Chitosan coated gold nanorod chelating gadolinium for MRI-visible photothermal therapy of cancer

Abstract

The quick development of photothermal therapy (PTT) affords great opportunities for cancer therapy owing to its minimally invasive nature and controllable treatment method. A facile imaging guidance for the treatment will facilitate its clinical application potential. A type of multifunctional hybrid nanoparticles (h-NPs) has been developed by loading gold nanorods (GNRs) into gadolinium (Gd)–DTPA-conjugated chitosan (Gd–DTPA–CS). Then, the vesicular surface was coated with a PGA-g-mPEG layer via electrostatic adsorption, resulting in GNR/Gd–DTPA–CS@PEG NPs with uniform size distribution, high colloidal stability and bio-safety. In vivo magnetic resonance (MR) imaging of tumor bearing mice revealed tumor accumulation of h-NPs administered by intravenous (IV) injection. Under the guidance of MR imaging, the photothermal therapy achieved an effective tumor ablation. Thus, this study showed the potential of h-NPs as a potent photothermal agent for MRI-guided cancer treatment.

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Introduction

Over the past decade, photothermal therapy (PTT) as an emerging therapeutic modality for cancer has drawn much attention owing to its advantages over traditional chemotherapy, e.g. focally ablating tumor tissue with minimal damage to normal tissue.^1,2 This strategy relies on the employment of externally applied PTT agents, which can convert the near infrared light (NIR) to thermal energy, a way well-known for tumor regional hyperthermia.^3,4 Up to now, gold nanorods (GNRs) have been widely explored as a PTT agent because of their superior photothermal properties. Compared with other PTT agents including gold nanocages,⁵ gold nanoshells,⁶ carbon nanotubes,⁷ graphene oxide,⁸ organic dye molecules and semiconductor nanocrystals,⁹ GNRs show tunable light absorption peaks and stronger photothermal conversion efficiency, which make them especially suitable for tumor regional therapy. However, the synthesis of GNRs usually involves the use of a cytotoxic surfactant, cetyltrimethylammonium bromide (CTAB), to stabilize the nanoparticles. The unbound CTAB is apt to self-aggregate and thus hard to completely eliminate from the GNR suspension via common purification methods like ultracentrifugation, which greatly hinders the clinical application potential of GNRs.¹⁰

On the other hand, imaging technologies currently used for disease diagnosis can be adopted to ensure that the NIR light irradiation covers the whole tumor sites, to offer straight information on delivery events of GNRs and to monitor the therapeutic response of tumor.^11,12 For example, the therapeutic outcome may be improved by optimizing irradiation time according to the tumor accumulation of photothermal agents detected by imaging tools.^13,14 Theranostic nanoplatforms combining therapy and imaging functions are highly desirable for focal ablation of tumor.¹⁵ At present, magnetic resonance imaging (MRI) is widely applied in tumor diagnosis. With the assistance of contrast agents, MRI may further improve its tissue resolution and detection accuracy.¹⁶ Unfortunately, the typical MRI contrast agent gadolinium complexes possess short circulation time because of their easy renal clearance. For example, the half-life of Gd–DTPA in the blood is only 4 minutes.¹⁷ The water-soluble chitosan with good biocompatibility has been used to conjugate the Gd–DTPA chelates, which prolonged the blood circulation time and meanwhile increased the relaxivity of the MRI contrast agents.¹⁸ However, the contrast agents still lack the ability to target tumor site. Integration of gadolinium into nanoplatforms with passive tumor targeting potential and prolonged blood circulation has been proved to be a feasible means for developing highly effective theranostic nanomedicines.^19,20 For example, the West group has developed gadolinium-conjugated gold nanoshells coating silica nanoparticles, which exhibited increased T₁ signal intensities for successfully imaging-guided thermal ablation.²¹ Another type of T₂ contrast agents such as SPIO-conjugated gold nanoshells or Fe₃O₄@Cu_2−xS core–shell NPs have been explored for MRI-guided PTT as well.²² The preparations of the above theranostic systems all need complicated multi-step synthesis in order to control the thickness of shells for light absorption in the NIR region. Besides, it was rather difficult to control the stability and uniformity of the nanostructures.^23,24 Therefore, it is necessary to fabricate novel theranostic nanoplatform via a simple and effective approach.

The present study aimed to develop multifunctional hybrid nanoparticles (h-NPs), denoted as GNR/Gd–DTPA–CS@PEG, via a simple method for MRI-guided photothermal therapy of cancer. To this end, we first prepared gold nanorods (GNRs) encapsulated with Gd–DTPA-conjugated chitosan via non-covalent counterion interaction. Then, a PGA-g-mPEG layer was coated onto the vesicular surface via electrostatic adsorption (Fig. 1). The Gd–DTPA-conjugated chitosan is expected to not only offer the h-NPs high MRI T₁ sensitivity, but also protect the GNRs from aggregation in bloodstream without the use of cytotoxic CTAB. Since the MRI scan may reveal the time course of the tumor accumulation of nanoparticles, the NIR laser can be applied to the tumor site in an appropriate post-injection time to improve the photothermal efficiency and meanwhile to lower the side effects. The murine CT-26 colon cancer-bearing mice were adopted to illustrate the theranostic potential of h-NPs combining MR imaging and photothermal therapy.

Fig. 1 The schematic illustration for the synthesis of GNR/Gd–DTPA–CS@PEG (h-NPs) as a theranostic nanoplatform.

Experimental

Materials

Chitosan (CS) with an average molecular weight (M_n) of 5000 was obtained from golden-shell Biomedical Company (Zhejiang, China). Diethylene triamine pentacetate acid (DTPA), gadolinium(III) chloride (GdCl₃·6H₂O), sodium borohydride (NaBH₄, 99%) and sodium oleate (NaOL, 98%) were purchased from J&K Science Ltd. N-Hydroxysuccinimide (NHS), 3-(3-dimethylaminopropyl)-1-ethylcarbodiimide hydrochloride (EDC) were obtained from Adamas Reagent, Ltd. Cetyltrimethylammonium bromide (CTAB), glutaraldehyde (GA), silver nitrate (AgNO₃, 99%) and 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide (MTT) were purchased from Sigma-Aldrich. Chloroauric acid (HAuCl₄) and ascorbic acid (VC) were purchased from Sinopharm Chemical Reagent Co. Ltd. Poly(γ-glutamic acid) (γ-PGA; 5000 Da) was purchased from Nanjing Sai Taisi Biotechnology Co. Ltd. (China). MeO–poly(ethylene glycol)–NH₂ (mPEG–NH₂; 3000 Da) was bought from Jenkem Technology Co. Ltd. The clinical MRI contrast agent gadopentetate dimeglumine were obtained from Beijing Beilu pharmaceutical Co. Ltd. The water used in all preparations was purified in Milli-Q synergy purification system. Unless otherwise stated, all reagents and chemicals were used as received without further purification.

Preparation of h-NPs

Fabrication of GNR. GNRs were synthesized with improved seed-mediated growth approach.²⁵ Briefly, the seed solution for gold nanorods was synthesized by injection of 0.6 mL ice cold 0.01 M NaBH₄ to a 10 mL aqueous solution containing 0.1 M CTAB and 0.25 mM HAuCl₄ under vigorous stirring. After the solution was stirred for 2 min, it was matured at room temperature for 30 min before further use.

To prepare growth solution, 1.4 g CTAB together with 246.8 mg sodium oleate were dissolve in 50 mL deionized water followed by sequential addition of 3.6 mL 4 mM AgNO₃, 50 mL 1 mM HAuCl₄ and 0.2 mL hydrochloric acid (37.5%). Next, 0.25 mL 0.064 M ascorbic acid as a mild reducing agent was injected into the mixed solution after stirring for 1 h. Finally, 0.08 mL seed solution was introduced to the growth solution. The resultant mixture was stirred for 30 seconds and left undisturbed at 30 °C for 12 h to obtain GNR.

Synthesis of Gd–DTPA conjugated chitosan. Coupling of DTPA with chitosan was performed by amide coupling reaction between amine groups of chitosan and carboxylic groups of DTPA using EDC/NHS activating system. In brief, 223.0 mg CS (5000 Da) and 78.7 mg DTPA were dissolved in PBS solution (pH = 5.8) respectively. The carboxyl groups of DTPA were activated by adding EDC and NHS (molar ratio of DTPA = 1.5 : 1) and stirring for 30 min at room temperature. The activated DTPA solution was added to the CS solution for 2 min, then allowed to react for further 24 h, after which Tris–HCl buffer (pH = 9.0) was added. The resulting product was dialyzed using 3000 Da cutoff membrane against deionized water for 3 days.

After lyophilization, a certain amount of DTPA–CS was mixed with GdCl₃ solution. The sample mixture was continuously stirred for 6 h. Excess Gd ions were removed by dialysis for 3 days, followed by lyophilization of the final product. The conjugation rate of DTPA–CS was calculated according to the ¹H-NMR spectrum. The powder of DTPA–CS dissolved in D₂O were measured with Bruker Avance III 400 MHz.

Synthesis of PGA-g-mPEG. Poly(glutamic acid)-gra-poly(ethylene glycol) (PGA-g-mPEG) was synthesized using the method reported in the literature.²⁶ In brief, PGA (5000 Da), MeO–poly(ethylene glycol)–NH₂ (mPEG–NH₂, 3000 Da) and N-hydroxysuccinimide (NHS) were dissolved in borate saline buffer (0.05 M, pH = 8.5) with different ratios. Then, N-(3-dimethylaminopropyl)-N-ethylcarbodiimide hydrochloride crystallite (EDC) was added to the mixture and the resultant solution was stirred for 8 h at room temperature. After that, the mixture was dialyzed using 5000 Da cutoff membrane against deionized water for 24 h. The resulting product was freeze-dried and stored at −20 °C. The conjugation rate of PGA-g-mPEG was calculated according to the ¹H-NMR spectrum. The powder of PGA-g-mPEG dissolved in D₂O were measured with Bruker Avance III 400 MHz.

Preparation of h-NPs. The GNR/Gd–DTPA–CS NPs were synthesized using a nonsolvent counterion complexation method.²⁷ In brief, unbound CTAB in GNRs was removed by several cycles of washing by deionized water. Firstly, 10 mg Gd–DTPA–CS and 0.48 mg DTPA were dissolved in 2 mL deionized water and then mixed with 0.5 mL concentrated GNRs solution (2.4 mg mL⁻¹). After sonication of the above mixture for 2 min, 5 mL ethanol was added dropwise into the resulting mixture under moderate stirring for 4 min. During this procedure, the final reaction mixture obviously turns to opalescent from clear, which indicates the formation of GNR/Gd–DTPA–CS NPs. Subsequently, the obtained hybrid nanoparticles were crosslinked by 10 μL GA solution (25%) for 12 h, followed by dialyzing to remove the excess DTPA, GA and ethanol. The empty hybrid nanoparticles were fabricated similarly without GNRs.

To prepare the PEG-modified GNR/Gd–DTPA–CS NPs, a certain amount of PGA-g-mPEG dissolved in 10% glucose solution was added to an equal volume of hybrid nanoparticles aqueous solution to adjust the glucose concentration to 5% (isosmotic solution). Then, the resulting solution was thoroughly blended by pipetting and incubated at 25 °C for 30 min to fabricate h-NPs. The h-NPs were stored at 4 °C for use.

Characterization

The structure and morphology of GNR/Gd–DTPA–CS and h-NPs were characterized by transmission electron microscopy (TEM-Hitachi, 7750, Japan). Briefly, 10 μL properly diluted sample was dropped on a carbon coated copper grid and stained with phosphotungstic acid. The hydrodynamic diameter, size distribution and surface charges of prepared h-NPs were measured by dynamic light scattering (DLS) method using Malvern Zetasizer Nano ZS 90 (Malvern Co., USA). The UV/Vis absorption spectra were recorded on a UV-Vis/NIR spectrophotometer (DU730, Beckman Co., USA). The gold and gadolinium concentrations were quantified by inductively coupled plasma-mass spectrometry (ICP-MS, IRIS-HR, Thermo Jarrell Ash Co., USA). The samples were digest in concentrated nitric acid and heated to 120 °C for 48 h before measurements. Thermogravimetric analysis curve of the samples were collected with a thermogravimetric analyzer (TGA, TGA/DSC-1, Mettler Toledo Co., USA) in a temperature range of 100–650 °C and at a heating rate of 10 °C min⁻¹ under nitrogen flow (flow rate: 40 mL min⁻¹).

Stability studies

The colloidal stability of GNR/Gd–DTPA–CS NPs and h-NPs was investigated by monitoring the size changes of nanoparticles under mimic physiological conditions. Typically, the nanoparticles incubated in PBS solutions (containing 10% fetal bovine serum) were mildly shaken at 37 °C. After a certain incubation time, the size distribution of nanoparticles was evaluated using DLS. All samples were measured in triplicate.

Measurement of MRI relaxation properties

Imaging of MRI phantoms and relaxivity measurement were performed with a 1.5 T system by using an 11 cm circular surface coil (C3; Philips Medical Systems) at room temperature. The h-NPs and gadopentetate dimeglumine were dispersed in water at various Gd concentration. The T₁ mapping-weighted images was acquired with an inversion recovery sequence and the following parameters: TR/TE = 3500/20 ms, NSA = 1, FOV = 60 mm, matrix size = 256 × 256, section thickness = 1.0 mm and no intersection gap. Relaxation rate, r₁ was then calculated as the inverse of relaxation time (1/T₁) according the following formula:

(1/T₁)_obs = (1/T₁)_d + r₁[M]

where (1/T₁)_obs and (1/T₁)_d are the relaxation rates of the protons in the presence and absence of paramagnetic species (Gd), respectively, and [M] is the concentration of the paramagnetic species (Gd).

Temperature elevation induced by NIR laser irradiation

The aqueous solution with various concentration of PEG-modified h-NPs in a polystyrene cuvettes (total volume of 4 mL) was irradiated by a NIR laser at 808 nm (BWT Beijing LTD, China) with output power of 2 W cm⁻² for 10 min. The temperature of the solution was measured by a digital thermometer every 10 s.

Cytotoxicity by MTT

The human cervical carcinoma cells (HeLa cells) and murine colon carcinoma cell line (CT-26 cells) were originally purchased from the Chinese Academy of Sciences (Shanghai, China). The HeLa cells and CT-26 cells were cultured in DMEM culture medium supplemented with 10% fetal bovine serum (FBS) and 1% penicillin/streptomycin at 37 °C under 5% CO₂.

The in vitro cytotoxicity of h-NPs was evaluated using MTT assay. HeLa cells and CT-26 cells were placed in 96-well plates at 4 × 10³ cells per well and incubated at 37 °C under 5% CO₂ for 12 h. After removing previous medium, fresh medium containing different concentrations of h-NPs was added and incubated continuously for further 24 h and 48 h respectively. Thereafter, MTT assay was carried out to determine the cell viability following washing cells by PBS.

Erythrocyte agglutination and hemolysis assay

For the erythrocyte agglutination study, erythrocytes were separated by centrifugation of human blood at 1500 rpm for 10 min and purified via four successive washes with sterile PBS (0.01 M, pH = 7.4). Then, 300 μL of the erythrocytes (2 × 10⁷ cells per mL) was added into 1 mL h-NPs solution with a concentration of 200 μg mL⁻¹ and 800 μg mL⁻¹, respectively. After gentle shaking, the mixtures were incubated at 37 °C for 3 h. Finally, 10 μL solution mixture was spread on a glass slide to examine erythrocyte agglutinating via a microscope.

For the hemolysis assay, 300 μL suspension of erythrocytes (4 × 10⁸ cells per mL) was added to 1.0 mL of 0.1% Triton X-100 (positive control), sterile PBS (negative control) and PBS buffer containing h-NPs with a concentration ranging from 50 to 800 μg mL⁻¹. Then, the mixture solutions were left to stand for 3 h at 37 °C. After the incubation period, the supernatants were collected by centrifugation (1500 rpm for 10 min) and measured at a wavelength of 541 nm using a microplate reader (BioTek Synergy 4, Gene Co., USA) to analyze the release of hemoglobin. The hemolysis percent of each sample was calculated by the following formula:

The percent hemolysis = (A_Sample − A_PBS)/(A_TX-100 − A_PBS) × 100%

In vitro photothermal therapy

To investigate the photothermal cell toxicity of h-NPs under laser irradiation, HeLa cells were incubated with various concentrations of h-NPs at 37 °C for 2 h. Then, the cells were exposed to near infrared laser using an 808 nm optical fiber-coupled diode (2 W cm⁻²) for 5 min and incubated for another 12 h. After that, a standard cell viability assay using MTT was adopted to evaluate the cancer cell killing efficiency. For calcein-AM/PI assay staining, HeLa cells were incubated with or without 800 μg mL⁻¹ h-NPs in 24 well plates (2 × 10⁵ cells per well). After exposure with or without the 808 nm laser (2 W cm⁻²) for 5 min, the cells were incubated for additional 1 h. The cells were then rinse by PBS and stained with 2.0 μM calcein-AM (Sigma-Aldrich, USA) and 1.5 μM propidium iodide (PI, Sigma-Aldrich, USA), the cells were then recorded via an a fluorescent microscope (LX71, Olympus Co., Japan).

Animal and tumor model in vivo MRI

Male Balb/c mice with the approximate average weight 20 g were purchased from laboratory animal center of Sun Yat-sen University. All animal experiments were approved and performed in accordance with the guidelines of the Institutional Animal Ethical and Welfare Committee of Sun Yat-sen University (approval number: IACUC-2014-0807). To develop the CT-26 tumor model, 5 × 10⁶ CT-26 cells in 100 μL PBS solution were subcutaneously injected into the right flank of mice. Animal experiments were carried on the 7–10th days after CT-26 cells implantation, once tumor had approached a size of 50–70 mm³.

For in vivo MR imaging, tumor-bearing mice were intravenous injected with h-NPs (200 μL, 800 μg mL⁻¹) and imaged using a clinical 3.0 T system (Achieva, Philips Medical Systems, NL) with a 50 mm × 69 mm linearly polarized birdcage radio frequency rat coil (Shanghai Chenguang Medical Technologies, China). Axial and coronal fast spin echo T₁-weighted imaging (TR/TE = 800/15 ms, NSA = 10) and T₁-map (TR/TE = 3500/20 ms, NSA = 1) were performed. Other parameters for these sequences were FOV = 90 mm, matrix = 56 × 256, section thickness = 2.0 mm and no intersection gap.

In vivo photothermal therapy

For photothermal therapy studies, the tumor bearing mice were randomly divided into 4 groups with 4 mice per group (1) control (saline); (2) saline with laser; (3) h-NPs without laser; (4) h-NPs with laser. 808 nm NIR laser irradiation was performed 6 h after intravenous injection at the power density of 2 W cm⁻² for 10 min. The thermo images of tumor tissue were recorded by an infrared thermal camera of group (2) and (4) (Ti27, Fluke Co., USA). Tumor size and body weight were measured every other day after treatment.

Histology analysis

One day after treatment, one mouse from each group were euthanized. Tumors were collected and fixed in 10% neutral buffer formalin. After the tumors were embedded in paraffin and sectioned at 5 mm, the slices were stained with hematoxylin and eosin (H&E) and examined with a microscope (BX53, Olympus Co., Japan). Forty days after injection of h-NPs, the mice form each group were sacrificed. Major organs were collected and fixed in 10% neutral buffer formalin. The slices were stained with hematoxylin and eosin (H&E) and examined with the same method.

Results and discussion

Preparation and characterization of h-NPs

To prepare the multifunctional h-NPs, DTPA was conjugated with chitosan. As shown in Fig. S1,† the successful synthesis of DTPA-CS was demonstrated by ¹H NMR analysis. The DTPA conjugation degree of CS amino groups was 10.8% based on the ratio of the characteristic resonance of DTPA at 3.6 ppm (–N–CH₂–C=O–) to that of chitosan at 2.0 ppm (NH–C=O–CH₃) in the ¹H NMR spectra. The signal at 2.0 ppm was attributed to the three proton of non-deacetylated monosaccharide units (representing 20% of the units).

The ¹H-NMR spectrum of the conjugate revealed the characteristic resonance of mPEG at 3.3 ppm (–OCH₃) and the absorption of PGA at 2.3 ppm (–CH₂CH₂COO–) (Fig. S2†). The conjugation rate of PGA-g-mPEG was 10.4%, which were determined according to the proton resonance absorptions and calculated as the percentage of PGA carboxylate groups that are PEGylated.

The morphology and nanostructure of the h-NPs (i.e. GNR/Gd–DTPA–CS@10% PEG) were characterized by transmission electron microscopy (TEM) and dynamic light scattering (DLS). As shown in Fig. 2A, the GNRs exhibit average dimensions of 17.7 ± 1.4 nm in diameter and 66.7 ± 2.4 nm in length (corresponding to an aspect ratio of ∼3.7). The morphology of the GNRs appear to be hemispherically capped cylinders and the zeta potential of GNRs was +33.3 mV. Moreover, TEM clearly showed that the GNRs/Gd–DTPA–CS nanoparticles prior to PGA-g-mPEG coating have a spherical morphology with a narrow size distribution around 80 nm (Fig. 2B). Interestingly, almost all of the h-NPs were doped with a single GNR. The TG analysis results (Fig. S3†) revealed the GNRs@Gd–DTPA–CS and Gd–DTPA–CS weight loss were estimated to be 59.3% and 68.3% respectively within a temperature range of 100–650 °C. The weight ratio of GNRs over Gd–DTPA–CS was calculated to be 1 : 6.6 (Fig. S3†). The mean hydrodynamic size of GNRs/Gd–DTPA–CS NPs detected by DLS was 117 ± 15 nm. In order to improve the colloidal stability in aqueous solution, the NPs were further functionalized with PGA-g-mPEG. Since the surface of GNRs/Gd–DTPA–CS nanoparticles was positive, anionic PGA-g-mPEG could attach to it by electrostatic interaction. After PGA-g-mPEG coating, zeta potential of the final h-NPs was changed to −16.3 mV from +22.1 mV. However, there was no obvious variation in morphology and size distribution after PEG coating (Fig. 2C). DLS measurement of the final h-NPs is supportive of the TEM results, i.e. only a negligible increase in hydrodynamic size (137 ± 15 nm vs. 117 ± 15 nm) was detected (Fig. S4†).

Fig. 2 TEM images of (A) GNRs; (B) GNR/Gd–DTPA–CS NPs and (C) GNR/Gd–DTPA–CS@10% PEG NPs (h-NPs); (D) stability for the GNRs/Gd–DTPA–CS with various PGA-g-mPEG (m/m) ratio incubated in PBS containing 10% FBS; data are given as mean ± SD (n = 3).

All h-NPs with various PEG contents from 10% to 30% remained their initial mean particle size during the whole incubation period (Fig. 2D), implying high colloidal stability in bloodstream. On the contrary, the mean particle size of GNRs/Gd–DTPA–CS NPs dramatically increased to 1290 nm from 119 nm during the first 2 hours. This quick aggregation of GNRs/Gd–DTPA–CS NPs may be due to their positively charged surface susceptible to protein adsorption. However, the PGA-g-mPEG coating not only increased the hydrophilicity but also reversed the surface to negatively charged, thus significantly enhancing the serum stability of the h-NPs. Admittedly, the high serum stability is especially critical for the h-NPs to be delivered to tumor site in vivo.

The T₁-weighted MR imaging sensitivity of h-NPs were evaluated on a 1.5 T MRI scanner (Fig. 3). The longitudinal relaxivity (r₁) value was 9.71 mM⁻¹ S⁻¹, which was obviously higher than that of the commercially available MRI contrast agent gadopentetate dimeglumine (r₁ = 3.77 mM⁻¹ S⁻¹). The high longitudinal relaxivity of h-NPs is attributed to the conjugated h-NPs which might have restrained the rotation of the Gd–DTPA part. Moreover, the MRI signal intensity was enhanced with the increase of concentration.

Fig. 3 (A) T₁-Weighted MR images of h-NPs and gadopentetate dimeglumine. (B) T₁ relaxation rate against Gd concentration of h-NPs and gadopentetate dimeglumine.

The photothermal properties were further investigated. As shown in Fig. 4A, the light spectra of GNRs had an absorbance peak in the NIR region, which is known to be essential for a desirable tissue penetration.²⁹ In contrast, the h-NPs doped with GNRs showed a slight red shift to 812 nm from 802 nm, most likely due to the high dielectric constant of the hybrid coating materials. Apparently, the h-NPs maintained the optical properties of GNRS. The potential of h-NPs as a PTT agent was then evaluated by determining the temperature variations of the h-NPs solutions at different concentrations which were irradiated by a NIR laser (808 nm, 2 W cm⁻²) for 10 min. In Fig. 4B, an obvious h-NPs concentration-dependent temperature was observed when the solutions were exposed to the laser, which demonstrating that the thermal energy was converted by the h-NPs. For example, at a low h-NPs concentration of 230 μg mL⁻¹, the solution temperature rapidly increased to 50.5 °C from 28.0 °C in 10 min. As a control, water displayed neglectable temperature increase upon NIR laser irradiation, which implies that h-NPs could act as a potent agent converting NIR light absorption to heat for photothermal therapy.

Fig. 4 (A) UV-vis spectra of gold nanorods (GNR), GNR/Gd–DTPA–CS and h-NPs; (B) temperature elevation in aqueous solutions containing of different h-NPs concentrations under NIR laser irradiation (808 nm, 2 W cm⁻²).

Cytotoxicity

As shown in Fig. 5A and B, h-NPs showed no obvious cytotoxicity in HeLa and CT-26 cells for 24 h. When the incubation time was prolonged to 48 h, cells incubated at high concentration of 800 μg mL⁻¹ still showed viabilities above 82.5% in both cell lines, indicating sufficient biocompatibility of h-NPs. Considering that the h-NPs as a photothermal therapy agent would enter the blood circulation after intravenous injection, erythrocyte agglutination and hemolysis experiments were conducted to reveal the blood compatibility. As shown in Fig. 5C, PBS and Triton X-100 were selected as the negative and positive control, respectively. There was no detectable erythrocyte agglutination induced by the h-NPs. In addition, the quantitative hemolysis analysis presented in Fig. 5D to 800 μg mL⁻¹, only approximately 3.5% hemolysis was detected at a prolonged incubation time of 3 h. Obviously, the good biocompatibility of h-NPs is favorable for their in vivo application as photothermal therapy agent.

Fig. 5 MTT assays of h-NPs on (A) HeLa cells and (B) CT-26 cells for 24 h and 48 h; (C) microscopic images (200× magnification) of the erythrocytes after treatment with PBS, Triton X-100 and the h-NPs at 200 μg mL⁻¹ and 800 μg mL⁻¹ (bar = 50 μm); (D) quantitative analysis on hemolytic activity of the h-NPs at different concentrations. Data are given as mean ± SD (n = 3).

The photothermal effect of h-NPs was assessed on HeLa cells. For the AM/PI staining, viable cells and dead cells were stained with calcein-AM and PI respectively to confirm the viability of cells (Fig. 6). Laser or h-NPs alone did not lead to obvious change in cell viability and density as compared with the negative control (Fig. 6A–C). In comparison, HeLa cells treated with h-NPs under NIR laser exposure exhibited substantial cell death (Fig. 6D). Therefore, the above results indicated that the h-NPs could exert photothermal effect to induce cell death only under NIR laser irradiation. MTT assay also showed that the presence of NIR laser irradiation (808 nm, 2 W cm⁻²) is essential for the h-NPs to have remarkable cytotoxicity. The cell viability gradually decreased with increasing concentration of h-NPs. Remarkably, after incubation with h-NPs (400 μg mL⁻¹) under irradiation for 5 min, almost all HeLa cells were effectively ablated (Fig. 6E). These results demonstrated that the h-NPs could serve as an effective photothermal agent to localized kill tumor cells in vitro.

Fig. 6 Photothermal ablation of HeLa cells: (A) control, (B) NIR laser only (808 nm, 2 W cm⁻²), (C) h-NPs only, (D) h-NPs combined NIR laser treatments (bar = 500 μm); (E) cell viability of h-NPs with or without NIR laser irradiation for 5 min.

MRI-guided photothermal therapy in vivo

To confirm the MR imaging capability in vivo, the T₁-weighed MR images of mice bearing CT-26 tumor were recorded before and after intravenous injection of h-NPs. Upon injection of h-NPs, a dramatic lighting effect in the tumor was observed, owing to the decreased spin-lattice relaxation time of proton caused by the paramagnetic Gd ion. As shown in Fig. 7, 6 h after IV injection, the maximal accumulation signal of h-NPs in tumor revealed the optimal time for implementation of tumor ablation. The gradually increased signal was attributed to the tumor accumulation of h-NPs through the EPR effect²⁸ and its good colloidal stability in the bloodstream. Interestingly, the tumor site revealed the high signal after 6 h, demonstrating longer retention of h-NPs compared with commercial MR contrast agent. Apparently, the tumor accumulation of h-NPs could be monitored with MRI to guide the NIR laser irradiation for photothermal therapy.

Fig. 7 (A) T₁-Weighted MR images of CT-26 tumor-bearing mice before and after IV injection of h-NPs for different time points, (B) the average grey values of T₁ weighted signals from the tumor at various time post h-NPs IV injection.

As h-NPs effectively accumulated in tumor site 6 h after IV injection, the mice were exposed to 808 nm laser irradiation for 10 min at this time point. As shown in Fig. 8, the average temperature of tumor in the h-NPs plus laser group rapidly increased to approximate 51.0 °C from 37.1 °C within 6 min, which was enough to ablate tumor in vivo. In contrast, saline group only increased the tumor area for 3.6 °C within 10 min, which further reveals that the NIR laser was harmless to the tested mice.

Fig. 8 (A) Infrared thermal images of CT-26 tumor-bearing mice with saline or with h-NPs (800 μg mL⁻¹, 6 h post IV injection) under 808 nm laser irradiation for 10 min; (B) heating curves in tumor site of tumor-bearing mice treated with saline and h-NPs.

As shown in Fig. 9A, tumors in the saline group, saline + laser group and h-NPs group grew similarly fast, and no animal survived longer than 23 days, suggesting that injection of h-NPs or laser exposure alone could not affect tumor growth. In comparison, the h-NPs with laser group displayed a remarkable decrease in tumor size and prolonged animal survival time. No death of mice or was recorded during the course of photothermal therapy over 40 days. In particular, complete tumor regression was observed after 40 days of treatment. As shown in Fig. S5,† laser exposure alone could not inhibit tumor growth, whereas the tumors were completely ablated in the treatment group, leaving black scars which recovered after approximately 14 days. In the hematoxylin and eosin (H&E) staining, apparent intensive nuclear shrinkage appeared in the tumor sections of animals from the laser with h-NPs group (Fig. 9B). In contrast, the histological sections of the three other groups all exhibited infiltrating tumor cells with pleomorphic nuclei, indicating limited benefit from h-NPs or laser treatment only.

Fig. 9 (A) Growth of CT-26 tumors in different groups of mice after various treatments (n = 4); (B) H&E stained images of tumors from different groups of mice 1 day post treatment; (C) body weight change of mice in different groups after treatment (n = 4).

To investigate the potential toxicity of h-NPs, mice from treated group and negative control group were sacrificed at the day 40 to collect major organs for H&E staining. As shown in Fig. S6,† H&E staining images of both the treatment and control groups revealed no obvious organ damage or inflammatory lesion in major organs, suggesting the safety of photothermal treatment in tumor-bearing mice. No obvious body weight loss was noted as well during 40 days (Fig. 9C). These results clearly indicated that the h-NPs might be an excellent theranostic agent combing good therapeutic effect and bio-safety vital for MR-guided photothermal therapy of tumor.

Conclusions

A multifunctional theranostic agent h-NPs has been successfully developed for MR-guided photothermal therapy using near-infrared (NIR) irradiation. The h-NPs possessed uniform size distribution, good physiological stability, enhanced MR imaging contrast, and efficient photothermal conversion. Under the guidance of MR imaging, an effective tumor ablation was achieved by applying the NIR light irradiation at proper time after the IV injection of h-NPs. Excitingly, almost a complete regression of CT-26-tumor after ablation was observed within 40 days. Moreover, no obvious side effect was detected at the investigated doses for the h-NPs. Thus, this work highlights the potential of h-NPs as a theranostic nanoplatform for MRI-guided photothermal therapy of cancer.

Acknowledgements

This work was supported by the National Natural Science Foundation of China (U1401242, 81101142, 81371607, 81571739), the Tip-top Scientific and Technical Innovative Youth Talents of Guangdong Special Support Program (2015TQ01R510) and Natural Science Foundation of Guangdong province, China (2016A030313324).

References

1. S. Lal, S. E. Clare and N. J. Halas, Acc. Chem. Res., 2008, 41, 1842–1851.
2. E. Boisselier and D. Astruc, Chem. Soc. Rev., 2009, 38, 1759–1782.
3. C. Yue, P. Liu, M. Zheng, P. Zhao, Y. Wang, Y. Ma and L. Cai, Biomaterials, 2013, 34, 6853–6861.
4. A. M. Gobin, M. H. Lee, N. J. Halas, W. D. James, R. A. Drezek and J. L. West, Nano Lett., 2007, 7, 1929–1934.
5. J. Chen, D. Wang, J. Xi, L. Au, A. Siekkinen, A. Warsen, Z.-Y. Li, H. Zhang, Y. Xia and X. Li, Nano Lett., 2007, 7, 1318–1322.
6. H. Liu, T. Liu, H. Wang, L. Li, L. Tan, C. Fu, G. Nie, D. Chen and F. Tang, Biomaterials, 2013, 34, 6967–6975.
7. P. Chakravarty, R. Marches, N. S. Zimmerman, A. D.-E. Swafford, P. Bajaj, I. M. Musselman, P. Pantano, R. J. Draper and E. S. Vitetta, Proc. Natl. Acad. Sci. U. S. A., 2008, 105, 8697–8702.
8. G. Yang, C. Zhu, D. Du, J. Zhu and Y. Lin, Nanoscale, 2015, 7, 14217–14231.
9. H. Chen, L. Shao, Q. Li and J. Wang, Chem. Soc. Rev., 2013, 42, 2679–2724.
10. I. P. Lau, H. Chen, J. Wang, H. C. Ong, K. C. Leung, H. P. Ho and S. K. Kong, Nanotoxicology, 2012, 6, 847–856.
11. L. Cheng, J. Liu, X. Gu, H. Gong, X. Shi, T. Liu, C. Wang, X. Wang, G. Liu, H. Xing, W. Bu, B. Sun and Z. Liu, Adv. Mater., 2014, 26, 1886–1893.
12. G. A. Sotiriou, F. Starsich, A. Dasargyri, M. C. Wurnig, F. Krumeich, A. Boss, J.-C. Leroux and S. E. Pratsinis, Adv. Funct. Mater., 2014, 24, 2818–2827.
13. C.-L. Peng, Y.-H. Shih, P.-C. Lee, T. M.-H. Hsieh, T.-Y. Luo and M.-J. Shieh, ACS Nano, 2011, 5, 5594–5607.
14. M. K. Yu, D. Kim, I.-H. Lee, J.-S. So, Y. Y. Jeong and S. Jon, Small, 2011, 7, 2241–2249.
15. Z. Cao, W. Zhu, W. Wang, C. Zhang, M. Xu, J. Liu, S. T. Feng, Q. Jiang and X. Xie, Int. J. Nanomed., 2014, 9, 5103–5116.
16. G. B. Frisoni, N. C. Fox, C. R. Jack, P. Scheltens and P. M. Thompson, Nat. Rev. Neurol., 2010, 6, 67–77.
17. H. J. Weinmann, R. C. Brasch, W. R. Press and G. E. Wesbey, Am. J. Roentgenol., 1984, 142, 619–624.
18. Y. Huang, B. Cao, X. Yang, Q. Zhang, X. Han and Z. Guo, Magn. Reson. Imaging, 2013, 31, 604–609.
19. K. S. Kim, W. Park, J. Hu, Y. H. Bae and K. Na, Biomaterials, 2014, 35, 337–343.
20. C. E. Smith, A. Shkumatov, S. G. Withers, B. Yang, J. F. Glockner, S. Misra, E. J. Roy, C.-H. Wong, S. C. Zimmerman and H. Kong, ACS Nano, 2013, 7, 9599–9610.
21. A. J. Coughlin, J. S. Ananta, N. Deng, I. V. Larina, P. Decuzzi and J. L. West, Small, 2014, 10, 556–565.
22. Q. Tian, J. Hu, Y. Zhu, R. Zou, Z. Chen, S. Yang, R. Li, Q. Su, Y. Han and X. Liu, J. Am. Chem. Soc., 2013, 135, 8571–8577.
23. E. V. Shevchenko, M. I. Bodnarchuk, M. V. Kovalenko, D. V. Talapin, R. K. Smith, S. Aloni, W. Heiss and A. P. Alivisatos, Adv. Mater., 2008, 20, 4323–4329.
24. D.-W. Wang, X.-M. Zhu, S.-F. Lee, H.-M. Chan, H.-W. Li, S. K. Kong, J. C. Yu, C. H. K. Cheng, Y.-X. J. Wang and K. C.-F. Leung, J. Mater. Chem. B, 2013, 1, 2934–2942.
25. X. Ye, C. Zheng, J. Chen, Y. Gao and C. B. Murray, Nano Lett., 2013, 13, 765–771.
26. Z. Yang, Z. Jiang, Z. Cao, C. Zhang, D. Gao, X. Luo, X. Zhang, H. Luo, Q. Jiang and J. Liu, Nanoscale, 2014, 6, 10193–10206.
27. R. Guo, L. Zhang, H. Qian, R. Li, X. Jiang and B. Liu, Langmuir, 2010, 26, 5428–5434.
28. R. Weissleder, Nat. Biotechnol., 2001, 19, 316–317.
29. A. K. Iyer, G. Khaled, J. Fang and H. Maeda, Drug Discovery Today, 2006, 11, 812–818.

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Footnote

† Electronic supplementary information (ESI) available. See DOI: 10.1039/c6ra23769j

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