Multifunctional Thermosensitive Liposomes Based on Natural Phase

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Biological and Medical Applications of Materials and Interfaces

Multifunctional Thermosensitive Liposomes Based on Natural Phase Change Material: Near-Infrared Light-Triggered Drug Release and Multimodal Imaging Guided Cancer Combination Therapy Yeneng Dai, Jinzhong Su, Kun Wu, Wenkang Ma, Bing Wang, Meixing Li, Pengfei Sun, Qingming Shen, Qi Wang, and Quli Fan ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.8b22748 • Publication Date (Web): 26 Feb 2019 Downloaded from http://pubs.acs.org on February 27, 2019

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Multifunctional Thermosensitive Liposomes Based on Natural Phase Change Material: Near-Infrared Light-Triggered Drug Release and Multimodal Imaging Guided Cancer Combination Therapy Yeneng Dai, Jinzhong Su, Kun Wu, Wenkang Ma, Bing Wang, Meixing Li, Pengfei Sun, Qingming Shen,* Qi Wang,* Quli Fan* Key Laboratory for Organic Electronics and Information Displays & Jiangsu Key Laboratory for Biosensors, Institute of Advanced Materials (IAM), Jiangsu National Synergetic Innovation Center for Advanced Materials (SICAM), Nanjing University of Posts & Telecommunications, Nanjing 210023, China. KEYWORDS: Liposomes, Thermosensitive, Phase change material, Multimodal imaging, Combination therapy

ABSTRACT: Multifunctional theranostic nanoplatforms (NPs) in response to environment stimulations for on-demand drug release are highly desirable. Herein, the near-infrared (NIR)absorbing dye, indocyanine green (ICG) and the antitumor drug, doxorubicin (DOX) were efficiently co-encapsulated into the thermosensitive liposomes based on natural phase change material (PCM). Folate and conjugated gadolinium chelate-modified liposome shells enhance 1 Environment ACS Paragon Plus

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active targeting and magnetic resonance (MR) performance of the NPs while maintaining the size of the NPs. The ICG/DOX loaded and gadolinium chelates conjugated temperature-sensitive liposomes nanoplatforms (ID@TSL-Gd NPs) exhibited NIR-triggered drug release and prominent chemo-, photothermal, photodynamic therapy properties. With the co-encapsulated ICG, DOX and the conjugated gadolinium chelates, the ID@TSL-Gd NPs can be used for triplemodal imaging (fluorescence/photoacoustic/magnetic resonance imaging, FL/PAI/MRI) guided combination tumor therapy (chemotherapy, photothermotherapy and photodynamic therapy, Chemo/PTT/PDT). After tail vein injection, the ID@TSL-Gd NPs accumulated effectively in subcutaneous HeLa tumor of mice. The tumor was effectively suppressed by accurate imaging guided NIR triggered phototherapy and chemotherapy, and no tumor regression and side effects were observed. In summary, the prepared ID@TSL-Gd NPs achieved multimodal imagingguided cancer combination therapy, providing a promising platform for improving diagnosis and treatment of cancer.

INTRODUCTION Multifunctional theranostic nanoplatforms (NPs) have offered great opportunities to improve the diagnosis accuracy and treatment efficacy of tumor.1-3 Compared to traditional sustained release, environmentally sensitive NPs can quickly respond to external stimuli (such as temperature, light, ultrasound and magnetic field) and internal stimuli (such as pH, enzymes and redox) to control the release of drugs at tumor site, improving tumor therapeutic efficacy and reducing harmful side effects.4-7 Among various types of stimuli, light is the most promising one due to its intrinsic merits, such as noninvasiveness, remote spatiotemporal control and facile application.8 With the deep tissue penetration capability of near-infrared (NIR) laser, the NIRresponsive systems have been broadly developed,9 most of which are built on the conformational

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changes associated with thermosensitive polymers to trigger anti-tumor drug release. Various thermal sensitive polymers have been developed, including poly(lactic-co-glycolic acid),10 poly(N-isopropylacrylamide),11 polysaccharides and their derivatives.12 However, these polymers have been limited to clinical applications mainly due to their intrinsic drawbacks such as complicated synthesis procedures, poor biodegradability, low degradability and considerable cytotoxicity. 9, 13, 14 Recently, the eutectic mixture of natural fatty acid, a novel phase-change material (PCM), was used as the thermal responsive gating material to construct NIR-responsive NPs for controlling drug release. As an ideal PCM, eutectic mixture of natural fatty acids have drawn great attention owing to its unique advantages such as low cost, good biocompatibility and chemical stability, as well as controllable melting temperature.15,

16

Under NIR irradiation, the photothermal agent

could generate a local hyperthermia, melt the eutectic mixture and release drugs quickly.17 Meanwhile, the photothermal effect generated by photothermal agents under NIR light irradiation has been demonstrated to be a highly efficient cancer therapeutic technique. As a promising alternative or supplement to traditional cancer therapy, phototherapy including photothermotherapy (PTT) and photodynamic therapy (PDT), has exhibited several unique advantages, such as noninvasiveness, precise spatial and temporal selectivity.18-21 Under irradiation, the light beam is precisely focused on the tumor area, and the light is absorbed by the photothermal agent or photosensitizer and converted into local hyperthermia or reactive oxygen species (ROS) to induce tumor apoptosis.22 Nevertheless, it is not enough to kill all tumor cells due to the uneven heat distribution over tumors from PTT technique. In comparison with monotherapy, the integration of PTT, PDT and chemotherapy into a single NP has shown enhanced therapeutic efficacy and reduced side effects. On the one hand, the photothermal effect


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induced can induce thermal ablation of cancer cells and increase the sensitivity of tumor cells to chemotherapy drugs. On the other hand, it can also improve the permeability of tumor vessels and the oxygenation status of the tumor microenvironment.15−17 Therefore, the PTT technique has been widely associated with PDT and chemotherapy to increase the tumor therapeutic efficacy in a combined manner. 23, 24 During tumor treatment, for precise, effective and safe treatment, it is important to determine the tumor size and location before treatment, monitor the distribution of drugs during treatment and subsequently evaluate therapeutic efficacy after tumor therapy.6–8 The imaging-guided therapy has aroused great attention due to its intrinsic advantages, such as providing location of tumor and monitoring therapeutic efficacy in vivo.17,18,25 Up to now, various imaging techniques have been developed while designing multifunctional theranostic nanoplatforms, such as positron emission tomography, magnetic resonance imaging (MRI), photoacoustic imaging (PAI) and fluorescence (FL) imaging.26-33 Integration of diagnostic and therapeutic functions into a single NPs is highly desirable for enhancing tumor therapeutic efficacy. Indocyanine green (ICG) is a clinical NIR dye approved by the U.S. Food and Drug Administration (FDA). It can not only convert NIR light energy into local hyperthermia and ROS for photothermal and photodynamic therapy, but also can be used for multimodal imaging including FL imaging and PA imaging.34-36 However, free ICGs are limited for further biomedical applications due to their intrinsic drawbacks, including concentration-dependent aggregation, poor photostability and thermal stability, rapid clearance in vivo, and the lack of tumor targeting specificity.37 Therefore, a series of nanocarriers have been developed to improve the stability and tumor-specificity of ICG, such as human serum albumin as a delivery system self-assembled with ICG for in vivo dual-modal imaging and combination therapy,38 poly(lactic-co-glycolic acid) (PLGA)-lecithin-



polyethylene glycol (PEG) nanoparticles encapsulating DOX and ICG for chemo-photothermal therapy and FL imaging,39 DOX/ICG-loaded temperature sensitive liposomes (DI-TSL) for chemo-photothermal combination therapy together with modulation of drug release.1 Herein, a biocompatible eutectic mixture of natural fatty acid composed of lauric acid and stearic acid was used as a PCMs core which co-encapsulated with ICG (phototherapeutic agent) and DOX (chemotherapeutic agent). When the weight ratio of lauric acid (m.p. = 44 °C) and stearic acid (m.p. = 69 °C) is 4:1, the eutectic mixture formulated by them exhibits a sharp melting point at 39 °C.17 The PCMs core, which exhibits excellent thermal sensitive properties, achieves on-demand drug release in response to NIR laser. The liposomal shell can not only improve the biocompatibility but also endow new functionality to the NPs. With the modification of folic acid-phospholipids and conjugated gadolinium chelates, the ICG/DOX loaded and gadolinium chelates conjugated temperature-sensitive liposomes NPs (ID@TSL-Gd NPs) were successfully synthesized. The prepared NPs not only possess passive tumor-targeting abilities via enhanced permeability and retention (EPR) effect but also showed good active tumor-targeting abilities by FA receptor-mediated transcytosis. The physicochemical properties and drug release mechanism of the NPs were systematically studied. With the co-encapsulated ICG, DOX and the conjugated gadolinium chelates, the ID@TSL-Gd NPs can be used for FL/PAI/MRI triple-modal imaging guided cancer combination therapy (Chemo/PTT/PDT). The drug release and metabolic distribution stimulated by NIR laser were monitored in HeLa cells, and the therapeutic efficacy of these NPs in vivo was further evaluated by intravenous injection.


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Scheme 1. The drug release process of ID@TSL-Gd NPs response to NIR laser. (a) The phase transition of the temperature sensitive liposome was triggered by NIR-induced hyperthermia. (b) ID@TSL-Gd entered tumor cells via the endocytic lysosomal pathway for multimodal imaginginduced combination therapy. MATERIALS AND METHODS



Materials and Apparatus. Diethylenetriaminepentaacetic Dianhydride (DTPA), stearylamine (SA), N, N-dimethylformamide (DMF), Gadolinium (III) chloride hexahydrate were purchased from Aladdin. DSPE-PEG-FA (MW:2000), lecithin were purchased from Shanghai Yare Biotech, Inc. Lauric acid, Stearic acid, Indocyanine Green (ICG), Doxorubicin (DOX) and Dimethyl sulfoxide (DMSO) were purchased from Sigma-Aldrich. Absorption spectra were examined by UV3600 UV/vis/NIR spectrophotometer (Shimadzu). A HT7700 transmission electron microscope (TEM) and a particle size analyzer (Brookhaven Instruments) were used to obtain the morphology and size of the NPs, respectively. Synthesis of Gd-DTPA-SA. Gd-DTPA-bis (stearylamide) (Gd-DTPA-SA) was prepared according to the method of previous report.46 Briefly, 0.054 g stearylamine was dissolved in 4 ml chloroform and then slowly added to DTPA dianhydride (0.04 g) in DMF (5 ml). After being stirred at the temperature of 40 °C for 2 h, the solution was cooled at 4 °C for 2 h to obtain white precipitate, which was filtered, washed with acetone, and dried overnight at 80 °C and finally crystallized in heated ethanol (40 ml). After being stirred at room temperature for 24 h, unbound DTPA and SA were removed by filtration with water and chloroform to obtain small crystals (DTPA-SA). To prepare Gd citrate solution, 0.0409 g GdCl3•6H2O was dissolved in 5 ml HCl, followed by addition of 5 ml aqueous solution of sodium citrate (0.0647 g). The Gd citrate solution was adjusted to pH 7.4 with HCl/NaOH, and then slowly added into a hot aqueous solution of DTPASA with a metal: chelating agent 1:1, stirring at room temperature for 1 h under N2, and the product was filtered and dried overnight to obtain about 0.1156 g Gd-DTPA-SA.


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Synthesis of ID@TSL-Gd NPs. The ID@TSL-Gd NPs are prepared by a method derived from previous report.17 Firstly, the PCMs based on two natural fatty acids, lauric acid and stearic acid (4:1 by weight) were dissolved in DMF at a concentration of 4 mg/ml. Then, 5 mg GdDTPA-SA, 5 mg DSPE-PEG2000-FA and 15 mg lecithin was dissolved in 3 ml of 4% ethanol aqueous solution and heated to 60 °C. Next, 2 ml fatty acid solution, 1 ml ICG solution (2.5 mg/ml in DMF) and 1 ml DOX solution (2.5 mg/ml in DMSO) were mixed together and then added dropwise to the preheated phospholipid solution under strong vortex. After homogenization, the mixed solution was cooled in a 4°C ice water bath for 10 minutes accompanied by intense sonication, followed by being heated to ambient temperature and vortexed vigorously for 10 minutes. Subsequently, the obtained liposome solution was washed three times with pure water through an ultrafiltration filter (10 kDa cut off, 10000 rmp) to remove unencapsulated ICG, DOX and organic solvent. Finally, uniform-size ID@TSL-Gd NPs were obtained through filtering using a 0.2 μm surfactant-free cellulose acetate membrane. ICG@TSL-Gd NPs and DOX@TSL-Gd NPs were prepared in the same manner as above, respectively. ID@TSL-Gd NPs without folate modification were prepared in the absence of DSPE-PEG2000-FA by the same method mentioned above. Characterization of ID@TSL-Gd NPs. The prepared NPs size distribution and polydispersity index were measured by dynamic light scattering (DLS) detector (ALV/CSG-3 laser light scattering spectrometers) at 25 °C. At the same time, the size stability of the particles was analyzed within ultrapure water, PBS and fetal bovine serum (FBS) after being stored for 5 weeks at 25 °C. The resultant NPs were negatively stained with 2% phosphotungstic acid before transmission electron microscopy (TEM) characterization to obtain particles’ morphology and size. UV–vis absorption spectra were obtained with a UV3600 UV/vis/NIR spectrophotometer



(Shimadzu). The amount of the DOX and the ICG in ID@TSL-Gd NPs was determined by UV– vis absorption spectra according to standard curve. The encapsulating efficiency (EE) and loading efficiency(LE) of the agents were calculated as follows: Drug loading efficiency (%) = weight of drug in NPs/weight of NPs×100% Drug encapsulating efficiency(%) = weight of drug in NPs/weight of total added drug×100% In Vitro NIR-triggered Drug Release. In vitro DOX release profile was determined by the variety of fluorescence of DOX in water. The release experiment is performed on the 2 mL of ID@TSL-Gd NPs samples (DOX concentration is 30 μg/ml) with and without 808 nm laser irradiation at 0.5 W/cm2 for 100 min. The morphologic changes of NPs with and without laser irradiation were also analyzed by TEM. At the indicated time points, 100 μl solution was taken out for fluorescence spectroscopy measurements using a microplate reader (Ex/Em =488/590 nm). Released DOX were calculated according to the expression (Ft–F0)/(F100–F0) × 100%. Herein, F0 represents the initial DOX FL intensity of ID@TSL-Gd NPs aqueous solution at the start point; Ft represents the DOX FL intensity at indicated time point, and F100 represents the DOX FL intensity of ID@TSL-Gd NPs after complete release. Measurement of Photothermal and Photodynamic Effects in Aqueous Solution. In order to evaluate the photothermal effect, ID@TSL-Gd NPs containing different concentrations of ICG (initial ICG concentration: 40 μg/ml) were exposed to 808 nm laser at 0.5 W/cm2 for 5 minutes respectively. The photothermal stability of the ID@TSL-Gd NPs solution and free ICG solution with the same ICG concentration of 20 μg/ml was explored within 56 minutes, ICG@TSL-Gd NPs, ID@TSL-Gd NPs , free ICG were exposed to 808 nm NIR Laser at 0.5 W/cm2 for 5 minutes, respectively and the changes of temperature were recorded using an IR image camera


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(FLIR E50; Estonia). The photodynamic effect of ID@TSL-Gd NPs was evaluated by using 1,3diphenylisobenzofuran (DPBF) as a selective probe to monitor the singlet oxygen produced by ID@TSL-Gd NPs. The absorption peak of DPBF at 414 nm was recorded at different irradiation time points. The singlet oxygen was monitored based on DPBF according to the previous literature.47,48 In Vitro Cellular Uptake and Target Analysis Mediated by Folate. HeLa cells and NIH3T3 cells were obtained from KeyGEN BioTECH of Nanjing and were cultured in MEM with 10% FBS and maintained at 37 °C with 5 % CO2 for cell, respectively. Hela cells (1×104 cells/well) were seeded in a confocal dish (Lab-Tek, Nunc, USA) for 24 h in dark. Then, the cells were washed with PBS for three times, the medium was replaced by the medium with free DOX (15μg/mL) or ICG (15 μg/mL), or FA modified ID@TSL-Gd NPs (containing 15 μg/mL DOX and 15 μg/mL ICG). After 6 h incubation, the cells treated with FA modified ID@TSL-Gd NPs were irradiated for two minutes. All groups were washed three times with PBS and fixed with 4% paraformaldehyde solution for 20 min and then washed three times with PBS, and finally the fluorescence of the mixed cells were observed using a Zeiss LSM 780 confocal microscope. (DOX: λex/λem = 488 nm/590nm and ICG: λex/λem = 633nm/710 nm). For the target uptake mediated by FA receptor, HeLa cells were seeded in two confocal dishes (1×104 cells/well). Another confocal dish was seeded with NIH-3T3 cells. Upon incubation for 24 h, two dishes containing HeLa cells were washed with PBS and replaced by fresh medium containing FA modified ID@TSL-Gd NPs and medium containing ID@TSL-Gd NPs without FA modification, respectively. The medium of NIH-3T3 cells was replaced by fresh medium containing FA functioned ID@TSL-Gd NPs and FA-free ID@TSL-Gd NPs. After 6 h incubation, fluorescence of DOX (λex =488 nm) was obtained by the confocal microscope. 10 Environment ACS Paragon Plus


NIR-triggered Drug Release in Cells. HeLa cells were seeded in 8-well chambered coverglasses (Lab-Tek, Nunc, USA) with a density of 1×104/well. After being incubated overnight at 37 °C, the medium was replaced by fresh medium containing ID@TSL-Gd NPs (20 μg/mL DOX). After another 6 h incubation, the cells were exposed to 808 nm NIR laser (0.5 W/cm2) for 10 minutes. The cells were washed by PBS three times and observed for DOX FL under the confocal microscope at 0 min, 5 min, 10 min, respectively. Detection of ROS in Vitro. The HeLa cells were seeded in 8-well chambered coverglasses. After 24 h incubation, the medium of each well was replaced with 1 mL fresh culture medium containing PBS, ID@TSL-Gd or ICG@TSL-Gd, respectively. The cells were further incubated for 2 h at 37 °C and 5 % CO2. After being washed three times with PBS, the cells were cultured with 1 mL carboxy-H2DCFDA (25 μM) for 10 min. Subsequently, every well was washed twice with PBS and exposed to 808 nm laser with 0.5 W/cm2 for 10 min for each well. Next, the cells were fixed by 4% paraformaldehyde solution for 10 min. Finally, the cells were washed 3 times with 1mL PBS. The fluorescence of DCF (λex = 495, λem = 529 nm) was immediately captured under the confocal microscope. Cellular Cytotoxicity Analysis. HeLa cells were seeded in 96-well plates. After being incubated for 24 h, the medium was replaced with 150 μL fresh medium containing ID@TSL-Gd NPs or ICG@TSL-Gd. After 6 h incubation, the cells were irradiated with/without 0.5 W/cm2 808 nm laser for 5 min. MTT assay was used to quantify the cell viability. The cell viability of PDT or PTT treatments was obtained by maintaining a constant temperature (25 °C) or treating cells with 100 mM NaN3, a well-known 1O2 scavenger avoiding the production of singlet oxygen during laser irradiation.38 In order to visualize the phototherapeutic efficacy, the cells irradiated with/without the laser were washed three times with PBS and fixed with 4% paraformaldehyde 11 Environment ACS Paragon Plus

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solution. The cells were stained with calcein-AM for live cells and PI for dead cells, followed by observation with biological inverted microscope (Olympus IX71, Japan). Flow cytometry was used to further evaluate the cytotoxicity of the NPs with/without 808 nm laser (0.5 W/cm2). Having been cultured overnight, HeLa cells were then treated with ID@TSLGd, ICG@TSL-Gd or DOX@TSL-Gd with/without laser (808 nm, 0.5 W/cm2, 5 min), respectively, and then incubated for 12 h. The medium containing drugs was immediately removed and washed with PBS. HeLa cells were harvested for cell apoptosis analysis using Alexa Fluor 488 Annexin V/PI Cell Apoptosis Kit and flow cytometry (Flowsight, Merck Millipore, USA). Herein, both Annexin V-positive and PI-negative cells were regarded as apoptotic, and both Annexin V-positive and PI-positive cells were regarded as late apoptotic/ necrotic cells. In Vivo Imaging and Biodistribution Analysis. All of the animal experiments were performed in compliance with the criterions of the National Regulation of China for Care and Use of Laboratory Animals. To set-up the tumor xenograft model, the female nude mice (6 weeks, KeyGEN Biotech.) were subcutaneously inoculated in the right forelimb with 1 x107 HeLa cells. When the tumor volumes approached about 150 mm3, imaging and cancer therapy were conducted. The mice were randomly divided into two groups, 200 μL free ICG or ID@TSL-Gd (both containing 200 μg/mL ICG) were injected via tail vein of mice. The IVIS Lumina K in vivo fluorescence imaging system (PerkinElmer, USA) was used to capture the in vivo fluorescence images and signals of ICG (λex = 633 nm, λem =710 nm) at 1, 6, 12, 24, and 48 h after injection.



The mice were sacrificed at 48 h to collect the heart, liver, spleen, lung, kidneys, and tumor for imaging and semiquatitative biodistribution analysis. Photoacoustic Imaging of FA-PC(IG)-Gd NPs In Vitro. The ID@TSL-Gd NPs solutions with different ICG concentrations (2, 4, 8, 15, 30 μg/ml) were added into the agarose tube as well as the ultrapure water as the control for the determine of PA images and signals using photoacoustic computerized tomography scanner (Endra Nexus 128, Ann Arbor, MI) with an excitation wavelength at 808 nm. Photoacoustic Imaging of ID@TSL-Gd NPs In Vivo. Female nude mice with subcutaneous HeLa cells in right forelimb were used as the animal model. The PA imaging experiment was performed when the tumor size reached ~150 mm3, PA imaging was obtained by the preclinical photoacoustic computerized tomography scanner (Endra Nexus 128, Ann Arbor, MI) after the injection of ID@TSL-Gd NPs ([ICG]=30 μg/ml, 200 μL) at indicated time points (Ex=808 nm). Magnetic Resonance Imaging In Vitro. The ID@TSL-Gd NPs with different Gd concentrations (0.0125～0.2 mM) were added into the agarose tube for the determination of MR images and T1 value , the ultrapure water was used as the control. Magnetic Resonance Imaging In Vivo. When the tumor of female nude mice with Hela cells located in a right forelimb reached ~150 mm3, after intravenously injected with ID@TSL-Gd (Gd concentration 0.5 mM), magnetic resonance imaging was performed using a 3.0 T clinical MR scanner (GE Healthcare, USA) equipped with a small animal imaging coil. In Vivo Combination Therapy of Chemotherapy and Phototherapy. HeLa tumor bearing mice randomly divided into 5 groups (5 per group) were intravenously injected with 200 μL


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PBS, PBS + laser, ID@TSL-Gd, ID@TSL-Gd + laser and ICG@TSL-Gd + laser, respectively. Thermal imaging was recorded by an infrared thermal imaging camera (E50, FLIR, USA). The changes of tumor volumes and mice body weight were recorded every two days. The mice were sacrificed by cervical dislocation under an anesthetic status after 14 days’ treatment and the major organs of mice were collected and stained with H&E for further analysis. RESULTS AND DISCUSSION Synthesis and Characterization of ID@TSL-Gd NPs. The synthesis protocol of ID@TSLGd NPs was shown in Scheme 1, which illustrated the anti-tumor therapeutic mechanisms guided by multimodal imaging including PAI, MRI and FL imaging. The ID@TSL-Gd NPs were fabricated according to an approach deriving from nanoprecipitation,40 similar to that of the previous report.17 Briefly, the eutectic mixture of lauric acid and stearic acid at a weight ratio of 4:1, was mixed with ICG and DOX, and then was added dropwise to the pre-heated phospholipid solution composed of DSPE-PEG-FA, lecithin and Gd-DTPA-SA (T1 magnetic resonance contrast agent). After that, the ID@TSL-Gd NPs were formed by the self-assembly process. Followed by vigorous vortex, the suspension was sonicated in a 4 °C water bath for 10 minutes to rapidly solidify the fatty acid eutectic and promote the encapsulation of phototherapy and chemotherapy drugs. Upon NIR irradiation, the NPs will be heated by the photothermal effect of the preloaded ICG. When the temperature of NPs is above the melting point of the fatty acid eutectic (39 °C), the NPs will melt and release encapsulated ICG and DOX quickly. The TEM image of ID@TSL-Gd NPs demonstrated that the NPs are monodisperse, spherical, and uniform in size (Figure 1a). According to the DLS measurement, the average particle size was approximately 95 nm, which would be easier to penetrate into deeper tissue and tumor areas



as well a more likely to be internalized by tumor cells through the enhanced permeability and retention effect (EPR).41 The color of ID@TSL-Gd NPs dissolved in ultrapure water, PBS and FBS all showed clear green (Figure S2a). The size of their particles was further evaluated within 5 weeks (Figure 1b), and the particle diameter of ID@TSL-Gd in FBS (about 97 nm) is slightly larger than that in PBS or in water (95 nm), which can be ascribed to the binding proteins. These results indicate that the ID@TSL-Gd NPs kept long-term stability without precipitation or phase separation in three solvents, which demonstrated good biostability. Compared with free ICG, the absorption peak of the ICG encapsulated in ID@TSL-Gd NPs shift from 780 nm to 800 nm (Figure 1c). Compared with ICG@ TSL-Gd, ID@TSL-Gd showed a more strong absorption at 808 nm, which can be ascribe to the aggregation of ICG induced by the addition of DOX, and thus will increase the photothermal effect of ID@TSL-Gd. The red shift of the absorption peak can be ascribed to the accumulation of ICG, which is more conducive to NIR light mediated photothermal and photodynamic therapy. According to the intensity of characteristic absorption peak in Figure 1c, the calculated encapsulation efficiency of ICG and DOX was 52.3±2.3% and 50.5±3.4%, respectively. Furthermore, the loading efficiency was calculated to be 3.6±0.5% and 3.3±0.6%, respectively, which were higher than the previously reported liposomes (Table S1).17 The high EE and LE may be due to the strong hydrophobic interaction between ICG and lipophilic fatty acid. At room temperature or physiological temperature (37 °C, below the eutectic point), the NPs exist in solid state, preventing the enveloped ICG and DOX from leaking out. The liposome melted and expanded quickly due to the thermal effect of ICG (Figure S1), and the particle size increased from 95 nm to 150 nm (Figure 1d). In order to evaluate the photothermal effect of ID@TSL-Gd NPs, we used an infrared thermal imaging camera to monitor the temperature changes of ID@TSL-Gd aqueous solution with different ICG


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concentrations under the 808 nm laser at 0.5 W/cm2 within 5 min. In case of 5, 10, 20 and 40 μg/ml of ID@TSL-Gd, after 5 min laser irradiation, the temperatures reached 45.7, 53.5, 58.5 and 61.2 °C, respectively (Figure 1e). These temperatures were all higher than that required for irreversible apoptosis of tumor cells (43 °C),42 melting the fatty acid eutectic and triggering drug release (39 °C). The ICG concentration in the ID@TSL-Gd aqueous solution was set at 20 μg/mL, and the laser was turned off after 5 min irradiation at 0.5 W/cm2, and the temperature of the solution gradually decreased from 58.5 °C to room temperature (28 °C). The infrared thermal images were shown in Figure 1f. After four cycles, the maximum temperature of aqueous solution of the NPs (containing 50 μg/mL ICG) still could reach 67 °C, which was much higher than that of free ICG with the same concentration of ICG, after four cycles (Figure S2b). It can be ascribed to the restriction of photobleaching for NPs under the protection of the lipid layer and fatty acid eutectic The photothermal effects of ID@TSL-Gd, ICG@TSL-Gd, and free ICG with the same ICG concentration (20 μg/mL) were also monitored. The temperature of ID@TSL-Gd and ICG@TSL-Gd increased rapidly to 57 and 55.7 °C within 5 min, while the temperature of free ICG and PBS (control) increased by only 16.5 and 3.8 °C (Figure 1g). The enhanced photothermal effect of ID@TSL-Gd NPs can be ascribed to the red shift of the absorption peak of the ICG encapsulated in the NPs. Additionally, the NPs provide a high local concentration of ICG, and a natural barrier that avoids the heat loss of the heated ICG which enhances the photothermal effect. The photothermal conversion efficiency of ID@TSL-Gd was further investigated according to the previously reported literature,49 and was determined to be ~9.181%, which is much higher than that of free ICG(~3.37%).36 Under the NIR laser irradiation, the hyperthermia of the heated photothermal dye ICG melted fatty acids (melting point of the eutectic is 39 °C), leading to quick release of chemotherapeutic drug-DOX and phototherapeutic



drug-ICG. The NPs were irradiated in an on-off formation within 100 minutes, resulting in a final DOX release of 55%, far higher than that without NIR irradiation (Figure 1h and S3a). The drug release behavior was been detected at room temperature and high temperature, respectively, to confirm the thermal sensitivity of natural phase change material. At 45 °C, the release rate of DOX could also reach 48.9% within 100 minutes, indicating the high sensitivity of fatty acids as phase change materials (Figure S3b). As shown in Figure S5, the single oxygen probe DPBF showed a relatively high degradation rate in the presence of ID@TSL-Gd under 808 nm laser (0.5W/cm2), The 1O2 generation efficiency (ΦΔ) of the NPs was calculated to be 0.294% by using free ICG (ΦΔ = 0.2%) as a reference,50-51 the higher 1O2 generation efficiency can be ascribe to the improved photostability of ID@TSL-Gd compared with free ICG. The results indicated that the photothermal agent ICG could not only be a good photothermal agent, but also a good photosensitizer which can produce ROS under NIR. Both could be used in tumor eradication.


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Figure 1. Characterization of ID@TSL-Gd NPs.(a) Diameter distribution and TEM image of ID@TSL-Gd NPs before the laser irradiation(Scale bar = 500 nm). (b) Size stability of ID@TSL-Gd NPs in ultrapure water, PBS and FBS within 5 weeks. (c) UV-vis absorption spectra of free ICG,free DOX, ICG@TSL-Gd and ID@TSL-Gd. (d) The diameter distribution of ID@TSL-Gd NPs after laser irradiation (808 nm laser, 0.5W/cm2). (e) Temperature rise profile of ID@TSL-Gd with different ICG concentrations(ranging from 40 to 2.5 μg/ml) within 300s under 808nm continuous laser irradiation at 0.5W/cm2. (f) Infrared thermographic maps of ID@TSL-Gd NPs at different time points under NIR laser radiation (NIR radiation for 5 min) ,recorded by an infrared thermal imaging camera. (g) Temperature rise profile of ID@TSL-Gd, 18 Environment ACS Paragon Plus


ICG@TSL-Gd, free ICG, all containing 20 μg/ml ICG, and PBS within 300s under 808 nm continuous laser irradiation at 0.5W/cm2. (h) Cumulative DOX release from ID@TSL-Gd NPs with or without 808-nm NIR irradiation (0.5W/cm2). In Vitro Cellular Uptake and Target Analysis Mediated by Folate Receptor. After 6 h of incubation in HeLa cells, the subcellular localizations of free DOX, free ICG, ID@TSL-Gd NPs and ID@TSL-Gd NPs+2 min laser irradiation were observed by confocal microscopy (Figure 2a). In case of free DOX and free ICG, the weak fluorescence intensity of DOX and ICG indicated that only a little amount of DOX or ICG was distributed in the cytoplasm due to the lack of targeting ability. However, in case of ID@TSL-Gd NPs, slightly stronger fluorescence indicated that more DOX or ICG entered the cells, owing to the targeting ability of folate. Moreover, much stronger fluorescence was found in the ID@TSL-Gd NPs+2 min laser irradiation group, which can be ascribed to the quick release of DOX or ICG and distribution in the cytoplasm under 808 nm laser irradiation. It indicated that the FA-modified liposome encapsulation could enhance the cellular uptake efficiency of the NPs, and the NIR laser irradiation could control the release of drugs. In absence of laser irradiation, ID@TSL-Gd NPs were mainly dispersed in the cytoplasm, and no fluorescence response was found in the nucleus, since the NPs prevent ICG and DOX from leaking out (Figure 2a). To verify the specific targeting ability of ID@TSL-Gd NPs, NIH-3T3 cells (a kind of normal cells) and HeLa cells were employed, in which the folate receptors were overexpressed on the surface of HeLa cells. Both HeLa and NIH-3T3 cells were co-cultured with FA functionalized ID@TSL-Gd, and the targeting analysis of ID@TSL-Gd NPs was obtained from the fluorescence intensity of DOX within cells. After 12 h of incubation, strong fluorescence of DOX was observed in HeLa cells treated with FA functionalized ID@TSL-Gd (Figure 2b, i), 19 Environment ACS Paragon Plus

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indicating that FA functioned ID@TSL-Gd NPs are more distributed in the cytoplasm of HeLa, because the overexpressed folate receptors could capture the folate modified NPs and guide them to enter tumor cells by endocytic lysosome. Furthermore, FA-free ID@TSL-Gd NPs were applied to incubate with HeLa cells. However, the fluorescence responses of DOX were quite weak (Figure 2b, ii), just like the normal cells (Figure 2b, iii and iiii). All of the above results indicate that folic acid plays a vital role during the orientation and uptake of NPs into tumor cells, which is consistent with the previous report.42,43

Figure 2. Endocytosis and uptake of ID@TSL-Gd NPs in HeLa cells before and after laser irradiation. (a) Subcellular localization of free DOX, free ICG, ID@TSL-Gd NPs and ID@TSLGd+2 min laser irradiation, after 6 h incubation in HeLa cells. Green fluorescence represents ICG and red fluorescence represents DOX. (b) Fluorescence images of HeLa cells incubated with ID@TSL-Gd NPs (i and ii, with and without folate modification), and fluorescence images of NIH 3T3 cells incubated with ID@TSL-Gd NPs for 6 h (iii and iiii, with and without folate



modification). The images were obtained at a NIR channel 550-610 nm with an excitation at 488 nm, all scale bars: 20 μm. NIR-Triggered Drug Release in Cells. The intracellular drug release under continuous laser irradiation was monitored by confocal microscopy. Before laser irradiation, NPs were mainly restricted to endolysosomes in cytoplasm, and almost no DOX entered the nucleus, as shown by the typical punctate distribution (Figure 3a). After NIR laser radiation for 5 min, DOX was released from the NPs after the melting of fatty acids eutectic, followed by slow diffusion from the lysosomes into the cytoplasm, and accumulation around the nucleus, as demonstrated by blurred and diverged fluorescence of DOX in HeLa cells. After 10 min NIR radiation, DOX was released continuously from NPs and dispersed homogenously in the cytoplasm, and then diffused into the nucleus (Figure 3a). Thus, as an external stimulus, NIR plays an important role to control the release of anticancer drugs in cells. The intracellular on-demand drug release is critical to increase therapy efficiency and reduce the side-effects, as well as suppress drug resistance.


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Figure 3. (a) Cumulative DOX release of ID@TSL-Gd NPs at the different time points in response to the continuous 808 nm laser at a power density of 0.5 W/cm2, scale bar: 30 μm. (b) Fluorescent changes of DCFH activated in HeLa cells incubated with PBS, free ICG, ICG@TSL-Gd, respectively, under 808 nm laser irradiation at a power density of 0.5 W/cm2, scale bar: 150 μm. Detection of ROS In Vitro. To evaluate the photodynamic effect of ID@TSL-Gd NPs in vitro, the reactive oxygen species (ROS, typically 1O2) in living cells were monitored using carboxy-H2DCFDA as a fluorescent indicator. This nonfluorescent molecule is readily converted to a green fluorescent carboxy-DCF when the acetate groups are removed by intracellular esterases and oxidation by singlet oxygen (1O2), enabling living cells to exhibit bright green fluorescence. HeLa cells were incubated with PBS, free ICG, and ICG@TSL-Gd, respectively, followed by addition of carboxy-H2DCFDA. After being continuously irradiated for 8 min at 808



nm laser (0.5W/cm2), the samples were tested at different irradiation time points (Figure 3b). The cells treated with PBS solution exhibited negligible fluorescence after the irradiation. However, cells treated with free ICG or ICG@TSL-Gd NPs showed bright green fluorescence after irradiation of 8 min. The fluorescence of the cells treated with ICG@TSL-Gd NPs was brighter than that of the cells treated with free ICG, which can be ascribed to the good targeting ability of the ICG@TSL-Gd NPs. Cellular Cytotoxicity Analysis. To evaluate the cytotoxicity of ID@TSL-Gd NPs and the effect of NIR triggered drug release, HeLa cells were incubated with ID@TSL-Gd NPs and ICG@TSL-Gd NPs for 6 h, and then irradiated with or without the 808 nm laser at 0.5 W/cm2, respectively. The obtained cell viability assay was shown in Figure 4a. The survival rate of HeLa cells treated with ICG@TSL-Gd NPs or ID@TSL-Gd NPs under laser irradiation were both gradually reduced with the increase of ICG concentration, and finally reached 15.5 % and 6.58 %, respectively. According to the results, chemo-photothermal-photodynamic combination therapy resulted in a greatly enhanced therapeutic efficiency compared with PTT, PDT or PTTPDT (Figure S6c and d). However, the survival rate of HeLa cells treated with ID@TSL-Gd NPs without the laser irradiation did not change significantly (90.2 %). That was because the DOX encapsulated within the NPs could not be released to induce tumor apoptosis without the hyperthermia generated by ICG under NIR irradiation. The dose dependent cytotoxicity of ID@TSL-Gd NPs was further evaluated with/without laser irradiation (Figure S6b). Under irradiation, the cell survival rate of HeLa was only 9 % at high concentration of ICG and DOX. However, there is no apparent dark cell toxicity for HeLa without laser irradiation, possibly due to the slight drug release from NPs. Apart from cancer cells, the dark cytotoxicity of ID@TSLGd NPs for normal cell of NIH-3T3 cells was also conducted. As shown in Figure S6a, the cell


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viability was over 80 % during the whole process, indicating that ID@TSL-Gd NPs did not have dark cell toxicity to 3T3 cells as NPs concentrations increased. To visualize the effects of drug release and phototherapy excited by NIR light, the staining experiments for living/dead cells were conducted. The cells were co-stained by calcein AM (living cells, green fluorescence) and propidium iodide (PI, dead/late apoptotic cells, red fluorescence), and then monitored with an inverted microscope. As shown in Figure 4b, the viability of cells treated with ID@TSL-Gd NPs with laser was lower than that of cells treated with ICG@TSL-Gd NPs under the same condition, owing to the combination therapy of chemotherapy, PTT and PDT. The therapeutic efficacy of PTT, PDT or PTT-PDT were also investigated. The fluorescence responses of the living/dead cells were shown in Figure S6d, which is consistent with the result of the MTT assay (Figure S6c). In the absence of NIR laser excitation, neither ID@TSL-Gd NPs nor DOX@TSL-Gd NPs exhibited apparent cytotoxicity. These results were confirmed by quantitative analysis by flow cytometry (Figure 4c). In comparison, the cells treated with ICG@TSL-Gd NPs and ID@TSL-Gd NPs with NIR laser irradiation both demonstrated a considerable increase ratio of apoptosis (89.1% and 94.5% for ICG@TSL-Gd NPs and ID@TSL-Gd NPs, respectively).



Figure 4. Increased cytotoxicity induced by NIR triggered drug release. (a) Quantitative detection of cell viability of HeLa cells incubated with ID@TSL-Gd, ICG@TSL-Gd with/without 808 nm radiation at 0.5 W/cm2 for 5 min. (**) p < 0.01, (***) p < 0.001. (b) Fluorescence images of Calcein AM/PI co-stained HeLa cells incubated with ID@TSL-Gd NPs, ICG@TSL-Gd NPs and DOX@TSL-Gd NPs, respectively, with/without 808 nm NIR radiation at 0.5W/cm2 for 5 min. Scale bar: 200 μm. (c) Flow cytometry analysis of HeLa cells after different treatments. Photoacoustic Imaging of ID@TSL-Gd NPs in Vitro. The photoacoustic properties of resultant ID@TSL-Gd solutions with different ICG concentrations (ranging from 2 to 30 μg/mL) were evaluated using a Nexus 128 PA imaging system. The obtained photoacoustic signal intensities were shown in Figure S7a. The photoacoustic imaging is mainly attributed to thermal


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expansion,44 where ICG absorbs laser pulses from the photoacoustic system, causing a part of the solution to be heated, generating rapid expansion of acoustic pressure waves. Quantitative analysis confirmed that the PA signal of the solution is linear with the concentration of ICG encapsulated in the NPs (Figure S7a), in which the PA signal was enhanced with the increase of ICG concentration. Magnetic Resonance Imaging in Vitro. The outermost layer of gadolinium-III (Gd3+) has 7 unpaired electrons, whose number of lone electrons is the highest among other elements, and its electron induction is the strongest, resulting in longer relaxation time. Thus, along with relatively low toxicity, it is a paramagnetic contrast agent which is currently widely used in clinic.45 GdDTPA-SA is an amphiphilic gadolinium complex prepared by attaching a polar gadolinium head to a long hydrophobic tail, whose structure and function are similar to phospholipid complexes and can be combined with phospholipids to form the shell of a drug-loaded liposome for detection of T1 relaxation time. The enhanced MR signals with the increase of the gadolinium concentration in the liposome were shown in Figure S7b. Quantitative analysis of the MR signals (Figure S7b) showed that the relaxation rate (1/T1) increased with probe concentration, and the r1 relaxivity of ID@TSL-Gd NPs was measured to be 31.25 mM-1S-1, much higher than that of the clinically approved T1-weighted contrast agent, Magnevist (4.29 mM-1s-1). The higher relaxivity of NPs can be ascribed to the strong hydrophobic interaction between DSPE-PEG2000-FA and Gd-DTPA-SA during the formation of liposomal shell that makes the part of Gd-DTPA-SA harder to rotate, resulting in the decrease of the rotational correlation time for the chelated part of the Gd3+, and ultimately increases the relaxivity.



Figure 5. In vivo images and biodistribution of HeLa tumor-bearing nude mice after intravenous injection with free ICG or ID@TSL-Gd NPs, respectively. (a) In vivo FL images of nude mice after the injection of ICG and ID@TSL-Gd NPs at the designed time points. (b) Quantified FL intensity of ICG around the tumor at different time points after intravenous injection. (c) Ex vivo FL images of major organs and tumors after injection of free ICG or ID@TSL-Gd NPs after 48 h. (d) Quantitative biodistribution of free ICG and ID@TSL-Gd NPs confirmed by the FL intensity of major organs and tumors after injection at 48 h. (*) p < 0.05, (**) p < 0.01. In Vivo Imaging and Biodistribution. It is well known that xenografted tumors are more invasive than in situ carcinoma, and possess a denser capillary network, which facilitates the enrichment of nanomedicine around the tumor by the enhanced permeability and retention effect


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(EPR). Hence, HeLa tumor-bearing nude mice were selected as tumor models to be injected with free ICG, ID@TSL-Gd NPs by tail vein, respectively, and the biodistribution of the drug was directly observed by FL imaging at different time points after injection.53 As shown in Figure 5a, after 6 h of intravenous injection, ID@TSL-Gd NPs started to accumulate at the tumor site and considerable fluorescence signals can be observed from the tumor. The NPs continuously accumulated at the tumor site, and reached the peak at 12 h postinjection, confirmed by the gradually increased FL intensity of ICG. The quantified FL intensities of ICG around the tumor were shown in Figure 5b at different time points. The signal to noise ratio of FL intensities reached to the maximum at 12 h. Due to the unique EPR effect and the auxiliary targeting property of ID@TSL-Gd NPs, the ID@TSL-Gd NPs treated mice still had strong fluorescence signals at the tumor site after 48 h. On the contrary, in the case of free ICG treated mice, the FL signal was distributed extensively in the liver after intravenous injection, in which the FL signal was much stronger than that in other organs. With elapsing of time, much of ICG was excreted from the body through the metabolic mechanism, and the FL signal at the tumor site became weaker than that in mice treated with ID@TSL-Gd NPs. The ex vivo FL images of major organs obtained at 48 h postinjection were shown in Figure 5b, in which bright fluorescence can be observed at the tumor of mice treated with ID@TSL-Gd NPs. However, free ICG tended to be distributed in liver, tumor, kidneys, lung and spleen (Figure 5c). The uptake of ICG by the liver and spleen may be attributed to reticuloendothelial system absorption, while the kidney uptake can be associated with possible renal excretion. According to the results, accumulation of ICG in the tumor site was considerably enhanced by ID@TSL-Gd NPs. As shown in Figure 5d, the FL intensity of tumor treated with ID@TSL-Gd was approximately twice the scale of mice treated with free 28 Environment ACS Paragon Plus


ICG, and much higher than the fluorescence intensity of other organs, indicating that NPs are ideal diagnostic reagents with tumor targeting, which avoid the damage of other healthy tissues.

Figure 6. Photoacoustic imaging of tumors in vivo. (a) Photoacoustic images of tumors of nude mice bearing HeLa cells at the indicated time points after intravenous injection with ID@TSLGd NPs. (b) Quantitative analysis of the photoacoustic intensity of the tumor corresponding to the images in (a). Photoacoustic Imaging of ID@TSL-Gd NPs in Vivo. In view of the remarkable PA performance of ID@TSL-Gd NPs, tumor tissue has a rich vascular network that can be observed through PA signals for further diagnosis and treatment. Hence, ID@TSL-Gd NPs solution containing 30 μg/mL ICG was injected into tail vein of the HeLa tumor-bearing nude mice. The PA signals of NPs in the tumor were monitored at different time points. As shown in Figure 6a, after intravenous injection, the PA intensity in the tumor tissue was gradually enhanced over time. At 4 h post-injection, ID@TSL-Gd NPs treated tumor exhibited a higher PA signal, allowing visualization of tumor microvessels with high contrast and resolution. The PA signal


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reached the maximum 12 h after injection, which was consistent with the result of the strongest FL signal at 12 h post-injection in FL imaging. These results indicated a longer blood circulation time of ID@TSL-Gd NPs, which may be due to the uniform diameter of small particle and good biocompatibility of the NPs that avoid the removal from the body by the liver. Quantitative analysis of PA signals at tumor sites revealed that the photoacoustic intensity of tumor tissues were enhanced by 1.87, 2.79, 3.65, and 4.05 folds at 2, 4, 8, and 12 h postinjection, respectively (Figure 6b). Even after 48 h, the signal value still remained 2.74 times higher than that at 0 h, which fully demonstrated the prolonged blood circulation time of ID@TSL-Gd NPs.

Figure 7. In vivo MR imaging. (a) T1-weighted MR images of nude mice bearing HeLa cells at different time after injection of ID@TSL-Gd NPs. (b) Quantitative analysis of the average T1MR signal in the tumor corresponding to the image of (a). Magnetic Resonance Imaging in Vivo. Based on the high r1 value of Gd3+, HeLa tumorbearing nude mice were injected with 0.5 mM ID@TSL-Gd solution by tail vein, followed by monitoring of tumor sites under a 3.0 T clinical MR scanner. Figure 7a showed cross-sectional



MR images of the mice at different time points after injection. The MR signal at the tumor site began to enhance, and the tumor did not appear to be significantly brightened 6 h after injection. At 12 h post-injection, the MR signal of the tumor reached a peak, about 4 folds that before injection (Figure 7b), which was consistent with the FL and the PA imaging. The MR signal then started to gradually reduce after 12 h. The MR signals of the coronal section of mice in different time points were shown in Figure S7c with time elapse. At 12 h post-injection, the MR signal at the tumor location reached the maximum, consistent with the results at the cross section. In view of the above results, the biodistribution of NPs and tumor in vivo could be presented by high sensitive FL imaging, and the tumor was located legibly by PA and MR imaging methods with high resolution and deep tissue penetration. Moreover, according to the signal intensity of FL, PA and MR, the best laser irradiation time could be confirmed at 12 h postinjection. Therefore, via this triple modal FL/PA/MR imaging technique, the tumor could be better orientated, and further treated with imaging guided tumor therapy. In Vivo Combination Therapy of Chemotherapy and Phototherapy. We constructed in vivo treatment on HeLa tumor-bearing nude mice based on the effective in vitro cell-based combination therapy and in vivo multimodal imaging assisted drug accumulation for the tumor therapy. HeLa tumor-bearing nude mice were divided into five groups and given the following treatments: (PBS, PBS+laser, ID@TSL-Gd, ICG@TSL-Gd+laser, and ID@TSL-Gd+laser). The mice were irradiated with 808 nm laser (0.8 W/cm 2) for 10 minutes 12 h after injection, and the surface temperature of the tumor was recorded by an infrared thermal camera (Figure 8a). As shown in Figure 8b, under NIR laser irradiation, the tumor surface temperature of mice injected with ID@TSL-Gd NPs rapidly increased to 61 °C within 3 min, providing sufficient hyperthermia for rapid release of DOX and ICG from NPs accompanied by significant PTT52 and 31 Environment ACS Paragon Plus

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PDT. The surface temperature of the tumor injected with ICG@TSL-Gd NPs could eventually reach 56.6 °C, which was also much higher than the apoptosis temperature of cancer cells. However, the temperature of the mice treated with PBS changed little, indicating that the NIR laser irradiation alone could not afford enough hyperthermia for tumor suppression.



Figure 8. In vivo combination treatment of HeLa tumor-bearing nude mice model. (a) The thermal images of the tumor under 808 nm laser irradiation (0.8 W/cm2) for 10 min. (b) Temperature profile of tumors surface at 808 nm laser irradiation (0.8 W/cm2) for 10 min. (c) The changes of tumor volume of HeLa tumor-bearing nude mice in the different treatment groups within 14 days. (*) p < 0.05, (**) p < 0.01. (d) The changes of body weight of mice in the different treatment groups within 14 days. (e) Representative photos of tumor anatomy after the treatment.

Figure 9. H&E staining images of the major organs from mice with different treatments after 14 days, scale bar: 50 μm. To further investigate the therapeutic effect, the changes in tumor volume and mice body weight were recorded every two days after treatment (Figure 8c, d). The tumors treated with


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PBS, PBS+Laser and ID@TSL-Gd grew gradually over time, indicating that HeLa cells could not be suppressed by ID@TSL-Gd and laser radiation, alone. Due to the phototherapy property of ICG under NIR laser irradiation, the growth of tumor treated with ICG@TSL-Gd+Laser was inhibited. In the case of ID@TSL-Gd+Laser treatment, the tumor volume gradually reduced over time, which indicated that tumor could be suppressed significantly via the participation of chemotherapy. Figure S9 showed the digital photos of the tumor growth over time. The tumor sizes of mice treated with ID@TLS-Gd NPs after laser irradiation were the smallest among other comparison groups (Figure 8e), and the specific values were shown in Figure S8a. In order to investigate the long-term toxicity of NPs, the mice were sacrificed after 14 days treatment to examine potential side effects on major organs. The major organs of the mice showed no appreciable abnormality according to H&E staining analysis (Figure 9), demonstrating that all treatments were biocompatible and safe to mice, which was consistent with the detection of cytotoxicity for normal cells in vitro. The tumor of mice treated with ID@TLS-Gd NPs and ICG@TSL-Gd under laser irradiation showed an irreversible damage to tumors according to the H&E staining of the tumors (Figure S8b). CONCLUSION In summary, multifunctional tumor targeting ID@TSL-Gd NPs were successfully prepared by co-encapsulating ICG and DOX into biocompatible eutectic mixture of natural fatty acid. Under the NIR laser irradiation, the resultant NPs achieved significant on-demand release of the payloads based on the phase transition of eutectic of nature fatty acids induced by hyperthermia of ICG. With the modification of the targeting agent FA and the conjugated gadolinium chelates, the ID@TSL-Gd NPs provides an effective approach to precise diagnosis of cancer through multimodal imaging (FL/PAI/MRI). At the same time, it achieved an outstanding effect of 34 Environment ACS Paragon Plus


cancer combination therapy (Chemo/PTT/PDT) guided by the precise orientation. This work provides a promising design strategy for research and development of multifunctional theranostic NPs for imaging guided synergetic cancer therapy. ASSOCIATED CONTENT Supporting Information. Thermal cycle curve of ID@TSL-Gd and free ICG aqueous solution, photodynamic effect of ID@ TSL-Gd aqueous solution, cell viability of NIH 3T3 cells treated with ID@TSL-Gd and PDT or PTT or PDT/PTT effect of ID@TSL-Gd NPs for HeLa cells, photoacoustic imaging and magnetic resonance imaging of ID@TSL-Gd NPs in vitro, H&E staining images of the tumors and representative growth photographs from different groups of mice. AUTHOR INFORMATION Corresponding Authors *E-mail: [email protected] (Q.M. Shen). *E-mail: [email protected] (Q. Wang). *E-mail: [email protected] (Q.L. Fan). Funding This work was financially supported by National Natural Science Foundation of China (21575069, 21602112 and 21674048), Ministry of Education of China (No. IRT1148), 333 project of Jiangsu province (BRA2016379), the program for Jiangsu Collaborative Innovation


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Center of Organic Electronics and Information Displays and a project funded by the Priority Academic Program Development of the Jiangsu Higher Education Institutions (BE2016770). Notes The authors declare no competing financial interest. ACKNOWLEDGMENTS We thank Prof. Deju Ye from the State Key Laboratory of Analytical Chemistry for Life Science, Nanjing University, for the MRI observation and analysis of the data. References (1) Zhao, P. F.; Zheng, M. B.; Luo, Z. Y.; Gong, P.; Gao, G. H.; Sheng, Z. H.; Zheng, C. F.; Ma, Y. F.; Cai, L. T. NIR-Driven Smart Theranostic Nanomedicine for on-Demand Drug Release and Synergistic Antitumour Therapy. Sci. Rep. 2015, 5, 14258. (2) Liu, D.; Yang, F.; Xiong, F.; Gu, N. The Smart Drug Delivery System and Its Clinical Potential. Theranostics 2016, 6, 1306-1323. (3) Blum, A. P.; Kammeyer, J. K.; Rush, A. M.; Callmann, C. E.; Hahn, M. E.; Gianneschi, N. C. Stimuli-Responsive Nanomaterials for Biomedical Applications. J. Am. Chem. Soc. 2015, 137, 2140-2154. (4) Meng, Z. Q.; Wei, F.; Wang, R. H.; Xia, M. G.; Chen, Z. G.; Wang, H. P.; Zhu, M. F. NIRLaser-Switched in Vivo Smart Nanocapsules for Synergic Photothermal and Chemotherapy of Tumors. Adv. Mater. 2016, 28, 245-253. (5) Epstein-Barash, H.; Orbey, G.; Polat, B. E.; Ewoldt, R. H.; Feshitan, J.; Langer, R.; Borden, M. A.; Kohane, D. S. A Microcomposite Hydrogel for Repeated on-Demand UltrasoundTriggered Drug Delivery. Biomaterials 2010, 31, 5208-5217. (6) Zhang, Y. Q.; Yu, J. C.; Bomba, H. N.; Zhu, Y.; Gu, Z. Mechanical Force-Triggered Drug Delivery. Chem. Rev. 2016, 116, 12536-12563. (7) Al-Ahmady, Z. S.; Al-Jamal, W. T.; Bossche, J. V.; Bui, T. T.; Drake, A. F.; Mason, A. J.;



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For Table of Contents (TOC) Only:



Figure 1 160x136mm (300 x 300 DPI)


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Figure2 160x93mm (300 x 300 DPI)



Figure 3 160x94mm (300 x 300 DPI)


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Figure 4 160x106mm (300 x 300 DPI)



Figure 5 160x100mm (300 x 300 DPI)


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Figure 6 160x75mm (300 x 300 DPI)



Figure 7 160x75mm (300 x 300 DPI)


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Figure 8 160x170mm (300 x 300 DPI)



Figure 9 160x115mm (300 x 300 DPI)


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Scheme 1 160x173mm (300 x 300 DPI)



TOC 80x41mm (300 x 300 DPI)


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Multifunctional Thermosensitive Liposomes Based on Natural Phase

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