Engineering Phototheranostic Nanoscale MetalâOrganic Frameworks

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Engineering Photo-Theranostic Nanoscale Metal-Organic Frameworks for Multi-Modal Imaging-Guided Cancer Therapy Wen Cai, Haiyan Gao, Chengchao Chu, Xiaoyong Wang, Junqing Wang, Pengfei Zhang, Gan Lin, Wengang Li, Gang Liu, and Xiaoyuan Chen ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.6b11579 • Publication Date (Web): 29 Dec 2016 Downloaded from http://pubs.acs.org on December 30, 2016

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ACS Applied Materials & Interfaces

Engineering

Photo-Theranostic

Nanoscale

Metal-Organic

Frameworks for Multi-Modal Imaging-Guided Cancer Therapy Wen Cai,

†, ‡, ǁ

Haiyan Gao, ‡

‡, ǁ

‡

‡

‡

Pengfei Zhang, Gan Lin, Wengang Li, Gang Liu †

‡

Chengchao Chu, Xiaoyong Wang, Junqing Wang, *, ‡

‡

§

and Xiaoyuan Chen

Institute of Medical Engineering, School of Basic Medical Sciences, Xi'an Jiaotong University

Health Science Center, Xi'an, Shaanxi 710061, China ‡

State Key Laboratory of Molecular Vaccinology and Molecular Diagnostics Center for Molecular

Imaging and Translational Medicine, School of Public Health, Xiamen University, Xiamen, Fujian 361102, China §

Laboratory of Molecular Imaging and Nanomedicine (LOMIN), National Institute of Biomedical

Imaging and Bioengineering (NIBIB), National Institutes of Health (NIH), Bethesda, Maryland 20892, United States

ABSTRACT: Many photo-responsive dyes have been utilized as imaging and photodynamic/photothermal therapy agents. Indocyanine green (ICG) is the only near-infrared region (NIR) organic dye for clinical applications approved by the United States Food and Drug Administration; however, the clinical application of ICG is limited by its poor aqueous solubility, low cancer specificity, and low sensitivity in cancer theranostics. To overcome these issues, a multifunctional nanoplatform based on hyaluronic acid (HA) and ICG-engineered metal-organic framework MIL-100(Fe) nanoparticles

(MOF@HA@ICG

NPs)

was

successfully

developed

for

imaging-guided, anti-cancer photothermal therapy (PTT). The synthesized NPs showed a high loading content of ICG (40%), strong NIR absorbance, and 1

ACS Paragon Plus Environment


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photostability. The in vitro and in vivo imaging showed that the MOF@HA@ICG NPs exhibited greater cellular uptake in CD44-positive MCF-7 cells and enhanced tumor accumulation in xenograft tumors due to their targeting capability, compared to MOF@ICG NPs (non-HA-targeted) and free ICG. The in vitro photothermal toxicity and in vivo PTT treatments demonstrated that MOF@HA@ICG NPs could effectively inhibit the growth of MCF-7 cells/xenograft tumors. These results suggest that MOF@HA@ICG NPs could be served as a new promising theranostic nanoplatform for improved anti-cancer PTT through cancer-specific and image-guided drug delivery. KEYWORDS: metal-organic frameworks (MOFs), indocyanine green, theranostic, bioimaging, photothermal therapy.

1. INTRODUCTION Recently, photothermal therapy (PTT), i.e., employing photo-irradiation and a photo-absorber together for heat production to elevate the temperature of target tumor tissues to kill cancer cells,

1-3

has attracted attention as a localized cancer treatment

due to its minimal invasiveness and easy implementation, compared to other conventional approaches, such as radiotherapy, chemotherapy, and surgery. 4, 5 Various nanomaterials, including gold nanostructures (nanoparticles (NPs),

6

nanorods,

7

nanocages, 8 nanoshells, 9 nanoprisms, 10 and nanocubes 11), palladium nanosheets, 12, 13

carbon nanomaterials, 14 copper sulfide, 15, 16 copper selenide, 17 molybdenum oxide

nanomaterials,

18

and molybdenum disulfide nanosheets, 2


19

have been extensively

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explored

as

PTT

agents;

however,

these

nanomaterials

are

typically

non-biodegradable and have the potential for long-term toxicity. In addition, the high cost of noble metal-based nanomaterials limits their wide application. It is therefore necessary to develop novel, biodegradable, structure-based NIR photo-absorbers for PTT with high photothermal conversion efficiencies. Many photo-responsive dyes with absorption maxima in the NIR wavelength range have been exploited as theranostic agents.

20

Indocyanine green (ICG), the

commercially NIR organic dye for clinical applications,

21

is suitable for noninvasive

bio-imaging and can convert absorbed NIR light into heat for PTT;

20, 22

however,

several major issues prevent the use of ICG in in vivo applications, such as ICG’s unstable optical properties,

23

concentration-dependent aggregation and poor aqueous

stability, 24 rapid elimination from the body (half-life of 2–4 min), 22 and proneness to photo-bleaching.

3

To address these issues, free ICG was incorporated into NP

delivery systems, including polyallylamine, micelles,

27

liposomes,

28

25

hybrid nanovesicles

calcium phosphosilicate NPs,

30

silica NPs,

31

29

perfluorocarbon,

26

polypeptide

and other inorganic NPs (e.g.,

gold NPs,

32

super paramagnetic iron

oxide NPs, 33 and graphene oxide nanosheets 34, 35) . Although imaging-guided cancer therapy has been widely exploited, it is currently difficult to select the proper imaging modality for clinical diagnostics. Medical imaging methods with high target sensitivity usually have poor resolution, while those providing perfect resolution are often limited by low sensitivity; for instance, the NIR fluorescence (FL) probes have high sensitivity, but suffer from poor resolution and 3



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shallow tissue penetration (in terms of depth). In contrast, magnetic resonance imaging (MRI) possesses high spatial resolution but relatively poor sensitivity.

36-38

Therefore, the development of multi-modal contrast agents, e.g., NPs that combine FL imaging and MRI in one system, for in vivo imaging is expected. Recently, metal-organic frameworks (MOFs), a new class of self-assembled porous materials with metal connecting points and organic bridging ligands, have been widely investigated for potential biomedical applications in various fields, such as drug delivery,

39-41

molecular imaging,

40, 41

and biological sensing.

42, 43

Compared

to traditional nanocarriers, including inorganic silica, zeolites, and organic polymers/lipids, nanoscale MOFs (NMOFs) exhibit many special properties, including adjustable composition and structure, easy functionalization, and biodegradability. For example, the Materials of Institute Lavoisier (MIL) family of materials, which are developed from the iron carboxylate NMOFs, have potential as theranostics with good T2-weighted MRI properties.

41

Among the MIL family, MOF

iron (III) trimesate (MIL-100(Fe)) NPs recently have attracted attention as drug carriers and bio-imaging agents. MIL-100(Fe) NPs exhibited porosity (adsorption surfaces SBET ≈ 2400 m2·g−1, pore volumes Vp ≈ 1.2 cm3·g−1) that are associated with two kinds of mesoporous cages (ϕ ≈ 25 and 29 Å) accessible by microporous pentagonal and hexagonal windows (ϕ ≈ 5.5 and 8.6 Å).

44, 45

Also, these pores have

been utilized for loading and delivering various drug molecules. 41 To combine FL, photoacoustic (PA), and magnetic resonance (MR) multi-modal imaging capabilities with anti-cancer PTT of cancer under NIR laser irradiation, we 4


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developed a novel nanoprobe by incorporating photo-responsive ICG molecules into the MOF MIL-100(Fe) NPs (MOF NPs) in this study, enhancing ICG’s tumor accumulation and imaging intensity. Hyaluronic acid (HA), which mediates the targeting recognition of CD44 over-expressing cancer cells, is a natural, highly biodegradable, and biocompatible biopolymer.

46

To generate MIL-100(Fe) NPs with

high selectivity and binding affinity for the surfaces of cancer cells,

34, 47-49

the MOF

NPs were firstly conjugated to HA, and subsequently loaded with ICG molecules; finally, the multi-modal imaging and PTT capabilities of MOF@HA@ICG NPs were investigated (Figure 1). Our HA-conjugated and ICG-loaded MOF NPs showed several unique advantages, compared to other therapeutic agent-loaded NP systems: (1) MOF NPs can be synthesized at the mild temperature of 50 oC, show suitable stability under physiological conditions

50

and can be eliminated in urine or feces

without being metabolized or showing substantial toxicity;

51

(2) the large pore sizes

(2.5 and 2.9 nm) of MOF NPs ensure that large amounts of various drug molecules, including ICG, can be successfully encapsulated into them; 38, 41, 52 (3) the Fe (III) ions in the MOF NPs can be utilized for T2-weighted MRI;

38, 41, 53

and (4) compared to

other core-shell NMOF systems designed for imaging-guided PTT,

38, 52, 53

our HA

and ICG-engineered MOF NP system is easy to synthesize and control and avoids the production of toxic materials, which may result from the decomposition of core-shell nanostructures. The prepared HA and ICG-engineered MOF NPs exhibited good NIR absorbance, low cytotoxicity, great cellular uptake in CD44-positive MCF-7 cancer cells and tumor accumulation in xenograft tumors, and good in vitro and in vivo 5



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capabilities in FL imaging, PAI, T2-weighted MRI, and PTT treatments. We also compared the performances of MOF@HA@ICG NPs with those of MOF@ICG NPs (non-HA-targeted) and free ICG. These results demonstrate that the MOF@HA@ICG NPs are a new promising theranostic nanoplatform for imaging-guided cancer therapy.

Figure 1. Schematic representation of the synthesis procedure, HA conjugation, ICG loading, and multi-modal imaging-guided PTT of MIL-100(Fe) NPs. 2. EXPERIMENTAL SECTION 2.1. Materials FeCl3·6H2O, 1, 3, 5-benzenetricaboxylic acid (trimesic acid, BTC), HA, ICG, eosin, hematoxylin, 3-(4, 5-dimethyl-thiazol-2-yl)-2, 5-diphenyltertrazolium bromide (MTT), 4’, 6-diamidino-2-phenylindole (DAPI), methanol, and deionized (DI) water were purchased from Sigma-Aldrich (USA). Penicillin−streptomycin, Dulbecco’s 6


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modified eagle medium (DMEM), phosphate-buffered saline (PBS, pH=5.5, 7.4), and fetal bovine serum (FBS) were purchased from GIBCO BRL (Gaithersburg, MD, USA). 2.2. Preparation of MOF, MOF@ICG, and MOF@HA@ICG NPs MIL-100(Fe) NPs were prepared using the method reported in the literature

52

with minor modifications. In a typical procedure: firstly, 40 mL of 7.5 mM FeCl3 solution (in methanol) and 40 mL of 7.5 mM 1, 3, 5-benzenetricaboxylic acid (trimesic acid, BTC) solution (in methanol) was mixed under strong stirring for about 10 minutes (min). Afterward, the mixture was transferred into a 50 oC oil bath and reacted for 90 min. The reaction product was collected by centrifugation and washed 3 times with ethanol and DI water. Finally, the product was stored for use after freeze-drying. In order to obtain the products with targeting ability, the MOF NPs were conjugated with HA by mixing MOF NPs with HA at weight ratio of 5: 1 in an aqueous solution for 24 hours (h). The MOF@ICG or MOF@HA@ICG NPs could be obtained by mixing ICG with MOF or MOF@HA NPs at weight ratio of 1: 1 in aqueous solutions, followed by about 24 h of shaking in the dark. The unloaded free ICG molecules were removed by centrifugation. 2.3. Characterizations of NPs X-ray diffraction (XRD) patterns of the sample were recorded using a Rigaku Ultima IV system. The energy-dispersive X-ray (EDX) spectrum was performed with a FEI Quanta FEG 200 instrument equipped with EDX spectrometer. The Fe element content in MOF NPs was detected by ICP-MS. The dynamic light scattering (DLS) 7



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size and surface zeta potential of NPs were determined using a SZ-100 Nano Particle analyzer (Horiba Scientific). Transmission electron microscopy (TEM) was performed by using a JEM-2100 microscope (JEOL, Japan) at 200 kV of accelerating voltage. Fourier transform infrared (FTIR) spectra were collected on a JASCO FT/IR-420 spectrometer. Ultraviolet-visible-near-infrared (UV-Vis-NIR) absorbance spectra were recorded on a microplate reader (SH-1000 Lab, Japan). The fluorescence spectra were measured using a LS-55 luminescence spectrometer (Perkin-Elmer, Inc., USA; ex: 740 nm). The photoacoustic signals of samples were measured at 780 nm with Endra Nexus 128 (Ann Arbor, USA). The photothermal capacity of samples was evaluated using an 808 nm NIR laser. MRI of the MOF@HA@ICG solutions with various NP concentrations was performed using a small animal MRI (9.4T) scanner (Bruker, Germany). 2.4. Tumor Cells MCF-7 and NIH 3T3 cells lines were purchased from the cell bank of the Chinese Academy of Sciences in Shanghai and utilized for cell studies. Cells were cultured in DMEM

medium

supplemented

with

10%

FBS

and

100

µg/mL

of

penicillin/streptomycin solution in a humidified atmosphere of 37 oC and 5% CO2. 2.5. Cell Viability Assay MCF-7 cancer cells were seeded onto 96-well cell culture plates and cultured overnight for various NP concentrations of MOF, MOF@HA, or MOF@HA@ICG NP solutions treatment for further 48 h. Then the MTT assay was adopted to measure MCF-7 cancer cell viability. 8


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2.6. Cell Internalization of NPs Confocal laser scanning microscopy (CLSM) imaging, photoacoustic imaging (PAI), fluorescence (FL) imaging, magnetic resonance imaging (MRI), and flow cytometry were used to confirm the cellular uptake of the MOF NPs. MCF-7 or NIH-3T3 (CD44-negative) cells were seeded onto corresponding cell culture dishes and grown to about 70% confluency before used. The cells were treated with free ICG, MOF@ICG, or MOF@HA@ICG solutions in dark for 8 h (final concentration of ICG for treatment was 10 µg/mL); for the blocking assay, MCF-7 cells were treated with HA for 1 h ahead of NPs treatment. Before following imaging progress, cells were washed with PBS for 3 times. CLSM imaging was performed on the confocal fluorescence microscope (Nikon C2) to investigate the location of NPs in cells. Cells cultured onto confocal dishes were fixed using 4 % paraformaldehyde. DAPI staining was performed and then free dye was removed by washing 3 times with PBS. For PAI, FL imaging and MRI in vitro, MCF-7 cells were cultured in 100 mm culture dishes (about 106/dish) cells were trypsin digested and collected in 0.2 ml centrifuge tubes. After centrifugation, cell pellets were imaging in the Endra Nexus 128 PAI system at 780 nm, the Carestream FX PRO fluorescence imaging system or the Bruker 9.4T small animal MRI scanner. The ICG fluorescence of cells was also detected using an Accuri C6 flow cytometer (BD, Ann Arbor, MI). 2.7. In Vitro Photothermal Effects MCF-7 cancer cells (about 1×104 cells/well) were cultured onto 96-well plates. 9



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Free ICG, MOF@ICG, or MOF@HA@ICG solutions (containing 15 µg/mL of ICG) were added to treat cells for 8 h at 37 oC under 5% CO2. Then, cells were washed with PBS, immersed in 100 µL of fresh DMEM culture medium, and exposed to an 808 nm laser at 1 W/cm2 for 10 min, followed by a further incubation of 37 oC under 5% CO2 for 6 h. Finally, the MTT assay was utilized to quantify the cell viability. 2.8. Animals and Tumor Model Female BALB/c mice (4–6 weeks old, weighed 18–20 g) were obtained from Shanghai SLAC laboratory Animal Co., Ltd. The research project has been approved by the Animal Management and Ethics Committee of Xiamen University, China. The xenograft model was established by the subcutaneous injection of 3×106 MCF-7 cancer cells into mice. Tumor volumes were calculated with the formula: the volume= length × width2/2. 2.9. Animal Imaging Studies The mice were subjected to tail vein injection of 125 µL of free ICG, MOF@ICG, or MOF@HA@ICG solutions which all contained 170 µg/mL of ICG (Injection Dosage[MOF NPs]=2.5 mg/kg), for in vivo imaging. The PA images in the tumor sites were recorded by the Endra Nexus 128 scanner (Ann Arbor, MI, USA). The working laser wavelength was 780 nm and 30 pulses were used on average. PA images were performed at the 0, 1, 24, 48, and 72 h time points. The FL signals were observed using Carestream FX Pro at the 0, 1, 24, 48, and 72 h time points. For biodistribution analysis, mice were sacrificed 48 h post-injection 10


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and organs (heart, liver, kidney, spleen, and lung) and tumor were collected for ex vivo FL imaging. The in vivo MR images were performed on a 9.4T MRI instrument at room temperature. T2-Weighted MR images were obtained with the following parameters: TR= 2500, TE= 33 ms, field of view (FOV) = 4 × 4 cm, slice thickness = 1 mm, and 9 contiguous slices. 2.10. In Vivo PTT The MCF-7 tumor- bearing mice were randomized into 6 groups (5 mice in each group): groups 1, 2 and 3 were intravenously injected with 125 µL of free ICG, MOF@ICG,

and

MOF@HA@ICG

solutions

(C[ICG]=170

µg/mL,

Injection

Dosage[MOF NPs]=2.5 mg/kg), respectively, and underwent laser irradiation. Groups 4, 5 and 6 were designated as the control groups. Group 4 was intravenously injected with 125 µL of PBS and underwent laser irradiation, group 5 underwent laser irradiation alone (no injection), and group 6 did not receive any treatment. The in vivo PTT was carried out by using an 808 nm NIR laser. The tumors of mice 48 h post-injection were irradiated at 1 W/cm2 for 10 min. Temperatures in tumors were monitored using the FLIR Ax5 infrared thermal imaging camera. The tumor volumes, body weights, and the survival rate of mice were recorded. The tumor volumes were measured by a caliper and calculated with the formula: the volume = length × width2/2. Hematoxylin and eosin (H&E) staining were applied for analyzing the potential toxicity of samples to tumor tissues and main organs. 3. RESULTS AND DISCUSSION 11



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3.1. Characterizations of MOF@HA@ICG NPs The MIL-100(Fe) NPs were synthesized by following the method reported in the literature

52

with minor modifications. XRD was used to identify the crystal structure

and phase composition of the products. Figure S1A in the Supporting Information presents the XRD pattern of the MIL-100(Fe) NPs before and after the encapsulation of ICG molecules. All the diffraction peaks in the XRD patterns could be identified as a cubic structure, which is in agreement with previous reports.

45, 52

The loading of

ICG did not change the crystalline structure of MIL-100(Fe) NPs but slightly decreased the reflection intensity. The EDX spectrum (Figure S1B in the Supporting Information) indicated the presence of Fe, O, and C in the MIL-100(Fe) NPs, further showing the high purity of the product. TEM images (Figure 2A) show that the size of MOF@HA@ICG NPs was around 100 nm. The DLS size measurement in DI water (Figure 2B) revealed that the average size of the MOF@HA@ICG NPs was 106.4 ± 21.3 nm. The polydispersity index of MOF@HA@ICG NPs, indicating the size distributions of NPs, was less than 0.05. DLS was further utilized to study the colloidal stability over time of MOF@HA@ICG NPs dispersed in water, PBS, or DMEM. The particle sizes of MOF@HA@ICG NPs after 12 days were similar and lacked apparent variation (Figure S2 in the Supporting Information), suggesting that these NPs possess good colloidal stability in biological medium. The surface properties of MOF, MOF@HA and MOF@HA@ICG NPs were characterized by zeta potential measurements in DI water. The HA conjugation and ICG loading was reflected by the change in the zeta potential of the NPs (Table S1 in the Supporting 12


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Information). Because of the presence of HA on MOF NPs, the zeta potential of MOF@HA NPs decreased from -16.3 mV to -23 mV, and then to -25.4 mV after encapsulation of ICG molecules, due to the SO3 groups on the surface of ICG. FTIR spectroscopy was also carried out to study the chemical structure of the products (Figure 2C). The peaks at 1676, 1601, and 3200–3600 cm−1 respectively originated from the amide I and amide II bands and the –OH group in HA. The three peaks that appeared at 3300–2500, 1720–1680, and 950–890 cm-1 were assigned to the absorption peaks of ν(C=O), ν(O–H) and δ(O–H) in BTC, respectively. Two apparent peaks at 1622 and 1380 cm-1, corresponding to the asymmetric and symmetric stretching vibrations of –COO– anions, were found in the MOF@HA NP spectra. These results further confirmed the formation of MIL-100(Fe) product in our experiment and that HA had been successfully conjugated with MOF NPs. 54-56 Figure S3 in the Supporting Information indicates that HA conjugation improved the stability of MOF NPs in PBS.

13



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Figure 2. (A) TEM images and (B) DLS of the MOF@HA@ICG NPs. (C) FTIR spectra of BTC, HA, and MOF@HA. (D) Absorbance spectra of HA, ICG, MOF, MOF@HA, MOF@ICG, and MOF@HA@ICG aqueous solutions and up solutions 14


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(the images of samples a and b after centrifugation at 12000 rpm for 10 min are shown in the insets: a) MOF mixed with 150 µg/mL of ICG for 24 h; b) 150 µg/mL of free ICG). (E) The fluorescence spectra of free ICG, MOF@ICG, and MOF@HA@ICG NPs. (F) The ICG loading amounts of MOF and MOF@HA NPs. (G) Photothermal capacities of water and MOF@HA@ICG solutions with various ICG concentrations, under the irradiation of an 808 nm laser with 1 W/cm2 for 10 min. (H) Photothermal curves of MOF matrix alone and free ICG, MOF@ICG, and MOF@HA@ICG solutions, each with 30 µg/mL of ICG, under the irradiation of an 808 nm laser with 1 or 0.5 W/cm2 for 10 min.

UV-vis-NIR spectroscopy was utilized to determine the ICG loading amounts of MOF NPs. The absorption spectra of HA, MOF, MOF@HA, free ICG, MOF@ICG, and MOF@HA@ICG aqueous solutions and up solutions (obtained after centrifugation of the mixture solutions of MOF or MOF@HA NPs with ICG) are shown in Figure 2D. No characteristic absorbance was observed for HA and MOF, while after the incorporation of ICG into MOF, the MOF@ICG and MOF@HA@ICG NPs and their up solutions had maximum absorption peaks around 780 nm in the NIR region due to ICG incorporation. The encapsulation of ICG into the MOFs led to a red shift of the characteristic maximum absorption peak of ICG molecules from 779 nm (curve F) to 785 nm (curve E), which might be attributed to the interactions between ICG molecules and MOF and among the ICG molecules.

22

The photos of sample a

(MOF NPs solutions mixed with 150 µg/mL of ICG for 24 h) and b (150 µg/mL of 15



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free ICG) after centrifugation at 12000 rpm for 10 min (the insets in Figure 2D) revealed that ICG had been successfully incorporated into MOFs. The emission peak of the ICG encapsulated into MOFs was also red-shifted 4–5 nm (Figure 2E). Interestingly, the MOF@HA@ICG NPs exhibited good fluorescence stability in water, PBS, and DMEM within 5 days (Figure S4 in the Supporting Information). Figure S4 also indicates that the MOF encapsulation could improve ICG’s fluorescence stability. The ICG loading amounts in MOF and MOF@HA NPs were about 44 and 42 wt%, respectively, as determined by UV-vis absorption spectra (Figure 2F). The mechanism of ICG’s incorporation into MOFs might be due to the small molecular dimension of ICG and the π－π interaction between the aromatic rings in ICG and MIL-100(Fe). 34, 57, 58

The ICG-loaded NPs with strong NIR absorbance have potential as new

theranostic nanoplatforms for bio-imaging and cancer therapy. 30-35 In our experiment, free ICG, MOF@ICG, and MOF@HA@ICG solutions were irradiated with an 808 nm NIR laser at a powder density of 1 W/cm2 to evaluate and compare their photothermal efficiencies. As the ICG concentration or irradiation time increased, the temperatures of MOF@HA@ICG solutions increased accordingly (Figure 2G). As shown in Figure 2H, after irradiation for 3 min, the temperature of free ICG, MOF@ICG, and MOF@HA@ICG solutions (30 µg/mL of ICG each) increased to 65.47, 70.12, and 70.34 oC, respectively, whereas water (Figure 2G) and MOF matrix alone (Figure 2H) only increased to 32.13 and 36.07 oC, respectively. The temperature of free ICG solution started to decrease after 3 min of irradiation (Figure 2H), suggesting that free ICG has poor photostability. In contrast, the temperature of MOF@ICG and 16


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MOF@HA@ICG solutions remained at around 72 oC even after 10 min of irradiation (Figure 2H). Even under irradiation with a lower laser powder density (0.5 W/cm2), the temperature of MOF@HA@ICG solutions remained around 55 oC after 10 min of irradiation (Figure 2H). These results indicated that the MOF NPs’ encapsulation could improve the photostability and photothermal efficiency of free ICG. We also found that there was no obvious variation in the size and morphology of MOF@HA@ICG NPs after irradiation (Figure S5 in the Supporting Information), demonstrating that these NPs were photothermally stable. The samples were dialyzed in PBS with various pH values to investigate the release of ICG from MOF NPs. The curves of ICG release over time from MOF NPs at pH values of 5.5 and 7.4 indicated that ICG release was hardly sensitive to the pH value (Figure S6 in the Supporting Information), which might be an important reason for the good photothermal capacity of ICG-loaded MOF NPs. These results demonstrated that MOF@HA@ICG NPs are suitable for PTT due to their heat generation capacity and photothermal stability under NIR light irradiation.

Figure 3. (A) The PA intensity and images of MOF@HA@ICG solutions with various concentrations of ICG. (B) Relaxation rate R2 (1/T2) versus concentrations of 17



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Fe ions (the inserts are the T2-weighted MR images of MOF@HA@ICG NPs with various Fe ion concentrations).

PAI is a non-invasive imaging modality that detects the pressure wave caused by the photoacoustic effect. Due to MOF@HA@ICG NPs’ strong NIR absorbance, we investigated their potential as a good PAI contrast agent. As expected, a higher PA signal was obtained as the ICG concentration increased in aqueous MOF@HA@ICG solutions (Figure 3A). Since there was no characteristic absorbance for MOF@HA NPs in the UV-vis-NIR range, the PA signal of MOF@HA@ICG solutions should have come from the ICG molecules. We also studied the potential of MOF@HA@ICG NPs for T2-weighted MRI. As shown in Figure 3B, darker MR images were obtained as the concentrations of NPs in solutions increased, and the transverse relaxivity (r2) of NPs was approximately 54 mM−1S−1. Although the r2 value of MOF@HA@ICG NPs was lower than that of the commercially available T2-MR contrast reagent resovist, e.g., Fe3O4 NPs (98.3 mM−1s−1), 59 it was comparable to those of other Fe-MOFs.

60, 61

The negative enhancement in T2-weighted images

indicated that MOF@HA@ICG NPs could be utilized as a MRI contrast agent. 3.2. Cell Internalization of MOF@HA@ICG NPs The cellular uptake of MOF@HA@ICG NPs was investigated through CLSM. After 8 h of incubation, the MOF@ICG and MOF@HA@ICG NP-treated MCF-7 cancer cells showed higher fluorescence (FL) intensity in the cytoplasm than those treated with free ICG, indicating that MOF encapsulation could enhance the cellular 18


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uptake of free ICG molecules (Figure 4A). It is known that CD44, which is frequently over-expressed on the surface of tumor cells, is a cell-surface receptor for HA. As expected, with the competitive HA blocking, the FL intensity of MOF@HA@ICG NPs decreased in the cytoplasm of HA treated cells as compared to MOF@HA@ICG NPs without HA block. The cancer cells targetability of MOF@HA@ICG NPs was also studied using NIH-3T3 cells (CD44-negative) as a control group. It was found that although some MOF@HA@ICG NPs were nonspecifically internalized by the CD44-negative cells, the amount was very small and almost no FL signal could be observed (Figure S7 in the Supporting Information). The flow cytometry analysis (Figure 4B) also showed a greater average FL intensity in the MOF@HA@ICG group than in the MOF@HA@ICG with HA blocking group, which further demonstrated that

the

CD44-positive

MCF-7

cancer

cells

effectively

internalized

the

HA-conjugated MOF NPs via CD44 mediation. These results indicated that MOF@HA@ICG NPs could target the tumor site by both HA-mediated targeting and the enhanced penetration and retention (EPR) effect.

Figure 4. In vitro cellular uptake: (A) Confocal images and (B) flow cytometry 19



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analysis of MCF-7 cells treated with free ICG, HA block, MOF@ICG, or MOF@HA@ICG solutions (solutions that contained 10 µg/mL of ICG, scale bar= 50 µm). 3.3. In Vitro Photothermal Effect of MOF@HA@ICG NPs The toxicology of ICG-loaded MOF NPs must be determined before they can have any biomedical use. We used the MTT assay to study the potential cytotoxicity of MOF@HA@ICG NPs by incubating MCF-7 cancer cells with NPs for 48 h. Figure 5A demonstrates that their cellular viability was greater than 80% with NP concentrations that ranged from 0 to 250 µg/mL, indicating that these NPs possessed low cell cytotoxicity. Inspired by the good photothermal performance of ICG-loaded MOF NPs, we investigated their in vitro photothermal effect on cancer cells. As shown in Figure 5B, after irradiation with NIR laser (808 nm, 1 W/cm2), 10% of MCF-7 cancer cells treated with free ICG (15 µg/mL) died, while 70% of the MOF@ICG and 90% of the MOF@HA@ICG groups’ cells died, compared to the nearly 100% viability of the control group without NIR laser irradiation, suggesting that encapsulation of ICG in MOF NPs could improve its effectiveness in PTT. The improved photothermal efficiency of MOF@HA@ICG NPs, compared to that of MOF@ICG NPs, was attributed to their enhanced cellular uptake by HA-mediated tumor cell targeting, which is demonstrated in Figure 4. These results revealed that MOF@HA@ICG NPs were highly biocompatible and could be employed for cancer specific and imaging-guided PTT.

20


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Figure 5. (A) Cell viabilities of MCF-7 cells incubated with MOF, MOF@HA, or MOF@HA@ICG solutions for 48 h. (B) Photothermal effects on the growth of MCF-7 cells treated with free ICG, MOF@ICG, or MOF@HA@ICG solutions (15 µg/mL of ICG, 1 W/cm2 for 10 min). 3.4. In Vitro PA, FL, and MR Imaging of MOF@HA@ICG NPs The MCF-7 cellular internalization of MOF@HA@ICG NPs was further investigated by PA, FL and MR imaging. Figure 6A, B show the PA images and PA intensity of MCF-7 cancer cells treated with ICG-encapsulated MOF NPs or free ICG solutions. As expected, a higher PA signal was obtained with increased ICG concentrations in solutions. Brighter images could be seen in the cells treated with MOF@HA@ICG than those treated with free ICG or MOF@ICG, indicating the HA-mediated targeting of CD44 over-expressing MCF-7 cells. Figure 6C demonstrates the FL images of MCF-7 cancer cells treated with MOF@HA@ICG solutions. As the ICG concentration increased in MOF@HA@ICG solutions, the FL intensity of cells increased accordingly. The in vitro T2-weighted MR images of cells treated with MOF@HA@ICG solutions of different NP concentrations (0, 1.0, 2.5, 5.0, and 10 µg/mL) were also acquired; as shown in Figure 6D, the MRI signal 21



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became hyperintense in T2-weighted MR images as MOF@HA@ICG concentration in solutions increased due to the dose-dependent cellular uptake.

Figure 6. In vitro MCF-7 cancer cell imaging: (A) PA images and (B) PA intensity of cells treated with free ICG, MOF@ICG, or MOF@HA@ICG solutions; (C) FL images and (D) T2-weighted MR images of cells treated with MOF@HA@ICG solutions.

3.5. In Vivo Imaging and PTT of MOF@HA@ICG NPs On the basis of the in vitro treatment results, we further applied the MOF@HA@ICG NPs to mice for FL imaging, PAI, MRI, and PTT in vivo. All the animal experiments were conducted according to protocols approved by the Laboratory Animal Center of Xiamen University. We investigated the accumulation of the ICG-loaded MOF NPs in tumors and organs using the NIR fluorescence of ICG molecules. The MCF-7 tumor-bearing mice were treated with free ICG, MOF@ICG, or MOF@HA@ICG solutions (125 µL, 170 µg/mL of ICG) via tail vein injection. The FL signals were collected at different time points (Figure 7A). The FL signals of 22


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the free ICG group 1 h after injection were extensive in the liver but less in the tumor, and these signals completely disappeared 24 h after injection. On the contrary, the FL signals in the tumor were progressively strengthened with increased time intervals (0, 1, 24, and 48 h) and reached a peak 48 h after MOF@ICG or MOF@HA@ICG solutions injection; the MOF@HA@ICG group showed stronger fluorescence than that of the MOF@ICG group, confirming the specific in vivo targeting of MOF@HA@ICG NPs. However, the FL intensity in the tumor sites decreased 72 h after injection of the MOF@ICG or MOF@HA@ICG solutions, though the tumor 72 h after injection of MOF@HA@ICG solutions still retained FL signals. These results of in vivo NIR FL imaging further demonstrated that the MOF encapsulation could effectively reduce the degradation and clearance of ICG in the living body. Ex vivo FL imaging was also performed to further investigate the NP distribution in various organs. The distribution of free ICG, MOF@ICG NPs, and MOF@HA@ICG NPs in major organs and tumors 48 h post-injection are shown in Figure 7B. The FL signal was clearly seen in the MCF-7 tumor site of the mouse treated with MOF@HA@ICG NPs, but a much weaker signal was observed in the MOF@ICG or free ICG groups (Figure 7B). The ex vivo FL intensity of the major organs and tumors 48 h post-injection of the MOF@HA@ICG, MOF@ICG, and free ICG solutions was consistent with our in vivo FL imaging results (Figure 7C). The in vivo imaging ability of MOF@HA@ICG NPs was further evaluated by PAI in mice bearing MCF-7 tumor. PAI was performed before and after intravenous injection of the MOF@HA@ICG solutions into tumor-bearing mice. The PA images 23



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around the tumor region obtained 48 h post-injection of the MOF@HA@ICG solution showed about 1.9 times stronger PA signals than the pre-injection images (Figure 7D). T2-weight MR images were also recorded before and after MOF@HA@ICG injection using a 9.4T MR instrument. The tumor area displayed a much darker image with a sharp

border than the

pre-injection image

(Figure

7E),

suggesting that

MOF@HA@ICG NPs were successfully delivered into MCF-7 tumors.

Figure 7. (A) FL images of MCF-7 tumor-bearing mice injected with free ICG, MOF@ICG, or MOF@HA@ICG solutions (each with 170 µg/mL of ICG; the tumors are highlighted by red circles). (B) ex vivo FL images and (C) intensity of organs (heart, liver, kidney, spleen, and lung) and tumors. (D) in vivo PA images and (E) T2-weighted MR images with MOF@HA@ICG treatment. The tumors are highlighted by red circles. 24


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MCF-7 tumor-bearing mice were then divided into 6 groups (5 mice in each group), i.e., control (without any treatment), PBS, laser only, free ICG, MOF@ICG, or MOF@HA@ICG groups, for further in vivo PTT experiments. One hundred and twenty-five µL of PBS, free ICG, MOF@ICG or MOF@HA@ICG solutions (170 µg/mL of ICG in ICG-containing solutions) were injected into the mice via their tail veins. According to the result of in vivo NIR FL imaging of ICG, PPT was performed at the time point of 48 h post-injection. Immediately after NIR laser irradiation, temperature changes in the tumor site were monitored in real-time using IR thermal imager. As expected and shown in Figure 8A, B, the tumors of the PBS or the laser only groups exhibited moderate increases in temperature, to around 36.5 or 37.5 oC during 10 min of laser irradiation, respectively, while the free ICG group showed no obvious temperature increase (38 oC). The temperatures of tumors treated with MOF@ICG or MOF@HA@ICG solutions gradually increased with the extension of irradiation time and reached about 43 or 52 oC after 10 min of laser irradiation, respectively. The tumor cellular cytotoxicity occurs when temperature reaches 41.5 oC, and the vascular destruction within tumor tissue will be induced if temperatures reach above 43 oC.

20

In our experiment, the temperatures of tumors treated with

MOF@HA@ICG (52 oC) were clearly greater than the threshold temperature that induced the irreversible damage of tumor.

62, 63

As shown in Figure 8C, the control,

PBS, or laser only groups showed rapid tumor growth, indicating that the tumor growth was not affected by the laser irradiation. The free ICG or MOF@ICG group 25



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Page 26 of 40

did not show significant tumor suppression, while the MOF@HA@ICG group caused complete ablation of the tumor after 14 days (Figure S8 in the Supporting Information). Additionally, the MOF@HA@ICG group exhibited an 80% survival rate after 20 days (Figure 8D). During the 14 days of treatment, no significant body weight change

was observed in all groups (Figure

8E),

showing that

MOF@HA@ICG NPs had no significant side effects on the treated mice in in vivo PTT.

Figure 8. (A) Photothermal effect and (B) the average temperature of the tumor site in all groups under the imaging-guided irradiation. (C) Tumor growth curves, (D) survival rates and (E) body weight curves of MCF-7 tumor-bearing mice after treatment with PBS, free ICG, MOF@ICG, or MOF@HA@ICG solutions.

26


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Figure 9. H&E stained images of the heart, liver, spleen, lung, kidney, and tumor from the mice in all groups.

The tumors and normal organs were analyzed by H&E staining to further investigate the PTT effect and potential toxicity of MOF@HA@ICG NPs.63 The MOF@HA@ICG group clearly showed tumor necrosis and destroyed blood vessels, while the MOF@ICG group induced moderate levels of tumor necrosis (Figure 9). By contrast, no apparent tumor destruction was observed in the other control groups (Figure 9). The H&E images of major organs (heart, liver, spleen, lung and kidney) from MCF-7 tumor-bearing mice in all groups showed that there was no noticeable organ damage or inflammation lesion induced by the applied laser or the PTT treatment (Figure 9). These results clearly demonstrate that the MOF@HA@ICG NPs 27



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have high biocompatibility and can be served as a nanotheranostic agent for imaging-guided PTT of cancer. 4. CONCLUSION In summary, we successfully developed a biocompatible and multifunctional HA and ICG-engineered MOF nanotheranostic agent for anti-cancer treatment. The MOF@HA@ICG NPs showed low cytotoxicity, high capacity for ICG loading, and the ability to target MCF-7 cells/xenograft tumors. The integration of FL, PA, and T2–weighted MR imaging-guided PTT using MOF@HA@ICG NPs induced significant MCF-7 cell death in vitro and efficiently suppressed the MCF-7 tumor growth in vivo. These results demonstrate the potential theranostic value of MOF@HA@ICG NPs for the imaging-guided PTT therapy of solid tumors and have opportunities to significantly contribute toward the clinical translation of MOF-based hybrid nanocomposites.

ASSOCIATED CONTENT Supporting Information: The XRD and EDX of MOF NPs; the stability of MOF@HA@ICG NPs in water, PBS, or DMEM; the zeta potential of samples; the photos of MOF and MOF@HA NPs dispersed in PBS; the fluorescence stability of MOF@HA@ICG NPs in water, PBS, or DMEM; the TEM image of MOF@HA@ICG NPs after irradiation with NIR laser; the release profiles of ICG over time from MOF NPs at different pH values; the confocal images of CD44-negative NIH-3T3 cells treated with MOF@HA@ICG NPs; the photos of 28


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tumor sites obtained on day 7 and 14 after different treatment. AUTHOR INFORMATIONS Corresponding Author *Tel:+86-592-2880648. E-mail: [email protected]. ǁ

W.C. and H.G. contributed equally to this work.

ACKNOWLEDGMENTS This project was financially supported by the Research Support Program for Young Teachers in Xi'an Jiaotong University (YX1K075), the Major State Basic Research Development Program of China (2014CB744503 and 2013CB733802), the National Natural Science Foundation of China (81422023, 51273165, and U1505221), the Fundamental Research Funds for the Central Universities (20720160065 and 20720150141), the Science Foundation of Fujian Province (2014Y2004 and 2014J05098), and the Program for New Century Excellent Talents in University, China (NCET-13-0502). REFERENCES [1] Huang X.; El-Sayed I. H.; Qian W.; El-Sayed M. A. Cancer Cell Imaging and Photothermal Therapy in the Near-Infrared Region by Using Gold Nanorods. J Am Chem Soc 2006, 128, 2115-2120. [2] Yang K.; Wan J.; Zhang S.; Tian B.; Zhang Y.; Liu Z. The Influence of Surface Chemistry and Size of Nanoscale Graphene Oxide on Photothermal Therapy of Cancer Using Ultra-Low Laser Power. Biomaterials 2012, 33, 2206-2214. [3] Zheng X.; Zhou F.; Wu B.; Chen W. R.; Xing D. Enhanced Tumor Treatment 29



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