Photothermal therapy enhance the anti-mitochondrial metabolism effect of lonidamine to renal cell carcinoma in homologous-targeted nanosystem

Jiyuan Chen, Fuzheng Ren, Wenfeng Cao, Zhengjie Wu, Guanqun Ju, Chengwu Xiao, Weijia Wu, Shen Gao, Chuanliang Xu, Yuan Gao
A Department of Clinical Pharmacy and Pharmaceutical Management, School of Pharmacy, Fudan University, Shanghai, China
B Department of Pharmacy, East China University of Science and Technology, Shanghai, China
C Department of Urology, Changhai Hospital, Second Military Medical University, Shanghai, China Revised 8 January 2021

Renal cell carcinoma (RCC) is a common malignant tumor of the urinary system with poor prognosis. Therapeutic drugs for RCC can easily develop resistance or have unignorable toxicity or limited efficiency. Here, the thermosensitive mitochondrial metabolism-interfering anticancer drug lonidamine (LND) was combined with the photothermal material polydopamine (PDA) to treat RCC. To delivery drugs accurately to RCC site, LND and PDA were loaded in stellate mesoporous silica nanoparticles (MSNs) with a large surface area and cloaked with RCC membranes (MLP@M). The results showed that MLP@M exhibited excellent tumor targeting ability. The synergistic effects of LND and PDA in MLP@M were greatly enhanced when triggered by an 808 nm laser. Moreover, the antiproliferative and tumor suppressing abilities were enhanced with good biocompatibility after MLP@M + laser treatment. Additionally, 80% of RCC tumor-bearing mice treated with MLP@M + laser did not relapse. Our study provides a potential therapeutic approach for RCC treatment.

Kidney cancer is a common malignant tumor in urology and accounts for approximately 2% of systemic malignancies worldwide.1 In 2020, it has been estimated that approximately 73,750 new cases of kidney cancer will be diagnosed in the United States with 14,830 deaths, and renal cell carcinoma (RCC) accounts for over 90% of kidney cancers.2,3 Vascular endothelial growth factor receptor (VEGFR) inhibitors and immune checkpoint inhibitors are the frontline treatments for RCC; however, these existing treatments could not avoid relapse and have poor prognosis.- 4 Although the overall survival of RCC has improved substantially with the development of RCC diagnosis and management, there are various reasons for the failure of RCC therapy, including drug resistance, loss of efficacy of a particular agent, and evidence for advanced progress and greater toxicity with combination therapy.5,6 Therefore, new combination therapies with better tumor-targeting ability for RCC treatment are urgently needed.
Lonidamine (LND), a thermosensitive mitochondrial metabo- lism inhibitor, has been used to treat many types of cancer with reduced toxicity.7–9 LND induces apoptosis of cancer cells by

inhibiting hexokinase and acting on mitochondrial adenine nucleotide translocase. The translocase can make mitochondria misfunction by triggering the opening of the mitochondrial permeability transition pore.10,11 Meanwhile, the intracellular ATP production is suppressed and the P-glycoprotein (a classic multidrug resistance efflux pump) inhibition is enhanced, re- sensitizing cancer cells to chemotherapy. Therefore, LND is widely used to enhance the therapeutic effects of chemotherapy as well as radiotherapy, photothermal therapy, and photodynamic therapy based on interfering of mitochondrial metabolism.12–14 LND can also inhibit glycolysis, reduce the activity of succinate dehydroge- nase and induce cellular reactive oxygen species (ROS), promoting apoptosis and cell death.15 Moreover, a phase II study showed that lonidamine was moderately effective against RCC with a 37.5% one-year survival rate.16 Therefore, LND could considerably enhance the photothermal therapy effects for RCC treatment and achieve a better outcome than other treatments.
Polydopamine (PDA) is a polymer of the dopamine (DA), which can oxidatively aggregate and self-assemble in alkaline aqueous solution, and has good biocompatibility, biodegrada- tion, reducibility, and especially photothermal conversion capability.17,18 PDA can be stimulated by a near-infrared (NIR) laser, which can quickly increase the local temperature to over 45 °C with superb photothermal properties. The increased temperature of tumor microenvironment could trigger the death of cancer cells, and nano-photothermal transduction agents could accumulate in the tumor site through the enhanced permeability and retention (EPR) effect.19,20 Therefore, the combination of LND and PDA may enhance the antitumor effects while showing good biocompatibility and biodegradability.
To delivery of LND and PDA to RCC, mesoporous silica nanoparticles (MSNs) were used due to their controllable size, large surface area (>1000 m2·g−1), tunable pore sizes (2-20 nm) and high pore volume.21 Moreover, the pore size and surface chemistry of MSNs can be easily modified to accommodate a variety of components (chemotherapeutic drugs, RNA, peptides, photosensitizers, etc.) with high loading capacities.22–24 How- ever, there are still some limitations for the efficient drug delivery of MSNs: 1) instability in blood, leading to leakage and premature release of the physically adsorbed drugs and 2) the limited circulation lifetime caused by the recognition of reticule endothelial system, resulting in sublethal tumor-targeted distri- bution of the therapeutic drugs.25 Therefore, it is necessary to find a simple and efficient strategy to overcome these deficiencies and achieve efficient RCC targeted delivery.
As a top-down approach, biomimetic cell membrane-coated nanoparticles bypass laborious group-modified engineering.26,27 Owing to the reserved antigens and cell membrane structure, biomimetic nanoparticles can acquire special functions, such as ligand recognition and targeting, long blood circulation, and immune escape, offering a promising nanoplatform for drug delivery, detoxification, and vaccination.28 A variety of cell membrane-cloaked nanosystems have shown excellent homolo- gous cancer targeting ability and biocompatibility.29–33 To our knowledge, biomimetic nanoparticles cloaked with RCC mem- branes have not been previously reported.
Herein, we designed a laser-responsive MSNs-supported cancer membrane-mimetic nanodrug delivery system withenhanced blood circulation and tumor-targeted drug release (as shown in Figure 1). The photosensitizer PDA and LND were coloaded in the pores of stellate MSNs (MLP), and RCC membranes were extracted to coat the surface of the MLP to form the final cancer cell membrane-coated MLP@M. We hypothesized that the membrane mimetic drug delivery system could stabilize MSNs in blood, protect the encapsulated drugs from leakage and enhance the circulation lifetime. In return, MSNs could suppress cancer membrane fluctuations and maintain a relatively stable system. In addition, as PDA can polymerize in alkaline solution, in return, it is biodegradable in tumor acidic microenvironment (TME) or lysosomal system.34 When stimulated by an 808 nm NIR laser, PDA can generate a high temperature and destroy the cancer cell membrane, leading to LND release and tumor-specific LND accumulation. We expected that cancer cell membrane-mimetic MSNs could integrate with PDA-based photothermal therapy and LND- based chemotherapy to completely suppress RCC.

All the materials and methods were provided in Supplementary Materials.

Results and discussion
Preparation and characterization of MLP@M
In this work, stellate MSNs with a special radial pore morphology were chosen as the nanovehicle. These stellate MSNs have a specific surface and a well-defined conical pore size of 2.9 nm with an interparticle void size of up to 80 nm, which has been reported for drug delivery (Supplementary Material 2).25,35 As shown in Figure 2, A, transmission electron microscopy (TEM) was applied to characterize the MSNs, which exhibited a stellate morphology. The drug-loaded MLP showed an accumulation of black particles with slight changes in size (Figure 2, B). After cloaking with the RCC membrane, a classic core-shell structure of the MLP@M could be witnessed (Figure 2, D). After the MLP and MLP@M were irradiated with an 808 nm NIR laser at a power of 1.0 W cm−2 for 2 min, the structures of MLP and MLP@M were obviously damaged (Figure 2, C, E). The dynamic light scattering results of the MSNs and modified MSNs were shown in Figure 2, F, G. After loading of LND, the size of the MLD increased from 107.5 ±2.0 nm to 116.8 ± 0.2 nm, and the zeta potential decreased from5.8 ± 1.3 mV to −1.6 ± 0.4 mV. When the MLD was further loaded with PDA, the size and zeta potential of the MLP were172.9 ± 4.2 nm and − 32.9 ± 1.8 mV, respectively. After cancercell membrane coating, the size of MLP@M increased to 211.1 ±9.4 nm, with a zeta potential of −15.1 ± 0.8 mV close to that of the membrane (−17.5 ± 0.3 mV), suggesting the success of the coating of the cancer cell membrane onto the surface of the MLP. Additionally, MLP@M were relatively stable in phos- phate buffer solution (PBS) for over 30 days (Figure S2). The drug loading (DL) and encapsulation efficiency (EE) were 49.7% and 96.0% for MLD, respectively, and those values for MLP@M were 12.4% and 24.6%, respectively (Figure S3). Thehigh loading capacity could be due to the high surface/volume ratio of MSNs.
For quantitative analysis, thermal gravimetric analysis (TGA) was performed (Figure S4). When the materials were heated to 300 °C, the weight loss of bare MSNs and MSNs- PDA was 7.304% and 12.66% respectively. The increase in weight loss was mainly due to the photothermal effects of the PDA films. Furthermore, the maximum ultraviolet (UV) absorption of LND at ~300 nm36 disappeared after coating with PDA and/or cell membrane films (Figure 2, H). Additionally, the MSN@M blank carrier had good biocom-patibility without obvious toxicity at concentrations up to 800 μg/mL (Figure S5).

In vitro drug release, cellular uptake, and homologous targeting ability of MLP@M
The in vitro drug release profiles of MLD, MLP and MLP@M at different pH values (pH = 7.4, 5.0) were investigated (Figure 3, A, B and Figure S6). The results showed that LND was typically released from all of the nanoparticles in a biphasic pattern. Under normal physiological condition(pH = 7.4), the MLD group exhibited 86.47% release of LND over 3 days. In contrast, for the MLP@M + laser and MLP@ M groups, the amount of released LND was 75.38% and 57.41%, respectively. The distinct release profile suggested that the PDA and cell membrane films coated onto the surfaces of the nanoparticles could block the pores of the MSNs and effectively suppress drug release. Moreover, when the pH value decreased from 7.4 to 5.0 to simulate the environment of late endosomes/lysosomes, the release rate notably increased. Both the MLP@M + laser and MLP@M groups exhibited remarkably higher LND release (87.80% and 67.72%, respectively). These results also suggested that the NIR laser helped the release of LND as it destroyed the PDA and cell membrane films. The conspicuous increase in LND release from the nanoparticles with the PDA film under acidic conditions suggested that upon decreasing the pH, the surface of the PDA film was partially depolymerized, which unlocked the gate for LND release. Interestingly, as the acidity increased, the LND release rate of non-PDA-coated nanopar-ticles (MLD) increased. This might be that acidic conditions could enhance the solubility of LND and therefore increase drug release.37
To evaluate the cellular uptake of MLP@M, Nile red (Nile) was used as the model drug. As shown in Figure 3, C, 786-O cells were incubated with Nile, Nile-loaded MSNs (MSNs-Nile), and Nile-loaded MSNs@M (MSNs-Nile@M) for 4 h. The red fluorescence of the free Nile group was weak, while the signals from the MSNs-Nile and MSNs-Nile@M groups were much stronger and the signal from the MSNs-Nile@M group was the strongest. Additionally, the results of confocal laser scanning microscopy (CLSM) images were in line with those of the flow cytometry (Figure 3, D and Figure S7).
Moreover, to evaluate the homologous targeting ability of MLP@M, MSNs-Nile@M (coated with a 786-O cell membrane) were co-incubated with 786-O, KETR-3, RWPE-1, C4-2B, PC- 3, U251, and DU145 cells for 2 h. The flow cytometry results showed that the cellular uptake by the RCC cell lines (786-O and KETR-3, especially 786-O) was much higher than that of theother cell lines, suggesting good homologous targeting ability of MLP@M (Figure 3, E).

Photothermal and antimitochondrial effects of MLP@M
PDA is capable of absorbing NIR irradiation (700-900 nm) efficiently, and its light-thermal conversion efficiency is up to 40%.37 To demonstrate the photothermal effects of MLP@M, we monitored the temperature change of the PBS (control group), MLP, and MLP@M groups using an 808 nm NIR laser (1.0 W cm−2, 5 min). The temperature quickly increased and then plateaued within 5 min (Figure 4, A, B). The finaltemperatures of the MLP and MLP@M groups were 56.6 °C and 47.2 °C, respectively. No obvious temperature changes were detected in the control group. Additionally, the photothermal effect was proportional to the concentration of MLP@M (Figure 4, C). These results demonstrated the significant photothermal effects of MLP@M.
As a mitochondrial metabolism inhibitor, LND can interfere with glycolysis in mitochondria and boost the generation of ROS in tumor cells.38 Therefore, the ROS probe 2,7-dichlorodihydro- fluorescein diacetate (DCFH-DA) was used to determine the intracellular ROS content, which can turn into a highly fluorescent product upon oxidation. 786-O cells were treatedwith PBS, LND, MLD, MLP, MLP + laser, MLP@M, and MLP@M + laser (laser: 808 nm, 1 W/cm2, 5 min). As shown in Figure 4, D, F and Figure S8, groups treated with the laser had a much stronger fluorescence intensity than other groups, and the signals increased from 8 to 24 h. Moreover, MLP@M irradiated with a NIR laser induced a much higher concentration of ROS than the other groups. Furthermore, the effects of LND on mitochondrial membrane potential were detected using rhoda- mine 123 as a probe. As shown in Figure 4, E, G, 786-O cells were incubated with each group for 8 h, and the fluorescence signals of the 6 drug groups weakened to different degrees, while those of the groups treated with laser decreased sharply. Moreover, there was almost no signal in the MLP@M + laser group. Taken together, these results suggest a synergistic effect of photothermal therapy and anti-mitochondrial metabolism therapy.

Antiproliferative effects of NIR laser irradiated MLP@M
The cytotoxicity effects were evaluated by a CCK-8 kit. As shown in Figure 5, A, B and Figure S9, MLP@M + laser displayed greater cytotoxicity than that of the other groups, and the MLP@M + laser cytotoxicity enhanced as the incubation time increased. This could be due to enhanced drug uptake and controlled release of the nanoparticles compared to the other groups. MLP + laser and MLP@M + laser showed significantly enhanced cytotoxicity to 786-O cells compared to MLP and MLP@M. The half maximal inhibitory concentration (IC50) values in 786-O cells after 48 h of incubation for MLP@M and MLP@M + laser were 10.3 μg/mL and 7.8 μg/mL, respectively, which were 1.5- and 1.3-fold lower than those of MLP (15.9 μg/ mL) and MLP + laser (10.1 μg/mL). This result was probably because the cloaked RCC membrane could prevent prematuredrug release of LND and PDA, leading to decreased efficiency. The MLP@M + laser group also showed the most effective cell killing ability compared to both the LND and MLD groups (in which the IC50 values showed little change after laser irradiation), as expected. The IC50 value of MLP@M + laser after 48 h of incubation was 4.7- and 2.7-fold lower than that of LND (36.7 μg/mL) and MLD (20.7 μg/mL), respectively. These results demonstrated that MLP@M could prevent premature drug release, trigger efficient intracellular LND distribution, and integrate photothermal therapy and chemotherapy to efficiently kill RCC cells under laser irradiation.
Moreover, 786-O cells were treated with LND, MLD, MLP, and MLP@M with or without an 808 nm NIR laser (1.0 W cm-−2, 5 min) for 24 h to detect the apoptosis rate. The resultsshowed that the apoptosis rate of the MLP@M group was much higher than that of the other LND-loaded groups or LND group (Figure 5, C). Additionally, the apoptosis rate increased sharply from 48.7% to 75.5% when MLP@M were irradiated by an NIR laser, which was in line with the results of the cytotoxicity study. All these results revealed the enhanced synergistic antiprolifer- ative effects of PDA and LND, as LND was sensitive to PDA- elevated temperatures.

In vivo biodistribution of MLP@M
To evaluate the in vivo distribution of MLP@M, 1,1- dioctadecyl-3,3,3,3-tetramethylindotricarbocyaine iodide (DIR) was used as a model drug. Free DIR, MSNs-DIR, and MSNs-DIR@M solutions were intravenously injected into nude mice bearing a subcutaneous tumor. As shown in Figure 6, A, there was a specific strong fluorescence signal in the tumor area of the MSNs-DIR@M group as early as 1 h post-injection that did not fade after 24 h. For the MSNs-DIR group, although the fluorescence signal of DIR increased from 4 to 24 h in the tumor area, this signal also accumulated in the liver area at 1 h after injection. In addition, the fluorescence signal of the free DIR group was observed mainly in the liver area without an obvious signal in the tumor area at all time internals. Furthermore,fluorescence images of the dissected major organs and tumors were detected (Figure S10). The results showed that the fluorescence signal of the tumors of the MSNs-DIR@M group was much stronger than that of the MSNs-DIR group, which was in line with the results of live imaging (Figure 6, B). In contrast, the fluorescence signal of the free DIR group mainly accumulated in the liver with rather weak exhibition in the tumor.

In vivo antitumor effects of MLP@M
The outstanding in vitro synergistic effects as well as good biocompatibility of MLP@M would signify the possibility of high therapeutic efficiency in vivo. To investigate the photo- thermal effects of MLP@M in vivo, mice were treated with PBS, LND, MLD, MLP, or MLP@M (LND or PDA: 10 mg/kg).
Eight hours after intravenous injection, all the mice were irradiated with a laser (808 nm, 2.0 W cm−2) at the tumor site for 5 min. As shown in Figure 7, A and Figure S11, the surface temperatures of the tumors treated with MLP and MLP@M increased to 53.2 °C and 55.7 °C within 5 min, respectively, while those of the PBS, LND, and MLD groups only increased to42.3 °C, 39.1 °C, and 42.0 °C, respectively. These results suggested the excellent photothermal effects of MLP@M with prolonged accumulation in tumor tissue. As shown in Figure 7, B, C and Figure S12, compared with RCC tumor-bearing mice treated with PBS, the tumor growth in mice treated with LND or LND-loaded formulations with or without laser irradiation was much slower. Moreover, in the MLP + laser and MLP@ M + laser groups, tumor progression was inhibited, and 80% of mice in the MLP@M + laser group did not relapse. In addition, only mice treated with free LND suffered body weight loss, while other groups gained weight steadily (Figure 7, D). Furthermore, the H&E staining results (Figure 7, E) showed that there was no obvious toxicity in the major organs of all groups except for the liver of the LND group. The nonspecific distribution of LND in the liver caused nuclear lysis, cell membrane damage and necrosis. In contrast, the MLP@M groupexhibited low toxicity, since the enhanced tumor-targeted ability of MLP@M reduced liver accumulation and alleviated the toxicity of LND. All these results supported the hypothesis of enhanced therapeutic efficacy and good biocompatibility byusing a biomimetic homologous targeted delivery nanosystem for the combined administration of LND and PDA.
In conclusion, we developed a targeted nanosystem MLP@M for combination of mitochondrial metabolism inhibitor and photo-thermal therapy for RCC. The large surface area of stellate MSNs provided a high drug loading and encapsulation efficiency. Meanwhile, MLP@M presented a pH-sensitive release behavior due to the pH sensitive attribute of PDA. The homologous targeting ability was demonstrated both in vivo and in vitro. When MLP@M was triggered by an 808 nm laser, MLP@M achieved enhanced anti- proliferation and tumor-suppressing abilities, with good biocompat- ibility. Moreover, there was almost no relapse in RCC tumor-bearing mice treated with MLP@M + laser. Therefore, this biomimetic nanosystem could be a potential approach for RCC treatment.

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