Fabrication of drugs loaded UiO-66 nanoparticles loaded core-shell nanofibers: investigation of antiproliferative activity and apoptosis induction in lung cancer cells

ABSTRACT

Nanotechnology-enabled drug delivery has been demonstrated to be a superior cancer treatment. Lung, liver, brain, breast, and cervical cancer cells have all demonstrated extreme cytotoxicity when exposed to electrospun nanofibers encasing anti-cancerous drugs. Herein, we developed Polyurethane (P)/Carboxymethyl chitosan (C)/Polyethylene oxide (P) (termed as PCP) core-shell nanofibers were loaded with gemcitabine (GEM) and folic acid (FA) for regulated GEM and FA release towards the apoptosis of A549 and H1299 cells. Analysis using scanning electron microscopy (SEM) and x-ray diffraction (XRD) were used to characterise the fabricated nanocarriers. Synthesised nanofibers have been studied for their drug-loading and release efficacy of GEM/MOFs and FA/MOFs. The cytotoxicity data suggested that PCP loaded with GEM and FA might be employed to treat a wide range of cancers. The cell morphological investigation was examined by the acridine orange and ethidium bromide staining, and the DAPI staining assessed the nuclear damage of the cells. The flow cytometry methods investigated the hallmark of apoptosis. There is an opportunity for promising GEM delivery for lung cancer treatment with the FA-GEM/UiO-66 loaded-PCP core-shell nanofibers.

KEYWORDS:

• Carboxymethyl chitosan
• gemcitabine
• core-shell nanofibers
• folic acid
• UiO-66
• metal-organic framework
• lung cancer

Introduction

Using a drug delivery system that responds to stimuli effectively combating cancer [1]. Several stimuli, such as changes in temperature, light, pH, reagent, etc., can trigger drug release at target areas. A pH-responsive system has attracted much attention as a potential source of stimulation [2]. Tumour cell environments are more acidic (pH 5.0–6.0) than normal tissues. Hence, a pH-mediated drug delivery approach, highly sensitive to a tiny pH shift, would release encapsulated drugs while passing via the tumour areas while remaining reasonably stable in healthy tissues to lessen adverse effects [3]. A pH-sensitive substance triggers the release of the drug in the pH-responsive drug delivery system. A polymer coating with good biocompatibility and pH sensitivity is often applied to a delivery vehicle [4]. Carboxymethylcellulose (CMC) is a natural anionic polymer that dissolves in water and is one of several pH-sensitive materials [5]. Excellent pH sensitivity, high hydrophilicity, non-toxicity, and high biocompatibility have led to its widespread application in the drug delivery system for pH-mediated drug release [6].

A novel class of materials, metal-organic frameworks (MOFs) are combined metal ions and ligands bound together in complex configurations [4]. Physical stability, high chemical simple surface modifications, cheap cost, and quick fabrication are all hallmarks of MOFs. Metal-organic frameworks (MOFs) have expanded and now include energy storage, catalysis, and pollution control. Because of their inexpensive cost and excellent biocompatibility, MOFs have found widespread use in the biotechnology industry. MOFs might be carriers for transporting functional materials like DNA, anti-tumour drugs, and enzymes to their intended destinations [7]. ZIF-8 MOFs, for instance, have been employed as a drug delivery carrier to distribute 5-fluorouracil (5-FU) through pH as a trigger, with impressive results [8]. Gemcitabine (GEM) is delivered to tumour locations for cancer treatment using UiO-66 MOFs containing Ag nanoparticles. One major drawback of MOFs, especially compared to porous silica and other drug delivery materials, is that their tiny specific surface area reduces the amount of drug that can be loaded onto the material [9]. Although MOFs have seen extensive usage in the drug delivery industry, only a small fraction of those designs aim to boost specific surface area. One technique to promote a material’s surface area is improving its porosity [10].

Strong candidates for anticancer drug delivery are the electrospinning core-shell nanofibrous frameworks provided by the coaxial electrospinning technique because of the constant release of drugs from nanofibers without any release (early release period and subsequent incessant release period). It’s also possible that the outer shell of core-shell nanofibers will serve as a shield for the drug-encapsulated core [11]. Since the drug is released in two phases, first from the core layer to the shell layer and then from the shell layer to the blood flow, suitable for killing tumour cells, core-shell nanofibers may provide an effective method for prolonged anticancer drug release. However, gemcitabine (GEM), folic acid, paclitaxel, 5-fluorouracil, and another pristine anticancer drug may not be as successful in treating cancer as local delivery of anticancer drugs from core-shell nanofibers [12]. Hydrophilic polymers such as polyvinyl alcohol, polyethylene oxide, gelatin, chitosan, and derivatives are commonly used for the core layer, where anticancer drugs are typically introduced [13]. Due to its biocompatibility, biodegradability, and inexpensive cost, chitosan has been employed in drug delivery systems (DDSs). Nonetheless, its application in DDSs is constrained by its adsorption capacity and low solubility (resulting from hydroxyl and amine groups in the chitosan backbone). Due to its high solubility, chitosan has been carboxymethylated and is used as a biocompatible pH-sensitive polymer in biomedical efforts as a DDS of anticancer drugs [14]. CMC is safe because it has a carboxymethyl group on the amino and hydroxyl positions of the glucosamine units that make up chitosan. Compared to chitosan, CMC excels in numerous key areas, including water solubility, stability, and antibacterial action [15].

Among the many types of cancer that may be treated with gemcitabine (GEM), lung cancer is among the most targeted [16]. The biggest drawback to using GEM in chemotherapy is that it is rapidly metabolised into an inactive metabolite. For this reason, cutting-edge treatment approaches are needed urgently [17]. Nanocarriers have been developed using various materials, from polymers and lipids to inorganic metals. A pH-stimuli-sensitive drug delivery system might improve drug accumulation at tumour locations. Doing so makes it possible to improve the therapeutic value of potent drugs while decreasing their harmful and unwanted side effects [18–20].

Active targeted delivery using tumour markers is possible thanks to the folic acid receptor (FR) [21]. Most tumour cells highly express the FR (or FOLR1) alpha isoform. However, active TAMs have an abundance of FR (or FOLR2). Targeted delivery into tumour tissues has been described using FA-conjugated micelles and liposomes. Fewer research has focused on FA-conjugated delivery methods for targeting FR-positive TAMs [22]. In contrast, much attention has been paid to tumour types with high FR expression in tumour cells. Tumour-associated macrophages (TAMs) are crucial in tumour development and progression, immunosuppression, metastasis, and angiogenesis. As revealed by Wei’s co-workers, overexpression of FR in tumour-associated macrophages (TAMs) of lung cancer was related to a poor prognosis [23–25]. They showed that FR-positive TAMs are a promising target for lung cancer treatment and fabricated a folate-modified lipoplex to target these cells specifically. However, active tumour targeting and efficient penetration using an FA-altered ultra-small nanocarrier have seldom been described [26–28].

Following the fabrication of the UiO-66, the GEM and FA were integrated into the UiO-66 MOFs. Polyurethane/Carboxymethyl chitosan/Polyethylene oxide (PCP) core-shell nanofibers. The fabricated nanofibers formulations for regulated release of FA and GEM towards A549 and H1299 cell killing were examined. Further, we examined the anticancer activity, drug release profile, and MOFs/drug encapsulated nanofibers’ capacity for efficient drug loading. The cell morphological investigation was examined by the acridine orange and ethidium bromide staining, and the DAPI staining assessed the nuclear damage of the cells. The flow cytometry methods investigated the hallmark of apoptosis.

Experimental section

Fabrication of GEM and FA loaded-MOFs

Zirconyl chloride octahydrate and terephthalic acid were reacted in dimethylformamide (DMF), yielding UiO-66 MOF nanoparticles. UiO-66 MOFs (10 mg/mL) were stirred in 100 mL of PBS for 50 minutes to load the FA and GEM into the MOFs. After 24 minutes of hand-stirring, the MOF suspension was centrifuged at 5000 rpm for 20 minutes after adding GEM and FA (each drug 100 mg). Then, distilled water was used for flushing the mixes. By relating the primary GEM and FA concentrations and the weight of FA and GEM in the residue following centrifugation, the final FA and GEM concentrations were determined using a UV-Vis set at 360 and 480 nm.

Fabrication of core-shell nanofibers

After immersing the CMC solution (5 wt%) in DI water with stirring for 5 h, the PEO/CMC core mixture was made by adding PEO solution (5 wt%) dissolved in DI water at a weight ratio of 2:8 PEO/CMC. The core-shell solutions consisted of PU (10%) in DMF. The dual capillary had an inside diameter of 1.270 mm and an outside diameter of 0.635 mm. Using two KD programmable syringe pumps. We adjusted the core solution flow rate to 0.5 mL/h and the shell solution flow rate to 0.3 mL/h, 0.5 mL/h, and 0.8 mL/h. The tip-collector distance of 15 cm and the applied voltage of 25 kV were held constant. We fabricated nanofibers on the collector using syringe pumps, a high-voltage DC power supply, and a spinning collector drum at 200 revolutions per minute. The conditions for all the electrospinning operations were constant: 25 ± 2°C and 50 ± 5% relative humidity.

Characterization tests

The crystallography of drug-loaded UiO-66 MOF was analysed using a mini-X-ray diffractometer with Cu Kα radiation (λ = 1.5418 Å, Japan). The morphology and structure of the as-fabricated drug-loaded UiO-66 MOF were examined using field emission transmission electron microscopes (TEM, Tecnai G2 20, Thermo) and field-emission scanning electron microscopy (SEM, ZEISS Sigma 300). UV-Vis spectroscopy (UH5300, Hitachi) determined the drug loading content at ex.480 nm. MTT assays were calculated by measuring the absorbance at 490 nm using a microplate reader (Biotek Cytation5, Biotek, United States). The cell morphology and damages were examined by a fluorescence microscope (Olympus CKX53, Tokyo, Japan).

GEM and FA loading, release, and encapsulation efficiency

The nanofibers were immersed in distilled water and DMF solutions. The GEM and FA absorbance was assessed using a UV-Vis at 360 and 480 nm to determine DLE and EE.

Fabricated nanofibers GEM and FA release behaviour was evaluated by 1 mL of PBS solution (pH 7.4 and 5.0). 2 mL of the solution was withdrawn at 1, 3, 6, 12, 24, 48, 72, 96, 120, 168, 336, 480, and 720-h to check the UV-Vis at 360 and 480 nm. But it was replenished with 2 mL of sterile PBS.

Cytotoxicity studies

The in vitro cell cytotoxicity was evaluated on A549 and H1299 lung cancer cells and the non-cancerous fibroblast cells (NIH3T3) cells using a standard MTT assay. Cells were seeded into 96-well plates at a density of 1 × 10⁴ per well. After incubation overnight, the previous medium was replaced with a fresh medium containing different concentrations of the samples and incubated for another 72 h. Afterwards, MTT solution (10 μL, 5 mg/mL) was added to each well. The medium containing MTT solution was discarded after 4 h, and dimethyl sulphoxide (DMSO,100 μL per well) was added to dissolved formazan crystals for 30 min with gentle shaking [29]. The multifunctional micropore detection board analysis system (Biotek Cytation5, Biotek, United States) measured the absorbance at 490 nm.

Morphological examination

The AO/EB staining technique has been widely applied to detect apoptotic cell death. A549 and H1299 (10⁴ cells/well) cells planted in a 6-well plate were treated with an IC₅₀ concentration of core-shell nanofibers, GEM-loaded core-shell nanofibers, and FA@GEM-loaded core-shell nanofibers and incubated for 24-h. Following the treatment, cells were stained with AO/EB (5 mg/ml) and incubated for 2 min. The unbound dye was removed by PBS wash, and the morphological variations were examined using fluorescence microscopy to examine the morphological changes of the cells [30].

The nuclear DAPI staining technique investigated the condensation or fragmentation commonly used to detect nuclear damage. A549 and H1299 (10⁴ cells/well) cells planted in a 6-well plate were treated with an IC₅₀ concentration of core-shell nanofibers, GEM-loaded core-shell nanofibers, and FA@GEM-loaded core-shell nanofibers and incubated for 24-h. Following the treatment, cells were stained with DAPI (5 mg/ml) and incubated for 2 min [31]. The unbound dye was separated by PBS wash, and the nuclear damage was observed using fluorescence microscopy to examine the nuclear damage of the cells.

Apoptosis staining

A549 and H1299 cells were seeded into 6-well dishes with a density of 2 × 10⁵ per well and incubated overnight. The cells were treated with core-shell nanofibers, GEM-loaded, and FA@GEM-loaded PCP nanofibers and incubated for 24-h. The flow cytometry apoptosis kit (Best BioScience, China). The cell death mechanism induced by NFs was investigated by Annexin V-FITC/PI Apoptosis Detection Kit according to the manufacturer’s instructions and analysed by FCM (BD) [32].

Statistical analysis

All data were presented as mean ± standard deviation (SD). The experimental results were analysed by one-way analysis of variance (ANOVA) by using GraphPad Prism software 8.0. P-values <0.05, 0.01, and 0.001 were considered statistically significant difference and marked with *, **, and ***, respectively.

Results and discussions

Fabrication and characterization of drug-loaded MOFs

UiO-66 MOFs were synthesised successfully, as evidenced by the presence of significant peaks at 2θ = 7.4°, 8.5°, and 25.7°, which correspond to the (111), (002), and (006) diffraction planes, respectively (Figure 1a). The diffraction bands in FA and GEM-loaded-MOFs were diminished compared to pure UiO-66 MOF bands. As a result of GEM and FA being trapped pores in the UiO-66, the X-ray difference between the pore cages and the porous framework is reduced, which in turn causes a drop in the diffraction bands of MOFs. The X-ray diffraction (XRD) profile of FA and GEM-loaded-MOFs, which showed amorphous molecules in the MOFs, showed no diffraction peaks. UV-Vis spectroscopy proved that GEM and FA were present in the GEM and FA loaded-UiO-66, respectively. The data showed that GEM and FA peaked sharply at 480 and 360 nm (Figure 1b). According to scanning electron microscopy (SEM) results, the morphologies of MOFs and GEM-loaded MOFs are comparable, with a particle size range of 50 to 100 nm (Figure 2a,b).

Figure 1. Morphological characterization of UiO-66 MOFs. (a) XRD profile of UiO-66 MOFs, UiO-66/GEM and UiO-66/FA. (b) UV–vis spectral analysis of UiO-66/GEM and UiO-66/FA.

Figure 2. Morphological investigation of MOFs. SEM images of UiO-66 MOFs (a) and GEM-loaded UiO-66 (b) Scale bar 100 nm.

Figure 3a–e shows the surface morphology of the nanofibers with diverse shell flow rates, including GEM/FA/PEO/CMC/UiO-66 nanofibers, PCP nanofibers, GEM/FA/UiO-66-loaded PCP nanofibers Electrospinning yielded uniform GEM/FA/PEO/CMC/UiO-66 nanofibers with a mean size of 220 nm (Figure 3a). The surface of some particles on the nanofibers may be attributable to GEM/FA/UiO-66. Homogenous nanofibers with a mean size of 270 nm were fabricated to fabricate PCP nanofibers (Figure 3b). Loading drug/MOFs into the nanofiber did not appreciably alter the nanofibers’ morphological surface. Therefore, the morphology of the nanofibers was not compromised by the encapsulation of drug/MOF in the core. The nanofiber diameter may grow due to physically immobilising UiO-66/GEM/FA inside the nanofibers. The shell thickness also rose when the shell flow rate was raised for PCP nanofibers. The shell flow ratio of 0.3 mL/h fabricated the thinnest GEM/FA/UiO-66-loaded PCP nanofibers with a mean size of 360 nm (Figure 3c). In comparison, 0.5 mL/h fabricated the following thickest nanofibers with a mean size of 420 nm (Figure 3d), and 0.8 mL/h fabricated the thickest nanofibers with an average diameter of 490 (Figure 3e). An increase in the shell’s flow rate led to a more excellent solution supply for the jet in the same amount of time, leading to the growth of coarser nanofibers.

Figure 3. Single electrospinning methods were used to fabricate the nanofibers. (a) SEM images of GEM/FA/PEO/CMC/UiO-66 nanofibers. (b) polyurethane (P)/Carboxymethyl chitosan (C)/Polyethylene oxide (P) (termed as PCP) core-shell nanofibers with various shell flow ratios of 0.3 mL/h (c), 0.5 mL/h (d), 0.8 mL/h (e). Scale bar 50 µm. (f) degradation ratio of fabricated nanofibers. Data are presented as mean ± standard deviation (SD) (n = 3).

The findings of a 30-day study comparing the nanofibers degradation ratio fabricated by the electrospinning method, GEM/FA/UiO-66-PCP, with different shell flow ratios are shown in Figure 3f. After 30 days, core-shell nanofibers made from GEM/FA/PEO/CMC/UiO-66 using electrospinning in flow levels of 0.3 mL/h, 0.5 mL/h, and 0.8 mL/h had mass losses of 60.1 ± 0.8%, 10.7 ± 0.5%, 7.4 ± 0.3%, and 5.7 ± 0.3%, respectively. Core-shell nanofibers may degrade more slowly than nanofibers made by single electrospinning because the hydrophobic nature of PU shields the core layer from moisture. Core-shell nanofibers that have been fabricated have long-term biological promise because of their resistance to degradation.

Drug loading and release

The effectiveness of synthetic carriers as drug encapsulation and drug loaders. The composite MOF-loaded nanofibers have encapsulation efficiencies of FA and GEM of more than 95%. The outcomes showed that the fabricated nanofibrous carriers had great promise as drug delivery devices [33]. Figure 4 shows the FA and GEM release patterns from GEM/FA/PEO/CMC nanofibers, GEM/FA/PEO/CMC/UiO-66, and GEM/FA/UiO-66-loaded PCP at pH 5.0 and 7.4. FA-GEM loaded-CMC/PEO nanofibers’ first burst released GEM and FA. It was because FA and GEM were becoming encased on the nanofibers’ surfaces. On the other hand, GEM/FA/PEO/CMC/UiO-66 nanofibers have released both GEM and FA drugs gradually over 72-h with no signs of a burst release. GEM and FA release rates were higher in an acidic environment compared to a physiological one. Faster drug release of GEM and FA was seen at pH 5 compared to pH 7.4, likely due to the amine groups protonation of CMC and its superior swelling in an acidic milieu. When comparing GEM and FA release characteristics, PCP nanofibers relied on release on the thickness of core-shell layers. GEM and FA have been released from core-shell nanofibers at pH 7.4 for 12 days, 21 days, and 30 days under a shell solution flow rate of 0.3 mL/h and at pH 5 for 10 days, 18 days, and 25 days under a shell solution flow rate of 0.5 mL/h, respectively (Figure 5). To reduce the drug release rate, the shell flow proportion of nanofibers was increased, which slowed the diffusion of drugs from the inner pores.

Figure 4. In vitro drug release profile. (a) GEM and (b) FA release profiles from fabricated PEO/CMC and PEO/CMC/UiO-66 nanofibers. Legends represent (A- GEM/FA/PEO/CMC-pH = 7.4; (B) GEM/FA/PEO/CMC/UIO-66-pH = 7.4; (C) GEM/FA/PEO/CMC-pH = 5.0, and (D) GEM/FA/PEO/CMC/UIO-66-pH = 5.0). Data are presented as mean ± standard deviation (SD) (n = 3).

Figure 5. In vitro drug release profile. (a) GEM and (b) FA release profiles from core-shell nanofibers. Legends represents (A-GEM/FA/UIO-66-PCP nanofibers-pH = 7.4, shell flow 0.3 mL/h, (B) GEM/FA/UIO-66-PCP nanofibers-pH = 5.0, shell flow 0.3 mL/h, (C) GEM/FA/UIO-66-PCP nanofibers-pH = 7.4, shell flow 0.5 mL/h, (D) GEM/FA/UIO-66-PCP nanofibers-pH = 5.0, shell flow 0.5 mL/h, (E) GEM/FA/UIO-66-PCP nanofibers-pH = 7.4, shell flow 0.8 mL/h, (F) GEM/FA/UIO-66-PCP nanofibers-pH = 5.0, shell flow 0.8 mL/h). Data are presented as mean ± standard deviation (SD) (n = 3).

Cell viability

Figure 6a shows the effects of the fabricated UiO-66 MOFs, GEM/FA/PEO/CMC/UiO-66 nanofibers, PCP nanofibers, GEM/FA/UiO-66-loaded PCP nanofibers on non-cancerous cells and A549 and H1299 regarding cell survival. Figure 6b,c demonstrate that after 168-h of incubation with UiO-66 MOFs, GEM/FA/UiO-66-loaded PCP nanofibers formed under a shell follow the rate, GEM/FA/PEO/CMC/UiO-66, and GEM/FA/UiO-66-loaded PCP, only 9.8%, 2.5%, 7.1%, and 5.2% of non-cancerous cells, respectively, were eliminated. The most remarkable cytotoxicity was seen in non-cancerous cells treated with UiO-66 NPs. This may have been caused by the comfortable diffusion of the pure MOF NPs across the accumulation and the cellular membrane of NPs in the cell nucleus. Slower anticancer drug release from core-shell nanofibers may explain why they are less hazardous to healthy cells when used with cancer drugs than GEM/FA/UiO-66-loaded PCP nanofibers fabricated by electrospinning. Single electrospun nanofibers degraded more quickly and had more pristine GEM and FA on their surface, leading to increased cytotoxicity against non-cancerous cells over 72 and 168-h. Hence, the shell layers enhanced the biocompatibility of the GEM/FA/UiO-66-loaded PCP nanofibers. Figure 6d,e demonstrate that the UiO-66 MOFs did not exhibit cytotoxicity against the A549 and H1299 cancer cells.

Figure 6. The in vitro cell cytotoxicity was evaluated on BT-474 and HeLa cervical cancer cells and the non-cancerous fibroblast cells (NIH3T3) cells using a standard MTT assay. (a) non-cancerous fibroblast cells (NIH3T3) cells cell viabilities of core-shell fibres, GEM/FA/UiO-66 loaded-nanofibers (A-MOFs; B-core shell nanofibers; C-GEM/FA/UiO-66 loaded single layer nanofibers; D-GEM/FA/UiO-66 loaded core-shell nanofibers). (b) HeLa cell viability of GEM/UiO-66 loaded nanofibers a-GEM/UiO-66 loaded single layer nanofibers; b-GEM/FA/UiO-66 loaded single layer nanofibers; c-GEM/UiO-66 loaded core-shell nanofibers; d-GEM/FA/UiO-66 loaded core-shell nanofibers; e-Free GEM; f-Free FA). (c) GEM/FA/UiO-66 loaded-PCP nanofibers with different shell flow ratios (GEM/FA/UiO-66 loaded core-shell nanofibers with different shell flow 0.3, 0.5, and 0.8 mL/h). (d) HeLa cell viability of GEM/UiO-66 loaded nanofibers. (e)1 GEM/FA/UiO-66 loaded-PCP nanofibers with different shell flow ratios. Data are presented as mean ± standard deviation (SD). (n = 3) (*p < 0.05, **p < 0.001, and ***p < 0.01).

Morphological and apoptosis examination of lung cells

Apoptosis is recognised as a unique form of genetically determined controlled death. AO/EB dual staining was employed to examine the apoptotic morphological alterations brought on by NPs [34]. In this study, A549 and H1299 cells treated with PCP nanofibers, GEM/FA/UiO-66, and GEM/FA/UiO-66-loaded PCP displayed apoptotic characteristics as apoptotic bodies, membrane blebbing, condensed nuclei, and at 24-h (Figure 7). Viable cells were identified as light green, early apoptotic cells as brilliant green fluorescence and condensed chromatin, late apoptotic cells as orange fluorescence, and nonviable cells as red fluorescence (Figure 7).

Figure 7. Morphological changes of HeLa and BT-474 cervical cancer cells investigated by AO/EB staining method. IC₅₀ concentration with core-shell nanofiber, GEM/UiO-66 loaded core-shell fibres, GEM/FA/UiO-66 loaded core-shell fibres treated with HeLa and BT-474 cervical cancer cells. Scale bar 100 µm.

DAPI stained to validate this finding in A549 and H1299 cells treated with IC₅₀ concentration of PCP nanofibers, GEM/FA/UiO-66, and GEM/FA/UiO-66-loaded PCP and compared the results to untreated controls (Figure 8). A significant factor determining a nanoparticle’s toxicity is cell apoptosis. There are many mechanisms by which it might cause apoptosis [31,35]. Using flow cytometry (FCM) and the Annexin V-FITC/propidium iodide (PI) double staining test at the IC₅₀ concentration of PCP nanofibers, GEM/FA/UiO-66, and GEM/FA/UiO-66-loaded PCP, the amount of apoptosis in A549 and H1299 cells was quantified (Figure 9).

Figure 8. Nuclear damages of HeLa and BT-474 cervical cancer cells investigated by AO/EB staining method. IC₅₀ concentration with core-shell nanofiber, GEM/UiO-66 loaded core-shell fibres, GEM/FA/UiO-66 loaded core-shell fibres treated with HeLa and BT-474 cervical cancer cells. Scale bar 100 µm.

Figure 9. Apoptosis investigation of HeLa and BT-474 cervical cancer cells investigated by AO/EB staining method. IC₅₀ concentration with core-shell nanofiber, GEM/UiO-66 loaded core-shell fibres, GEM/FA/UiO-66 loaded core-shell fibres treated with HeLa and BT-474 cervical cancer cells.

Discussion

The molecular weight of the applied PVP exceeded that of the PCL. Consequently, the augmentation of PVP content in the polymer solution resulted in an elevation of the viscosity of the electrospinning solution. As a result, the nanofibers in PVP core samples had a larger diameter than PCL core samples due to the increased viscosity of the solutions containing more PVP [36–38]. Increased viscosity and more chain entanglements enhance the viscoelastic forces during the electrospinning process, leading to an increase in the diameter of nanofibers. The nanofiber diameter grew as the PVP content in the shell composition was raised, similar to the core effect. Other studies have also reported an increase in fibre diameter when the fraction of PVP in the PVP/PCL blend is increased up to 40 wt% [39].

Typically, a higher flow rate led to a larger diameter of the nanofibers. The flow rate of the shell solution had no discernible effect on the diameters of the nanofibers. Therefore, the changes in diameter were caused mainly by the chemical makeup of the nanofibers [40–42]. The release profiles exhibited a rapid beginning gradient that transitioned into a gradual gradient curve as time progressed. The initial quick gradient could be attributed to the deterioration of the PVP within the core and shell structures. The results indicate that the quantity of PVP in the shell composition is responsible for the magnitude of the initial burst discharge. The gradual decrease in the release profile is associated with the release of the medication from the remaining PCL in the nanofibers [43,44].

When both GEM and FA were present, the viability of A549 and H1299 cells was much more significant on nanofibers and core-shell nanofibers than on nanofibers or core-shell nanofibers loaded with GEM alone. Because of the rapid release of FA and GEM from GEM/FA/UiO-66-loaded PCP nanofibers made by electrospinning, they could kill more A549 and H1299 cells after 24-h than core-shell nanofibers. After 72 and 168-h, there was no discernable difference in the GEM and GEM/FA/UiO-66-PCP cell mortality percentages. This action occurred because the nanofibers began to degrade after 72-h, which decreased the medium’s drug concentration. In contrast, A549 and H1299 cell viability were much worse after 168-h than at 72-h due to the continual delivery of GEM and FA from GEM/FA/UiO-66-PCP nanofibers.

The highest cell death was found after being treated with FA@GEM-loaded core-shell nanofibers compared with the core-shell nanofibers and GEM-loaded core-shell nanofibers, in conjunction with our current findings, human lung cancer cells treated with IC₅₀ concentration of core-shell nanofibers, GEM-loaded core-shell nanofibers, and FA@GEM-loaded core-shell nanofibers similarly displayed comparable apoptotic morphological features following AO/EB staining. DAPI stained to validate this finding in A549 and H1299 cells treated with IC₅₀ concentration of PCP nanofibers, GEM/FA/UiO-66, and GEM/FA/UiO-66-loaded PCP and compared the results to untreated controls. Contrary to normal nuclei, apoptotic cells’ nuclei will have broken or compacted chromatin and glow brightly under the fluorescence microscope. It is evident that FA@GEM-loaded core-shell nanofibers considerably increase compared with core-shell nanofibers, and GEM-loaded core-shell nanofibers cause nuclear condensation in A549 and H1299 cells, a hallmark of apoptosis. Early apoptosis is represented by cells labelled with Annexin V-FITC alone, whereas necrosis is shown in cells stained with PI. With fluorescent dyes, late apoptotic cells are highlighted. GEM/FA/UiO-66 nanofibers and PCP nanofibers loaded with GEM/FA/UiO-66 May all cause apoptosis. With GEM/FA/UiO-66-loaded PCP nanofibers, there were more apoptotic cells than in the control group.

Conclusion

Controlled drug release and therapy of A549 and H1299 lung cancer cells were accomplished by loading GEM and FA into the UiO-66 MOF and then incorporating the MOFs into electrospun PEO/CMC nanofibers fabricated through electrospinning GEM/FA/UiO-66-PCP. The GEM/FA/UiO-66 XRD patterns confirmed the effective fabrication of UiO-66 MOFs and the amorphous condition of GEM in the MOF solution. Their slower rate of deterioration suggested nanofibers’ longer-lasting durability. FA and GEM drug loading efficiencies from nanofibers exceeded 95%. FA and GEM were continuously released from PCP nanofibers at 0.3, 0.5, and 0.8 mL/h shell flow ratios for 12, 21, and 30 days, respectively, at a pH of 7.4 and 10, 18, and 25, under a pH of 5. After 168-h, 240-h, and 240-h of exposure to shell flow rates, 87 0.5%, 83 0.5%, and 82 0.7% of cells in the fabricated GEM/FA/UiO-66-loaded PCP nanofibers had killed effectively in cancerous cells without affecting non-cancerous cells. The staining assay findings showed that the co-delivery of GEM/FA/UiO-66-loaded PCP nanofibers might considerably increase apoptosis and be broadly employed for treating different cancers, including lung cancers.

Disclosure statement

No potential conflict of interest was reported by the author(s).

Data availability statement

All data generated or analysed during this research are included in this published article.

Additional information

Funding

This project was supported by Researchers Supporting Project number [RSP2024R383], King Saud University, Riyadh, Saudi Arabia.