Redox-Responsive Polymeric Nanocarriers for Targeted Delivery of Bruceantin to Pancreatic Cancer

Abstract

Bruceantin (BCT), a natural product with promising anticancer properties, has limited clinical utility due to its hydrophobicity, low systemic bioavailability, and dose-limiting toxicity. Here, we report a redox-responsive polymeric nanoplatform based on a disulfide-linked diblock copolymer, poly (ethylene glycol)-block-poly(ε-caprolactone) (PEG-SS-PCL), functionalized with luteinizing hormone-releasing hormone (LHRH) peptides for the targeted delivery of BCT to pancreatic cancer cells overexpressing LHRH receptors. The nanoplatform exhibits favorable physicochemical characteristics, including optimal particle size (45.7 nm), zeta potential (+3.5 mV), and high drug encapsulation efficiency (74.8%). Notably, it demonstrates redox-responsive drug release under glutathione-rich conditions that mimic the tumor microenvironment. In vitro assays highlight the superior cytotoxicity of the LHRH-targeted BCT formulation (IC₅₀, 224.1 μg/mL) compared to the non-targeted formulation (IC₅₀, 279.5 μg/mL). Biodistribution studies in an orthotopic pancreatic cancer murine model revealed substantial accumulation of LHRH-targeted nanoparticles primarily in tumors following systemic administration, with minimal hepatic deposition. Furthermore, BCT-loaded nanoplatforms reduced tumor growth by ~72.4% relative to controls, with no detectable adverse effects. Collectively, this work highlights the potential of LHRH-functionalized nanoplatforms as a promising strategy for the targeted treatment of pancreatic cancer with BCT, offering enhanced therapeutic efficacy with minimal systemic toxicity and paving the way for future clinical translation of this natural product.

1. Introduction

Natural products have long served as a cornerstone of drug discovery, with more than 60% of anticancer drugs derived from natural sources.^{1, 2} Bruceantin (BCT), a quassinoid isolated from the stem bark of Brucea antidysenterica, a plant traditionally used in Ethiopia for cancer treatment, has demonstrated potent activity against lymphoma, leukemia, and myeloma cell lines.^3–6 Its anticancer effect arises primarily from inhibition of protein synthesis through binding to the ribosomal peptidyl transferase center, thereby blocking peptide bond formation. ^5–7 Despite these promising properties, the clinical application of BCT is severely restricted by hydrophobicity, low bioavailability, systemic toxicity, and non-specific organ distribution.^{8, 9}

Drug delivery systems (DDS) offer a powerful strategy to overcome these limitations by improving aqueous solubility, enhancing tumor targeting, enabling controlled release, and optimizing biodistribution of anticancer drugs.^8,10 DDS performance can be further augmented through ligand-mediated targeting, which allows preferential accumulation in malignant tissues.^10,11 Among available platforms, polymeric nanoparticles are particularly effective for hydrophobic drugs.^12–14 Amphiphilic block copolymers, such as poly(ε-caprolactone)-block-poly(ethylene glycol) (PEG-PCL), with, self-assemble in aqueous environments to form stable nanoparticles with high drug-loading capacity, biodegradability, and biocompatibility.^15,16 Furthermore, PEG-PCL chemistry enables conjugation with targeting ligands, restricting drug exposure in normal tissues while improving the therapeutic index.¹⁵ Redox-responsive nanoplatforms represent a promising strategy for cancer drug delivery by exploiting the elevated intracellular glutathione (GSH) levels in tumor cells ¹⁷. Nanocarriers incorporating redox-sensitive linkages, such as disulfide bonds, remain stable during systemic circulation but undergo rapid cleavage in the reductive tumor microenvironment, enabling triggered drug release. This selective intracellular activation enhances tumor drug accumulation, improves therapeutic efficacy, and reduces systemic toxicity¹⁸.

Pancreatic cancer (PC) remains one of the deadliest malignancies, with a 5-year survival rate of approximately 4%.^19,20 Its poor prognosis results from late-stage diagnosis, stromal desmoplasia that‘ creates physical barriers to drug penetration, and intrinsic chemoresistance driven by genetic and signaling abnormalities.^21–25 Novel therapeutic approaches that increase intratumoral drug delivery and overcome stromal resistance are urgently needed. ^26,27 Plant-derived compounds are promising candidates for the treatment of pancreatic cancer. Structural analogues of BCT, including brusatol, bruceine D, and bruceine H, have demonstrated activity against pancreatic cancer cell lines. ^28,29 However, the anticancer potential of BCT in pancreatic cancer tumors has not been investigated.

The luteinizing hormone-releasing hormone receptor (LHRH-R) is overexpressed in multiple cancers, including prostate (86%), ovarian and endometrial (80%), renal (80%), breast (50%), and pancreatic cancers (32–50%).^11,30–34 LHRH peptides are attractive targeting ligands, offering advantages over antibodies such as small size, low immunogenicity, efficient tumor penetration, and ease of synthesis.³⁵ Maleimide-functionalized PEG-PCL nanoparticles conjugated with LHRH peptides can significantly enhance tumor cell uptake, promoting deeper drug infiltration and stronger therapeutic responses.^{11, 33, 36}

Here, we developed a redox-responsive nanoplatform for targeted BCT delivery using a disulfide-linked PEG-SS-PCL copolymer and maleimide-PEG-PCL conjugated to LHRH peptides. The nanoplatform was systematically characterized in terms of particle size, morphology, drug loading, and release profile. Its cytotoxicity and uptake were assessed in KPC (WT 8069) pancreatic cancer cells, and in vivo biodistribution, therapeutic efficacy, and safety were evaluated in an orthotopic pancreatic cancer mouse model.

2. Experimental

2.1. Materials

Bruceantin (BCT) was purchased from MedChem Express (NJ, USA). Maleimide-PEG(5K)-PCL (10K) and methoxy-terminated PEG(5K)-SS-PCL (10K) were obtained from Ruixi Biotech (Xi’an, China). The LHRH peptide (QHWSYK(C)LR; purity >90%) was sourced from Pepmic Co. Ltd. (Suzhou, China), and Nile Red was acquired from MP Biomedicals (Irvine, CA, USA). Silicon 2,3-naphthalocyanine bis(trihexylsilyloxide) (SiNc) was supplied by Alfa Chemistry (Ronkonkoma, NY, USA). Dimethyl sulfoxide (DMSO, >99.9%), tetrahydrofuran (THF), and acetonitrile (99.9%) were purchased from Sigma-Aldrich (EMD Millipore, Burlington, MA, USA). Cell culture reagents, including Dulbecco’s Modified Eagle Medium (DMEM), fetal bovine serum (FBS), phosphate-buffered saline (PBS), and penicillin-streptomycin, were obtained from Gibco (Gaithersburg, MD, USA). Saline and Milli-Q water were used consistently throughout all experimental procedures. Additional chemicals and reagents required for cell culture, animal studies, and high-performance liquid chromatography (HPLC) analyses were procured as needed.

2.2. Functionalization of Mal-PEG-PCL with LHRH Peptides

The LHRH-targeted polymer was synthesized by conjugating the LHRH decapeptide to maleimide-terminated PEG-PCL (Mal-PEG-PCL) following a previously established protocol.³⁷ Briefly, 30 mg of Mal-PEG-PCL (2 μmol) was dissolved in 1.5 mL of acetonitrile, and 3.36 mg of LHRH peptide (2.5 μmol) was dissolved in 150 μL of DMSO. The solutions were combined, mixed using an Eppendorf MixMate, and incubated overnight at room temperature to enable maleimide-thiol conjugation. The reaction mixture was then added dropwise into 45 mL of pre-chilled diethyl ether (−20 °C) to precipitate the conjugate. The precipitate was collected by centrifugation at 7000 rpm for 10 min at 4 °C, dried under vacuum overnight, and stored at −20 °C until use.

2.3. Preparation of Non-Targeted and LHRH-Targeted BCT-Loaded Nanocarriers

Non-targeted nanocarriers (NT-NC) were prepared using the solvent evaporation method. Briefly, 50 mg of methoxy-terminated PEG-SS-PCL was dissolved in 1.5 mL of tetrahydrofuran (THF) and mixed with 5 mg of BCT dissolved in 500 μL of THF. Then, 2 mL of saline was added, and the mixture was stirred for 5 min. The THF was removed under reduced pressure using a rotary evaporator for 15 min, and the resulting dispersion was left at room temperature overnight to remove residual solvent. LHRH-targeted nanocarriers (LHRH-NC) were formulated by mixing methoxy-terminated PEG-SS-PCL and LHRH-PEG-PCL at a 9:1 ratio and following the same procedure. The final suspensions were centrifuged at 5500 rpm for 5 min using Amicon^® Ultra centrifugal filters (100 kDa MWCO; Millipore, St. Louis, MO, USA), to remove the free drug and filtered through a 0.2 μm nylon membrane to remove aggregates. For biodistribution and cellular uptake studies, BCT was replaced with fluorescent dyes (Nile Red or SiNc). The free BCT solution was prepared by dissolving 2 mg of the drug in 100 μL of dimethyl sulfoxide (DMSO). The resulting stock solution was subsequently diluted with saline to obtain a final volume of 2 mL and vortexed vigorously to ensure complete solubilization. ^{15, 38}

2.4. Characterization of Nanocarriers

Nanocarriers were characterized using attenuated total reflectance fourier-transform infrared spectroscopy(ATR-FTIR), proton nuclear magnetic resonance (¹H NMR), and X-ray photoelectron spectroscopy (XPS) (Thermo Fisher Scientific, Waltham, MA, USA) to confirm LHRH conjugation. Hydrodynamic size, polydispersity index (PDI), and zeta potential were measured using dynamic light scattering (DLS, Nano-ZS, Malvern Instruments, UK). Morphology was examined by cryogenic transmission electron microscopy (cryo-TEM, Thermo Fisher Scientific). BCT optical properties were analyzed using UV–visible spectroscopy (Shimadzu, Japan), and drug loading and release profiles were quantified by high-performance liquid chromatography (HPLC, Shimadzu, Japan).

2.5. Determination of BCT Loading in Nanocarriers

The absorbance spectrum of BCT in acetonitrile was recorded from 200 to 800 nm, with a maximum at 278 nm (Figure S1A) was observed. The BCT content in NT-NC and LHRH-NC was quantified using high-performance liquid chromatography (HPLC; Shimadzu, Japan) equipped with an Agilent ZORBAX C18 column (3.5 μm, 4.6 × 75 mm). Separation was performed under gradient elution of 70% acetonitrile and 30% water with 0.1% trifluoroacetic acid at a flow rate of 1 mL/min.¹³ Calibration curves were generated using BCT standard solutions in the concentration range of 50–350 μg/mL (Figure S1B). For assay, 50 μL of each formulation (NT-NC or LHRH-NC) was dissolved in 300 μL of acetonitrile to completely disrupt the nanocarriers and release the encapsulated drug. The amount of BCT was quantified by HPLC using the corresponding calibration curve (Figure S1C). Encapsulation efficiency (%) was calculated as the amount of BCT loaded in nanocarriers (mg)/initial amount of BCT (mg) × 100 %.

2.6. In Vitro Drug Release Study

BCT release from LHRH-NC was evaluated using a dialysis bag method (3.5 kDa MWCO, Spectrum Laboratories, USA) in 20 mL PBS (pH 7.4), with or without 5 mM GSH, at 37 °C under stirring (100 rpm) for 72 h. At set intervals, 0.5 mL aliquots were collected and replaced with fresh PBS. BCT concentration was determined by HPLC at 278 nm under the same conditions as described above.¹⁵ A GSH concentration of 5 mM was selected because it represents a commonly used value within the typical intracellular GSH range observed in cancer cells (1–10 mM).¹⁷ This concentration provides a conservative test of the system’s redox-responsiveness, demonstrating that triggered drug release can be achieved even at moderate GSH levels rather than only under optimal high-GSH conditions.¹⁷

2.7. In Vitro Cytotoxicity Studies

KPC-WT 8069 cells were cultured in DMEM with 10% FBS and 1% penicillin-streptomycin at 37 °C with 5% CO₂. Cells were seeded in 96-well plates and treated with blank nanocarriers, free BCT, NT-NC, or LHRH-NC for 24 h. Viability was assessed using the resazurin assay, and absorbance was measured at 570 and 660 nm.³⁹

2.8. In Vitro Cellular Internalization Studies

Nanocarriers were labeled with Nile Red (50 μg/mL) (Figure S2A) and incubated with KPC-WT 8069 cells seeded in 6-well plates (150,000 cells/well) for 24 h. After washing, intracellular fluorescence was quantified by microscopy (Keyence BZ-X700) and flow cytometry (BD Symphony A5). Nuclei were stained with DAPI for qualitative visualization of internalization.³⁸

2.9. In Vivo Biodistribution Studies

All animal procedures were performed in accordance with the Guidelines for Care and Use of Laboratory Animals of Oregon Health and Science University and approved by the Institutional Animal Care and Use Committee (IACUC) of Oregon Health and Science University (IP00000033).

Orthotopic pancreatic tumors were established in female C57BL/6 mice by injecting 1 × 10⁵ KPC-WT 8069 cells in 20 μL of PBS/Matrigel (1:1) mixture. After 10 days, mice were injected IP with NT-NC or LHRH-NC encapsulating SiNc (50 μg/mL) (Figure S2B). Mice were euthanized after 24 h, organs collected, and fluorescence imaged using the Pearl Impulse system. To quantitatively compare the biodistribution of NT-NC and LHRH-NC, fluorescence signals from excised tumors and major organs were analyzed using ImageJ software. The corrected total fluorescence (CTF) was calculated by subtracting background fluorescence from the integrated fluorescence intensity according to the equation: CTF = Integrated Density – (Area of selected organ × Mean Background Fluorescence).^37,40,41

2.10. In Vivo Therapeutic Studies in Orthotopic KPC Model

Mice were randomly divided into three groups (n = 4): saline, free BCT (3 mg/kg), and LHRH-NC (3 mg/kg). Treatments were administered IP once daily for nine doses starting three days post-inoculation. On day 14, mice were euthanized, and tumors were collected and weighed. Blood samples were collected via cardiac puncture for comprehensive hematological and biochemical analysis at the IDEXX laboratory (Portland, OR, USA), including complete blood count and serum chemistry parameters.¹⁵

2.11. Statistical Analysis

Data are presented as mean ± SD. One-way ANOVA was used for group comparisons using GraphPad Prism (GraphPad Software, CA, USA) and Origin Pro 2025 (OriginLab Corporation, Northampton, MA, USA). A p-value of less than 0.05 was considered statistically significant (*p < 0.05); p-values below 0.01 (**p < 0.01), 0.001 (***p < 0.001), and 0.0001 (***p < 0.0001) were considered highly significant.

3. Results and Discussion

3.1. Preparation and Characterization of BCT-Loaded Targeted Drug Delivery System

The anticancer activity of BCT was previously investigated in phase I and II clinical trials in patients with malignant melanoma and metastatic breast cancer.^{7,42, 43} Although BCT showed promising therapeutic efficacy, its clinical development was discontinued due to dose-limiting toxicities, including nausea, vomiting, fever, and hypotension.⁷ Recently, BCT has been re-evaluated in preclinical models of various cancers, but its unfavorable physicochemical properties, particularly hydrophobicity and systemic toxicity, remain significant barriers to clinical translation.⁴⁴

To address these limitations, we developed a self-assembled amphiphilic block copolymer nanoplatform for targeted BCT delivery. The system is based on a reduction-responsive, mid-disulfide-functionalized diblock copolymer composed of poly(ethylene glycol)-block-poly(ε-caprolactone) (PEG(5K)-SS-PCL(10K)), chosen for its biodegradability, biocompatibility, non-immunogenicity, and ability to passively accumulate in tumors.^15,38 In this design, the hydrophobic PCL segment forms the inner core of a nanoparticle to encapsulate BCT, while the hydrophilic PEG segment introduces the outer shell, providing aqueous stability and prolonged systemic circulation. ⁴⁵ To enable active targeting, nanoparticles were functionalized with LHRH peptides via maleimide-thiol conjugation, as illustrated in Figure 1A. ³⁷

Figure 1.

Schematic illustration of the LHRH-targeted nanocarrier loaded with BCT. (A) Conjugation of LHRH-targeting peptide to maleimide-terminated PEG-PCL (Mal-PEG-PCL) via thiol–maleimide coupling (PEG chain is shown in blue and the PCL chain in red). (B) encapsulation of BCT into polymeric nanocarriers composed of 10% LHRH-PEG-PCL and 90% methoxy-PEG-SS-PCL, resulting in redox-sensitive, tumor-targeted nanoparticles.

The LHRH-PEG-PCL conjugate was synthesized by covalent coupling of maleimide-terminated PEG(5K)-PCL(10K) (Mal-PEG-PCL) with the thiol group of the D-cysteine residue in the LHRH decapeptide, following our previously reported protocol.³⁷ (Figure 1A). Successful conjugation was confirmed using ATR-FTIR, ¹H NMR, and XPS. ATR-FTIR spectra of LHRH, Mal-PEG-PCL, and the LHRH-PEG-PCL conjugate are shown in Figure 2A. Mal-PEG-PCL exhibited characteristic peaks at 1720 cm⁻¹ and 1180 cm⁻¹ (C=O and C–O stretching, respectively), as well as peaks at 2949 cm⁻¹ (C–H stretch of CH₂) and 1239 cm⁻¹ (C–N bond). After conjugation, LHRH-PEG-PCL displayed absorption bands corresponding to both LHRH and Mal-PEG-PCL, including a distinct peak at 1525 cm⁻¹ (C–N stretching vibration of the amide bond), an intensified shoulder at 1630 cm⁻¹ (C=O stretching of the peptide backbone amide), and a peak at 798 cm⁻¹ (C–S bond), all confirming successful peptide conjugation. ⁴⁴

Figure 2.

(A) ATR-FTIR spectra of Mal-PEG-PCL (red), LHRH peptide (green), and the LHRH-PEG-PCL conjugate (blue). XPS elemental deconvolution spectra of (B) N1s for Mal-PEG-PCL, (C) S1s for LHRH-PEG-PCL, and (D) N1s for LHRH-PEG-PCL.

¹H NMR spectra further supported the presence of conjugation (Figure S3A). Mal-PEG-PCL exhibited methylene proton peaks of the caprolactone segment at 1.5, 1.7, 2.4, and 4.12 ppm, as well as PEG-related peaks at 3.48 ppm (methoxy, OCH₃) and 3.61 ppm (methylene, CH₂).^{46, 47} The characteristic maleimide proton signal at 6.7 ppm disappeared after conjugation with LHRH (Figure S3B), indicating successful covalent coupling.¹⁵ XPS analysis confirmed these findings. Full survey scans of Mal-PEG-PCL showed peaks corresponding to C1s, N1s, and O1s (Figure S4C), while LHRH-PEG-PCL displayed an additional S1s peak from the sulfur-containing peptide (Figure S4F). Deconvoluted N1s spectra revealed C–N and O=C–N bonds at 399.9 and 400.9 eV, respectively, in Mal-PEG-PCL (Figure 2B). In contrast, LHRH-PEG-PCL exhibited an additional peak at 401.3 eV, corresponding to C=N–C from the conjugated peptide (Figure 2D).³⁷ Fine analysis of the S1s spectrum of LHRH-PEG-PCL further confirmed the presence of C–S bonds at 168 eV, originating from the peptide (Figure 2C).

Non-targeted nanocarriers (NT-NC) were prepared using only methoxy-terminated PEG-SS-PCL and loaded with BCT via the solvent evaporation method. Targeted nanocarriers (LHRH-NC) were formulated by mixing LHRH-PEG-PCL with methoxy-terminated PEG-SS-PCL at a 1:9 ratio (Figure 1B). PEG-SS-PCL enables efficient encapsulation of hydrophobic drugs within the PCL core, even at relatively low polymer concentrations.⁴⁸ The amount of encapsulated BCT was quantified using high-performance liquid chromatography (HPLC) (Figure S3).⁹ Both formulations demonstrated high encapsulation efficiency, averaging 72.2% for NT-NC and 74.8% for LHRH-NC, underscoring the robust drug-loading capacity of the system. These values are in good agreement with our previous reports demonstrating ~75% encapsulation efficiency for hydrophobic drug molecules such as zimlovisertib in PEG-PCL nanoparticles.^15,38. Dynamic light scattering (DLS) measurements showed that NT-NC had an average hydrodynamic diameter of 38.5 nm with a polydispersity index (PDI) of 0.157 ± 0.005, indicating a uniform particle population. LHRH-NC exhibited a slightly larger size of 45.7 nm with a PDI of 0.181 ± 0.003, likely due to peptide conjugation (Figure 3C). The zeta potential shifted from –1.72 ± 0.23 mV for NT-NC to +3.5 ± 0.49 mV for LHRH-NC, consistent with the addition of positively charged LHRH peptides to the nanoparticle surface.³⁷

Figure 3.

(A) Cryo-TEM image of LHRH-NC (inset: magnified view of the region indicated by the yellow dashed box). (B) Particle size distribution of LHRH-NC determined from cryo-TEM analysis. (C) Dynamic light scattering (DLS) profiles showing the hydrodynamic diameters of NT-NC and LHRH-NC. (D) In vitro BCT release kinetics from LHRH-NC under physiological conditions (pH 7.4, 37 °C) in the presence (5 mM) and absence of glutathione (GSH). Data are presented as mean ± standard deviation (n = 3) and analyzed by two-way ANOVA *p< 0.0001.

Cryogenic transmission electron microscopy (cryo-TEM) confirmed the spherical morphology of LHRH-NC (Figure 3A). Particle size analysis by cryo-TEM showed a uniform distribution with an average diameter of 18.5 nm (Figure 3B), smaller than the hydrodynamic size measured by DLS. This discrepancy likely reflects the absence of the hydration layer in cryo-TEM measurements, which contributes to larger sizes in DLS.⁴⁹ Importantly, LHRH conjugation did not significantly alter particle size, zeta potential, or encapsulation efficiency, confirming that surface modification preserved the fundamental physicochemical properties of the nanoplatform.

In vitro cumulative drug release was evaluated under physiological conditions, leveraging the redox-responsive nature of the disulfide-functionalized PEG-SS-PCL diblock copolymer. Intracellular glutathione (GSH) levels (1–10 mM) are much higher than extracellular/plasma levels (2–5 μM). ⁵⁰ This redox gradient is widely exploited to design GSH-responsive nanocarriers that remain stable in circulation but rapidly disassemble and release their cargo in the high-GSH intracellular environment. ^{17,18, 50,51} The high amount of intracellular GSH can effectively trigger BCT release from the nanoplatform, enabling selective drug release within tumor cells. To assess this behavior, BCT release was measured under sink conditions in phosphate-buffered saline (PBS, pH 7.4) at 37 °C.¹⁵ In the presence of 5 mM GSH, LHRH-NC released 80.7% of BCT within the first 6 h and reached approximately 90% after 12 h. In contrast, in the absence of GSH, drug release was slower and more sustained, with 48.0% and 68.5% released at 6 and 12 h, respectively (Figure 3D).

Drug release profiles from the formulations with and without the treatment were statistically compared over the 72-h study period. Two-way analysis of variance (ANOVA) indicated a significant effect of time and formulation type on cumulative drug release (p < 0.0001, n = 3), demonstrating that drug release increased progressively with increasing time. In addition, the formulation containing GSH exhibited significantly higher drug release compared to the formulation without GSH at most sampling points (p < 0.0001, n=3). However, no statistically significant difference was observed at the initial time point (0.25 h) and at the final plateau phase (72 h) (p > 0.05, n=3). Overall, these results confirm the presence of GSH significantly enhanced drug release behavior (Figure S5A). These results demonstrate that LHRH-NC can respond to the reductive intracellular environment, enabling accelerated drug release in GSH-rich tumor cells. This redox-sensitive release profile highlights the potential of LHRH-NC as a controlled, targeted drug delivery platform capable of site-specific therapeutic release.^{50, 51}

The stability of the formulations was also investigated, as it is a critical parameter influencing biological performance and applicability. Both non-targeted and LHRH-targeted nanocarriers were stored at 4 °C for 4 weeks, and their physicochemical properties, including drug content, particle size, and zeta potential, were monitored (Figure S6). No significant changes were observed in any of these parameters during the study period. BCT content remained stable, indicating efficient retention within the nanoparticles. Similarly, particle size and zeta potential remained consistent, confirming the maintenance of colloidal stability and surface properties over time.¹⁵ These results demonstrate that both NT-NC and LHRH-NC are physically stable, supporting their suitability for subsequent therapeutic applications.

3.2. In Vitro Evaluation of BCT-Loaded Drug Delivery System

The in vitro anticancer activity of BCT was evaluated in a pancreatic cancer cell line (KPC-WT 8069) cultured in 96-well plates (Figure 4A). Cell viability was assessed using the resazurin assay to compare the cytotoxic effects of blank nanoparticles (LHRH-PEG-PCL/PEG-SS-PCL), free BCT, NT-NC, and LHRH-NC, with equivalent BCT concentrations across all groups.³⁹

Figure 4.

Experimental design for evaluating the anticancer efficacy of BCT-loaded nanoparticles in pancreatic cancer. (A) In vitro studies using KPC-WT 8069 pancreatic cancer cells. (B) In vivo studies in mice bearing orthotopic KPC tumors. (C) Designed for Cell-Penetrating Drug Delivery.

The blank nanoparticles exhibited no detectable cytotoxicity at concentrations up to 1000 μg/mL, confirming their biocompatibility (Figure S7). In contrast, free BCT and BCT-loaded nanocarriers (NT-NC and LHRH-NC) showed dose-dependent cytotoxicity. Statistical comparison with the control groups showed that even the lowest tested concentration (5 μg/mL) produced a significant decrease in viability (p<0.01) for Free BCT. In contrast, significant cytotoxic effects were first observed at concentrations of 25 μg/mL (p<0.01) for LHRH-NC and 50 μg/mL (p < 0.05) for NT-NC. Furthermore, all higher concentrations resulted in highly significant reductions in cell viability compared with the control (p < 0.001), specifically at concentrations ≥25 μg/mL for Free BCT, ≥50 μg/mL for LHRH-NC, and ≥100 μg/mL for NT-NC. These findings indicate a strong dose-dependent inhibition of cell survival, with Free BCT exhibiting the earliest and most pronounced cytotoxic effect among the tested formulations.

Among the formulations, LHRH-NC displayed enhanced cytotoxicity compared to NT-NC, with IC₅₀ values of 224.1 μg/mL (Figure 5C) and 279.5 μg/mL (Figure 5B), respectively. This difference was statistically significant (p < 0.001), indicating the benefit of LHRH surface modification in improving therapeutic efficacy (Figure S5B). This improvement is likely due to the enhanced cellular uptake mediated by LHRH surface modification, leading to greater intracellular accumulation of BCT.¹¹ Free BCT demonstrated the highest cytotoxicity, with an IC₅₀ of 203.4 μg/mL (Figure 5A), which was significantly lower than those of both LHRH-NC (p = 0.01) and NT-NC (p < 0.0001), consistent with its immediate availability in solution. In contrast, nanocarrier-based formulations rely on controlled drug release, which may delay the onset of cytotoxic effects. While free BCT provides rapid activity in vitro, nanocarriers, particularly LHRH-NC, offer the added advantages of targeted delivery and potential reduction in systemic toxicity in vivo. ⁵²

Figure 5.

Cell viability of KPC-WT 8069 cells treated with free BCT, NT-NC, and LHRH-NC at drug concentrations ranging from 5 to 1000 μg/mL for 24 h. (A) Free BCT, (B) NT-NC, (C) LHRH-NC. (Statistical significance was determined using one-way ANOVA; *p < 0.05, **p<0.01, ***p < 0.001. ****p < 0.0001) (D) Cellular uptake of Nile Red–labeled NT-NC and LHRH-NC in KPC-WT 8069 cells was analyzed by flow cytometry after 24 h incubation,

To further evaluate the role of LHRH targeting in intracellular delivery, cellular uptake was examined using Nile Red–labeled nanoparticles (see Table S1 for colloidal properties and encapsulation efficiencies of Nile Red loaded NPs). After 24 h of incubation, fluorescence microscopy revealed minimal uptake of NT-NC, evidenced by weak red fluorescence. In contrast, LHRH-NC showed ~4.2-fold higher intracellular fluorescence intensity than NT-NC, with signals predominantly localized in the perinuclear region, indicating efficient cytoplasmic delivery. Nuclear co-staining with DAPI confirmed that the nanoparticles did not penetrate the nucleus (Figure 6).

Figure 6.

Representative fluorescence microscopy images of KPC-WT 8069 cells incubated with Nile Red–labeled NT-NC (top) and LHRH-NC (bottom) for 24 h. Red fluorescence corresponds to Nile Red–labeled nanoparticles, and blue fluorescence indicates DAPI-stained nuclei. Scale bar: 75 μm.

Quantitative uptake analysis by flow cytometry supported these findings. NT-NC produced a modest increase in fluorescence relative to untreated controls, whereas LHRH-NC demonstrated a pronounced ~3-fold increase in intracellular fluorescence compared to NT-NC, consistent with receptor-mediated endocytosis (Figure 5D). Together, these results demonstrate that LHRH surface modification significantly enhances nanoparticle internalization in KPC cells and improves the anticancer efficacy of the delivered BCT.

3.3. In Vivo Evaluation of BCT-Loaded Drug Delivery System

To evaluate the tumor accumulation and therapeutic efficacy of BCT-loaded LHRH-NC, we employed the KPC (LSL-Kras^12D/+; LSL-Trp53^R172H/+; Pdx1-Cre) pancreatic cancer murine model. This model closely recapitulates the clinical features of human pancreatic ductal adenocarcinoma (PDAC), including its highly aggressive nature, associated cachexia, and, critically, the desmoplastic stroma.^{24, 25, 27, 53, 54} The stromal barrier, which is a major obstacle to drug penetration in PDAC, makes this model particularly relevant for testing nanocarrier-based drug delivery systems. By maintaining these tumor microenvironmental challenges, the KPC model provides a stringent and clinically meaningful platform to determine whether our LHRH-targeted nanoparticles can overcome stromal resistance, enhance intratumoral accumulation of BCT, and improve therapeutic efficacy. The model was established by orthotopic implantation of KPC (WT 8069) cells into the pancreas of immunocompetent C57BL/6 mice (Figure 4B and C).^55,56

To assess the biodistribution and tumor accumulation efficiency of non-targeted nanocarriers (NT-NC) and LHRH-functionalized nanocarriers (LHRH-NC), both formulations were loaded with the near-infrared (NIR) fluorescent dye silicon naphthalocyanine (SiNc) (see Table S1 for colloidal properties and encapsulation efficiencies of SiNC loaded NPs).³⁸ Following a single intraperitoneal (IP) injection, mice were euthanized after 24 h, and fluorescence imaging of resected tumors and major organs was performed. As shown in Figure 7A, both NT-NC and LHRH-NC accumulated at the tumor site, likely due to the enhanced permeability and retention (EPR) effect facilitated by their nanoscale size. However, LHRH-NC exhibited nearly a 2.5-fold higher fluorescence intensity in tumors compared to NT-NC (Figure 7B), and this difference was statistically significant (p < 0.05), underscoring the contribution of active targeting through LHRH peptides. Importantly, LHRH-NC displayed near-complete clearance from major organs at 24 h post-injection, with only minimal signal detected in the liver. In contrast, NT-NC showed markedly higher liver accumulation, with fluorescence intensity approximately 2.4-fold greater than that observed for LHRH-NC (p<0.05), indicating predominant non-specific hepatic uptake. Collectively, these results highlight the superior tumor-targeting capability of LHRH-conjugated nanocarriers and their potential to enhance intratumoral drug delivery while reducing systemic exposure and off-target toxicity.

Figure 7.

(A) NIR fluorescence images of resected KPC tumors and major organs collected 24 h after IP administration of SiNc-loaded NT-NC and LHRH-NC. (B) Quantitative analysis of corrected total fluorescence intensity in tumors and major organs from both treatment groups, measured using ImageJ. Data are presented as mean ± SD (n = 3), analyzed by one-way ANOVA *p< 0.05.

Since LHRH-equipped nanocarriers demonstrated enhanced accumulation in KPC tumors compared to non-targeted ones, we next evaluated the therapeutic efficacy of BCT delivered by LHRH-NC in comparison with a free drug. Three days after orthotopic inoculation of KPC cells into the pancreas, mice received daily IP injections of either free BCT or BCT-loaded LHRH-NC at a dose of 3 mg/kg for nine consecutive days (Figure 8A). Control animals were treated with saline. On day 14 post-inoculation, mice were euthanized, and tumors and blood were collected for analysis.

Figure 8.

(A) Schematic timeline of the treatment protocol: mice received IP injections of saline (control), free BCT (3 mg/kg), or LHRH-NC (3 mg/kg) once daily for nine consecutive days, starting three days after orthotopic tumor implantation. (B) Representative images of tumors collected on day 14 post-implantation. (C) Quantitative analysis of tumor weights for each treatment group. Data are presented as mean ± SD (n = 3). Statistical significance was determined using one-way ANOVA; ***p < 0.001, ****p < 0.0001.

Both free BCT and LHRH-NC treatment groups exhibited tumor growth inhibition relative to the control group (Figures 8B and S6). However, tumor weight analysis revealed a markedly pronounced antitumor effect in the LHRH-NC group, with a 72.4% reduction in tumor weight compared to controls, whereas free BCT achieved only a 38.0% reduction (Figure 8B). Consistently, tumors treated with LHRH-NC exhibited nearly two-fold greater suppression compared to free BCT at equivalent doses (Figure 8C and S8). Notably, unlike control mice, animals treated with LHRH-NC showed no signs of ascitic fluid accumulation, a common complication in advanced pancreatic cancer that reflects severe tumor burden and peritoneal inflammation.⁵⁷ The absence of ascites in treated mice highlights an additional therapeutic advantage of the targeted nanocarrier system beyond tumor shrinkage (Figure S9).

The systemic toxicity of free BCT and BCT-loaded LHRH-NC was evaluated by monitoring body weight changes and performing comprehensive blood analyses, including complete blood count (CBC) and serum chemistry.^{37, 51} As shown in Figure 9A, a slight, transient weight loss was observed in mice treated with both free BCT and LHRH-NC at the beginning of treatment; however, no significant differences in overall body weight were detected compared to the saline control. To further assess systemic safety, blood samples collected at euthanasia were analyzed for hematological and biochemical parameters, including liver function markers (ALP, AST, ALT), kidney function markers (BUN, creatinine), muscle injury marker (creatine kinase, CK), complete blood counts (RBC, WBC), differential leukocyte counts, plasma proteins, and electrolytes (Figure 9B–E). No statistically significant alterations were observed in any of these parameters across the control, free BCT, and LHRH-NC groups. These findings demonstrate that LHRH-NC not only delivers superior antitumor efficacy but also maintains a favorable systemic safety profile, supporting its potential as a well-tolerated and effective therapeutic strategy.

Figure 9.

(A) Body weight changes of mice during treatment with saline, BCT, or BCT-loaded LHRH-NC. (B) Serum levels of total protein, albumin, and globulin. (C) Kidney function markers: blood urea nitrogen (BUN) and creatinine. (D) Electrolyte levels. (E) Liver function markers, including alanine aminotransferase (ALT), alkaline phosphatase (ALP), aspartate aminotransferase (AST), and muscle injury marker creatine kinase (CK). (F) Complete blood counts (RBC, WBC). Data in panels B–F are presented as mean ± SD (n = 3) and were analyzed using one-way ANOVA.

4. Conclusion

In this study, we developed a redox-responsive, LHRH-targeted polymeric nanoplatform for the delivery of the natural anticancer agent BCT to pancreatic tumors. The nanocarriers exhibited favorable physicochemical properties, high drug encapsulation efficiency, and redox-triggered release under intracellular conditions rich in glutathione. In vitro studies demonstrated enhanced cytotoxicity and cellular uptake of LHRH-targeted nanocarriers in KPC pancreatic cancer cells, confirming the effectiveness of receptor-mediated targeting. In vivo, LHRH-NC achieved superior tumor accumulation, preferentially delivered BCT to orthotopic KPC tumors, and elicited a significantly greater antitumor effect than free BCT, while maintaining a favorable systemic safety profile. Notably, the targeted delivery system also reduced ascitic fluid accumulation, highlighting an additional therapeutic benefit. Collectively, these findings demonstrate that LHRH-functionalized, redox-responsive nanocarriers provide a potent and well-tolerated platform for targeted pancreatic cancer therapy, offering a promising strategy to overcome the challenges of poor drug bioavailability, off-target toxicity, and stromal barriers, and paving the way for potential clinical translation of BCT.

Data Availability Statement

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