A New Photosensitized Oxidation-Responsive Nanoplatform for

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Biological and Medical Applications of Materials and Interfaces

A new photosensitized oxidation-responsive nanoplatform for controlled drug release and photodynamic cancer therapy Huan-Pu Yeh, Andrea C. del Valle, Ming-Chen Syu, Yu Qian, Yu-Cheng Chang, and Yu-Fen Huang ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.8b05205 • Publication Date (Web): 04 Jun 2018 Downloaded from http://pubs.acs.org on June 4, 2018

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ACS Applied Materials & Interfaces

A new photosensitized oxidation-responsive nanoplatform for controlled drug release and photodynamic cancer therapy Huan-Pu Yeh, Andrea C. del Valle, Ming-Chen Syu, Yu Qian, Yu-Cheng Chang, and Yu-Fen Huang*

Department of Biomedical Engineering and Environmental Sciences, National Tsing Hua University, Hsinchu, Taiwan, ROC

The ﬁrst two authors contributed equally to this work.

Keywords: Controlled drug release, Photodynamic therapy, Photosensitized oxidation, Stimuli-responsive polymer, Multidrug resistance

ABSTRACT

Abnormal biochemical alteration such as unbalanced reactive oxygen species (ROS) levels has been considered as a potential disease-specific trigger to deliver therapeutics to target 1

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sites. However, in view of their minute variations in concentration, short lifetimes, and limited ranges of action, in situ generation of ROS with specific manipulations should be more effective for ROS-responsive drug delivery. Here we present a new delivery nanoplatform for photodynamic therapy (PDT) with on-demand drug release regulated by light irradiation. Rose bengal (RB) molecules, which exhibit a high yield of ROS generation, were encapsulated in a mixture of chitosan (CTS), polyvinyl alcohol (PVA), and branched polyethylenimine (bPEI) with hydrophobic iron oxide nanoparticles through an oil-in-water emulsion method. The as-prepared magnetic nanoclusters (MNCs) with a tri-polymer coating displayed high water dispersibility, efficient cellular uptake, and the cationic groups of CTS and bPEI were effective for RB loading through electrostatic interaction. The encapsulation efficiency of RB in MNCs could be further improved by increasing the amount of short bPEI chains. During the photodynamic process, controlled release of the host molecules (i.e. RB) or guest molecules (i.e. paclitaxel) from the bPEI-based nanoplatform was achieved simultaneously through a photooxidation action sensitized by RB. This approach promises specific payload release and highly effective PDT or PDT combined therapy in various cancer cell lines including breast (MCF-7 and multidrug resistant MCF-7 subline), SKOV-3 ovarian and Tramp-C1 prostate. In in vivo xenograft studies, the nanoengineered light-switchable carrier also greatly augments its PDT efficacy against multidrug resistant MCF-7/ADR tumor as compared with free drugs. All the above findings suggest that the substantial effects of enhanced drug distribution for efficient cancer therapy was achieved with this smart nanocarrier capable of on demand drug release and delivery, thus exerting its therapeutic activity to a greater extent.

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INTRODUCTION

Smart polymers that are susceptible to stimuli and intrinsic to tumor environments have emerged as efficient drug and gene carriers in recent years. Most materials have been developed with specific responses to biological abnormalities, such as a low extracellular tumor pH value and aberrant expression of protease. Reactive oxygen species (ROS)-sensitive materials are still relatively new, but they have recently begun gaining importance.1-4 Since excessive ROS production and elevated oxidative stress are critical biochemical alterations that greatly increase cancer cell growth, tumor cells generate ROS, including H2O2, hydroxyl radical, and superoxide, at a higher rate than healthy cells.5 ROS are considered to be unique cancer-related stimuli that could potentially mediate intracellular therapeutics delivery. The first ROS-sensitive biomaterial was developed by Hubbell et al. in 2004 for drug delivery applications.6 A block copolymer comprised of hydrophilic polyethylene glycol (PEG) and hydrophobic polypropylene sulfide (PPS), can undergo the phase transition from hydrophobic sulfide to hydrophilic sulfoxide or sulfone in an oxidative environment. The rapid solubility change of PPS in response to ROS makes it an ideal material for oxidation-dependent drug release and vaccine applications.7-9 In addition, a new class of ROS-responsive cationic water-soluble polymer has been developed for gene delivery by the Murthy group.10 This polymer is composed of biodegradable thioketal linkages that are readily cleavable in ROS-abundant conditions. The degradation property of thioketal specific to high levels of ROS allowed great improvement in gene transfection and intracellular trafficking.11 Despite the advantages mentioned above, an efficient triggered response remains a major challenge, especially for the minimal amount of ROS produced endogenously. Stimuli that can be controlled externally are thereby increasingly studied for biomedical applications. 3



Photodynamic therapy (PDT) is a technology that utilizes light to activate photosensitizer (PS) in the target site, leading to a signification accumulation of ROS.12 It was hypothesized that the combination of PS and an ROS-responsive polymer could be developed as an effective carrier to facilitate intracellular payload release in response to light irradiation.9, 13-15 On the other hand, ROS have a short lifespan and limited diffusion distance. The subcellular localization of PS is therefore, critical in PDT-mediated antitumor activities.16 An activatable nanoplatform that is designed for efficient payload release and subsequent intracellular redistribution during light irradiation is a promising strategy to elicit effective PDT action at the target site. In light of an efficient photooxidation of tertiary amines sensitized by Rose Bengal (RB),17 we reported herein the discovery of a branched polyethylenimine (bPEI) as an innovative ROS-responsive material. RB, which exhibits high ROS yield, is a well-known anionic photosensitizer in PDT.18

bPEI, which contains primary, secondary, and tertiary

amines in a ratio of 1:2:1, provides multiple positive charges over a wide pH range (pKa values: 4.5, 6.7, and 11.6).19 At physiological conditions, a notable amount of RB was adsorbed onto bPEI electrostatically. The strength of the interaction was, however, greatly reduced as the ROS-mediated oxidation gives less positively charged products. bPEI, which serves as a photoinducible switch, drives the specific release of RB molecules in response to photosensitized oxidation. To the best of our knowledge, this is the first study presenting a photooxidation process sensitized by RB that can be used to control the release of loaded cargos from a bPEI-based nanoplatform.

RESULTS AND DISCUSSION

To demonstrate the photoswitchable concept, a core-shell nanocarrier consisting of magnetite nanoparticle clusters coated with polymer blends was prepared via an oil-in-water emulsion approach (Figure 1A). Iron oxide nanoparticles (Fe3O4 NPs) with an average size of 4


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7.5 ± 1.5 nm were synthesized initially through a thermal decomposition process in an oil phase (Figure S1, Supporting Information). The X-ray powder diffraction (XRD) pattern of the resulting NPs also revealed a cubic spinel structure of magnetite. The polymer blends of chitosan (CTS), polyvinyl alcohol (PVA), and low molecular weight bPEI (1.8 kDa) constituting the shell layer can serve as an effective coating to stably encapsulate Fe3O4 NPs into the emulsion droplets. After the gradual removal of residual organic solvent by evaporation, magnetic nanoclusters (MNCs) were successfully produced. The size distributions and morphologies of the resulting MNCs in aqueous solution were characterized by dynamic light scattering (DLS) and transmission electron microscopy (TEM) analysis. As depicted in Figure 1B, CTS-functionalized MNCs (a), which exhibited a uni-modal size distribution of 325.2 ± 75.7 nm, tend to aggregate on the TEM grid (Figure 1C), owing to the high viscosity and low water solubility of CTS. To enhance the dispersion capability of MNCs, a hydrophilic and water-soluble polymer, PVA was chosen as a co-emulsifier to provide better colloidal stability. MNCs (b) with a narrow size distribution of 100.7 ± 16.7 nm were obtained when the weight ratio of CTS to PVA was 10:2. A decrease in surface charges of MNCs from 47.0 ± 6.9 to 28.7 ± 6.6 mV is depicted in Table 1 and occurred upon raising the weight percentage of PVA. To achieve a high entrapment efficiency of our model drugs in MNCs, short bPEI chains were also used to compose the polymeric shells. As compared to those lacking bPEI, the surface charges of MNCs coated with tri-polymer blends become more positive; the resultant MNCs (c) at a specific weight ratio of CTS, PVA, bPEI, and Fe3O4 (10:2:4:1) exhibited a high zeta potential value of 42.8 ± 6.3 mV. In addition, the increase in the hydrodynamic size (152.3 ± 5.8 nm) of the produced MNCs was likely attributable to the increasing numbers of magnetite nanoparticles flocculated into cluster structures. As confirmed by TEM (Figure 1C), the average diameters of MNCs (b) and (c) were 35.9 ± 5.7 and 55.4 ± 7.0 nm, respectively. It should be noted that MNCs coated with a 5



polymer mixture of PVA and bPEI (d) were also prepared; however, they were found to be relatively unstable over a long period of time (data not shown). Based on the above results, MNCs (c) prepared at a weight ratio of 10:2:4:1 were chosen for further studies. It also worthy to note that the polydispersity index (PDI) value of MNCs (c) is approximately 0.2 (Table 1), indicating a nearly monodisperse size distribution, as required for an effective drug delivery nanoplatform. Next, the loading capability of photosensitizers encapsulated into different MNCs was characterized by UV-Vis spectra in Figure 2A. After removal of excess reagents by washing twice with phosphate buffer (PB, pH 7.4), a greater absorbance of RB molecules was observed in RB-loaded MNCs (c) than those lacking bPEI (b). The increase in absorption at shorter wavelengths also indicates the existence of colloidal MNC structures. Considering the more positive zeta potential of MNCs (c) than MNCs (b), the above results imply that the incorporation of RB into the polymeric coating was primarily through the electrostatic interaction. The subsequent loading of drug molecules into MNCs (c) and (b) also led to a decline in the positive value of the zeta potential (37.1 ± 3.4 and 21.8 ± 6.1 mV, respectively). In addition, the original absorption maximum of RB is at 545 nm at physiological pH. An observation of a bathochromic shift of approximately 15 to 20 nm, related to RB molecules after entrapment, further confirms the contribution of electrostatic effects originating from the surrounding environment.20 Finally, the encapsulation efficiency (EE, %) of RB into MNCs (b) and MNCs (c) was also compared with different washing conditions. As displayed in Figure S2A (Supporting Information), MNCs (b) revealed a more pronounced drug loading capacity at pH 5.3 (EE% = 81 ± 5%) than at pH 7.4 (EE% = 39 ± 3%). However, the EE% value of MNCs (c) with tri-polymer coating, remained almost unchanged (approximately 87%). These findings, suggest that bPEI is, in fact, the leading

6


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cause of effective drug loading at physiological pH. This is consistent with the relatively higher charge densities of bPEI (pKa = 7.4−8.5) in comparison to CTS (pKa ~ 6.3).21-22 The stability of RB-loaded MNCs (c) [RB: MNCs (c)] in complex biological conditions was further studied by DLS measurements. Except for those in a DMEM culture medium, RB: MNCs (c) displayed high stability in various physiological environments (Figure S2B, Supporting Information). Agglomeration of drug nanocarriers was not observed even after incubation for 72 h in both Dulbecco's phosphate-buffered saline (DPBS) and a cell culture medium containing 10% fetal bovine serum (FBS/DMEM) (data not shown). Also, there has been a remarkable reduction in the hydrodynamic sizes of MNCs (c) after drug encapsulation; the sizes of MNCs (c) and RB: [MNCs (c)] in PB at pH 7.4 were 169.8 ± 15.6 nm and 108.6 ± 11.7 nm, respectively. The condensation of polymeric shells through the incorporation of trapping molecules of opposite charges further confirms the long-term stability of RB: [MNCs (c)] in high salinity conditions. Similarly, MNCs (c) were found to be more stable in serum containing culture medium. The adsorption of serum proteins onto the surface of the nanoclusters also greatly improved the colloidal dispersibility in complex DMEM. It should be noted that the residency of Fe3O4 NPs can promote the emulsification or dispersion of the oil droplets into an aqueous phase, and therefore the production yield of nanoparticles constituted by RB: [MNCs (c)] was significantly improved compared with its non-Fe3O4 NPs counterpart, RB: [CTS/PVA/bPEI]. A larger agglomerate but a less intense scattering intensity (represented as the derived count rate) was detected from RB: [CTS/PVA/bPEI] (Table S1, Supporting Information). A reduced RB encapsulation efficiency revealed by a light purple color of the corresponding redispersion (b) in Figure 2B also suggests that the encapsulation of Fe3O4 NPs is beneficial to the success of nanoformulations for effective photooxidation-mediated drug delivery. Moreover, the 7



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field-dependent magnetization curves of MNCs (c) was recorded using a superconducting quantum interference device (SQUID) magnetometer, displaying a superparamagnetic behavior with no hysteresis loop observed at 300 K (Figure S3A, Supporting Information). The saturation magnetization (Ms) values of MNCs (c) and the oleylamine-coated Fe3O4 NPs were nearly the same (59.43 and 61.43 emu/g, respectively), while the average polymer content constituting to MNCs (c) was approximately 48.7% examined by thermogravimetry (Figure S3B, Supporting Information). As expected, MNCs (c) could also be used as a negative

contrast

agent

in

magnetic

resonance

(MR)

imaging,

providing

a

concentration-dependent darkness in signal intensity (Figure S3C, Supporting Information). A remarkably higher transverse relaxivity r2 value (241 s-1 mM-1 Fe) was observed for MNCs (c) over FDA-approved contrast agent (Feridex, 142 s-1 mM-1 Fe), further suggesting a substantial advantage of MNCs (c) as a more effective contrast enhancer in T2-weighted MR imaging, particularly at reduced doses. In order to investigate the photoresponsive characteristics of the developed nanoplatform, RB-loaded MNCs (c) in a DPBS were subjected to light illumination with their energy transmitted through a red filter with a center wavelength of 632 nm. The supernatant of each colloidal sample with different exposure periods was collected and analyzed by fluorescence spectrophotometry

to determine the concentration of RB in the release media. As shown in

Figure 3A, an obvious release behavior of loaded drugs was observed for RB: [MNCs (c)] irradiated with red light (15 mW/cm2). About 30% of RB was released within 2 h. In contrast, the leakage of drugs was less than 8% for the sample that was kept in dark conditions for the same period of time. These results confirmed that the constructed nanoplatform was able to reduce non-specific drug action but allow an on-demand release of therapeutic payloads upon light irradiation. In addition, bPEI, which constitutes the shell layer of MNCs (c), was further compared with its linear counterpart (PEI, 2.0 kDa). Unlike with bPEI, a limited intensity 8


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change under a similar irradiation condition was obtained from the supernatants of RB-loaded MNCs coated with CTS/PVA/PEI [RB: MNCs (e)]. It should also be noted that the amount of loaded drug was approximately the same in both nanocarriers, but the linear PEI led to a greater extent of RB leakage. The results reported herein provide additional support for the conclusion that bPEI is crucial to the controlled release process triggered by light illumination. We confirm that the triggered release of loaded drugs was activated by the photodynamic reaction. ROS generation was monitored using a oxidation-sensitive fluorescent probe, 2’,7’-dichlorodihydrofluorescein diacetate (DCFH-DA). As demonstrated in Figure S4 (Supporting Information), ROS rapidly accumulated in RB-loaded MNCs (c) following a 60-min irradiation. However, no appreciable fluorescence change was observed in the control groups without light treatment. Additionally, no production of ROS was expected for MNCs lacking photosensitizers after light exposure (data not shown). Since the electrostatic attraction is considered to be the driving force to encapsulate RB into MNCs (c), a significant difference in surface charge density would be expected during the photodynamic action (Figure 3B). Upon light illumination, an apparent decrease in normalized zeta potential of RB-loaded MNCs (c) occurred (Figure S5A, Supporting Information), whereas the values of those lacking drugs remained at the original level. Additionally, the presence of bPEI also led to a ROS-dependence of zeta potential determined from the drug nanosuspensions. An obvious loss of surface charge density had been observed for RB: [MNCs (c)], while the relative values of those without bPEI (b) decreased slightly with the subsequent addition of H2O2 (Figure S5B, Supporting Information). It is widely accepted that H2O2 can react with Fe(II) originating from magnetite nanoparticles to generate ROS via a Fenton reaction.23 In accordance with the previous observations in Figure 2B, these results collectively suggest

9



that ROS can facilitate the release behavior of drugs from bPEI-containing MNCs through the alteration of electrostatic interactions. Subsequently, UV-Vis spectroscopy was further conducted to study the release mechanism involved in the photodynamic process (Figure S6A, Supporting Information). The spectrum features of RB molecules revealed obvious blue shifts in the nanosuspensions of RB: [MNCs (c)], confirming the strength of electrostatic interactions varied during light irradiation. No change was observed in control sample with free drugs under the same illumination conditions (data not shown). Moreover, the photoproducts of the exposed sample mixture including RB and bPEI (1.8 kDa) were corroborated by proton nuclear magnetic resonance (1H-NMR) (Bruker Avance 500 NMR spectrometer, Bruker, Germany). As shown in Figure S7A (Supporting Information), there was an appreciable decrease in the integrated peak intensity of signals appearing around 2.5–2.7 ppm,24 which corresponding to the CH2 groups next to primary, secondary and tertiary amines of RB-treated bPEI followed by light irradiation. However, those signals of bPEI itself remained almost unchanged under similar exposure condition (Figure S7B, Supporting Information). A noticeable decrease in the signal intensity was also observed for bPEI reacted with H2O2 (sample c in Figure S7B, Supporting Information), indicative of the degradation of amines in bPEI through an oxidation-mediated cascade reaction. This result is consistent with the previous findings that amino groups, especially tertiary amines are susceptible to photooxidation sensitized by RB.17 The exposed amines, which give possible products of carbonyl compounds and imines, could greatly inhibit the electrostatic attraction between drugs and nanocarriers, leading to a controllable payload release in response to light irradiation. To ensure a more effective PDT performance, light-assisted drug release study was further carried out using laser irradiation at a wavelength of 532 nm, which is in close proximity to the maximum absorbance of RB.16, 25-27 As shown in Figure 3C, the release rate can be 10


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greatly accelerated (approximately 85% within 15 min) when switching to a 532-nm green laser light at the same irradiation power (15 mW/cm2). UV-Vis spectra in Figure 3D also confirmed that the irradiated (+) RB: [MNCs (c)] was featured with a blue shift in absorption peak as compared to its non-irradiated counterpart (-) RB: [MNCs (c)], suggesting the involvement of electrostatic interactions in the cargo entrapment/release behavior.

Next,

both samples were submitted to centrifugation to separate the supernatant liquid (sup) from the precipitate (pre). The UV-Vis (Figure 3E) and fluorescence (Figure 3F–3G) spectra corresponding to the individual components were identified. As shown in Figure 3F, the fluorescence signal of RB quenched by MNCs (c) was restored followed by light activation. The sup and pre collected from the irradiated (+) RB: [MNCs (c)] also gave a steep fluorescence increase, whereas those of non-irradiated (-) RB: [MNCs (c)] produced only a limited fluorescence signal (Figure 3G). Similarly, irradiated (+) RB: [MNCs (c)] was found to be more active for ROS generation in response to additional light irradiation (Figure 3H). Altogether, these results corroborate the advantage of photo-activatable payload release in subsequent ROS production and it is necessarily beneficial for enhanced PDT-mediated therapeutic outcome. To illustrate the PDT process inside living cells (MCF-7, human breast cancer cells), the intracellular ROS formation was then studied using DCFH-DA upon red light exposure. After 1 h of irradiation, cells treated with RB-loaded MNCs (c) displayed an obvious fluorescence increase (3.2-fold) versus non-irradiated counterpart (column 4 in Figure 4A). However, no statistical significant difference in ROS production was observed inside cells incubated with MNCs lacking photosensitizers (p = 0.11, column 3). Additionally, both untreated cells (p = 0.15, column 1) and RB-treated cells (p = 0.06, column 2) exhibited negligible fluorescence changes following light irradiation. This result suggests a superior performance of the uptake drug nanocarriers in photodynamic action than that of free drugs. A similar phenomenon was 11



also observed in the corresponding fluorescence microscopic imaging (Figure S8, Supporting Information). Hydrogen peroxide (H2O2, 300 µM) was used as a positive control to induce intracellular ROS generation. Moreover, the effectiveness of intracellular drug release triggered by PDT was also investigated by flow cytometry (Figure 4B). Since the fluorescence intensity of RB molecules was significantly quenched by MNCs, the recovery of RB signals was utilized to monitor the in vitro drug release. As compared with all other controls, a considerable increase in RB fluorescence (1.2-fold, p < 0.01) was detected from cells incubated with RB: MNCs (column 4) after light treatment. In addition, fluorescence images also suggest an abruptly increasing brightness of RB signals inside cells under the same experimental conditions, whereas no obvious fluorescence change was observed in the case of RB-treated cells with red light irradiation (Figure S9, Supporting Information). Collectively, these results made it quite evident that RB molecules could be liberated from the loaded nanocarriers in response to the photodynamic process. A similar phenomenon could also be observed for treated cells submitted to a green laser irradiation (data not shown). It is also worth noting that in Figure 4B, a more intense RB signal (10.4 a. u.) was obtained from cells treated with drug nanocarriers than from those incubated with free RB (3.9 a. u.). This finding correlates well with the substantial differences in red color fluorescent spots between two treated cells (Figure S9, Supporting Information). The apparent blue granules, visualized by cells labeled with RB: [MNCs (c)] followed by Prussian blue staining (BF channel), also suggest an enhanced cellular uptake of hydrophilic drugs through MNC formulation. Detailed information about intracellular RB liberation from the drug nanocarriers under light exposure was further deconvoluted using confocal microscopy (Figure 5A). After 6 h of incubation, the fluorescence of RB was primarily detected within the endosomal compartment, colocalized with fluorescent transferrin (transferrin-Alexa633) in cultured 12


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MCF-7 cells. This result agrees well with previous findings that nanocarriers in the size range of 100–200 nm are preferably taken up by cells through the endosomal pathway.28 When exposed to red light, the fluorescence signals emitted from RB molecules were greatly enhanced and spread out into the cytosol inside the cells. Since RB is relatively membrane-impermeable, the observation of a remarkable difference in drug distribution further implies the occurrence of an intracellular trafficking event during the photodynamic reaction. In addition, the therapeutic efficacy of our developed RB-loaded MNCs (c) toward MCF-7

cell

lines

was

evaluated

by

performing

the

3-(4,5-dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)-2-(4-sulfophenyl)-2H-tetrazolium) (MTT) assay. Before light exposure, a minute toxic effect (< 12 ± 11%) was found in cells incubated with drug nanocarriers (Figure 5B). However, a dramatic increase in cell death up to 72 ± 3% was observed, followed by 1 h of light illumination. In addition, the cytotoxic effects of bare MNCs (c) and RB-loaded MNC (e) were also assessed at equal Fe and RB concentrations. No cellular damage had been involved in both treatments under the same illumination conditions, indicating the potency was mainly attributable to the phototoxic events generated by RB released from the bPEI-based formulation. Moreover, the relative viability of cells incubated with free RB and light treatment was also examined. It was observed that less than 21 ± 4% of MCF-7 cells were killed at the maximum concentration tested. We also note that a similar phenomenon could be found in MCF-7 cells after exposure to a green laser light (Figure S10A, Supporting Information). The effectiveness of the developed nanocarrier in PDT was furthermore verified with two additional cell lines, including mouse prostate cancer cell lines Tramp-C1 and human ovary cancer cell line SKOV-3 (Figure S10B–10C, Supporting Information). As expected, RB: [MNCs (c)] without irradiation showed nonsignificant effects on both cells up to the highest tested dose, whereas the cell viabilities of treated cells were dramatically reduced after an exposure to 13



green laser (22.6 ± 8.1 % of Tramp-C1 and 31.1± 14.8% of SKOV-3 cells, respectively). Based on the above mentioned results, our drug nanocarrier suggests a superior cytotoxicity profile than the parent PDT agent against several types of cancer. Next, RB: [MNCs (c)] was encapsulated with a model drug paclitaxel (PTX) to explore their application potential for controlled drug release as well as a combined photodynamicand chemo-therapy in response to a photooxidation action sensitized by RB. PTX was chosen in view of its established efficacy of first-line chemotherapy for cancer, but its poor aqueous solubility and indiscriminate accumulation in normal tissues remain clinical challenges.29 Basically, PTX was successfully encapsulated into the nanocarrier via the same emulsion/solvent evaporation procedure. The encapsulation efficiency of RB and PTX in the dual drug-loaded nanocarrier remained equivalent to its original drug counterpart and was found to be 74.8 ± 1.6% and 40.2 ± 4.6%, respectively (Table S2, Supporting Information). No negative effects on the physicochemical properties (e.g. UV-Vis) were observed for the resulting nanoformulation that carries two types of therapeutic payloads (Figure S11A, Supporting Information). In response to a green laser irradiation (15 mW/cm2), approximately 85% cumulative RB release was reached for RB/PTX: [MNCs (c)] after 15 min light exposure, offering a similar release profile to that observed with RB: [MNCs (c)] (Figure 3C). More importantly, irradiated nanocarrier also allowed its guest drug cargos, PTX to be released in a controlled manner (Figure 11B, Supporting Information). A 31.5 ± 0.7% PTX release was detected from RB/PTX: [MNCs (c)], whereas the drug leakage was negligible in the dark. Nonspecific payload release was not present in PTX: [MNCs (c)] (+/-) during the observation period either. Taken together, the developed nanoplatform was efficient in stimuli-responsive release of encapsulated drugs, ascribing to the photodynamic action sensitized by RB. In vitro studies also confirmed the cytotoxic effects of a combination of PDT and PTX on MCF-7 cells (Figure 5C). The result indicates that the cell viability in 14


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RB/PTX: [MNCs (c)] (+) group was significantly decreased (22.2 ± 9.8%) when compared with that of RB: [MNCs (c)] (+) (57.5 ± 9.6%) and PTX: [MNCs (c)] (+) (95.4 ± 7.5%) at the highest tested iron concentration. The observation of a superior therapeutic efficacy of RB/PTX: [MNCs (c)] (+) over the non-irradiated counterpart (95.9 ± 8.9%) further suggests that the current approach is promising for on-demand dual-payload release, enabling simultaneous photodynamic and chemo treatments for successful eradication of cancer cells. In clinics, multidrug resistance (MDR) is one major limiting factor for effective chemotherapy against many types of cancer. It has been reported that the MDR phenotypes are particularly resistant to conventional anticancer drugs such as doxorubicin and cisplatin, owing to the positive correlation with the over-expression level of drug efflux transporters such as the P-glycoprotein.30 Drug-resistant human breast cancer cells, MCF-7/MDR, were therefore chosen for the subsequent in vivo study to demonstrate the PDT-based strategy is promising for overcoming MDR. As shown in Figure 6A, RB: [MNCs (c)] displayed a similar cytotoxic action to MCF-7/MDR cells as that towards the parental MCF-7 cells (Figure S10A, Supporting Information) when cells were exposed to a green laser light (100 mW/cm2, 5 min). Moreover, in agreement with the aforementioned result, a remarkable fluorescence recovery of RB signals was also found in MCF-7/MDR cells after irradiation (Figure S12, Supporting Information). These results collectively suggest the potential effectiveness of RB: [MNCs (c)] (+) to combat tumor growth in MDR xenografts. As expected, a significant growth arrest of tumor was observed in nude mice after single intratumoral injection of the developed therapeutic nanoagent, followed by a 10-min green laser irradiation (100 mW/cm2) at 4 h post injection (Figure 6B). RB: [MNCs (c)] (+) significantly suppressed tumor growth (p < 0.05) over PBS (-) control tumors, as compared to the other treatment of RB: [MNCs (c)] (-), [MNCs (c)] (+) and RB (+), respectively. PBS (+) compared with PBS (-) also revealed a neglectable change in tumor size, suggesting the 15



induced non-specific photothermal effect was insufficient in the current regimen. Moreover, the tumors were photographed (Figure 6C), harvested and weighed (Figure 6D) at day 38 from the tumor-bearing mice administered different therapeutics. Accordingly, our findings support that the mice who received photodynamic treatment using RB: [MNCs (c)] (+) produced more significant tumor inhibition (approximately 85%) versus PBS (-) than the other groups. We also note that the free RB (+) at an equivalent dose of 124 µg/mouse demonstrated a limited efficacy in delaying tumor growth; it can be explained by the poor cellular uptake efficiency as well as the low retention ability of free drug molecules at tumor sites.31-32 The appearance of a superior anti-tumor activity of RB: [MNCs (c)] (+) over free RB (+) also emphasized the significance of nanoformulation for effective PDT-mediated treatments against MDR tumors. Lastly, the variation of the body weights of mice after intratumoral administration of therapeutic nanoagents is displayed in Figure 6E; no apparent change was observed in the groups treated with RB: [MNCs (c)] (+) in comparison with the PBS-treated controls (p > 0.05). Moreover, pathological assessment of hematoxylin and eosin (H&E)-stained sections of different organs (heart, lung, liver, spleen and kidney) obtained from RB: [MNCs (c)] (+)-treated mice at day 38 also confirmed no obvious tissue damages while an incision of the tumor revealed multiple scattered necrotic regions (Figure S13, Supporting Information). Collectively, these results suggest that the developed RB: [MNCs (c)] (+) represents a promising nanotool for future nanomedicine with high curative effect and good biocompatibility. To further understand the inhibition mechanism for PDT treatment, the TUNEL assay was conducted to detect cell apoptosis of tumor tissues harvested from mice treated with RB: [MNCs (c)] at 24 h post PDT (Figure 7A). Severe cells apoptosis (TUNEL-positive cells) was observed in irradiated groups (+) versus non-irradiated counterpart (-), revealing that the RB: [MNCs (c)] (+)- mediated PDT can effectively inhibit the tumor growth by inducing 16


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cells apoptosis in the tumor. Additionally, representative images in Figure 7A also showed RB fluorescence (green) was highly correlated with the apoptotic region (red), indicating that RB: [MNCs (c)] was able to efficiently release its payloads and induce effective apoptosis in the surrounding area after laser irradiation. In accordance with the abovementioned results, elevated areas of tissue damage were visualized for irradiated tumor, while the other organs showed no detrimental effects (Figure 7B). The observation of multiple scattered necrotic regions within the tumor at locations distinct from the loci where the MNCs were found (indicated by arrow), further suggesting that the drug payloads had been released by RB: [MNCs (c)] and had exerted profound cytotoxic actions toward a wide range of surrounding tumor cells after light activation.

CONCLUSIONS

In summary, a new class of delivery vehicle has been designed to attain a site-specific controlled release in response to light irradiation. The bPEI-based MNCs allows efficient encapsulation and protection of loaded drugs with negative charges through electrostatic interaction. On-demand payload release based on reduction of attractive forces was realized by photooxidation of bPEI sensitized by RB. Meanwhile, the generation of ROS further facilitated intracellular drug relocalization, which was ascribed to the photochemical reactions proceed in endocytic membranes. In comparison with free drugs, the nanoengineered light-switchable carrier augments its PDT activity against various cancer cell lines owing to its efficient cellular uptake and intracellular trafficking. An improved therapeutic efficacy of RB: [MNCs] co-equipped with additional chemodrug has also been successfully demonstrated in vitro through a precise control over drug release and subsequent combination therapy. Notably, the designed therapeutic nanoagent appeared to suppress tumor growth more effectively than free drugs in drug-resistant mouse xenograft model. This 17



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approach promises highly effective in vivo PDT treatment mediated by RB using a novel photooxidation-responsive nanovesicle capable of on demand drug release and delivery. This new nanoplatform could be further exploited as a safe and theranostic carrier for next-generation cancer therapy in the future.

EXPERIMENTAL SECTION

Chemicals. Poly(vinyl alcohol) (PVA) and branched polyethylenimine (bPEI) were purchased from Sigma-Aldrich (St. Louis, MO, USA) and Alfa Aesar (Ward Hill, MA, USA) with Mw of 9,000~10,000 and 1,800 Da, respectively. Low molecular weight of chitosan was from Sigma-Aldrich (St. Louis, MO, USA) with 75%~85% degree of deacetylation and viscosity of 20~300 cps in 1% acetic acid. Iron(III) acetylacetonate (Fe(acac)3) was purchased from Alfa Aesar (Ward Hill, MA, USA). Benzyl ether, oleylamine, β-estradiol BioReagent,

2’,7’-Dichlorofluorescin

diacetate

(DCFH-DA),

paclitaxel

and

4,5,6,7-Tetrachloro -2’,4’,5’,7’ -tetraiodofluorescein disodium salt (Rose Bengal) were purchased from Sigma-Aldrich (St. Louis, MO, USA). Transferrin, Alexa Fluor® 633 conjugate was obtained from Invitrogen (Carlsbad, CA, USA). Hexane (99.5%), ethanol (99.5%), Eosin-Y alcoholilc solution, Hematoxylin (MAYERS) solution and Ultra kit (Mounting medium) were obtained from J.T.Baker (Center Valley, PA, USA). Chloroform was purchased from Merck schuchardt (Hohenbrunn, Germany). Matrigel matrix high concentration was purchased from BD Biosciences (Franklin Lakes, New Jersey, USA). TUNEL Assay (Alexa Fluor 647) was purchased from Thermo Fisher (Massachusetts, USA). Fetal bovine serum (FBS) and penicillin/streptomycin were obtained from GIBCO (Grand Island, NY, USA). Dulbecco’s phosphate-buffered saline (DPBS) was purchased from Biosource (Camarillo, CA, USA). Tissue-Tek OCT cryo gel was purchased from Sakura 18


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Finetek (Torrance, CA, USA). SuperFrost Plus Microscope slides were purchased form Thermo Scientific (New Hamisphire, USA). Deionized water (18.2 MΩ．cm) was used to prepare all of the aqueous solutions. For the cellular experiments, all of the reagents, buffers and culture medium were sterilized by steam autoclave (121 °C, 40 min) or filtration (0.22 µm pore size, Millipore), and maintained under a sterile condition.

Synthesis of Fe3O4 NPs. Monodispersed Fe3O4 NPs were synthesized by thermal decomposition approach.33 Fe(acac)3 (3 mmol) was dissolved in 15 mL of benzyl ether and 15 mL of oleylamine. The solution was dehydrated at 110 ˚C for 1 h under N2 atmosphere, then quickly heated to 300 ˚C at a heating rate of 20 ˚C/min, and aged at this temperature for 1 h. After the reaction, the solution was allowed to cool down to room temperature. The Fe3O4 NPs were precipitated upon the addition of 50 mL of ethanol, followed by centrifugation at 16,000 g for 10 min. The pellet was washed with hexane twice and the yield of the purified Fe3O4 NPs was approximate 200 mg.

Synthesis of CTS/PVA/PEI-blended (tri-polymeric) magnetic nanoclusters (MNCs). CTS solutions of 0.5 wt% were prepared by dissolving chitosan in 0.5% aqueous acetic acid solution at room temperature with vortexing. PVA and bPEI were both dissolved in water to form 2 wt% polymer solutions. The three polymer solutions were then mixed carefully at various ratios; the weight fraction of bPEI was different to obtain a series of blends with 0 to 0.1 wt% bPEI in the resulting solution. MNCs were produced by emulsion/solvent evaporation method. A total of 1 mg oleylamine-coated Fe3O4 NPs (7.5 ± 1.5 nm) were dispersed in chloroform (2 mL), followed by the addition of aqueous solution (4 mL) containing different polymeric mixtures. The weight ratios of CTS, PVA, and bPEI to magnetite nanoparticles in different samples were 19



listed in Table 1. To form oil-in-water (O/W) emulsion, ultrasonication was then applied to the solution with an ice bath for 10 min using a tip-type Qsonica sonifier (pulse mode with on: 10 sec, off 10 sec, and output of 85). After solvent evaporation, the remaining polymer and solvent were removed by centrifugation (20,000 g, 20 min) and washed by phosphate buffer (PB, 10 mM) twice to obtaine the purified polymeric MNCs.

Characterization of the physiochemical properties of polymeric MNCs. The hydrodynamic diameter and zeta potential values of the constructed polymeric MNCs were measured by a dynamic light scattering (DLS) instrument (Malvern Instruments, United Kingdom). The morphology and size were also confirmed by transmission electron microscopy (Hitachi, Tokyo, Japan). The content of polymers within the MNCs was measured using a thermogravimetric analyzer, TGA/SDTA 851e (Mettler Toledo, Zurich, Switzerland) with a heating rate of 10 °C/min under a flow of helium gas. The iron concentration of the samples was quantified by inductively coupled plasma mass spectrometry (ICP-MS, Agilent 7700 Series ICP-MS, USA). The magnetization properties of oleylamine-coated Fe3O4 NPs and MNCs were analyzed with superconducting quantum interference device (SQUID) magnetometer (MPMS5, Quantum Design, U.S) at 300 K.

Preparation of rose bengal (RB)-loaded polymeric MNCs. To fabricate RB-incorporated nanoparticles, RB (10 mM, 16 µL) was added to the aqueous polymer blend solution (4 mL) and then mixed with oleylamine-coated Fe3O4 NPs dispersed in chloroform (2 mL), followed by ultrasonication (10 min, on/off intervals of 10 sec each) and vortex overnight. The resulting solution was subjected to three centrifuge/wash cycles (20,000 g, 20 min) to remove excess drug and re-suspended in 10 mM PB. The amount of unbound drug molecules in the supernatant was calculated from the emission intensity of RB at 560 nm (excitation at 500 20


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nm) using microplate reader (Tecan Group AG, Basel, Switzerland). To fabricate RB/PTX incorporated nanoparticles, PTX of 200 µM was pre-mixed with Fe3O4 NPs in chloroform (2 mL), followed by adopting the same synthetic strategy described above. The amount of unbound PTX molecules in the supernatant were quantified by high performance liquid chromatography (HPLC) with a detector wavelength of 227 nm. Elution was made by a mobile phase composed of A: deionized water and B: acetonitrile in a gradient mode from 30-90%, increasing B by 5% each minute. Stop time: 12 min. Post time: 5 min. Encapsulation efficiency (EE, %) = [residual amount of drugs in the nanoparticles/feeding amount of drugs] × 100.

Magnetic Resonance Imaging (MRI). Images were acquired on a Bruker 7T MRI system. The T2 relaxation rate of MNCs was measured by resuspending MNCs at series concentration in DPBS with a field of view of 40 mm, a slice thickness of 2 mm and an imaging matrix size of 128 by 128. TR was 3,000 ms and TE was 60 ms.

Cell lines and buffers. Tramp-C1 (transgenic adenocarcinoma of the mouse prostate), and MCF-7 (HTB-22 breast adenocarcinoma) cells were obtained from American Type Culture Collection (ATCC, Manassas, VA, USA). The resistant cell line MCF-7/MDR to adriamycin was kindly provided by Professor Hsin-Cheng Chiu, National Tsing Hua University, Taiwan. The cisplatin resistant cell line SKOV-3 was kindly provided by Professor Ching-Wei Luo, National Yang Ming University, Taiwan. Cells were cultured in suspension in DMEM medium supplemented with 10% FBS and 1% penicillin-streptomycin (Invitrogen, Carlsbad, CA, USA) at 37° in a balanced air humidified incubator with an atmosphere of 5% CO2. The cells were passaged every 2–3 days. Cell density for every experimental assay was

21



determined using a hemocytometer; purity of cell density was determined by visual microscopic inspection of the nuclei stained by trypan blue.

ROS assay. ROS production of RB-loaded polymeric MNCs was measured using the fluorescence probe DCFH-DA. 2 µL of DCFH-DA (2 mM) was added to RB-loaded MNCs solution (0.2 mL) prior to light exposure (632 nm, 15 mW/cm2) for 30 and 60 min. The Fe3O4 concentration of RB-loaded polymeric MNCs was maintained constant at a level of 25 µg/mL for all measurements. The increase in fluorescence signal as a consequence of ROS generation was acquired using microplate reader at an excitation wavelength of 450 nm and an emission wavelength of 535 nm. Intracellular ROS were further detected by flow cytometry using DCFH-DA. Briefly, MCF-7 cells were seeded at a density of 3 × 104 cells per 48-well plate for 12 h attachment. Cells were then incubated with RB-loaded MNCs (62.5 µg/mL) in complete culture medium for 6 h and washed twice in DPBS. 5 µM DCFH-DA prepared in DPBS was added to the cells for 15 min at 37 °C. Following 1 h red light exposure, cells were trypsinized and collected in the tube by centrifugation (1,000 g for 5 min) and resuspended in 200 µL washing buffer [4.5 g/L glucose and 5 mM MgCl2 in DPBS] for flow cytometry analysis (excitation = 488 nm; emission = 530 nm). For each analysis, at least 10,000 events were counted.

Release of drugs from polymeric MNCs upon light irradiation. The release study was conducted as follows: RB-loaded MNCs or RB/PTX-loaded MNCs (125 µg/mL) was dispersed in DPBS buffer (pH 7.4) at 37 ˚C, with stirring at a rate of 150 rpm. After red light or green laser light (15 mW/cm2) illumination for various time periods, spinning down the samples and measuring fluorescence intensity of the supernatants. The amount of RB release 22


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was quantified from the respective calibration curves. The release amount of PTX was quantified by HPLC.

Intracellular uptake of RB-loaded polymeric MNCs. To observe the cellular uptake of RB-loaded MNCs, MCF-7 cells were seeded at a density of 2 × 104 cells on 10 ×10 mm sterile cover glasses inserted into 48-well plates for 12 h attachment. Cells were then incubated with RB-loaded MNCs (62.5 µg/mL) in complete culture medium for 6 h and washed twice in DPBS. For microscopic imaging, cells were fixed with 4% paraformaldehyde for 10 min and monitored by confocal laser scanning microscopy (C2 plus Confocal system, Nikon, Tokyo, Japan). Endosomes were stained with transferrin, Alexa Fluor

633

conjugate

(200

nM)

for

30

min.

Nuclei

were

stained

with

4',6-diamidino-2-phenylindole (DAPI, 1.0 µM) for 15 min. To quantify the extent of MNCs uptake by MCF-7 cells, Prussian blue staining (0.5% potassium ferrocyanide with 1.2 M HCl) was performed on fixed cells for 10 min prior to microscopic analysis.

Cytotoxicity assay. MCF-7, Tramp-C1, SkOV-3 or MCF-7/ADR cells were seeded at a density of 7.5 × 103 cells per 96-well plate for 12 h. Cells were washed once and incubated with various concentrations (0 to 250 µg/mL) of RB-loaded or RB/PTX-co encapsulated MNCs at 37 °C, respectively. Following 6 h treatment, cells were washed twice in DPBS. In PDT studies, treated cells were exposure to a red light (15 mW/cm2) for 60 min or a green laser light (100 mW/cm2) for 5 min. After irradiation, cells were kept in complete culture medium for an additional 48 h at 37 °C in a 5% CO2 atmosphere. For cytotoxicity measurement, 10 µL Cell Titer reagent (Promega, Madison, WI, USA) was added to each well and incubated for 2 h. The absorption was recorded at 570 nm and 600 nm using a plate

23



reader, respectively. The percentage of cell viability was determined by comparing treated cells with the untreated control.

In Vivo PDT Treatment. The animal use protocol in this work has been approved by the Institutional Animal Care and Use Committee of National Tsing Hua University, Taiwan. Briefly, MCF-7/MDR cells were suspended in a PBS:matrigel solution (1:1) at a concentration of 5×107 cells/mL. 0.1 mL of cancer cells was injected subcutaneously into the right thigh of the seven-week-old BALB/c nude mice. 8 mg/L of β-estradiol was then administered in daily drinking water to induce the growth of the hormone-dependent MCF-7/MDR cells. The hormone treatment was stopped 1 week before therapy. Mice bearing MCF-7/MDR tumors were randomly assigned into six groups (n = 3), when the tumor grew up to a diameter of 200 mm3; each group were treated with saline, free RB, [MNCs (c)] or RB: [MNCs (c)], (124 µg and 0.25 mg per mouse of RB and Fe3O4 NPs, respectively) via intratumor injection. After 4 h post injection, the mice in the experimental groups were treated with different therapeutics: (1) non-irradiated or (2) PDT induced 532 nm light (100 mW/cm2, 10 min). Saline-treated mice without light irradiation were studied as control. The therapeutic efficacy of RB-loaded MNCs on mice bearing a MCF-7/ADR tumor was assessed by measuring tumor volume and body weight of mice in each group every day. Tumor size was measured using a vernier calliper. Tumor volume was calculated as length × (width)2 × 1/2. All mice were sacrificed and dissected to remove the tumors and several main organs (heart, liver, spleen, lung, and kidney) at 38 days after treatment. Tumor inhibition ratio (%) = (Wc–Wt)/Wc × 100, where Wc and Wt are the average tumor weight of control group and treatment group, respectively.

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Histology Examination. The histopathological analysis was performed by staining organs and tumor tissues, using hematocylin and eosin stain (H&E stain). After treatment, mice were euthanized; tumor tissues and main organs (heart, lung, liver, spleen, kidney) were harvested. Next, tissues were embedded in OCT (Sakura Finetek, Torrance, CA, USA) and cryosectioned at 10 µm thickness onto Superfrost Plus Slides (Thermo Scientific, New Hamisphire, USA) (n = 3). The sliced organs were then stained with H&E, covered with multimedium and examined under a light microscope (Olympus, Center Valley).

In Vivo TUNEL Assay. TUNEL assays were performed to detect DNA fragmentation that is characteristic of apoptotic cells. The TUNEL assay begins with incorporation of modified dUTP at the 3’-OH end of the fragmented DNA. The nude mice bearing tumor of the size approximately 200 mm3 were first administrated with RB-loaded MNCs intratumorally (124 µg and 0.25 mg per mouse of RB and Fe3O4 NPs, respectively) and irradiated with 532 nm laser (100 mW/cm2, 10 min) at 4 h post injection. A control group without laser irradiation was also performed at the same time. One day after, the nude mice were euthanized under CO2. After the tumor tissues were harvested, embedded in OCT gel and cryo-sliced; tissues were fixed with 100% cold methanol for 10 min, followed by paraformaldehyde (4%) fixation for 15 min at 37oC. Next, samples were washed twice with PBS and treated with proteinase K for 30 min. Afterwards, samples were washed with 1% BSA in PBS for three times (5 min each) and rinsed with deionized water. Click-iT Plus TUNEL (Thermo Fisher Scientific Inc.) assay with Fluor Alexa 647 was stained for 30 min at 37oC in the dark. Click-iT Plus TUNEL reaction cocktail was removed by washing twice each slide with 3% BSA in PBS for 5 minutes. Finally, DAPI was added to label cell nuclei. Images were examined by a fluorescence microscopy (Olympus, Center Valley).

25



Figure 1. (A) Schematic drawing showing the fabrication of tri-polymer (CTS/PVA/bPEI) coated magnetic nanoclusters (MNCs) as RB nanocarriers through an oil-in-water (O/W) emulsion method. (B) Hydrodynamic size distributions and (C) TEM images of MNCs at weight ratios (a) 10:0:0:1, (b) 10:2:0:1, and (c) 10:2:4:1 of CTS:PVA:bPEI:Fe3O4 in ddH2O (scale bar: 100 nm).

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27



Figure 2. (A) UV-Vis spectra of RB-loaded MNCs at weight ratios of 10:2:0:1 [MNCs (b)] and 10:2:4:1 [MNCs (c)] of CTS:PVA:bPEI:Fe3O4. The absorption profiles of RB (2.5 µM) and [MNCs (c)] were also recorded individually. The concentration of magnetite NPs was fixed at 31.3 µg/mL (0.125×). (B) Digital images of the as-prepared aqueous suspensions of (a) free RB, (b) RB-loaded tripolymer CTS/PVA/bPEI (10:2:4) nanoblends and (c) RB: [MNCs (c)] after centrifugation at 20,000 g for 20 min. Furthermore, the resulting supernatant is discarded, and the pellet is suspended in DPBS, respectively.

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Figure 3. (A) Photoinduced drug release from different RB-loaded MNC suspensions as a function of irradiation time. The RB: [MNCs] were diluted with DPBS to 125 µg/mL (0.5×) for measurement. bPEI, which constitutes the shell layer of [MNCs (c)], was compared to its linear counterpart, [MNCs (e)], with respect to the capture and release processes of loaded cargos, RB under red light exposure (632 nm, 15 mW/cm2). (B) Schematic drawing showing the possible mechanism of the controlled release of loaded cargos from [MNCs (c)] in response to photooxidation sensitized by RB. (C) Drug release behaviors from RB: [MNCs 29



(c)] diluted with DPBS to 125 µg/mL (0.5×), induced by 532-nm green laser and 632-nm red light lamp (15 mW/cm2), respectively. (D) UV-Vis spectra of RB-loaded [MNCs (c)] before (-) and after (+) green light irradiation. (E) UV-Vis absorption spectra of the sup and pre collected from the irradiated (+) RB: [MNCs (c)] and non-irradiated (-) RB: [MNCs (c)], respectively. (F) Fluorescence spectra of RB-loaded [MNCs (c)] before (-) and after (+) green light irradiation (RB, Ex = 540 nm; Em = 550 nm). (G) Fluorescence spectra of the sup and pre collected from the irradiated (+) RB: [MNCs (c)] and non-irradiated (-) RB: [MNCs (c)], respectively. Samples in (D–F) were irradiated under green laser light at 100 mW/cm2 for 5 min. The absorption and fluorescence profiles of RB (3.3 µM) were recorded individually. The concentration of magnetite NPs was fixed at 31.3 µg/mL (0.125×). (H) ROS generation signals of irradiated (+) RB: [MNCs (c)], non-irradiated (-) RB: [MNCs (c)] and their respective supernatant. Prior to ROS measurement, RB: [MNCs (c)] was diluted to an RB concentration of 1.0 µM. DCFH-DA (20 µM) was added to the suspension, followed by a second irradiation (532 nm, 100 mW/cm2, 5 min).

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Figure 4. Flow cytometry analysis of (A) ROS generation and (B) RB signals of non-treated MCF-7 cells, MCF-7 cells treated with RB (7.5 µM), 0.25× [MNCs (c)], 0.25× RB: [MNCs (c)] (RB = 7.5 µM), and H2O2 (300 µM) in a culture medium (10% FBS) at 37 °C for 6 h. For ROS measurements, DCFH-DA (5.0 µM, 15 min) was added to each treated cell suspension, followed by red light exposure for 1 h. *p < 0.05, **p < 0.01 versus non-irradiated counterpart.

31



Figure 5. (A) Confocal microscopic images of treated cells before and after light exposure. MCF-7 cells were incubated with 0.25× RB: [MNCs (c)] containing 7.5 µM RB in a culture medium (10% FBS) at 37 °C for 6 h. Cells were then rinsed with DPBS, followed by red light irradiation for 1 h. For microscopic analysis, cells were reacted with 4% paraformaldehyde solution for 10 min and stained with DAPI (1.0 µM, 15 min) and transferrin-Alexa633 (200 nM, 30 min). The fluorescence images were monitored for DAPI 32


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(nucleus), transferrin (endosome), and RB, and an overlay of these three channels is also shown. The scale length of each image is 20 µm. (B) Cytotoxicity assays of MCF-7 cells treated with various conditions: (a, b) RB: [MNCs (c)], (c) [MNCs (c)], (d) RB, and (e) RB: [MNCs (e)] in culture medium (10% FBS) at 37 °C for 6 h. After drug treatment, cells (b–e) were exposed to a red light (632 nm, 15 mW/cm2) for 1 h. Following a recovery period of 48 h in a fresh medium (10% FBS), the cytotoxicity was measured by an MTT assay. (C) Cytotoxic activity of RB/PTX: [MNCs (c)] (1×) against MCF-7 cells. The corresponding concentration of RB and PTX in 1× nanocomposites was 32 µM and 50 µM, respectively. The experiment is carried out in the same way except the laser irradiation was proceeded through a green laser (100 mW/cm2) for 5 min.

33



Figure 6. (A) Cytotoxic analysis of RB: [MNCs (c)] in MCF-7/MDR cancer cells. Cells were incubated with samples suspended in culture medium (10% FBS) at 37 °C for 6 h. After drug treatment, cells were irradiated with green laser light laser (100 W/cm2) for 5 min. Cells were 34


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left to recover for 48 h before MTT assay. (B) Tumor growth curve of mice after receiving different treatments (RB = 124 µg and Fe3O4 NPs = 0.25 mg per mouse) through intratumoral administration [1st drug dose + irradiation (532 nm, 100 W/cm2, 10 min) at day 0; 2nd irradiation (532 nm, 100 W/cm2, 10 min) at day 1]. (C) Representative digital images and (D) tumor weight obtained from mice after 38 days of different treatments. (E) Percentage body weight change of MCF-7/MDR tumor-bearing nude mice receiving different treatments over a period of 38 days. No significant difference was observed between the treated and the control groups. Error bars show SD for triplicate samples (n = 3). *p < 0.05 versus PBS (-) control.

35



Figure 7. (A) Fluorescence microscopic images of RB signals and TUNEL staining of apoptotic cells were observed within the tumor tissue section harvested from MCF-7/MDR tumor-bearing nude mice injected intratumorally with RB: [MNCs (c)] (RB = 124 µg and Fe3O4 NPs = 0.25 mg per mouse), before and after a green laser irradiation (100 W/cm2, 10 min). (B) Histopathological analysis of tumor and major organs obtained from RB: [MNCs (c)]-treated mice. The mice were sacrificed and dissected to remove the tumors and several main organs at 24 h after irradiation. Scale bar = 50 µm.

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ASSOCIATED CONTENT Supporting Information. Synthesis and characterization of the materials: UV-Vis, TEM images, XRD patterns, magnetic hysteresis analysis, TGA analysis, T2-weighted MR images, zeta potential, hydrodynamic size, drug encapsulation and release analysis, 1H NMR analysis; ROS generation evaluation; fluorescence microscopic images; cytotoxicity assay and H&E staining. The supporting Information including Figure S1 to Figure S13 is available free of charge.

AUTHOR INFORMATION Corresponding Author Prof. Yu-Fen Huang E-mail: [email protected] Notes Any additional relevant notes should be placed here.

ACKNOWLEDGMENT We appreciate financial support from the Ministry of Science and Technology (105-2113-M-007 -021 -, 106-2113-M-007 -008 -, 105-2627-M-019 -001 -, 106-2627-M-019 -001 -) of Taiwan, ROC. We also appreciate Dr. Hsin-Hung Chen for his guidance and constant advice to perform MDR animal model experiment.

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A New Photosensitized Oxidation-Responsive Nanoplatform for

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