Caco-2 cell line (ATCC CRL-2102) was purchased from the American Type Culture Collection and cultured in Dulbeccos's Modified Eagle medium (DMEM, Thermo Fisher Scientific, Waltham, MA, USA) supplemented with 10% (v/v) FCS (Lonza, Verviers, Belgium) and 1% (v/v) nonessential amino acids (NEAA) at 37 °C, 5% CO₂ and 95% relative humidity. The medium was changed every 2 to 3 days. Passages 50–65 were used in subsequent studies.

**In Vitro Cell Viability Testing:** The potential cytotoxicity of unloaded and curcumin-loaded hybrid NCs on Caco-2 cells was assessed using the 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide (MTT) assay. 30000 Caco-2 cells per well were seeded in a 96-well plate and cultured for 48 h. Subsequently, the cells were exposed to different concentrations of hybrid NCs for 24 h at 37 °C with 5% CO₂. Following the NC incubation period, the cells were rinsed with 1x phosphate-buffered saline (PBS), and 1 mg/ml of MTT reagent

was added to each well. After a 4-hour incubation at 37 °C and 5% CO₂, the supernatant was removed, and the formazan crystals were dissolved using dimethyl sulfoxide. Absorbance was measured at 570 nm (Spark multimode microplate reader, Tecan, Männerdorf, Switzerland). Cells treated solely with DMEM or 1% TritonX100 (diluted in DMEM) were used as the negative and positive control, respectively. All experiments were conducted in triplicate. Cell viability was calculated according to Equation 5.

$$ \text{Cell viability (\%)} = \frac{\text{Absorbance}_{\text{cells incubated with nanocarriers}} - \text{Absorbance}_{\text{positive control}}}{\text{Absorbance}_{\text{negative control}} - \text{Absorbance}_{\text{positive control}}} \times 100 \quad (5) $$

**In Vitro Model of the Inflamed Intestine:** 50000 Caco-2 cells cm⁻² were seeded on the apical compartment of 24-Transwell inserts (Falcon, Corning, NY, USA). The cells were cultured in DMEM supplemented with 10% FCS (v/v) and 1% (v/v) NEAA for 21 days. The medium was changed every 2 to 3 days. The obtained *in vitro* intestinal models were inflamed by adding 37.5 ng/ml tumor-necrosis factor α (TNFα), 37.5 ng mL⁻¹ interferon γ (IFNγ), and 18.75 ng mL⁻¹ interleukin 1β (IL1β) to the basolateral compartment.[34] The inflammation process was started 24 h before the NC treatment and was repeated during the treatment with NCs in the apical compartment (Figure 1). The integrity of the cell layer was determined before, during, and after experiments by measurement of the transepithelial electrical resistance (TEER) using a Volt-Ohm meter equipped with chopstick electrodes (Millicell ERS-2, MilliporeSigma, Burlington, MA, USA).

**Immunofluorescence Staining:** The uptake of NCs into the inflamed intestinal models and their subcellular localization were investigated using immunofluorescence staining. After the permeation experiments, the inflamed intestinal models were washed thrice with PBS for 5 min to remove NCs adsorbed to the cellular layer. After fixation of the models with 4% (v/v) formaldehyde solution for 15 min at RT, the Transwell insert (Falcon, Corning, NY, USA) membranes were cut out and washed thrice with washing buffer (PBS with 0.05% Tween20). The cells were treated with blocking buffer (3% bovine serum albumin in PBS) for 60 min at RT, followed by another washing step with washing buffer. Subsequently, the primary antibody (LAMP1 IgG2b, 1:200) was administered to the cells, and incubation was conducted at 4 °C overnight. Again, the cells were washed thrice and further incubated with the secondary antibody (goat anti-mouse IgG2 coupled to Alexa Fluor 633, 1:1000) for 60 min. Following another washing step with 0.05% Tween in PBS, the cytoskeleton was stained with Alexa Fluor 488 Phalloidin (0,165 µM in PBS) for 60 min. The samples were covered in FluorSave Reagent and imaged using a confocal laser scanning microscope (Zeiss LSM 900, Carl Zeiss AG, Oberkochen, Germany). The ImageJ Plugin JACoP was used to evaluate and quantify the co-localization of the NCs via Manders' coefficient fraction 1 (M1) and 2 (M2).[35,36] Manders' coefficient was thereby a statistical measure to quantify the degree of co-localization between 2 fluorescent signals in a micrograph. It was used to quantitatively assess the spatial
<figure><img src="image_4.png" /><figcaption>Figure 1. Cultivation, stimulation, and treatment regime of the in vitro models of the inflamed human intestine.</figcaption></figure>

correlation of 2 differently labeled structures within a sample. In this study, M1 represents the overlap of LAMP1 signals with cur-cumin signals relative to the total amount of LAMP1 signals. M2 represents the co-localization of curcumin signals with LAMP1 signals relative to the total number of curcumin signals. Quan-tification measurements were conducted in triplicate.

**In Vitro Permeation Analysis:** Using the inflamed intestinal models described above, the permeation of curcumin-loaded hy-brid NCs was evaluated (Figure 1). The cell culture medium was aspirated, and the cells were washed multiple times with pre-warmed Hank's balanced salt solution (HBSS). HBSS was added into the basolateral compartment, whereas 0.1 mg mL⁻¹ of each NC type, related to the polymer mass and diluted in HBSS, was applied in the apical compartment. The inflamed intestinal mod-els were incubated for 6 h, and samples were drawn from the basolateral compartment at 1, 2, 4, and 6 h. The removed sam-pling volume in the basolateral compartment was replaced with the same amount of pre-warmed HBSS. The concentration of curcumin-loaded NCs was assessed via fluorescence intensity measurements at 400 nm excitation and 535 nm emission (Spark multimode microplate reader, Tecan, Männerdorf, Switzerland). The measured fluorescence intensities were corrected for blanks of the same permeation setup in the absence of nanocarriers. Per-meation studies were performed thrice with different NC batches and inflamed in vitro intestinal models. All samples were evalu-ated in triplicate.

**Human IL8-Specific Enzyme-Linked Immunosorbent Assays (ELISAs):** To quantify the release of interleukin 8 (IL8) 6 and 24 h after treatment with hybrid NCs, supernatants of untreated and treated inflamed intestinal models and noninflamed intesti-nal models were analyzed via ELISA. Supernatants were collected and centrifuged (13000×g, 15 min, 4 °C) to separate cell debris. The cell-free supernatant was stored at -80 °C until further use. IL8 release was quantified using uncoated ELISA kits according to the manufacturer protocol (Invitrogen, Waltham, MA, USA). Absorption measurements were performed at 450 nm with a reference wavelength of 570 nm (Spark multimode microplate reader, Tecan, Männerdorf, Switzerland). Obtained IL8 concen-trations were further analyzed according to Equation 6.

$$ \text{Cytokine release (\%)} = \frac{\text{Cytokine conc.}_{\text{Hybrid treated IBD models}} - \text{Cytokine conc.}_{\text{blank control}}}{\text{Cytokine conc.}_{\text{Medium treated IBD models}} - \text{Cytokine conc.}_{\text{blank control}}} \times 100 \quad (6) $$

<table><caption>Table 1. Loading capacity (LC) and encapsulation efficiency (EE) of PLGA, CDOC₆, and hybrid NCs. All data are expressed as mean ± SD.</caption><thead><tr><th>Formulation</th><th>Loading Capacity [%]</th><th>Encapsulation Efficiency [%]</th></tr></thead><tbody><tr><td>PLGA (P100:C0)</td><td>9.16 ± 0.09</td><td>91.60 ± 0.86</td></tr><tr><td>P80:C20</td><td>8.35 ± 0.31</td><td>83.45 ± 3.13</td></tr><tr><td>P50:C50</td><td>7.82 ± 0.09</td><td>78.24 ± 0.85</td></tr><tr><td>P20:C80</td><td>7.81 ± 0.28</td><td>78.09 ± 2.78</td></tr><tr><td>CDOC<sub>6</sub> (P0:C100)</td><td>7.01 ± 0.15</td><td>70.13 ± 1.48</td></tr></tbody></table>

All results were expressed relative to the IL8 secretion level of an inflamed and untreated control.

**Statistical Analysis:** Dynamic and electrophoretic light scat-tering data, release profiles in the different media, and Manders' coefficients were analyzed for statistical significance using One-Way ANOVA followed by the Bonferroni posthoc test. The sig-nificance of ELISA data was evaluated via One-Way ANOVA and Dunn-Sidak posthoc test. Data were considered significant if the *p*-value was below 0.05 with *p* < 0.05, **p* < 0.01, ***p* < 0.001.

## 3. Results

### 3.1. Drug Loading and Physicochemical Characterization of Nanocarriers

Hybrid NCs composed of different ratios of poly(lactic-co-glycolic acid) (PLGA) and heptakis-6-O-hexanoyl-β-cyclodextrin (CDOC₆) were prepared by nanoprecipitation. Visualization of the differ-ent NC types via scanning electron microscopy revealed solid, spherical NCs with a smooth surface morphology (Figure S1A–J, Supporting Information). Loading capacity (LC) and encapsula-tion efficiency (EE) were evaluated for the curcumin-loaded NCs. Based on a theoretical LC of 10%, all NC types showed high encapsulation efficiencies of at least 70% (Table 1). The LC in-creased with higher amounts of PLGA, leading to the highest EE of more than 90% for pure PLGA NCs.

In blank formulations, PLGA NCs showed a significantly larger hydrodynamic diameter (265.75 ± 8.82 nm) than the hy-brid and pure CDOC₆ NCs, which displayed sizes below 225 nm without significant differences from each other (Figure 2A). Cur-cumin loading increased the hydrodynamic diameter of NCs to 285.62 ± 41.49 nm (PLGA NCs) and 249.98 ± 5.95 nm (CDOC₆ NCs), respectively. In contrast, the size of drug-loaded hybrid NCs decreased with an increasing amount of
<figure><img src="image_5.png" /></figure>

<figure><img src="image_6.png" /></figure>

<figure><img src="image_7.png" /><figcaption>**Figure 2.** A) Hydrodynamic diameter, B) PDI, and C) ζ potential of un-loaded and drug-loaded PLGA, CDOC₆, and hybrid NCs. All data are expressed as mean ± SD of 3 independent experiments. Statistical analysis was performed using One-Way ANOVA with Bonferroni post hoc test. \*p < 0.05, \*\*p < 0.01, and \*\*\*p < 0.001.</figcaption></figure>

PLGA with a significantly larger hydrodynamic diameter for P20:C80 NCs (224.10 ± 22.96 nm) compared to P50:C50 NCs (194.92 ± 2.88 nm). Blank hybrid formulations self-assembled into monodisperse NCs with a polydispersity index (PDI) below 0.2, whereas the PDI of CDOC₆ was substantially lower (0.08 ± 0.02), and the pure PLGA NCs showed significantly higher values (0.28 ± 0.04) (Figure 2B). After the curcumin load-

ing, the PDI increased for CDOC₆ NCs (0.16 ± 0.01), reaching values comparable to the hybrid NCs, whereas curcumin-loaded PLGA NCs showed the most significant increase in their PDI (0.34 ± 0.07). Unloaded pure CDOC₆ (−35.8 ± 0.62 mV) and PLGA NCs (−31.42 ± 11.32 mV), as well as the different hybrid NC types (up to −40 mV), had a negative surface charge with no prominent differences between the individual NC types (Figure 2C). Loading of curcumin led to significantly lower ζ potentials of PLGA NCs (−28.84 ± 0.64 mV) compared to the other curcumin-loaded NCs (≈−35 mV).

## 3.2. Understanding Material Interactions

When different materials are combined in a nanocarrier, unique chemical interactions can occur at their molecular interface, influencing the overall nanocarrier's stability, drug release behavior, and interactions with biological fluids and cells. Raman spectroscopy was employed to identify the vibrational, rotational, and other low-frequency modes of PLGA and CDOC₆ within the hybrid NCs to analyze these interactions, providing insight into the physical and structural properties of the system. The characteristic peak for C—C—O deformation at 490 cm⁻¹ increased with higher proportions of CDOC₆, while peaks at 1750 cm⁻¹ (ester stretching) and 2950 cm⁻¹ (C—H stretching) decreased (Figure 3A). Physical mixtures and the NCs showed similar shifts in peak intensities for the different polymer proportions; however, the spectra were not identical. Interestingly, at 1118–1132 cm⁻¹ (C—O—H deformation) and 1445–1455 cm⁻¹ (CH₂ symmetric deformation), 2 peaks each were linearly correlated with measures of certainty above 0.98, respectively (Figure 3B,C).

Principal component analysis (PCA) was employed to reduce the dimensionality of the Raman spectra and to investigate differences between hybrid NCs and physical mixtures of the respective components. The loading for principal component 1 (PC1) was mainly used to guide the interpretation of the PCA results since principal component 2 (PC2) explained only 3.32% of the variance, while PC1 constituted 90.91%. In the loading plot, peaks that displayed positive correlations with PC1 exhibited values above 0, while those with negative correlations displayed values below 0 (Figure 3D). The analysis revealed that the peaks indicating C—C—O deformation, C—O—H deformation, and CH₂ symmetric deformation were negatively correlated to PC1, while other peaks stemming from C—O—H deformation, CH₂ symmetric deformation, ester stretching, and C—H stretching were positively correlated to PC1. The PC loading plot indicated the same spectral features observed in the Raman spectra. Furthermore, the score plot for PC1 and PC2 delineated the populations corresponding to the NCs and the physical mixtures of pure NCs in the respective proportions (Figure 3E). PC1 effectively differentiated between these groups, whereas PC2 accounted for less variation, primarily capturing intragroup variability. The data points corresponding to the hybrid NCs were found between the aforementioned clusters, showing a shift toward lower values of PC1 with increasing proportions of CDOC₆. The intragroup variability decreased with increasing amounts of CDOC₆. PCA revealed major differences between hybrid NCs and the physical mixtures of pure PLGA and CDOC₆ NCs, which showed a robust clustering with the signals of pure CDOC₆ NCs. The data points for the
<figure><img src="image_8.png" /><figcaption>**Figure 3.** A) Raman spectra of NCs and the respective physical mixtures of pure NCs. Peak ratios of B) peaks at 1118 cm⁻¹–1132 cm⁻¹, and C) at 1445–1455 cm⁻¹ for the NCs. D) Loadings of the first principal component of the spectra and E) score plot of the first 2 principal components.</figcaption></figure>

hybrid NCs were located between the previously mentioned clusters, with a noticeable shift toward lower PC1 values as the proportion of CDOC₆ increased. Additionally, intragroup variability decreased with higher amounts of CDOC₆. Thus, PCA revealed significant differences between the hybrid NCs and the physical mixtures of pure PLGA and CDOC₆ NCs, as the latter exhibited strong clustering with the signals of pure CDOC₆ NCs.

These investigations were further complemented with attenuated total reflectance Fourier-transform infrared (ATR-FTIR) spectroscopy and differential scanning calorimetry (DSC). Several peaks of the ATR-IR spectra varied based on the polymer proportion employed. Specifically, peaks occurring from 3600–3100 cm⁻¹ (OvH), 3000–2750 cm⁻¹ (C–H), and at 1020 cm⁻¹ (C–O–C) increased with higher proportions of CDOC₆ (Figure S2A, Supporting Information). In contrast, peaks at 1740 cm⁻¹ (C=O) and 1080 cm⁻¹ (C–O) decreased with higher proportions of CDOC₆. No peak alterations in the spectra of the curcumin-

loaded NCs with respect to unloaded NCs were visible, whereas spectral components of curcumin could not be detected due to the limit of detection (Figure S2B, Supporting Information). Differential scanning calorimetry (DSC) thermograms of pure PLGA NCs showed a melting point of 314.59 °C, similar to the melting point of the PLGA control at 313.39 °C (Figure S2C, Supporting Information). The PLGA control, PLGA NCs, and all hybrid NCs except for P20:C80 NCs showed an endothermic event at 45 °C. With increasing proportions of CDOC₆, the PLGA melting point shifted toward higher temperatures with peaks at 360.86 °C for P80:C20 NCs and 359.94 °C for P50:C50 NCs. The melting point of PLGA could no longer be detected for P20:C80 NCs. Instead, the decomposition of CDOC₆ was determined at 278.98 °C for P20:C80 NCs and 269.54 °C for pure CDOC₆ NCs. The decomposition temperatures observed for the hybrid NCs aligned with those of the CDOC₆ control. In the thermal analysis of curcumin-loaded NCs, no distinct melting point for curcumin
<figure><img src="image_9.png" /><figcaption>**Figure 4.** Cumulative release profiles of curcumin from NCs in A) FeSSIF (pH 5) and B) acidified FeSSIF (pH 2) at 37 °C. All data were expressed as mean ± SD of 3 independent experiments.</figcaption></figure>

was detected for any of the loaded NCs, in contrast to the melting point of pure curcumin at 177.35 °C (Figure S2D, Supporting Information). Additionally, the melting point shift of PLGA seen for hybrid NCs was no longer present in the curcumin-loaded formulations, and the endothermic events for PLGA closely resembled those of the PLGA control.

### 3.3. In Vitro Drug Release

Especially during active phases of IBD, prominent pH shifts along the intestinal tract are observed. While the pH value of the small intestine remains unchanged compared to healthy conditions, the colonic lumen becomes more acidic during active inflammation (pH 2.3-5.5).[37,38] Thus, the curcumin release from the NCs was investigated in Fed State Simulated Intestinal Fluid (FeSSIF) (pH 5) and acidified FeSSIF (pH 2). While FeSSIF pH 5 was employed to represent the colonic medium in resting phases, acidified FeSSIF represented the colonic inflamed tissue environment during acute phases. All formulations showed a basal release at pH 5, likely through taurocholate-mediated solubility enhancement and concentration-driven diffusion of curcumin (Figure 4A). NC formulations with higher amounts of PLGA exhibited faster and higher drug release. CDOC₆ NCs released only ≈7% of the payload within the first hour, while PLGA NCs released ≈35%. Hybrid NCs released ≈35% (P80:C20), 23% (P50:C50), and 19% (P20:C80). After this initial burst release phase, sustained drug release was observed, with the release rate dependent on the CD:PLGA ratio. After 24 h, PLGA NCs demonstrated the highest and CDOC₆ NCs the lowest curcumin release, with ≈50% and 20%, respectively. The release profile from hybrid NCs fell between the individual pure NCs' release curves caused by the material's interactions revealed above. Acidification of FeSSIF led to a significantly increased release rate for all formulations, most probably due to accelerated cleavage of the ester bonds of PLGA and the cyclodextrin ester CDOC₆ (Figure 4B; Table S1, Supporting Information). Accelerated release behavior was observed as early as 1 h, with PLGA NCs releasing ≈39% of curcumin, while CDOC₆ NCs released only ≈8%. Correspondingly, hybrid NCs released 40% (P80:C20), 31% (P50:C50), and 26% (P20:C80) of the drug. After 5 h, the release profiles mirrored those seen at pH 5, with P80:C20 NCs exhibiting the highest release (≈57%), followed by PLGA NCs at ≈52%. The remaining

formulations released ≈44% (P50:C50), 38% (P20:C80), and 23% (CDOC6). However, in acidic conditions over 24 h, notable differences emerged. P50:C50 NCs released significantly more curcumin (≈49%) compared to P20:C80 NCs (≈41%), while P80:C20 NCs (≈60%) released nearly the same amount as pure PLGA NCs (≈59%). These findings demonstrate that combining PLGA and CDOC₆ into hybrid NCs enables precise control over drug release and adaptable drug levels, addressing the distinct requirements of resting and active IBD phases characterized by substantial pH fluctuations.

### 3.4. Mathematical Description of Drug Release Mechanisms

The Siepmann & Hopfenberg equations were fitted to the experimentally observed drug release data at the respective time points for each sample analyzed during the dissolution tests to investigate the underlying drug release mechanisms from hybrid nanoparticles of varying compositions. These fitted curves are presented in Figures 5 and 6, while the corresponding standard errors of the regression (S) values are summarized in Table 2. They quantify the average deviation of observed data points from the regression line, serving as an indicator of the fit's accuracy. Lower S values indicate that the experimental data points are closer to the fitted regression line, reflecting a more accurate fit. The fit's accuracy is further illustrated in a plot comparing calculated drug release to experimentally observed drug release (Y/Y plot, Figure S3, Supporting Information). In an ideal scenario, the calculated Y values would align perfectly with the experimentally obtained values, resulting in a linear trendline with a regression coefficient (R²) equal to 1. The equation proposed by Siepmann & Siepmann for diffusion-controlled drug release from spheres assumes solid amorphous dispersions with the drug being homogenously dispersed within the polymeric matrix.[28] When this equation (Equation 3a) was fitted to experimental drug release profiles of the hybrid nanoparticles, estimated curcumin release curves were obtained (Figure 5). It becomes evident that the fitted curves generally underestimated drug release during the first 3 h and generally overestimated drug release at 24 h. Nevertheless, the pH of the dissolution medium and the standard error of the regressions (S) were relatively high and not close to zero. The best fit of curcumin release was obtained with nanoparticles based on cyclodextrin when exposed to FeSSIF pH 5. When
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Sample 1 ■ Sample 2 ■ Sample 3

<caption>**Figure 5.** Experimentally obtained cumulative curcumin release (symbols) for each sample analyzed and the fit of drug release, calculated (full line) via the diffusion-based Equation 3a. A, B) P100:C0, C, D) P80:C20, E, F) P50:C50, G, H) P20:C80, and I, J) P0:C100 NCs in FeSSIF or acidified FeSSIF, respectively. The values of the standard error of the regression (S) can be found in Table 2.</caption>
<figure><img src="image_20.png" /><figcaption>**Figure 6.** Experimentally obtained cumulative curcumin release (symbols) for each sample analyzed and the fit of drug release, calculated (full line) using the erosion-based Equation 3b. A, B) P100:C0, C, D) P80:C20, E, F) P50:C50, G, H) P20:C80, and I, J) P0:C100 NCs in FeSSIF or acidified FeSSIF, respectively. The values of the standard error of the regression (S) are summarized in Table 2.</figcaption></figure>
<table><caption>**Table 2.** Values of the standard error of the regression (S).</caption><thead><tr><th>Analyzed Medium</th><th>S Values</th><th>Diffusion</th><th>Erosion</th></tr></thead><tbody><tr><td rowspan="5">FeSSIF</td><td>PLGA (P100:C0)</td><td>0.144 ± 0.033</td><td>0.248 ± 0.036</td></tr><tr><td>P80:C20</td><td>0.148 ± 0.029</td><td>0.253 ± 0.034</td></tr><tr><td>P50:C50</td><td>0.104 ± 0.026</td><td>0.184 ± 0.033</td></tr><tr><td>P20:C80</td><td>0.072 ± 0.012</td><td>0.142 ± 0.018</td></tr><tr><td>CDOC<sub>6</sub> (P0:C100)</td><td>0.056 ± 0.006</td><td>0.069 ± 0.012</td></tr><tr><td rowspan="5">Acidified FeSSIF</td><td>PLGA (P100:C0)</td><td>0.168 ± 0.009</td><td>0.263 ± 0.013</td></tr><tr><td>P80:C20</td><td>0.174 ± 0.022</td><td>0.269 ± 0.024</td></tr><tr><td>P50:C50</td><td>0.151 ± 0.012</td><td>0.256 ± 0.008</td></tr><tr><td>P20:C80</td><td>0.135 ± 0.037</td><td>0.230 ± 0.044</td></tr><tr><td>CDOC<sub>6</sub> (P0:C100)</td><td>0.071 ± 0.011</td><td>0.127 ± 0.011</td></tr></tbody></table>

fitting the Hopfenberg equation (Equation 3b) to the experimentally observed curcumin release (Figure 6), it is clearly visible that this erosion model resulted in underestimation during the first 8 h of drug release and a subsequent generous overestimation of curcumin amounts released at 24 h. The best fit of curcumin release was again observed with nanoparticles based on cyclodextrin only, exposed to FeSSIF. In summary, in vitro and in silico analyses showed that hybrid NCs exhibit matrix-dependent release profiles, enabling basal release under physiological conditions and enhanced release at lower pH, addressing the severity of inflammation. Notably, P50:C50 demonstrated the highest pH sensitivity, highlighting its potential for disease-adaptive drug delivery.

### 3.5. Composition Analysis of Intestinal Biomolecular Corona

The interactions between the NCs and human intestinal fluids (HIFs) were investigated to retrieve information on their impact on colloidal NC stability and in vivo behavior. PEG-PLGA NCs were included in these experiments, representing noncorona forming controls.[39,40] The in situ measurement of the hydrodynamic diameter revealed different behaviors of the NC in Fasted state Human Intestinal Fluid (FaHIF) and Fed state Human Intestinal Fluid (FeHIF) (Figure 7A). While after incubation with FaHIF, none of the carrier types substantially changed their hydrodynamic diameter, an increase in diameter of 50 to 60 nm could be observed for all carrier types in FeHIF. Interestingly, there were no significant differences between any carrier type after incubation with the respective HIF. Further, the surface charge was influenced in all carrier types, showing a consistent decrease in ζ potential, independent of the type of HIF (Figure 7B). While for NCs incubated with FaHIF, significant differences were only found between P100:C0 and P50:C50, P50:C50, and P0:C100, as well as PEG-PLGA and P0:C100, the incubation with FeHIF resulted in several significantly different decreases in surface charge. P100:C0 showed the lowest difference in ζ potential, while the highest effect was observed for P20:C80. Since FaHIF and FeHIF contained proteins, lipids, lipid digestion products, and bile salts (Figure S4A, Supporting Information), the attachment of the different compounds to NCs was subsequently analyzed. All investigated formulations showed protein corona formation to varying extents (Figure 7C). While

association factors in FaHIF ranged between 5% and 15% for all NC types, PEG-PLGA NCs showed an association below 4%. In FeHIF, association was highest for P50:C50 NCs, while increasing amounts of either PLGA or CDOC<sub>6</sub> lowered association. PEG-PLGA NCs showed an association below 2%. Regarding lipids and lipid digestion products, no biomolecule attachment could be found in FaHIF. In contrast, the association of free fatty acids (FFA) to all carrier types in FeHIF could be detected (Figure 7D). While P100:C0, P20:C80, and P0:C100 showed association factors above 50%, in hybrid NCs, the attachment significantly decreased with increasing proportions of PLGA. Thereby, P20:C80 and P50:C50 showed comparably low association factors, similar to PEG-PLGA NCs with an association below 20%. The analysis of different bile salts showed that in FaHIF, significant amounts of glycochenodeoxycholic acid were present in the biomolecular corona of all hybrid NCs (Figure 7E). Additionally, notable amounts of taurochenodeoxycholic acid were observed in the corona of P20:C80 carriers. While the attachment of glycodeoxycholic acid was detected for all NC types, it was not significantly different from the control. PEG PLGA NCs exhibited the lowest attachment for all bile salts. No attachment of taurocholic acid, glycocholic acid, or their derivatives was observed in FeHIF (Figure S4B, Supporting Information). In further experiments, the respective pure formulations were not analyzed due to their aforementioned insufficient physicochemical properties (P100:C0) and release behavior (P0:C100).

### 3.6. In Vitro Permeation and Local Anti-Inflammatory Effect

Besides pH-responsive drug release, spatial control over drug liberation to the colon is necessary to control inflammation. Thus, NCs not only need to be taken up by the intestinal cells, but limited permeation is also important. After confirmation of the biocompatibility of the curcumin-loaded hybrid NCs (Figure S5A,B, Supporting Information), their in vitro uptake into cells of inflamed intestinal mucosa was confirmed via confocal laser scanning microscopy. After 6 h incubation, signals of curcumin (yellow), lysosomal marker LAMP1 (lysosomal-associated membrane protein 1, red), and actin (green) were observed within the epithelial cells (Figure 8A–C) with no macroscopic differences in uptake of the hybrid NC formulations. The subcellular localization of the respective NCs within the epithelial cells was
<figure><img src="image_21.png" /><figcaption>**Figure 7.** Interactions between NCs and biomolecules after incubation with human intestinal fluids for 1 h at 37 °C. A) Hydrodynamic diameter and B) ζ potential of the formulations. C) Association factors of protein, D) free fatty acids, and E) bile salts to the different NCs in a fasted state (□) or fed state (○) human intestinal fluids. All data were expressed as mean ± SD of 3 independent experiments. Statistical analysis was performed by One-Way ANOVA with Bonferroni posthoc test. In (E), statistical differences are depicted, referring to the difference between the negative control and the sample. \*p < 0.05, \*\*p < 0.01, and \*\*\*p < 0.001.</figcaption></figure>

investigated by analyzing the co-localization of curcumin with the lysosomal marker LAMP1 using Mander's coefficients 1 and 2 (M1, M2) (Figure 8D). M1, which expressed the relative amount of LAMP1 signals that co-localized with curcumin signals relative to the total number of LAMP1 signals, was significantly higher for P50:C50 NCs compared to P80:C20. No significant difference was found regarding the co-localization of P20:C80 and P50:C50 NCs. The subsequent analysis of curcumin signals co-localizing with LAMP1 signals related to the total curcumin signals (M2) revealed significant variations within the hybrid NCs. Co-localization increased significantly with higher proportions of CDOC~6~. In permeation testing, all hybrid NCs showed an apparent permeability coefficient (P~app~) below $5 \times 10^{-8} \text{ cm}^{-1}\text{s}$ over the inflamed intestinal barrier (Figure 8E). Although P20:C80 hybrid NCs permeated at a higher rate, no significant difference was revealed in comparison to the other formulations. Thus, formulations can be regarded as lowly permeable, and local drug liberation can be achieved. Determination of the formulations' local anti-inflammatory effect on the in vitro model was conducted via enzyme-linked immunosorbent assays (ELISA). All curcumin-loaded hybrid NCs significantly decreased the release of the pro-inflammatory cytokine IL8 after 6 h, lowering its secretion to the level of the uninflamed control (Figure 8F). There was no distinguishing effect between the different hybrid NCs. Blank hybrid NCs did not significantly influence the IL8 secretion. After the initial treatment of 6 h, the NCs were removed, and IL8 release was re-evaluated at 24 h post-treatment, revealing similar IL8 con-

centrations compared to the inflamed control for all tested formulations (Figure 8G).

## 4. Discussion

Aiming to achieve disease-adaptive local drug release and anti-inflammatory effects, hybrid nanoparticles (NPs) combining PLGA and the amphiphilic cyclodextrin derivative CDOC~6~ were developed as a novel drug delivery system. Using a ratio-nal approach, various PLGA:CDOC~6~ ratios (P100:C0, P80:C20, P50:C50, P20:C80, P0:C100) were formulated and thoroughly characterized for their physicochemical properties, drug release kinetics, and interactions with ex vivo human intestinal fluids and in vitro inflamed intestinal mucosa models. Incorporating cyclodextrins into a PLGA matrix is known to influence nanoparticle properties, such as size, encapsulation efficiency, and release kinetics.

Our findings showed that the physicochemical properties of hybrid NCs strongly depend on the PLGA:CD ratio (Figure 2). For both, blank and drug-loaded NCs, increasing the PLGA proportion led to larger sizes and higher polydispersity, potentially based on the cyclodextrin-driven entrapment of the PLGA chains.[41] Pure CDOC~6~ NCs formed the smallest carriers due to their compact, self-encapsulating matrix structure.[42] Curcumin loading slightly increased the size and PDI of pure PLGA, CDOC~6~, and P20:C80 NCs, while P80:C20 and P50:C50 NCs remained unaffected, likely due to synergistic encapsulation of the drug
<figure><img src="image_22.png" /><figcaption>**Figure 8.** Confocal laser scanning micrographs of inflamed intestinal models incubated with A) P80:C20, B) P50:C50, and C) P20:C80 hybrid NCs loaded with curcumin (Cur) visualized by staining of lysosomal LAMP1 (red, Alexa Fluor 633), curcumin-labeled NCs (yellow), and actin (green, Alexa Fluor 488). Scale bars depict 20 µm. D) Manders' coefficients depicting the co-localization of curcumin in hybrid NCs and LAMP1-stained lysosomes were calculated relative to the total amount of LAMP1 (M1) and curcumin (M2) from micrographs after immunofluorescent staining. E) Apparent permeability coefficient (P<sub>app</sub>) of curcumin-loaded hybrid NCs over the inflamed in vitro models corrected for blank medium samples. IL8 secretion after F) 6 h and G) 24 h of the inflamed intestinal models quantified via ELISA. Data in (F) and (G) were normalized against the cytokine release of an inflamed, untreated control. Data in (D), (E), (F), and (G) were expressed as mean ± SD. (D), (F), and (G) show the results of 3 independent experiments. (E) was calculated based on 6 representative micrographs. Statistical analysis was conducted via One-Way ANOVA with Dunn-Sidak posthoc test. \*p < 0.05. \*\*p < 0.01, and \*\*\*p < 0.001.</figcaption></figure>

within CDOC₆ cavities and PLGA chains.[43] The high lipophilicity of curcumin (logP 3.3) facilitated interactions with the amphiphilic CDOC₆, as well as the hydrophobic PLGA, facilitating its encapsulation.[44,45] Further, the encapsulation efficiency was dependent on the PLGA:CD ratio, resulting in increased curcumin entrapment for hybrid NCs with higher PLGA proportions (Table 1). All formulations displayed negative ζ-potentials greater than -30 mV, ensuring colloidal stability. The negative ζ-potential results from the terminal carboxylic groups and hydroxyl anions attached between the sugar-based OH-groups of PLGA and CDOC₆, respectively.[46,47] Partial curcumin attachment onto the PLGA NC surface, in contrast, might have contributed to a decrease in surface charge in drug-loaded NCs.[48] Adding CDOC₆ resulted in additional curcumin encapsulation into the cyclodextrin's hydrophobic cavities, preventing the decrease of ζ-potential and leading to more stable NCs. Size and surface charge can significantly impact interactions with cells and tissues, particularly intestinal absorption. Smaller carrier systems generally show higher intestinal permeation rates than larger entities, while a low-magnitude negative surface charge has further been linked with adequate oral absorption rates.[49] Accordingly, permeation experiments revealed negligible transport of the hybrid NCs through the intestinal mucosa. The determined apparent permeability (P<sub>app</sub>) values below 5 × 10<sup>-6</sup> cm<sup>-1</sup>s (fraction ab-

sorbed <50%) indicate low permeability, thus supporting the intended local treatment strategy mechanistically.[50,51]

Raman and ATR-IR spectroscopy and DSC were further employed to understand the interactions of the excipients and the drug within the hybrid NC structure. Raman analysis proved the presence of hybrid NCs composed of both polymers (Figure 3). The PCA showed subtle right shifts of physical blends made from pure NCs had higher similarities to the score of CDOC₆, whereas the hybrid NC signals were distributed between the pure NCs. These findings indicated that the characteristic Raman signals of CDOC₆ were attenuated in hybrid NCs due to chemical interactions with PLGA. Similarly, strong interactions between PLGA and CDOC₆ were verified by DSC, where the prevailing van-der-Waals interactions between cyclodextrin and PLGA increased the energy necessary to divide both molecules, which led to a higher melting temperature visible in the respective thermograms (Figure S2, Supporting Information). After adding curcumin, the PLGA chains in the cyclodextrin cavity were partially displaced, decreasing the melting point. The melting point of curcumin in the loaded NCs could not be assessed via DSC, indicating the molecular distribution of the drug within the NCs. The DSC analysis revealed that curcumin, PLGA, and CDOC₆ form a coherent matrix, leading to chemical interactions between the substances, such as the competitive
displacement of PLGA chains, and explaining the observed
changes in physicochemical properties as well as drug encapsu-
lation and liberation.

The pH-responsive drug release observed in hybrid NCs is es-
sential for tailoring drug delivery during active and resting phases
of IBD, characterized by fluctuations in tissue pH. In vitro, exper-
iments demonstrated a biphasic release profile: an initial burst
release, likely from surface-adsorbed curcumin, followed by a
slower phase controlled by drug diffusion and matrix erosion
(Figure 4, Table S1, Supporting Information). Further, release ki-
netics could be controlled by adjusting the ratio of PLGA and
CDOC6. While the initial burst release of curcumin may be at-
tributed to superficially adsorbed drug molecules, the slower re-
lease phase could be based on the diffusion of entrapped drug
molecules through the carrier or to erosion of the matrix. In sil-
ico curve fitting indicates complex interference of diffusion and
erosion phenomena during drug release from the hybrid NCs
(Figures 6 and 7 and Table 2). Both excipients are potentially
subject to swelling as well as ester bond cleavage.[52,53] While
pure CDOC6 NCs followed a predominantly diffusion-driven
release, the hybrid formulations exhibited a combined influence
of PLGA susceptibility to ester bond cleavage and CDOC6 stabil-
ity. The employed Hopfenberg erosion model poorly described
the drug release mechanisms of hybrid NCs, likely due to lo-
cal phenomena such as swelling or drug recrystallization, which
traditional models do not capture. Polymer degradation or bulk
erosion is generally negligible at these very early time points of
drug release.[54] However, the water uptake and, thus, swelling
of the nanoparticles can explain the deviation of experimentally
obtained drug release from the diffusion-assuming theoretically
predicted release rates.[55] Logically, nanosized polymeric deliv-
ery devices that rely on diffusion-controlled drug release should
exhibit a very rapid release rate, given that the surface area-to-
volume ratio is significantly higher in nanospheres.[56] The fact
that nanoparticles based on cyclodextrin only (P0:C100) generally
show best fits with low standard error of regression (S) suggests a
likelihood of diffusion as well as surface erosion playing partly a
role within the total mass transport mechanisms with this partic-
ular matrix system when exposed to FeSSIF pH 5. However, the
susceptibility of PLGA outweighed the diffusion of CDOC6 in the
hybrid NCs. Accordingly, ester bonds of CDOC6 were most likely
cleaved in the acidic medium as well, but regarding their lower
number and the stability-providing acetal structure, CDOC6-rich
NCs showed highly controlled, diffusion-driven curcumin re-
lease. Following that premise, pure CDOC6 NCs released cur-
cumin only via diffusion control, underlining the importance of
the pH-dependent degradation and disease-adaptiveness of the
hybrid NC formulations. Combining in vitro release studies with
mathematical modeling revealed that hybrid NCs enable matrix-
dependent, pH-responsive drug release. This behavior supports
disease-adaptive drug delivery, with a basal release under physi-
ological conditions and enhanced release under acidic, inflamed
environments.[20,57,58] Among the hybrid formulations, P50:C50
demonstrated the most pronounced pH sensitivity, making it par-
ticularly promising for tailored IBD therapy.

After administration, NCs encounter different body fluids with
a heterogeneous composition of biomacromolecules, including
proteins, lipids, and salt molecules. Protein association with
NCs has been previously observed after intravenous injection

and the interaction with blood, but is less understood in case of
oral administration.[59–63] Initial studies with different simulated
intestinal fluids indicate that the intestinal biomolecular corona
plays a critical role in the drug-release behavior and uptake by
epithelial cells.[64] In line with previous studies, protein associ-
ation with PEG-PLGA NCs, which were used as a control, was
generally lower than for other carrier types, which did not have
pegylated groups mediating a stealth effect (Figure 7).[65] The
hydrodynamic diameter of all NC types remained essentially
unchanged after incubation in Fasted state Human Intestinal
Fluid (FaHIF), indicating minimal biomolecule attachment.
In contrast, incubation in Fed state Human Intestinal Fluid
(FeHIF) led to an increase in hydrodynamic diameter for all
carriers, suggesting the formation of a biomolecular corona. The
reduction in negative surface charge in both fluids confirmed
adsorptive processes, where molecules in FaHIF and FeHIF
shielded the surface charge or created a stealth effect on the NCs’
charged end groups. Notably, ζ potential reduction in FeHIF dif-
fered significantly between carrier types, with P80:C20 showing
the most pronounced decrease. These variations in ζ potential
and size likely reflect distinct biomolecule adsorption patterns
influenced by the PLGA:CDOC6 ratio. The biomolecular corona
composition depended on both, the NC matrix and the intestinal
fluid type: proteins dominated in FaHIF (Figure 8C), while lipids
played a more significant role in FeHIF (Figure 8D). In FaHIF,
pure PLGA NCs exhibited significantly higher protein associa-
tion than P80:C20, P20:C80, and PEG-PLGA NCs. This may be
due to differences in matrix composition and steric orientation,
where CDOC6 entrapped PLGA side chains, promoting the out-
ward orientation of longer PLGA chains and reducing protein
attachment. Consequently, for pure NCs (P100:C0, P0:C100), the
attachment of proteins was therefore higher. In FeHIF, P100:C0
and P0:C100 NCs showed lower protein attachment compared to
FaHIF, likely due to the competitive influence of free fatty acids
(FFAs), which form a major component of the biomolecular
corona in the fed state. While protein association factors for
P100:C0 NCs were below 5%, similar to PEG-PLGA NCs, FFA
association exceeded 70%, indicating that FFAs replaced proteins
on the NC surface. This trend was also observed in hybrid NCs,
where lower FFA attachment correlated with higher protein as-
sociation. In contrast, PEG-PLGA NCs exhibited minimal attach-
ment to both FFAs and proteins, underscoring the stealth effect
of pegylation. No FFA attachment occurred in FaHIF, which is
consistent with the low FFA concentrations in these fluids. Inter-
estingly, protein attachment was not concentration-dependent:
despite FaHIF containing significantly lower protein levels than
FeHIF, the proteins present were sufficient to saturate the NC
surfaces.[33,66] The quantification of bile salts revealed that the
biomolecular corona of the NCs contained significant amounts,
with glycochenodeoxycholic acid as the most relevant bile salt
associated with the NCs by mass, although glycocholic acid is
also found in high concentrations in HIFs.[31] While glycodeoxy-
cholic acid and taurochenodeoxycholic acid are also attached in
notable quantities, the overall association of bile salts was low
compared to the other biomolecules. Furthermore, in FeHIF, no
attachment of bile salts could be observed, likely based on the
high amount of FFAs associated with the NCs in the fed state.
Having similar amphiphilic characteristics as bile salts, they
most likely competitively replaced the bile salts in the corona.
After treatment of human in vitro models of IBD, NCs were taken up by the epithelial cell in substantial amounts and showed strong co-localization with lysosomal compartments (Figure 8A–D). Using Manders’ coefficient 1 (M1), a higher cellular uptake of CDOC₆-rich NCs was determined, probably mediated by the synergistic effects of biocompatible PLGA and sugar-based cyclodextrins and their interactions with cell-surface receptors and glycoproteins.[67] The findings suggested that the entrapment of the NCs in lysosomes hindered the permeation of the formulation through the cell layer (Figure 8E). In lysosomal compartments, where curcumin can only escape to the cytosol after its release from the NCs, higher M2 values indicated lower release from the NCs.[68–70] In contrast, lower co-localization hinted toward higher proportions of released drugs and successive lysosomal escape. The amount of curcumin co-localization increased with higher proportions of CDOC₆, proportional to the findings of the in vitro release study in acidified FeSSIF. Increasing proportions of PLGA showed higher susceptibility against acids in the in vitro release experiments. Consequently, the same effect elevated curcumin release in the acidic lysosomal environment for P80:C20 and P50:C50 NCs.[70–72] These findings supported the polymer-dependent release kinetics of the hybrid formulations in biorelevant conditions.

The observed high lysosomal uptake, entrapment, and decomposition of biodegradable NCs in the cellular barrier led to curcumin release, which allowed the drug to reach its site of action in the cell.[73,74] All curcumin-loaded NCs mediated a highly significant anti-inflammatory effect after 6 h of incubation due to the intracellularly released curcumin, as the hybrid NCs lowered the IL8 release to the basal level of the non-inflamed and untreated control (Figure 8F). The absence of significant differences between each curcumin-loaded NC type could be based on the significant anti-inflammatory effect and the comparable penetration capabilities of all presented formulations in combination with a limited overall difference between IL8 secretion in inflamed and non-inflamed tissue. However, these findings do not negate the necessity of different matrix ratios, as inflammatory events in vivo are not completely mimicked in the cellular model. The large mucosal surface, multiple sites of local inflammation, and different severity states demand flexible platforms, as presented here. Because of the mimicked intestinal transit time, an anti-inflammatory effect was observed for neither loaded nor blank NCs after 24 h, underlining the time-limited drug release of the formulation, also observed in the dissolution experiments. This was most likely based on the degradation of curcumin in the cells and the one-time treatment.[75] Based on the data obtained from in vitro models of the inflamed intestine, the formulation penetrated the tissue without permeation, leading to local drug liberation. This mechanism enabled a profound anti-inflammatory local effect in the intestinal tissue models of all tested hybrid NC formulations.

## 5. Conclusion

In conclusion, this study introduced an innovative, disease-adaptive hybrid nanocarrier system for oral delivery with significant potential for treating intestinal inflammatory diseases, such IBD. The combination of PLGA and CDOC₆ into hybrid NCs not only enabled tailor-made control over physicochemical properties

but also local and pH-responsive drug release tailored to the inflammatory state of intestinal tissue. Analysis of chemical interactions within the hybrid NCs and in silico modeling confirmed the critical role of excipient ratios in drug release mechanisms. The formation of a biocorona affected surface charge and colloidal stability, yet the nanocarriers maintained their integrity in human intestinal fluids. NCs showed mucosal accumulation and reduced cytokine secretion when applied to human in vitro models of the inflamed intestine. These results show the potential of PLGA:CD hybrid NCs as delivery systems for the oral administration of anti-inflammatory drugs, allowing the on-demand adaptation of the liberated drug dosage dependent on symptom severity in chronic inflammation disorders, such as IBD.

## Supporting Information

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

## Acknowledgements

J.S. and F.E.B.B. These authors contributed equally to this work. The authors thank Jannick Schicktanz for his contributions to the drug release experiments. This work was partly funded by the Cluster4Future PROXIDRUGS (BioDEL 03ZU1109GA, funded by the Federal Ministry of Education and Research, BMBF), by the German Research Foundation (DFG, Grant 414985841), and by the Cluster project ENABLE funded by the Hessian Ministry for Science and the Arts. This project has received funding from the European Union's Horizon 2020 research and innovation program under the Marie Skłodowska-Curie Grant Agreement No. 956851. The EUbOPEN project has received funding from the Innovative Medicines Initiative 2 Joint Undertaking under grant agreement No 875510. This Joint Undertaking receives support from the European Union's Horizon 2020 research and innovation program, EFPIA companies, and Associated Partners: KTH, OICR, Diamond, and McGill. This communication reflects the views of the authors and neither IMI nor the European Union, EFPIA, or any Associated Partners are liable for any use that may be made of the information contained herein. This work was further supported by a fellowship of the German Academic Exchange Service (DAAD).

Open access funding enabled and organized by Projekt DEAL.

## Conflict of Interest

The authors declare no conflict of interest.

## Author Contributions

J.S. performed conceptualization, data curation, formal analysis, funding acquisition, investigation, methodology, validation, visualization, writing – original draft, writing – review & editing. F.E.B.B. performed conceptualization, data curation, formal analysis, investigation, methodology, validation, visualization, writing – original draft, writing – review & editing. S.S. performed data curation, formal analysis, investigation, methodology, and writing – original draft. A.H. performed data curation, formal analysis, investigation, methodology, and writing – original draft. R.M. performed data curation, formal analysis, investigation, methodology, and validation. N.J. performed data curation, formal analysis, investigation, methodology, validation, visualization, writing – original draft, writing – review & editing. S.G. performed writing – review & editing. S.V.-K. performed conceptualization, methodology, writing – original draft, writing – review & editing.
S.M. performed data curation, formal analysis, investigation, methodology, validation, and writing – original draft. P.A. performed conceptualization, project administration, resources, supervision, writing – review & editing. M.W. performed conceptualization, funding acquisition, project administration, resources, supervision, writing – original draft, writing – review & editing.

## Data Availability Statement

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

## Keywords

amphiphilic cyclodextrins, biomolecular intestinal corona, disease-adaptive release, inflammation control, nanoparticles, pH-dependent release

Received: August 18, 2024

Revised: December 18, 2024

Published online: January 19, 2025

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