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Back to Journal »International Journal of Nanomedicine» Volume 16

Externally triggered new fast-release sonosensitive folic acid modified liposomes for gemcitabine: development and characteristics

Author: Omar MM, Hasan OA, Zaki RM, Eleraky NE

Published on January 28, 2021, Volume 2021: 16 pages, 683-700 pages

DOI https://doi.org/10.2147/IJN.S266676

Single anonymous peer review

Editor approved for publication: Prof. Dr. Anderson Oliveira Lobo

Mahmoud M Omar,1,2 Omiya Ali Hasan,1,2 Randa Mohammed Zaki,3,4 Nermin E Eleraky5 1 Department of Pharmacy and Industrial Pharmacy, Deraya University School of Pharmacy, Minia, 61768, Egypt; 2 Sohag University School of Pharmacy Department of Pharmacy and Clinical Pharmacy, Sohag, 82524, Egypt; 3 Department of Pharmacy, Faculty of Pharmacy, Prince Harjisatam Bin Abdulaziz University, Saudi Arabia; 4 Faculty of Pharmacy, Beni Suf University, Egypt Department of Industrial Pharmacy; 5 Assiut University, Faculty of Pharmaceutics, Faculty of Pharmaceutical, Assiut, Assiut, 71526, Egypt Mailing address: Mahmoud M Omar Department of Pharmaceutics and Industrial Pharmacy, Faculty of Pharmacy, Deraya University, Deraya Square Street, Minia, 61768 , Egypt Tel +2039321099 Email [Email protected] Purpose: To develop an externally triggered rapid release targeting system for the treatment of ovarian cancer, gemcitabine (GMC) is wrapped in sonosensitive (SoS) folic acid (Fo) modification Liposome (LP). Methods: Using membrane hydration technology to prepare GMC-loaded LP (GMC LP), GMC-loaded Fo-targeted LP (GMC-Fo LP), and GMC-loaded Fo-targeted SoS LP (GMC-SoS Fo LP) and according to The particle size, zeta potential and percentage of retained drug are evaluated. Flow cytometry was used to quantify the cellular uptake of the fluorescent delivery system in ovarian cancer cells expressing Fo. Finally, the tumor targeting ability, in vivo evaluation and pharmacokinetic studies were carried out. Results: GMC LPs, GMC-Fo LPs and GMC-SoS Fo LPs were successfully prepared, with a size of <120.3±2.4 nm, 39.7 mV ζ-potential and a drug embedding rate of 86.3%±1.84%. The cellular uptake of GMC-SoS Fo LPs is 6.51 times higher than that of GMC LPs (under ultrasound irradiation-p <0.05). However, the cellular uptake of GMC-Fo LP was only 1.24 times higher than that of GMC LP (p>0.05). Biodistribution studies have shown that the concentration of GMC in tumors treated with GMC-SoS-Fo LPs (ultrasound) is 2.89 times that of free GMC (p <0.05). In vivo, GMC-SoS Fo LPs show the highest anti-proliferative and anti-tumor effects on ovarian cancer. Conclusion: These findings indicate that the externally triggered fast-release SoS Fo modified LP is a promising system for delivering fast-release drugs to tumors. Keywords: sonosensitive liposome, gemcitabine, folate modified liposome, external trigger, ovarian cancer

Cancer is one of the most dangerous anomalies threatening human life today. Among Egyptian women, ovarian cancer is the fifth most common cause of death. 1 Since ovarian cancer cannot be detected and diagnosed early, it will spread in the internal organs. Therefore, most patients with ovarian cancer cannot be diagnosed, leading to advanced stages (stage III or IV). 2 Late diagnosis of ovarian cancer may lead to death within 5 years. Surgical removal of ovarian tumors and chemotherapy are the basic methods of standard treatment. Chemotherapy is an important factor in the treatment of cancer. There are many anticancer drugs, such as doxorubicin, topotecan, and gemcitabine (GMC). 3,4

Doxorubicin may damage topoisomerase II-mediated DNA repair by intercalating DNA, or cause damage to important cell components (such as cell membranes, DNA, and proteins) by generating free radicals. 5 The importance of doxorubicin for treatment Due to the rapid development of drug resistance, the incidence of ovarian cancer has become limited. Topotecan is a derivative of camptothecin, which interferes with DNA replication by binding to the topoisomerase I-DNA complex and prevents DNA from reconnecting, thereby interfering with DNA replication and causing cell death. 6 It has an effective effect on metastatic ovarian cancer after the failure of platinum therapy. It has many side effects, such as nausea, vomiting, bleeding, difficulty breathing, blood in the urine and extreme fatigue. 7 Platinum-based chemotherapy drugs, such as cisplatin, carboplatin, and oxaliplatin, are used to treat cancer. They mainly act on guanine N7 to form DNA 1,2-strand cross-links or DNA-protein cross-links, and inhibit DNA repair and/or DNA synthesis in cancer cells. According to reports, compared with carboplatin, cisplatin is an analogue superior to carboplatin, which is based on its therapeutic effect on a series of tumors and its toxicity characteristics, especially the loss of hematological toxicity. 8 Paclitaxel is a microtubule-targeted chemotherapy drug used to treat cancer. It leads to the stabilization of dynamic microtubule polymerization, leading to mitotic failure. In addition, it changes other cellular functions involving microtubules, such as intracellular signal transduction and organelle transport. 9

GMC is an attractive option for the treatment of ovarian cancer,10 because of its well-known efficacy in platinum/paclitaxel-resistant cancer, no cross-resistance to platinum compounds, and easy combination with other chemotherapeutics to treat recurrence Sexual ovarian cancer. 11 can be used to treat ovarian cancer, lung cancer and pancreatic cancer. 12 GMC penetrates the cell membrane and is converted into GMC diphosphate and GMC triphosphate by deoxycytidine kinase. Ribonucleotide reductase is inhibited by GMC diphosphate, causing the synthesis of deoxynucleotide triphosphates required for DNA synthesis to stop. At the same time, GMC triphosphate is integrated into DNA instead of endogenous deoxynucleoside triphosphate. 13 GMC-triphosphate metabolites are also synthesized by RNA, inhibiting it. 14 However, the hydrophilicity, short plasma half-life and rapid metabolism of GMC lead to an increase in the dose required for the effective plasma concentration of cancer cell death, leading to many side effects. For example, hematological neutropenia is more common than thrombocytopenia. 15 GMC is a hydrophilic drug (dissolved in water at 25 mg/mL) with an octanol-water partition coefficient (logP) of 1.123. In addition, compared with free drugs, adding GMC to liposomes (LP) can enhance its cytotoxicity. 16

In order to reduce the side effects of GMC, many delivery systems have been used, such as nanoparticles, metastases, and LP. 17-19 LP has biocompatibility, biodegradability and safety, and can be used to improve efficacy, reduce side effects, and enhance the stability of encapsulated drugs. The phospholipid bilayer of LP can be used as a membrane, and the hydrophobic part can be inserted into it to target the LP to the desired site. 20,21 LP is widely used as a delivery system because of the relatively long circulation time, enhanced permeability and retention effect (EPR)), and enhance the accumulation of drugs in cancer tissues, thereby improving the therapeutic effect and reducing toxic effects. 22 In addition, LP has the ability to capture hydrophobic (in the double layer) and hydrophilic (in the core) drugs. They are non-toxic and will not activate the immune system. 21 Many attempts have been made to enhance the targeting ability of LPs by modifying the surface, such as PEGylated LPs and Fo-modified LPs. 23,24 To improve cell specificity and intracellular delivery, LP can be further combined with targeting ligands, such as Fo.25. Currently, Fo modified LPs are used based on the expression of Fo receptors, which is very high in most types of cancer, but low in healthy tissues. 26 These modified LPs have sustained release characteristics. 27 However, there is an urgent need to develop a delivery system that has an enhanced targeting index and an external trigger for rapid release of drugs in ovarian cancer tissues. 28

Since Fo-targeted LP has been proven to be very effective in many medical applications, LPs that are easily destroyed are rarely studied. The fast release ability of Fo-coupled LP containing Chlorin e6 (Ce6) after binding to cancer cells can be studied. 29 Encapsulating drugs in long-cycle PEGylated LPs overcomes the limitations of other non-PEGylated LPs. 30 In addition, proper PEGylation is also required because the initial (and overall) tumor accumulation still relies on passive extravasation. 31 Active targeting (such as binding to Fo) is used to reduce the disadvantages associated with passive targeting (such as decorative PEG binding and long circulation). 32

Due to low specificity, most anti-tumor drugs are limited in their use in the treatment of cancer tissues. 33 To solve this problem, anti-tumor drugs can target cancer tissues. In addition, high levels of anti-tumor drugs are required in cancer tissues to kill abnormal cells. To achieve this goal, an externally triggered rapid-release drug carrier should be developed.

This study aims to achieve rapid release of loaded drugs in tumors. The formulated LPs should be sensitive to external stimuli, such as ultrasound, and once applied, the drug should be released. This study aims to target drugs to the ovaries using a new type of fast-release sono-sensitive (SoS) Fo modified LP that is externally triggered. SoS Fo modified LP is characterized by measuring particle size, zeta potential, and encapsulation efficiency (EE). In addition, ultrasound sensitivity, cell correlation using flow cytometry, anti-tumor efficacy in vivo, and pharmacokinetics of liposomal drugs were also examined.

Soybean phospholipids (L-α-phosphatidylcholine [PC, purity ≥99%), folic acid-polyethylene glycol-stearylamine (FA-PEG-SA) and cholesterol (purity ≥99%) were purchased from Xi’an Ruixi Biological Technology Co., Ltd. Ce6 was purchased from Frontier Scientific (Salt Lake City, Utah, USA). The CAOV3 cell line comes from the Chinese Academy of Sciences (Shanghai, China). RPMI 1640 medium and FBS were purchased from Sigma Chemical.

All reagents are high performance liquid chromatography or analytical grade. Sprague Dawley rats were purchased from the animal house of the Medical School of Assiut University in Egypt.

GMC LPs (GMC LPs) are prepared by a pH gradient method using PC and cholesterol (PC:cholesterol 5.5:4.5). 15,34 In short, LPs are prepared using thin film hydration. To prepare non-Fo modified LP (500 mg lipid), the lipid composition of PC and cholesterol (55% and 45% molar ratio) was completely dissolved in a mixture of ether and chloroform (1:3 v:v). Using a rotary evaporator (Büchi), the organic solution was evaporated into a sterile Erlenmeyer flask under vacuum to form a dry film. Phosphate citrate buffer (50 mM, 10 mL, pH 2.2) is used to hydrate the membrane at 55°C (50 mg/mL) to produce a colloidal solution, and then use disodium hydrogen phosphate to raise the pH to 7.2. The colloidal solution was extruded 12 times through a 100 nm polycarbonate filter (micro extruder, Avanti), aimed at optimizing the particle size, as shown in Figures S1 and S2 (Supplementary Section). Then GMC (concentration 0.9–1.6 mg/mL) was incubated in LP at 65°C for 4 hours to obtain an active load. The separation of free GMC was performed using extensive dialysis against buffer (20 mM HEPES, pH 7.4, containing 150 mM NaCl and 0.1 mM EDTA). GMC-Fo LPs were prepared by the same procedure, except that FA-PEG-SA was added to the lipid mixture at a molar ratio of 0.1%. To prepare GMC-SoS Fo LP, the lipid composition of PC and cholesterol (55% and 45% molar ratio) and FA-PEG-SA (0.1%) is completely dissolved in a mixture of ether and chloroform (1:3 v: v). Dissolve the lipophilic sonosensitizer Ce6 ester (3%, 6% or 9%) in the prepared organic phase. Using a rotary evaporator, evaporate the organic solution into a sterile Erlenmeyer flask under vacuum to form a dry film. Then, all preparation steps use the same procedure mentioned earlier. These procedures are performed under sterile conditions and vertical laminar flow (vertical laminar flow hood, BZ series, Germfree).

After the LP was purified by ultrafiltration in a centrifuge tube using an ultrafilter (Amicon Ultra, 30 kDa molecular weight cut-off, Millipore), the GMC-EE and drug loading (DL) were calculated. Quantification is performed by spectrophotometry. 35 In short, place the LP preparation (0.5 mL) in the upper part of the centrifuge tube, and then centrifuge at 6,000 rpm at 4°C for 30 minutes to separate the free drug from the encapsulated LP. Repeat the last step until all LP quantities are obtained. The amount of free drug in the filtrate was estimated at 267 nm by spectrophotometry (LC20A, Shimadzu). EE and DL are calculated using formulas 1 and 2:

Among them, Atotal, Afree and Aliosomes represent GMC, free GMC and recovered total LP, respectively.

Use a transmission electron microscope to check the morphology of GMC LP, GMC-Fo LP and GMC-SoS-Fo LP. In short, an aliquot was placed on a hydrophobic copper mesh and dried, and then negatively stained with uranyl acetate (50 µL, 2.5% w:v). Malvern Zetasizer Nano ZS Zen 3600 is used to study colloidal dispersion to determine the zeta potential, particle size and polydispersity index (PDI) value. 36 In short, the LP sample (200 μL) was diluted with deionized water (800 μL) to be analyzed using Zetasizer based on dynamic light scattering.

GMC-SoS-Fo LPs were diluted eight times with PBS (0.1 M, pH 7.4) and kept at 37°C for 10 minutes, then transferred to an ultrasonic ultrasonic processing system (2 MHz, 1.5 W/cm2 power intensity, 10 seconds ultrasonic irradiation 10 seconds rest for a total of 0-150 seconds). The temperature of the LP formulation was fixed during the ultrasonic irradiation using a water bath containing a thermocouple. The drug release was measured and recorded after each ultrasound irradiation. The colloidal dispersion is allowed to stand for 10 minutes and then purified by ultrafiltration in an Amicon centrifuge tube35. According to the determination of EE, quantification was carried out under 267 nm spectrophotometry (Shimadzu LC20A).

The GMC release of GMC-SoS Fo LP was measured under the same procedure, and the molar ratio of the LP was 3%, 6% or 9% Ce6 of the total lipid. However, the total time of ultrasonic irradiation was changed to 130 seconds.

A dialysis bag was used to determine the in vitro GMC release of GMC-loaded LPs. 37 In short, put 5 mL of GMC-loaded LP into a cellophane bag (molecular weight cut-off 12-14 kDa), and then insert 100 mL PBS (pH 7.4, 0.1 M, 37°C). Gently agitate the dissolving medium and treat it with ultrasonic irradiation (2 MHz, 1.5 W/cm2 power intensity, 10 seconds of ultrasonic irradiation, 10 seconds of rest, total 0-130 seconds) or no treatment. An aliquot (2 mL) replaced with an equal volume of freshly prepared PBS was removed from the dissolution medium for spectrophotometric analysis at λmax = 267 nm to measure the released GMC. The measurement was repeated 3 times and expressed as the mean ± SD.

In order to check the stability of the colloidal dispersion, the LP dispersion was kept at different temperatures (4°C±2°C, 25°C±2°C, 37°C​​±2°C, 65% relative humidity). Months. Take aliquots every month to study particle size, zeta potential, and EE.

CAOV3 cells (adherent type) were cultured in RPMI 1640 medium (50 mL) containing FBS (10%, 5 mL) 1%, streptomycin (0.1 mg/mL) and penicillin (200 IU, 1 mL), At 37±2°C/5% carbon dioxide. The cells are re-seeded and suspended every 4 days for the 5th-20th generation of the study.

To determine the intercellular uptake of GMC LP, GMC-Fo LP, and GMC-SoS Fo LP, flow cytometry was used based on measuring the average fluorescence intensity. 38 In short, grow cells at a density of 6×105 cells/well in a 48-well plate. After 24 hours, the cultured cells were removed from the culture medium, then washed with PBS, re-seeded and suspended in fresh culture medium at 37°C±2°C/5% CO2 for 4 hours. The cell suspension was incubated with Calcein LP, Calcein-Fo LP and Calcein-SoS Fo LP for 48 hours. Subsequently, they were subjected to ultrasonic irradiation (2 MHz, 1.5 W/cm2 power intensity, ultrasonic irradiation for 10 seconds, 10 seconds of rest, 0-130 seconds in total) or left untreated, and then incubated at 37°C for 30 minutes. ±2°C/5% carbon dioxide. Thereafter, the medium was removed and the adhered monolayer was rinsed with sterile PBS. Trypsin (10 mL, 0.25% v:v) is used to release the cells from the adhered state. The cells were suspended in sterile PBS and analyzed by flow cytometry (Thermo Fisher Scientific) at an excitation wavelength of 495 nm and an emission wavelength of 516 nm. The LP dispersion was prepared following the same procedure described above to encapsulate calcein instead of GMC.

Cell suspensions incubated with calcein-free LPs are irradiated with ultrasonic waves (2 MHz, 1.5 W/cm2 power intensity, 10 seconds ultrasonic irradiation and 10 seconds rest time, total 0-130 seconds) or not under the above conditions Next treatment. The first sample (incubated with an empty LP) was treated with ultrasonic irradiation, so it was considered as a control for the test sample treated with ultrasonic irradiation (Control 1). The second sample (incubated with empty LP) was not irradiated with ultrasound and was therefore regarded as a control for the test sample that was not irradiated with ultrasound (Control 2). These controls are used to subtract the effects of autofluorescence and ultrasound radiation.

After that, the cells were washed twice with PBS and suspended in PBS, and then subjected to flow cytometry to calculate cell binding. The measurement is repeated 3 times.

To examine the cell killing effect of SoS Fo LP, free GMC, GMC LP, GMC-Fo LP, GMC-SoS Fo LP (without ultrasonic irradiation) and GMC-SoS Fo LP (ultrasound irradiation), cytotoxicity was evaluated. The preparation was checked by MTS assay in CAOV3 cells. In short, the cells were seeded in a 96-well plate at 105 cells/well and cultured at 37°C and 5% CO2 for 48 hours. Thereafter, the cells were treated with different concentrations (0–150 µM) of GMC from the preparation for 72 hours. After 48 hours of incubation, only the last formulation was irradiated with ultrasound for 150 seconds, and then incubated for 24 hours (total incubation time is 72 hours). Subsequently, add 20 µL of MTS reagent (tetrazolium internal salt) to each well and incubate for another 4 hours at 37°C in 5% CO2. Use a multi-mode microplate reader to measure the absorbance at 490 nm to calculate cell viability:

Among them, S[Test], S[Control] and S[Blank] represent the absorbance of the treated group, untreated control group and blank medium group, respectively. In addition, IC50 is calculated using an online calculator (https://www.aatbio.com/tools/ic50-calculator).

The rat xenograft ovarian cancer model obtained from the Assiut University School of Medicine was produced by surgical orthotopic implantation. 39 In short, by using 0.25 mL of dexamethasone acetate (4 mg/mL) daily intramuscular injection. The immunosuppressive treatment was also started 5 days before the injection of CAOV3 cells and was maintained throughout the experiment.

The resulting subcutaneous tumor was grown to about 0.5 cm3, and then separated into small pieces (1 mm3). One piece was hatched into the ovary by surgery. Monitor the development of ovarian cancer through ultrasound imaging. 40 The study was carried out under the permission of the Ethical Approval Committee of Assiut University School of Medicine (Assiut 155-020, January 20, 2020), and was carried out in accordance with the nursing guidelines of "The Use of Laboratory Animals" published by the National Institutes of Health ( Eighth edition, revised in 2011).

In order to evaluate the in vivo targeting efficiency of GMC LP, GMC-Fo LP and GMC-SoS-Fo LP, biodistribution and pharmacokinetic studies were performed using 40 rat ovarian cancer xenograft models. The rats were randomly divided into fife groups: the first group (n=8) was injected intravenously with free GMC (2mL, 4mg/mL), the second group (n=8), the third, fourth and fifth groups were injected intravenously with free GMC LPs GMC LPs, GMC-Fo LPs and GMC-SoS-Fo LPs (without ultrasound) and GMC-SoS-Fo LPs (with ultrasound) at a concentration of 12.5 mg/kg. 30 minutes after the injection, the rats treated with GMC-SoS-Fo LPs were irradiated with ultrasound (2 cm2 ultrasound area, 2 MHz ultrasound, 1.5 W/cm2 power intensity, 10 seconds) ultrasound irradiation on the ovaries on both sides of the abdomen for 10 seconds. 0-150 seconds in total). One hour after the injection, 3 rats in each group were sacrificed humanely, and the ovaries, kidneys, lungs, spleen, liver, and heart were collected. As reported in the literature, the drug was measured in plasma with some modifications. 41 The collected organs are stored at -20°C until further analysis. To detect the GMC concentration, the collected organs were homogenized in four times the volume of double-distilled water. The homogenate (2 mL) added with 100 µL of fluorouridine (20 ng/mL) as an internal standard was treated and vortexed with acetonitrile (1 mL) and centrifuged (2,000 rpm, 3 minutes) to precipitate the protein. The supernatant was evaporated under vacuum at 50°C. The dried residue was re-dissolved with 2 mL of mobile phase, which consisted of 40 mmol/L ammonium acetate buffer solution (pH 5.5): acetonitrile (97.5:2.5, v:v). Inject an aliquot of the reconstituted solution (50 µL) into the chromatographic column (C18, 250 mm × 4.6 mm ID, 5 µm), and set the flow rate to 0.8 mL/min to detect the drug at λmax=268 nm.

Twelve female Sprague Dawley rats weighing 250–280 g were used in this pharmacokinetic study and were randomly divided into four groups (n=3). The rats were kept under standard conditions, 25°C, relative humidity 55%±5%, and free drinking water. The pharmacokinetic studies of GMC LPs, GMC-Fo LPs and GMC-SoS-Fo LPs were performed using a single dose of 4 mg/kg GMC (injected via the tail vein, diluted to 1 mL with a 0.09% w:v isotonic saline solution) . At predetermined time intervals (5, 15, 30, 45, 60, 120, 180, 240, and 300 minutes) blood samples (0.3 mL) are drawn from the tail vein to separate plasma. For analysis, 100 µL of fluorouridine solution was added as an internal standard (20 ng/mL) to each plasma sample. By adding 1 mL of acetonitrile to each sample, then vortexing for 10 minutes and centrifuging at 2,000 rpm for 4 minutes, the plasma proteins in all samples were precipitated. Remove the supernatant and evaporate in vacuo at 50°C. The dried residue was reconstituted with 2 mL mobile phase (40 mmol/L ammonium acetate buffer solution, pH 5.5): acetonitrile, 97.5:2.5, v:v). Analysis based on biodistribution studies.

Rat xenograft models of ovarian cancer were divided into different groups, each with ten rats. The first intravenous injection of LPs without GMC (control group) and ultrasound irradiation. The second, third, fourth and fifth were injected intravenously with 4mg/kg free GMC solution, GMC LPs, GMC-Fo LPs and GMC-SoS-Fo LPs. The sixth group injected with GMC-SoS-Fo LPs was irradiated with ultrasound. The dosing schedule is 4 mg drug equivalent intravenously on days 1, 7, 10, 13, 16 and 20. The tumors in each group were measured every 3 days until the 27th day, using a commercial three-dimensional sonar (Artida, Toshiba Medical Systems). 42 The survival time of each group (from tumor inoculation to death) was recorded and plotted. A Kaplan-Meier survival curve was drawn for each group. The body weight of each mouse is measured every day.

All findings are expressed as mean ± SD. Apply one-way analysis of variance to calculate the difference between any two groups. Significance was regarded as p<0.05, and paired comparisons were performed using Tukey-Kramer multiple assessment or two-sided Student's t-test (GraphPad Prism 6.0, GraphPad, San Diego, CA, USA).

Targeting drug delivery systems to tumors is very important. At present, there is a great demand for rapid drug release after targeting the tumor site.

The physical properties of the prepared LP were studied by measuring the particle size, EE and DL (Table 1). The average particle diameters of GMC LP, GMC-Fo LP and GMC-SoS-Fo LP are 120.3±2.4, 119.1±3.5 and 109.5±2.9 nm, respectively. Although a micro extruder (100 nm polycarbonate filter) is used to optimize the particle size, the average particle size is> 100 nm. These results can be explained by the elasticity of LPs: they are squeezed through polycarbonate filters under high pressure. After the LP passes through the extruder, they reach their final size. In addition, the PDI values ​​of all LPs are 0.209-0.272, indicating a high degree of homogeneity. The EE of the prepared LPs was 82.8%-86.3%, confirming that the conjugate of LPs and Fo part (GMC-Fo LPs) or Fo and Ce6 (GMC-SoS-Fo LPs) had no significant effect on EE (p>0.05). As shown in Table 1, the zeta potential of all LP preparations is greater than >34.52±1.01 mV, confirming the stability of all preparations. Finally, the DL of LP is 3.45%±0.12%–3.73%±0.27%, which indicates that the conjugation of LP has no effect on DL, as shown in Table 1. The slight difference between GMC LP, GMC-Fo LP and GMC-SoS-Fo LPs can be explained based on the same lipid composition, while other ingredients-Fo part (GMC-Fo LPs) or Fo and Ce6-to be ignored Proportion preparation. In addition, most of the Fo parts are outside the phospholipid bilayer membrane, so they may not interfere or have little effect on the phospholipid bilayer membrane. Table 1 Physical properties of nanocarriers

Table 1 Physical properties of nanocarriers

The zeta potential of LP may have a great influence on the accumulation of drugs in cells. The accumulation of cationic LPs is high after intravenous injection. Cationic macromolecules show higher glomerular permeability than anionic macromolecules of similar molecular weight. Due to various factors, larger cationic macromolecules accumulate in the kidney and liver. The main charge on the cell surface is a negative charge, which causes cationic molecules to bind to the cell. The interaction of the negatively charged components in the blood with the cationic macromolecules causes the embolism of the aggregates. In addition, cationic LPs are excellent as gene delivery carriers and liver, lung and tumor targeting. Cationic LP can be used as a carrier targeting the nucleus. In addition, cationic LPs cross the blood-brain barrier more easily than anionic LPs. Although cationic LPs have many advantages over viral vectors in gene delivery, the transfection efficiency of DNA cationic LPs (lipid complexes) is too low, and they are more toxic than engineered viral vectors. It has been reported that cationic LP can selectively target tumor vasculature due to an inherent but unexplainable mechanism. 43 In addition, extensive scientific research has focused on formulating cationic LP to improve its vascular targeting efficiency and reduce toxicity-related reactions.

The TEM of GMC-SoS-Fo LPs showed the formation of spherical single-bilayer vesicles (monolayer LPs), as shown in Figure 1, the size of which was consistent with the size obtained using Zetasizer. Figure 1 Transmission electron micrograph of GMC LP stained with 10% uranyl acetate.

Figure 1 Transmission electron micrograph of GMC LP stained with 10% uranyl acetate.

The release of GMC in GMC-SoS-Fo LPs was studied under ultrasonic irradiation at different times (2 MHz, 1.5 W/cm2 power intensity, 10 seconds of ultrasonic irradiation, 10 seconds of rest, total 0-150 seconds), such as As shown in Figure 2. The GMC release increased with time, reaching 90.7%±3.7% at 130 seconds (10 seconds of ultrasonic irradiation, 10 seconds pause). However, the release of GMC did not increase significantly after 150 seconds (92.4%±3.9%), so the ultrasound irradiation time was adjusted to 130 seconds in the following experiment. Figure 2 Release of GMC from GMC-SoS-Fo LPs (6% Ce6) in vitro using different ultrasonic irradiation times in PBS (0.1 M, pH 7.4) at 37°C. The results are expressed as mean ± SD (n=3).

Figure 2 Release of GMC from GMC-SoS-Fo LPs (6% Ce6) in vitro using different ultrasonic irradiation times in PBS (0.1 M, pH 7.4) at 37°C. The results are expressed as mean ± SD (n=3).

In order to confirm the role of Ce6 as an activator of SoS GMC LPs, GMC-Fo LPs and SoS LPs, different concentrations of Ce6 (3%, 6% and 9%) incubated at 37°C for 20 minutes were checked under specific conditions. ) Ultrasonic irradiation conditions are 2 MHz, 1.5 W/cm2 power intensity, 10 seconds of ultrasonic irradiation and 10 seconds of standing for a total of 0-130 seconds), as shown in Figure 3. All untreated LPs were recorded to the minimum GMC release ultrasound exposure. Under the same conditions, the GMC release in conventional liposomes (GMC LP) and GMC-Fo LP did not change, as shown in Figure 3. However, the GMC release in GMC-SoS-Fo LP and the GMC-SoS-Fo LPs of 3% Ce6, GMC-SoS-Fo LP 6% Ce6 and 9% Ce6 were 48.7%±3.9% and 89.7%±4.2%, respectively And 90.1%±2.7%. The results show that Ce6 content is the basic factor for the release of GMC from related LPs, and compared with the release from GMC-SoS-Fo LP with 6% Ce6, an increase in Ce6 content to 9% results in an increase in GMC release (p<0.05). Ultrasound Irradiation conditions (2 MHz, 1.5 W/cm2 power intensity, 10 seconds of ultrasound irradiation and 10 seconds of rest, total 0-130 seconds) may not be sufficient to fully activate SoS LPs GMC-SoS-Fo LPs 9% ce6. However, in Under the same ultrasonic irradiation conditions, 6% Ce6 is the optimal concentration to produce sonodynamic activity, which leads to a burst of GMC-SoS-Fo LPs. These results may be attributed to the ability of the SoS activator (Ce6) to form pores or irreparably damage the LP bilayer. 44,45 The acoustic sensitivity of LPs containing DSPE can be attributed to the possibility that DSPE may form an inverted hexagonal frame in the liposome bilayer, such as the response to high temperature or high pressure. 46 Due to ultrasonic irradiation, DSPE in the lamellar liquid crystal transforms into a hexagonal phase, forming a cavity, because the polar head group occupies less space than the non-polar part. 47 Figure 3 The in vitro release of GMC comes from conventional LP and traditional LP with different SoS activator percentages (Ce6 0, 3%, 6%, and 9%) and Fo-targeted LP, which were irradiated by ultrasound (2 MHz, 1.5 W/ cm2 power intensity, 10 seconds of ultrasonic irradiation) treated or untreated in PBS (0.1 M, pH 7.4) at 37°C for 10 seconds, a total of 0-130 seconds). The results are expressed as mean ± SD (n=3).

Figure 3 GMC of conventional LP and Fo-targeted LP with different SoS activator percentages (Ce6 0, 3%, 6%, and 9%) treated or untreated by ultrasound irradiation (2 MHz, 1.5 W/cm2 power) In vitro release intensity, ultrasonic irradiation for 10 seconds and rest for 10 seconds, 0-130 seconds in total), in PBS (0.1 M, pH 7.4) at 37°C. The results are expressed as mean ± SD (n=3).

The in vitro release of GMC LPs, GMC-Fo LPs and GMC-SoS-Fo LPs with or without ultrasound treatment is shown in Figure 4. The GMC release of GMC-SoS-Fo LPs shows ultrasound dependence: if GMC-SoS-Fo LPs are exposed to ultrasound, they are destroyed and GMC is released quickly. However, after 50 minutes of incubation, the GMC release rate of GMC-SoS-Fo LP without ultrasonic irradiation was 9.5%±1.5%. In addition, GMC LPs and GMC-Fo LPs represent ultrasound independence, and minimal release changes were observed, as shown in Figure 4. Figure 4 The in vitro release of GMC LPs, GMC-Fo LPs and GMC-SoS-Fo LPs were treated with ultrasonic irradiation in isotonic PBS (pH 7.4, 37°C). The results are described as mean ± SD (n=3).

Figure 4 GMC LPs, GMC-Fo LPs, and GMC-SoS-Fo LPs are released in vitro in GMC without treatment or ultrasonic irradiation in isotonic PBS (pH 7.4, 37°C). The results are described as mean ± SD (n=3).

GMC LP, GMC-Fo LP and GMC-SoS-Fo LP incubated for 24 hours in isotonic PBS (pH 7.4, 37°C) released very little GMC (<23%). However, 25 hours of ultrasound irradiation of GMC LP, GMC-Fo LP, and GMC-SoS-Fo LP resulted in explosive release (93.6%±3.9) using only GMC-SoS-Fo LP, which was observed with GMC The insignificant release changes of LPs and GMC-Fo LPs are shown in Figure 5. The results confirmed that the release of GMC-SoS-Fo LPs was very low until the LPs were in contact with ultrasound irradiation: burst release was achieved, indicating that almost no drug was released in the blood to the desired site (ovarian tumor) before the LPs and received ultrasound examination . GMC LPs, GMC-Fo LPs and GMC-SoS-Fo LPs showed a strong stabilization trend without ultrasound exposure. In the absence or presence of ultrasonic irradiation, the drug release of GMC LP or GMC-Fo LP does not change. These findings can be explained by the insensitivity of GMC LPs and GMC-Fo LPs to ultrasound irradiation. However, GMC-SoS-Fo LPs are destroyed and release encapsulated drugs once they are stimulated by sound energy. The sensitizer Ce6 is the key ingredient, and the destruction of the LP membrane may be due to the vibration level of Ce6 being sufficient to destroy the double-layer membrane or form enough holes to escape the encapsulated drug. SoS LP for ultrasonic triggered release has been developed. 48 With the application of ultrasonic radiation, bubbles oscillate and burst in the medium, so the introduction of high mechanical pressure will significantly increase the release. This can be explained by the formation and collapse of small air nuclei in the hydrophobic region of the lipid bilayer during exposure to ultrasound radiation, thereby inducing the formation of transient pores through which the drug is released. 44 Figure 5 In vitro release of GMC from GMC. After incubation in isotonic PBS (pH 7.4, 37°C) for 24 hours, LP, GMC-Fo LP and GMC-SoS-Fo LP were untreated or treated with ultrasound. The results are expressed as mean ± SD (n=3).

Figure 5 GMC in vitro release of GMC LPs, GMC-Fo LPs and GMC-SoS-Fo LPs after 24 hours of incubation in isotonic PBS (pH 7.4, 37°C). These LPs were untreated or treated with ultrasound. The results are expressed as mean ± SD (n=3).

GMC LP, GMC-Fo LP and GMC-SoS-Fo LP are physically and chemically stable under the conditions examined (4 C, up to 3 months), as shown in Table 2. The EE changes of GMC LP, GMC-Fo LP, and GMC-SoS-Fo LP) are trivial. The increase in particle size of LPs, the decrease in zeta potential and the little change in PDI value prove that the monodispersion of LPs is stable. Finally, the ultrasonic sensitivity of GMC-SoS-Fo LPs continues to remain unchanged during storage. Table 2 The effect of particle size, PDI, packaging efficiency and ultrasonic irradiation on the release of GMC LP, GMC-Fo LP and GMC-SoS-Fo LP over 3 months

Table 2 The effect of particle size, PDI, packaging efficiency and ultrasonic irradiation on the release of GMC LP, GMC-Fo LP and GMC-SoS-Fo LP over 3 months

Flow cytometry analysis is used to quantify the cellular uptake of GMC LPs, GMC-Fo LPs, and GMC-SoS-Fo LPs with or without ultrasound irradiation. The green fluorescence recognizes the cell population that binds and absorbs the released calcein. Calcein acetoxymethyl molecule does not have green fluorescence, but is converted into green fluorescence calceinacetoxymethyl through intracellular metabolism. The cell uptake rates of GMC LPs, GMC-Fo LPs and GMC-SoS-Fo LPs were 12.5%±5.7%, 15.85%±3.9% and 16.12%±4.9%, respectively. In addition, the cell uptake of GMC LPs and GMC-Fo LPs was incubated with the cell line and treated with ultrasonic irradiation (2 MHz, 1.5 W/cm2 power intensity, 10 seconds of ultrasonic irradiation and 10 seconds of rest, totaling 0–130 seconds), They are 13.4%±4.5% and 15.45%±5.2%, respectively. However, the cell uptake of SoS-targeted LPs (GMC-SoS-Fo LPs) and cells treated with ultrasound (2 MHz, 1.5 W/cm2 power intensity, 10 seconds of ultrasound irradiation and 10 seconds of total rest time) Incubate the lines together for 0-130 seconds), which is 81.4%±8.5%, as shown in Figure 6. Compared with the cell uptake achieved with GMC LPs that have not been irradiated with ultrasound, the cell uptake achieved with GMC-SoS-Fo LPs that have been irradiated with ultrasound is improved by 6.51 times. These findings can be explained based on the intercellular uptake of Fo ligands, which are then released as bursts in response to ultrasound radiation. The Ce6 molecule inserted into the SoS targeting LP vibrates under the action of ultrasonic irradiation, creating pores or breaking the nanocarrier. In addition, the acetoxymethyl calcein released outside the cell is easily swallowed by the cell, making it green fluorescence. These results prove that SoS-targeted LPs can strongly bind to cells expressing Fo and rapidly release cargo due to ultrasound irradiation. Figure 6 Flow cytometry results of GMC LP, GMC-Fo LP and GMC-SoS-Fo LP) when incubated with CAOV3 cell line and treated or untreated with US-IR. Abbreviation: US-IR, ultrasonic irradiation.

Note: The estimated measured value after 30 minutes of incubation is the average ± SD of the three results.

Figure 6 Flow cytometry results of GMC LP, GMC-Fo LP and GMC-SoS-Fo LP) when incubated with CAOV3 cell line and treated or untreated with US-IR.

Note: The estimated measured value after 30 minutes of incubation is the average ± SD of the three results.

The cytotoxic effects of SoS Fo LP, free GMC, GMC LP, GMC-Fo LP, GMC-SoS Fo LP (without ultrasonic irradiation) and GMC-SoS Fo LP (ultrasound irradiation) on CAOV3 cells are shown in Figure 7. The CAOV3 cell line was selected for this study because its viability is not affected by the chemotherapeutic drug cisplatin at concentrations >200 µM and GMC >1,000 nM. 49 The drug-free carrier system alone (SoS Fo LPs) has no effect on cell viability, confirming the biocompatibility of SoS Fo LP. CAOV3 cells were treated with different concentrations of GMC. Free GMC, GMC LP, GMC-Fo LP and GMC-SoS Fo LP (untreated) reduce cell viability in a dose-dependent manner, up to 80 µM. However, GMC-SoS Fo LPs (treated with ultrasonic irradiation) reduced cell viability in a dose-dependent manner, up to 40 µM. GMC-SoS Fo LPs (treated with ultrasound, 40 µM drug equivalent) reported the highest cytotoxicity. Our results show that compared with GMC-SoS Fo LPs (treated with ultrasound, 40 µM drug equivalent), GMC-SoS Fo LPs (treated with ultrasound, 80 μM drug equivalent) has increased cytotoxicity Not obvious. Figure 7 Cytotoxicity of GMC. CAOV3 cells were grown in different liposome preparations and treated with increasing concentrations of GMC. Assess cell viability with MTS reagents (see method). Values ​​are expressed as the mean ± SEM of at least three different experiments in triplicate.

Figure 7 Cytotoxicity of GMC. CAOV3 cells were grown in different liposome preparations and treated with increasing concentrations of GMC. Assess cell viability with MTS reagents (see method). Values ​​are expressed as the mean ± SEM of at least three different experiments in triplicate.

The pharmacokinetics and biodistribution of the preparation were evaluated to check the efficiency of ovarian cancer targeting in vivo. As shown in Figure 8, the concentration of GMC in plasma was measured after intravenous injection of LP. The detection limits are 0.1617 and 0.4903 µg/mL, respectively, calculated using the calibration curve according to the guidelines of the International Coordinating Committee. The accuracy and precision of analysis and determination are realized within its linear range.

The plasma concentration of GMC reaches its peak (24.1±1.2 µg/mL) within five minutes after a single GMC injection. The rapid decrease to about 8.78% of the peak value within 1 hour indicates rapid elimination. The rapid plasma elimination of GMC after intravenous injection of pure drugs can be explained by the rapid metabolism and conversion of GMC to its inactive metabolites. These metabolites are more easily eliminated, which is confirmed by the increase in GMC concentration in the kidney, as shown in Figure 8.

The plasma concentrations of GMC LP, GMC-Fo LP and GMC-SoS-Fo LP are not significantly different. These results confirm that the Ce6 component does not interfere with the long circulation in LP (Table 3). Table 3 Pharmacokinetic parameters of GMC after intravenous injection of different dosage forms (free drugs, GMC LPs, GMC-Fo LPs and GMC-SoS-Fo LPs)

Table 3 Pharmacokinetic parameters of GMC after intravenous injection of different dosage forms (free drugs, GMC LPs, GMC-Fo LPs and GMC-SoS-Fo LPs)

However, the pharmacokinetic parameters of free GMC—t½, AUC0–t, AUC0–∞, mean residence time and clearance rate—are significantly different from GMC LP, GMC-Fo LP and GMC-SoS-Fo LP (p<0.05 ,table 3). Figure 8 The relationship between plasma concentration of GMC and time after intravenous injection of free drugs, GMC LP, GMC-Fo LP and GMC-SoS-Fo LP in rats (n=3). The results are expressed as mean ± SD (n=3).

Figure 8 The relationship between plasma concentration of GMC and time after intravenous injection of free drugs, GMC LP, GMC-Fo LP and GMC-SoS-Fo LP in rats (n=3). The results are expressed as mean ± SD (n=3).

The concentration of GMC in the tumor after intravenous injection of GMC LPs, GMC-Fo LPs and GMC-SoS-Fo LPs without ultrasound and GMC-SoS-Fo LPs confirmed the selective targeting of the drug to the tumor, as shown in Figure 9. The GMC concentration of GMC-SoS-Fo LPs ultrasound treatment of tumors is 2.89 times that of free GMC, while the GMC concentration of GMC-SoS-Fo LPs ultrasound treatment of tumors is 1.98 times that of GMC LPs and 1.44 times that of GMC-Fo LP. GMC-SoS -1.42 times of Fo LP without ultrasound. The highest concentration of GMC is obtained in tumors treated with GMC-SoS-Fo LPs ultrasound, which not only improves the specificity of the drug and kills cancer cells, but also reduces side effects. These benefits of using GMC-SoS-Fo LP in combination with ultrasound encourage further applications of SoS delivery systems to treat tumors.

These findings can be explained by the targeting efficiency of GMC-SoS-Fo LP and its externally triggered rapid release, which produces high concentrations in tumor tissues. The dual mechanism of GMC-SoS-Fo LPs may support the anticancer efficiency of GMC in vivo. In order to confirm these results clinically, the anti-tumor activity of various LPs was evaluated in vivo. Figure 9 The biodistribution of GMC in heart, liver, spleen, lung, kidney and ovarian cancer 30 minutes after intravenous injection of free GMC, GMC LPs, GMC-Fo LPs and GMC-SoS-Fo LPs and GMC-Fo LPs with ultrasound irradiation GMC-SoS-Fo LP. The results are expressed as mean ± SD (n=10). * p<0.05.

Figure 9 The biodistribution of GMC in heart, liver, spleen, lung, kidney and ovarian cancer 30 minutes after intravenous injection of free GMC, GMC LPs, GMC-Fo LPs and GMC-SoS-Fo LPs and GMC-Fo LPs with ultrasound irradiation GMC-SoS-Fo LP. The results are expressed as mean ± SD (n=10). * p<0.05.

As shown in Figure 10, the inhibition of tumor growth after intravenous injection of free GMC, GMC LP, GMC-Fo LP and GMC-SoS-Fo LP was evaluated, with or without ultrasound irradiation. To determine the tumor suppressor effect, the ovarian volume at the end of the experiment was measured to the ovarian volume at the beginning of the experiment. The higher the ovarian growth ratio, the lower the tumor suppressor effect. The ovarian growth ratios of free GMC, GMC LP, GMC-Fo LP and GMC-SoS-Fo LP with or without ultrasound irradiation were 3.46, 2.76, 2.25, 1.75, 1.71, and 1.21, respectively. As predicted, ultrasound-irradiated GMC-SoS-Fo LP had the lowest tumor growth ratio-1.21. This can be attributed to the targeting efficiency of the drug delivery system, the relatively long cycle time of the nanocarrier, and the external triggering of ultrasound irradiation, resulting in the rapid release of the drug cargo. Figure 10 Anti-tumor activity of GMC LPs, GMC-Fo LPs and GMC-SoS-Fo LPs without ultrasound irradiation and GMC-SoS-Fo LPs ultrasound irradiation on isotonic saline and free drugs in xenograft ovarian cancer rat models . The results are expressed as mean ± SD (n=10). The arrow shows the number of days the treatment was given. *p<0.5. Figure 11 Kaplan-Meier survival curve of rats treated intravenously with no GMC LPs (control group), 4 mg/kg free GMC solution (G1), GMC LPs (G2), GMC-Fo LPs (G3), GMC-SoS -Fo LPs (G4) and GMC-SoS-Fo LPs (G5) are exposed to ultrasound. The dose of drug administered was 4 mg/kg (n=3 for all groups).

Figure 10 Anti-tumor activity of GMC LPs, GMC-Fo LPs and GMC-SoS-Fo LPs without ultrasound irradiation and GMC-SoS-Fo LPs ultrasound irradiation on isotonic saline and free drugs in xenograft ovarian cancer rat models . The results are expressed as mean ± SD (n=10). The arrow shows the number of days the treatment was given. *p<0.5.

Figure 11 Kaplan-Meier survival curve of rats treated intravenously with no GMC LPs (control group), 4 mg/kg free GMC solution (G1), GMC LPs (G2), GMC-Fo LPs (G3), GMC-SoS -Fo LPs (G4) and GMC-SoS-Fo LPs (G5) are exposed to ultrasound. The dose of drug administered was 4 mg/kg (n=3 for all groups).

As for the Kaplan-Meier survival curve, the median survival rate of GMC-SoS-Fo LPs (G5) and ultrasound-irradiated rats was compared with all other groups. Compared with the untreated group and other groups, a significant improvement in median survival was observed in the group treated with GMC-SoS-Fo LPs (G5) and ultrasound irradiation, thereby slowing tumor recurrence (Figure 11).

Finally, the body weight of each group was measured to evaluate the safety of the dose. As shown in Figure 12, the groups injected with GMC LPs, GMC-Fo LPs, and GMC-SoS-Fo LPs recorded insignificant weight changes with or without ultrasound irradiation. However, the control group and the intravenous free GMC group were characterized by a sharp drop in body weight due to cancer cachexia and the toxicity of non-targeted dosage forms. Inhibition of tumor growth inhibited weight loss.

In living organisms with cancer, weight loss and specific loss of skeletal muscle and adipose tissue are recorded, which is called cachexia. This is highly correlated with pancreatic cancer, ovarian cancer, stomach cancer, lung cancer and bowel cancer. Weight loss may be due to processes related to metabolic changes mediated by excessive release of pro-inflammatory cytokines and increased sympathetic nervous system activity. Both catecholamines and pro-inflammatory cytokines promote catabolic processes. 36 In addition to reducing the efficacy of growth hormone, pro-inflammatory cytokines also act as inflammatory mediators in the central nervous system and act as catabolic factors that stimulate the proteolytic pathway, leading to muscle atrophy and fat gain. 50 Pro-inflammatory cytokines, including IL6, IL10, IL1β and TNFα, in addition to promoting adipose tissue waste, inhibiting adipocyte differentiation, stimulating lipolysis and increasing apoptosis in adipocytes. 51 Figure 12 Weight loss recorded in the group treated with control (isotonic saline solution), free GMC, GMC LP, GMC-Fo LP, and GMC-SoS-Fo LP, without ultrasound irradiation and GMC-SoS-Fo LP treatment with ultrasound radiation ) In a rat model of xenograft ovarian cancer. The results are expressed as mean ± SD (n=10).

Figure 12 Weight loss recorded in the group treated with control (isotonic saline solution), free GMC, GMC LP, GMC-Fo LP, and GMC-SoS-Fo LP in xenografts without ultrasound irradiation, GMC-SoS-Fo LP treated with ultrasound irradiation) ovarian cancer rat model. The results are expressed as mean ± SD (n=10).

Finally, these agents successfully control ovarian cancer because they accumulate in tumors through passive (GMC LPs) or active (GMC-Fo LPs and GMC-SoS-Fo LPs) targeting. It is reported that the accumulation of nanocarriers in tumors is due to the EPR effect, 52 due to the leakage of the vascular system and lymphatic drainage in the cancer tissue. The physical and chemical properties of nanocarriers, such as particle size, surface charge, and cycle time (lifetime), will affect their accumulation in tumors. The particle size of the nanocarrier is an important parameter that affects the EPR effect of the nanocarrier. 53 In order to achieve the extravasation of drugs in cancer tissues, the average particle size of the drug carrier should be about 100 nm. At the same time, EE should be enough. These technical requirements are important for effective drug delivery in vivo. 54,55

Due to the presence of carboxylic sugar and sulfuric acid groups, the negative charges on the surface of the blood vessel lumen interact with the positively charged nanocarriers, preventing them from recirculating. 56 However, the high positive charge may cause the opsonization and clearance of the nanocarrier 57 and may be the reason for the insignificant difference in tumor accumulation found between GMT LP and GMT-Fo LP.

Due to the large number of channels achieved through the goal, longevity can affect the passive and active accumulation of nanocarriers. In order to prevent the capture and removal of nanocarriers, space protection should be achieved by grafting polymers (such as PEG) on the surface of nanocarriers. Space protection can be explained based on shielding surface charges, improving hydrophilicity, and forming a polymer layer covering the surface of the nanocarrier, resulting in repulsion between blood components and the nanocarrier. 58

Active targeting can be achieved through two techniques. The first is to decorate the surface of the nanocarrier with ligands (such as Fo) that interact with specific receptors. Fo receptors are widely expressed in tumor tissues. The choice of targeting ligand is an important factor in the design of targeted nanocarriers. 59 The second technique is achieved by formulating stimulus-sensitive nanocarriers that respond to internal stimuli (such as changes in pathological areas or external stimulus ultrasound radiation).

Ultrasound is a mechanical wave with periodic continuous vibration with a frequency of 20 kHz or higher. They are safe; however, they have excellent tissue penetration without causing a lot of energy attenuation. 60 The impact of ultrasound can be localized by irradiating specific areas and selecting SoS drug delivery systems with cancer affinity. 61 In this study, the sensitizer is a key component. The distribution and cellular uptake of Ce6 are the most important factors in its therapeutic effects, because certain free radical products (such as peroxides from sensitizers and produced after their cellular uptake) have a short life span and a very short diffusion distance. 62 Li et al. reported that due to the accumulation of Ce6 in the mitochondria, Ce6 enhanced cytotoxicity and apoptosis, leading to injury and apoptosis. 63 Subsequently, the damaged mitochondria release many chemical factors, such as cytochrome C. The released factors stimulate caspase 9 zymogen to activate its active form, leading to cell apoptosis. 63

In this study, ultrasound irradiation was used to enhance the delivery of GMC to GMC-SoS-Fo LP for ovarian cancer. This technology uses ultrasonic irradiation to increase the permeability of the phospholipid membrane, induce cancer cells to release GMC from nanocarriers, and increase the absorption of drugs by cancer cells.

The targeted SoS LP (GMC-SoS-Fo LP) is spherical and has the best zeta potential in nanometer size. Compared with non-targeted LPs (GMC LPs) and targeted non-ultrasound sensitive LPs (GMC-Fo LPs), targeted SoS LPs successfully release drug cargo as an external stimulus in response to ultrasound radiation. In addition, targeting SoS LP enhanced intracellular uptake, thereby significantly inhibiting tumor growth.

Research data shows that GMC-SoS-Fo LPs are a promising anti-cancer drug delivery system. In order to improve the targeting efficiency of SoS LP, the Fo ligand is inserted into the prepared LP (active targeting), and SoS LP (passive targeting) is prepared in the nanometer range.

In the future, we need to develop SoS therapeutic diagnostic LP to explore the uses of these drug delivery systems, not only for tumor treatment, but also for diagnosis.

GMC, gemcitabine; GMC LP, GMC-loaded LP; GMC-Fo LP, GMC-loaded Fo target LP; GMC-SoS Fo LP, GMC-loaded Fo target SoS LP; PC, L-α-phosphatidylcholine; EE , Retention efficiency); DL, drug loading; PDI, polydispersity index; EPR, enhanced permeability and retention; Ce6, chlorin e6.

The authors report no conflicts of interest in this work.

1. Ibrahim A et al. Cancer incidence in Egypt: the results of the national population-based cancer registry program. J Cancer Epidemiology. 2014; 2014: 18. doi:10.1155/2014/437971

2. Goff BA et al. Diagnosis of ovarian cancer. cancer. 2000;89(10):2068-2075. doi:10.1002/1097-0142(20001115)89:10<2068::AID-CNCR6>3.0.CO;2-Z

3. Ten Bokkel Huinink W et al. Comparison of topotecan and paclitaxel in the treatment of recurrent epithelial ovarian cancer. J Clinical Oncology. 1997;15(6):2183–2193. doi:10.1200/JCO.1997.15.6.2183

4. Gershenson DM et al. Single-agent cisplatin in the treatment of advanced ovarian cancer. Obstetrics and Gynecology. 1981;58(4):487–496.

5. Ozols RF, etc. Phase I and pharmacological study of intraperitoneal injection of adriamycin in patients with ovarian cancer. Cancer research. 1982; 42(10): 4265-4269.

6. Coleman RL. The new role of topotecan in the first-line treatment of ovarian cancer. Oncologist. 2002;7(90005):46-55. doi:10.1634/theoncologist.7-suppl_5-46

7. Rave-Fränk M et al. The combined effect of topotecan and radiation on in vitro survival and chromosomal aberration induction. Strahlentherapie and Onkologie. 2002;178(9):497-503. doi:10.1007/s00066-002-0959-y

8. Lokich J. What is the "best" platinum: cisplatin, carboplatin or oxaliplatin. Cancer investment. 2001;19(7):756–760. doi:10.1081/CNV-100106152

9. Alexander JRM, etc. New effect of paclitaxel on cancer cells: bystander effect mediated by reactive oxygen species. Cancer research. 2007;67(8):3512–3517. doi:10.1158/0008-5472.CAN-06-3914

10. Ozols RF. The current role of gemcitabine in ovarian cancer. At the oncology seminar. Elsevier; 2001.

11. Le TN et al. Retrospective evaluation of gemcitabine/platinum regimen in the treatment of recurrent ovarian cancer. Gynecological Oncology Research and Practice. 2017;4(1):16. doi:10.1186/s40661-017-0053-x

12. Tomita Y et al. The safety and effectiveness of cisplatin combined with gemcitabine in the treatment of recurrent ovarian cancer. Int J Clin Oncol. 2014;19(4):662–666. doi:10.1007/s10147-013-0599-5

13. Alvarellos M et al. PharmGKB abstract: gemcitabine route. Pharmacogenomics. 24(11): 564-574. doi:10.1097/FPC.0000000000000086

14. Mackey JR, Mani RS, Selner M, etc. The influx of gemcitabine and the manifestation of toxicity in cancer cell lines require functional nucleoside transporters. Cancer research. 1998;58(19):4349-4357.

15. Immordino ML et al. Preparation, characterization, cytotoxicity and pharmacokinetics of liposomes containing lipophilic gemcitabine prodrugs. J Controlled release. 2004;100(3):331–346. doi:10.1016/j.jconrel.2004.09.001

16. Celano M et al. Cytotoxicity of liposomes loaded with gemcitabine in human anaplastic thyroid cancer cells. BMC cancer. 2004;4(1):1-5. doi:10.1186/1471-2407-4-63

17. What VV and so on. Magnetic nanoparticle drug delivery system for tumor targeting. Applied nanoscience. 2014; 4(4): 385–392. doi:10.1007/s13204-013-0216-y

18. Kim A et al. In vitro and in vivo transfection efficiency of a new type of ultra-deformable cationic liposome. biomaterials. 2004;25(2):305-313. doi:10.1016/S0142-9612(03)00534-9

19. Tahover E, Patil YP, Gabizon AA. Emerging delivery systems to reduce adriamycin cardiotoxicity and increase therapeutic index: focus on liposomes. Anti-cancer drugs. 2015;26(3):241–258. doi:10.1097/CAD.0000000000000182

20. Perche F. The latest trend of multifunctional liposomal nanocarriers for enhanced tumor targeting. J Drug delivery. 2013; 2013: 705265. doi:10.1155/2013/705265

21. Vice President of Deshpande PP, Biswas S, Torchilin. The current trend of using liposomes to target tumors. Nanomedicine. 2013;8(9):1509–1528. doi:10.2217/nnm.13.118

22. Corvo ML et al. Liposomes are used as a delivery system for Sn (IV) complexes for cancer treatment. Medical research. 2016;33(6):1351–1358. doi:10.1007/s11095-016-1876-6

23. Allen TM. Long circulating (sterically stable) liposomes for targeted drug delivery. Trends in Pharmaceutical Sciences. 1994;15(7):215-220. doi:10.1016/0165-6147(94)90314-X

24. Li X, etc. In vitro and in vivo evaluation of docetaxel-loaded folate receptor targeting amphiphilic copolymer modified liposomes. International J Nanomedicine. 2011; 6:1167.

25. Chen Y, etc. Baicalin is loaded in folic acid-PEG modified liposomes to improve stability and tumor targeting. Colloidal surfing B biological interface. 2016;140:74-82. doi:10.1016/j.colsurfb.2015.11.018

26. Fasehee H et al. Delivering disulfiram to breast cancer cells using PLGA-PEG nanoparticles targeting folate receptors: in vitro and in vivo studies. J Nano Biotechnology. 2016;14(1):32. doi:10.1186/s12951-016-0183-z

27. Lukyanov AN et al. Tumor-targeted liposomes: Long-circulating liposomes loaded with adriamycin modified with anti-cancer antibodies. J Controlled release. 2004;100(1):135–144. doi:10.1016/j.jconrel.2004.08.007

28. Kumar P, Huo P, Liu B. Folate targeting liposome formulation strategy and its biomedical application. pharmaceutics. 2019;11(8):381. doi:10.3390/pharmaceutics11080381

29. Zhang Ping, et al. Iron oxide nanoparticles are used as nanocarriers to improve the sonosensitivity of chlorin e6 in sonodynamic therapy. The drug Devel Ther. 2018; 12:4207. doi:10.2147/DDDT.S184679

30. Ms. Newman et al. Comparison of pharmacokinetics, tissue distribution and therapeutic effects of cisplatin encapsulated in long-circulating pegylated liposomes (SPI-077) in tumor-bearing mice. Cancer Chemistry and Pharmacy. 1999;43(1):1-7. doi:10.1007/s002800050855

31. Yamada A et al. The design of folic acid linked liposome doxorubicin has anti-tumor effect on mice. Clinical cancer research. 2008;14(24):8161-8168. doi:10.1158/1078-0432.CCR-08-0159

32. Torchilin Vice President. Passive and active drug targeting: Take drug delivery to tumors as an example, in drug delivery. Springer; 2010:3-53.

33. Monteiro LOF et al. Folic acid-coated long-circulation and pH-sensitive liposomes loaded with paclitaxel as a potential drug delivery system: a biodistribution study. Biomedical drug treatment. 2018;97:489-495. doi:10.1016/j.biopha.2017.10.135

34. Gulati M et al. Lipophilic drug derivatives in liposomes. Int J Pharm. 1998;165(2):129-168. doi:10.1016/S0378-5173(98)00006-4

35. Liu Rui et al. Liquid crystal nanoparticles as an ophthalmic drug delivery system for tetrandrine: development, characterization, and in vitro and in vivo evaluation. Nano Res Lett. 2016;11(1):254. doi:10.1186/s11671-016-1471-0

36. Egelhaaf SU et al. Determination of the size distribution of lecithin liposomes: a comparative study using freeze-fracture, cryo-electron microscopy and dynamic light scattering. J Microsc. 1996;184(3):214-228. doi:10.1046/j.1365-2818.1996.1280687.x

37. Omar MM, Hasan OA, El Sisi AM. Preparation and optimization of lidocaine transfer gel containing penetration enhancer: a promising method to enhance skin penetration. International J Nanomedicine. 2019; 14: 3705-3727. doi:10.2147/IJN.S201356

38. Salem HF et al. Targeting brain cells with nanoliposomes modulated by glutathione: in vitro and in vivo studies. The drug Devel Ther. 2015; 9:3705. doi:10.2147/DDDT.S85302

39. X Fu, RM Hoffman. A human ovarian cancer metastasis model constructed in nude mice by orthotopic transplantation of histologically complete patient specimens. Anticancer Research 1993;13(2):283-6. Available from: https://pubmed.ncbi.nlm.nih.gov/8517640/

40. Left L, Roberts WJ, Evans KD. Diagnostic ultrasound imaging of mouse diaphragm function. J Vis Exp. 2014; 86: 51290.

41. Lin NM et al. High pressure liquid chromatography and diode array detector were used to determine gemcitabine and its metabolites in human plasma. Journal of Pharmaceuticals. 2004;25:1584-1589.

42. Pflanzer R et al. Advanced 3D ultrasound imaging is used as a precise technique for evaluating tumor volume. Translation oncology. 2014; 7(6): 681–686. doi:10.1016/j.tranon.2014.09.013

43. Lila ASA, Ishida T, Kiwada H. Cationic liposomes are used to target anticancer drugs to the tumor vasculature. Medical research. 2010;27(7):1171–1183. doi:10.1007/s11095-010-0110-1

44. Schroeder A, Kost J, Barenholz Y. Ultrasound, liposomes, and drug delivery: The principle of using ultrasound to control the release of drugs from liposomes. Chemical and physical lipids. 2009;162(1-2):1-16. doi:10.1016/j.chemphyslip.2009.08.003

45. Suzuki R et al. The development of ultrasound-mediated gene delivery systems using nanometers and microbubbles. J control release. 2011;149(1):36-41. doi:10.1016/j.jconrel.2010.05.009

46. ​​Evjen TJ et al. Distearoylphosphatidylethanolamine-based liposomes for ultrasound-mediated drug delivery. European J Pharmaceutics Biopharmaceutics. 2011;75(3):327–333. doi:10.1016/j.ejpb.2010.04.012

47. Kusube M, Matsuki H, Kaneshina S. Thermally induced and positive pressure phase transition of N-methylated dipalmitoylphosphatidylethanolamine bilayer. Biochimica Et Biophysica Acta (BBA)-biofilm. 2005;1668(1):25–32. doi:10.1016/j.bbamem.2004.11.002

48. Huang SL, McDonald RC. Acoustic active liposomes for drug encapsulation and ultrasound-triggered release. Biochimica Et Biophysica Acta (BBA)-biofilm. 2004;1665(1–2):134–141. doi:10.1016/j.bbamem.2004.07.003

49. Kawaguchi H et al. As a molecular targeting agent, gemcitabine can block the Akt cascade in platinum-resistant ovarian cancer. J Ovarian Research. 2014;7(1):1-11. doi:10.1186/1757-2215-7-38

50. Steinman J, DeBoer, MD. Treatment of cachexia: melanocortin and growth hormone releasing peptide intervention in vitamins and hormones. Elsevier; 2013:197-242.

51. Rania S Abdel-Rashid, Doaa A Helal, Mahmoud M Omar and Amani M El Sisi; nanogels loaded with surfactant-based nanovesicles to enhance the ocular delivery of acetazolamide. International J Nanomedicine. 2019; 14: 2973-2983.

52. Matsumura Y, Maeda H. New concept of macromolecular therapy in cancer chemotherapy: the mechanism of promoting tumor accumulation of proteins and anti-tumor agents smancis. Cancer research. 1986; 46 (12 Part 1): 6387–6392.

53. Maeda H, Nakamura H, Fang J. EPR effect of macromolecular drug delivery to solid tumors: improve tumor uptake, reduce systemic toxicity, and different tumor imaging in vivo. Adv Drug Deliv Rev. 2013; 65(1): 71–79. doi:10.1016/j.addr.2012.10.002

54. COSCO D et al. In vivo activity of pegylated small unilamellar liposomes loaded with gemcitabine on pancreatic cancer. Cancer Chemistry and Pharmacy. 2009;64(5):1009-1020. doi:10.1007/s00280-009-0957-1

55. Celia C et al. Innovative nanocarriers loaded with gemcitabine and GEMZAR: biodistribution, pharmacokinetic characteristics and anti-tumor activity in vivo. Expert opinion on drug delivery. 2011;8(12):1609-1629. doi:10.1517/17425247.2011.632630

56. Ding HM. Use pH-sensitive polymers to control the cellular uptake of nanoparticles. Scientific Representative 2013; 3:2804. doi:10.1038/srep02804

57. Xiao Ke, wait. The effect of surface charge on the in vivo biodistribution of PEG-oligocholic acid-based micellar nanoparticles. biomaterials. 2011;32(13):3435–3446. doi:10.1016/j.biomaterials.2011.01.021

58. Gabizon A, Papahadjopoulos D. The effect of surface charge and hydrophilic groups on the clearance of liposomes in vivo. Biochimica Et Biophysica Acta (BBA)-biofilm. 1992;1103(1):94–100. doi:10.1016/0005-2736(92)90061-P

59. Bertrand N, Leroux JC. The journey of the drug carrier in the body: an anatomical and physiological perspective. J Controlled release. 2012;161(2):152-163. doi:10.1016/j.jconrel.2011.09.098

60. Jensen JRA. Model of the propagation and scattering of ultrasound in tissues. J Acoust Soc Am. 1991;89(1):182-190. doi:10.1121/1.400497

61. FERIL, Jr. LB, Kondo T. The biological effects of low-intensity ultrasound: the mechanisms involved and their impact on the treatment and the biological safety of ultrasound. J Radiation Research. 2004;45(4):479–489. doi:10.1269/jrr.45.479

62. Yumita N. Membrane lipid peroxidation as a mechanism of sonodynamic induction of red blood cell lysis. Int J Radiat Biol. 1996;69(3):397-404. doi:10.1080/095530096145959

63. Li Qiang, etc. The effect of chlorin e6 mediated acousto-photodynamic therapy on 4T1 cells. Cancer Biother Radiopharm. 2014;29(1):42-52. doi:10.1089/cbr.2013.1526

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