Ligand-Switchable Micellar Nanocarriers for Prolonging Circulation

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Ligand-Switchable Micellar Nanocarriers for Prolonging Circulation Time and Enhancing Targeting Efficiency Tangjian Cheng, Yumin Zhang, Jinjian Liu, Yuxun Ding, Hanlin Ou, Fan Huang, Yingli An, Yang Liu, Jianfeng Liu, and Linqi Shi ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.7b18137 • Publication Date (Web): 17 Jan 2018 Downloaded from http://pubs.acs.org on January 17, 2018

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ACS Applied Materials & Interfaces

Ligand-Switchable Micellar Nanocarriers for Prolonging Circulation Time and Enhancing Targeting Efficiency Tangjian Cheng,

†

Yumin Zhang,

‡

Jinjian Liu, ‡ Yuxun Ding,

†

Hanlin Ou,

†

Fan

Huang,‡ Yingli An, † Yang Liu, *,† Jianfeng Liu,*, ‡ Linqi Shi*,†

†Key Laboratory of Functional Polymer Materials of Ministry of Education, State Key Laboratory of Medicinal Chemical Biology and Institute of Polymer Chemistry, College of Chemistry, Nankai University, Tianjin, 300071, China

‡Tianjin Key Laboratory of Radiation Medicine and Molecular Nuclear Medicine, Institute of Radiation Medicine, Chinese Academy of Medical Science & Peking Union Medical College, Tianjin, 300192, P.R. China. ABSTRACT Targeted drug delivery of nanomedicines offered a promising strategy to improve the tumor accumulation and reduce the side effects of chemotherapeutics. However, undesired recognition of the targeting ligands on the surface of nanocarriers by immune system or normal tissues decreased the circulation time and reduced the targeting efficiency. Here we developed a ligand-switchable micellar nanocarrier that can hide the targeting ligands during circulating in the bloodstream and expose them on the surface when entering the tumor microenvironments. With the ligand-switching capability, the nanocarrier achieved 66% longer blood circulation half-life and 23%

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higher tumor accumulation than the nanocarrrier with targeting ligands on the surface. This targeting strategy could serve as a universal approach to improve targeting efficiency for nanomedicines.

KEYWORDS: Switchable targeting ligands, mixed shell micelles, immune recognition, blood circulation, tumor accumulation

1. INTRODUCTION Cancer is one of the world’s deadliest diseases of human beings1,2. Conventional small-molecule chemotherapies share several limitations including the lack of sufficient selectivity to the tumors and consequent toxicity to healthy tissues3,4, low accumulation at the tumor tissues hence requiring high drug dose, as well as the frequent emergence of multidrug resistance5-7. Nanomedicine has emerged as a promising alternative to small-molecule chemotherapies aimed at specifically targeting the therapeutic payload to tumors, in order to enhance the therapeutic effects and reduce the adverse side-effects8-10. Targeted nanoparticles could be achieved by surface conjugation of the nanocarriers with targeting ligands including saccharides, folic acid, peptides and aptamers, which specifically recognized the overexpressed receptors on the tumor cell membrane11-14. However, undesired recognition of these targeting ligands by immune systems, organs and tissues leads to the acceleration of the nanoparticle clearance from the body, which significantly limits the overall targeting efficiency15-17. Such effects have been observed by several studies15,18-20, indicating that the conjugation of targeting ligands on the surface increased the 2


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clearance of nanoparticles by the reticuloendothelial system and decreased the circulation time in the blood, resulting in no significant enhancement in tumor accumulation compared to non-targeted ones. This is still a big challenge for targeted nanomedicines. Better strategies for the surface decoration of nanoparticles with targeting ligands need to be developed to realize the full potential of the targeting ligands and enhance the overall targeting efficiency of the nanoparticles.

To avoid the undesired immune-recognition of the targeting ligands without compromising their targeting capabilities, one potential strategy could hide the ligands during the circulation, while expose them when they are approaching their targets. Such switch-on/off mechanism of targeting ligands has been investigated in the field of nanomedicine and resulted in several advanced targeted-delivery nanocarriers21-26. Fundamentally, these approaches achieve the temporary deactivation of the targeting ligands by covering them with detachable structures such as caged molecules21-23 or degradable polymer coatings24-26, and then reactivate them under certain stimuli (photo, specific enzyme or acid condition). Although these approaches did reduce the undesired recognition and extend the circulation time of the targeted nanoparticles21,24, the reactivation mechanisms of these approaches often rely on chemical-bond cleavage or polymer degradation, which are usually slow in kinetics24. Such intrinsic shortcoming leads to the incomplete removal of the coverage structures when the nanoparticles approaching to the tumor sites, which prevents the ligands from realizing their full targeting capability. Moreover, such reactivation processes are usually irreversible, resulting in the ligands exposure to the immune systems again in 3



case the nanoparticles fail to bind to their targets and flow back into the bloodstream27.

Aiming to further enhancement in the targeting efficiency of cancer Nanotherapeutics, we have herein developed a polymeric micelle-based nanocarrier capable of hiding and exposing the targeting ligands conjugated on its surface in response to the change of its surrounding microenvironment. As illustrated in the Figure 1a, the nanocarrier has a core-shell structure with a hydrophobic core of poly(ε-caprolactone) (PCL) and a stimuli-responsive shell composed of PAE28-30 and poly(ethylene glycol) (PEG). Different from traditional core-shell micelles, the shell of the micelles here possesses a mixed shell structure that is composed of two different types of polymers. Such unique shell structure allows us switching the surface properties of the micelles significantly in responsive to certain stimuli without compromising their stability in the aqueous solution. Exemplified here with PAE/PEG mixed shell, since PAE undergoes reversible phase transition29 from deprotonated hydrophobic state at pH 7.4 to protonated hydrophilic state at pH 6.5, the collapse and stretch of the PAE chains on the micelle shell can be controlled by tuning the pH of the solution. By conjugating targeting ligands to be the terminal of the PAE chains, this mixed shell structure can hide the targeting ligands by pulling them inside the shell and exposing only PEG chains on the surface at the physiological pH (OFF state), whilst expose the ligands by pushing them back to the surface when the micelles entering an acidic tumor microenvironment (ON state). Due to the shielding effect of PEG, interactions between the targeting ligands and immune system are 4


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inhibited in the OFF state, leading to the enhancement in circulation stability of the micelles when in the blood stream (Figure 1a). Once the micelle reached tumor sites, the acidic tumor microenvironment (pH 6.5) will trigger the exposure of the targeting ligands, which activates the targeting capability and enhances the accumulation of the micelles at the tumor site (Figure 1b, ON state). Moreover, the state-switching process is rapid and reversible, which is essential for tumor-targeted delivery systems. This is because a significant amount of the micelles flow back to blood from the tumor microenvironment due to the high pressure of tumor interstitial fluid31. Such reversible process can effectively hide the targeting ligands again to avoid the immune clearance when the micelles flowing back to bloodstream. By reducing the clearance, there is a higher probability for the micelles to reach the tumor microenvironment again and bind to their targets, which enhances the overall targeting efficiency eventually.

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Figure 1. Schematic of polymeric micelle-based nanocarrier with the hiding/exposing of targeting ligands strategy. (a) The structure of mixed shell polymeric micelles: PCL as the core, PEG/PAE as the mixed shell. The targeting ligand is conjugated at the terminal of PAE chain under the coverage of PEG, which suppresses its targeting capability (OFF state). (b) Targeting ligand is exposed under an acidic tumor microenvironment (ON state). I, hidden targeting ligands inside the shell of micelles at blood pH reducing undesired immune recognition and clearance. II, accumulation of mixed shell micelles at tumor site with targeting ligands exposed on the surface. III, micelles flow back to the blood under high tumor interstitial fluid pressure and hide the targeting ligands again to avoid the immune clearance. IV, specific binding to receptor on the tumor cells. (c)Targeting ligands exposed on the surface of micelles are recognized by immune system, reducing the targeting efficiency.

2. MATERIALS AND METHODS 2.1. Materials. All chemicals were purchased from J&K chemical Ltd., USA unless otherwise stated. HOOC-PEG5000-OH and BOC-NH-PEG2000-OH were purchased 6


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from Shanghai YareBio Ltd., China. Doxorubincin-HCl was purchased from Jingyan chemical Ltd., China. Cyclic (Arg-Gly-Asp-D-Phe-Lys) (cRGDfk) was purchased from Shanghai GL Biochem Ltd., China. Cy5 NHS ester was purchased from Lumiprobe Co., USA.

2.2. Synthesis and Characterization of Micelles. Detailed protocols for the synthesis of all the block copolymers were described in the Supporting Information (). The micelles were prepared through nanoprecipitation technique. Briefly, the block polymers were respectively dissolved in THF with the concentration of 5 mg/ml. For the preparation of different kinds of micelles, certain amount of polymer solutions were mixed together first, and then added rapidly into phosphate buffer (pH ~ 5.5, 20 mM) under vigorous stirring, following by sonication for 10 min to form the micelle solution. Then, the solution was evaporated to remove the THF, and then transferred into a dialysis bag (MWCO 3500) and dialyzed against phosphate buffer (pH 7.4, 10 mM) for 2 days. The final micelle solution was concentrated by ultrafiltration before use. The formulation of the five types of micelles was shown in Table S1. DOX-loaded micelles were prepared in a similar method. Briefly, the solution of DOX in DMSO (w/ 2 eq. TEA) and the corresponding polymer solution in THF were mixed and added into phosphate buffer (pH ~ 5.5, 20 mM) under vigorous stirring, followed by sonication for 10 min to form DOX-loaded micelles. Cy5-labeled micelles

(Cy5-SSPM,

Cy5-SSPMRGD,

Cy5-MSPM,

Cy5-MSPMRGD

and

Cy5-MSPMSRGD) were prepared by the introduction of certain amount of Cy5-PEG2k-b-PCL2.3k (see in Supporting Information) to the polymer solution. These 7



Cy5-labeled micelles were concentrated via ultrafiltration for the studies of the blood clearance kinetics and the biodistribution of the micelles in vivo. Dynamic light scattering (DLS) measurements were performed on a laser light scattering spectrometer (BI-200SM) equipped with a digital correlator (BI-9000AT) at a 90° scatter angle under required temperature. The zeta potential values were measured on a Brookheaven ZetaPALS (Brookheaven Instrument, USA) using phosphate buffer (PB) solution (0.01 M) with a pH range from 5.0 to 7.4 as the background buffer. And the pH value of the micelle solution was adjusted by addition of nitric acid solution or sodium hydroxide solution. TEM samples were prepared by dropping the micelle solution onto a carbon-coated copper grid and dried slowly at required temperature. The stability of MSPMSRGD was evaluated in 10 mM PBS (pH 7.4) at 37 ℃, at the set time, the mean size of micelles was recorded by DLS. The data was shown in supporting information. 2.3. Investigation on the Phase Transition of PAE in the Micelles. To investigate the protonation of PAE at acidic pH (< 6.5), the 1H NMR spectra of micelles were detected under different pH condition. MSPMSRGD was prepared as described above using D2O as the solution. Deuterium chloride (20 wt. % solution in D2O for NMR) and sodium deuteroxide solution (30 wt. % solution in D2O for NMR) were used to adjust the pH of micelle solution. Then, observations of MSPMSRGD were performed at pH 7.4 and 6.5. Finally, the MSPMSRGD solution was dialyzed against phosphate buffer (pH 7.4), then against distill water. The NMR observation of MSPMSRGD was performed after lyophilisation. 8


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2.4. Cellular Uptake. HepG2 cells were seeded into a 96-well plates at a density of 0.6×104 cells per well in 100 µL RPMI-1640 medium/PBS. After an incubation of 24 hours, the culture medium of each well was replaced with 500 µL of fresh medium with different pH (pH 7.4 and 6.5), then the DOX loaded micelles (SSPM/DOX, SSPMRGD/DOX, MSPM/DOX, MSPMRGD/DOX and MSPMSRGD/DOX) were added to each well. After another 2-hour incubation, the culture medium was removed and cells were washed three times with 500 µL PBS buffer. The cellular uptake of micelles was observed with an inverted fluorescence microscope (DMI6000B, Leica, Wetzlar, Germany). For quantitative analysis of cellular uptake, micelle-treated HepG2 cells were detached by 0.02% (w/v) EDTA and 0.25% (w/v) trypsin solution, and then dispersed in 0.25 mL of PBS for flow cytometric measurement. Cells treated with PBS were used as control.

2.5. In Vitro Drug Release. 2 ml of micelle solution (pH 5.0 and 7.4) with the DOX concentration of 50 µg/ml was added to dialysis bag (MWCO 7000), and dialyzed against 40 ml of 10 mM buffer solution with the corresponding pH value under vigorous stirring. At set time points, 4 ml of dialysis fluid was taken out for fluorescence measurement (excitation at 490 nm) and an equal volume of fresh buffer was added in. The amount of released DOX was determined by measuring the intensity of the emission (590 nm) using free DOX as standard. The release experiments were performed in triplicate, and the results presented were the average data.

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2.6. In Vivo Blood Clearance Kinetics Assay. The animal studies were performed in accordance with the Regulations for the Administration of Affairs Concerning Experimental Animals (Tianjin, revised in June 2004) and adhered to the Guiding Principles in the Care and Use of Animals of the American Physiological Society. Standard curves of the Cy5-labeled micelle solution (Cy5-SSPM, Cy5-SSPMRGD, Cy5-MSPM,

Cy5-MSPMRGD

and

Cy5-MSPMSRGD)

was

established

before

administration. The rats (n = 3) were injected intravenously via the tail vein with Cy5-labeled micelle at a dose of 5 mg/kg, respectively. At predetermined time points (10 min, 30 min, 1 h, 2 h, 4 h, 8 h, 12 h, and 24 h), blood samples were collected into heparinized tubes. The blood samples were centrifuged at 4000 rpm for 10 min to remove the blood cells, the plasma fractions were collected and measured by a fluorescence spectrometer (excitation at 646 nm and emission at 660 nm). The amount of the micelle remaining in blood was determined according to the standard curves. The percentage of the injected dose (%ID) was calculated by comparing the amount of micelle remaining in blood with the total injected dose. Blank samples from mice administrated with physiological saline were analyzed as the background fluorescence of the plasma.

2.7. Imaging of Tumor Accumulation. HepG2 tumor-bearing nude mice (BALB/c mice) were used for tumor accumulation studies. When the tumors reached about 200 mm3, the mice were randomly divided into two groups and intravenously injected with Cy5- MSPM, Cy5-MSPMRGD and Cy5-MSPMSRGD with the same level of Cy5. At 1h, 6h, 24h post-injection, the mice were sacrificed, and the major organs were 10


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harvested. ex vivo imaging was conducted by the Kodak IS in vivo FX imaging system. 2.8. In vivo Antitumor Efficacy. When the tumor volume reached about 150 mm3, the tumor bearing mice were divided into four groups. From Day 0, the mice were weighed and injected with free DOX solution, DOX loaded MSPMRGD and DOX loaded MSPMSRGD and PBS (control) at set time points (Day 0, 3, 5, 7). The dose of DOX was fixed at 5 mg/kg body weight. Weight of mice and tumor volume were measured at determined time points for three weeks. The estimated tumor volume was calculated by the formula. Tumor volume (mm3) = Length × Width2/2. At Day 21, one tumor-bearing mouse in each group was anesthetized with 8% Chloral hydrate and sacrificed. Tumors were collected. Tumor samples were fixed for 24 h in 4% paraformaldehyde, embedded in paraffin, and cut into 8-µm-thick sections for hematoxylin/eosin (H&E) assays according to the manufacturer’s instructions. The photos were taken using optical microscope (Leica DMI6000 B). 3. RESULTS AND DISCUSSION 3.1. Fabrication and Characterization of Targeted Polymeric Micelles. The targeted micelles with switchable targeting ligands were synthesized from three types of block copolymers including PEG-b-PCL, PAE-b-PCL and cRGDfk-PAE-b-PCL (Figure S1, S2, and S3 in Supporting Information). cRGDfk (cyclic RGD, denoted as RGD) peptide was employed as a model targeting ligand and conjugated to the end of PAE. The mixed-shell-polymeric micelles with switchable RGD (denoted as MSPMSRGD) were successfully constructed through self-assembly of the three types of block copolymers in an acidic solution. The feeding ratio of PEG/PAE/RGD-PAE segments was 5:4:1 (w/w/w). As shown in Figure 2a, PAE exhibited hydrophilic state

11



at pH 6.5, exposing the RGD on the surface of MSPMSRGD. Adjusting the pH to 7.4 leads to the phase transition of PAE from hydrophilic state to hydrophobic state, which effectively pull the RGD to the inner layer of the shell and hide them under the coverage of PEG. The ON/OFF-switching process was firstly observed by measuring the hydrodynamic diameters of MSPMSRGD using DLS (dynamic light scattering), indicating an average diameter of 82±2 nm at pH 6.5 and 86±2 nm at pH 7.4 (Figure 2b). Further observation with transmission electron microscopy (TEM) (Figure 2c) confirmed these results. The slight differences in size between pH 6.5 and 7.4 indicate that the RGD-switching process occurs in the shell of MSPMSRGD and does not substantially change the size of micelles.

12


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Figure 2. Fabrication and characterization of targeted polymeric micelles. (a) Schematic representation of micelles with different formulations. (b) The hydrodynamic diameters of MSPMSRGD at pH 7.4 and 6.5. (c) The TEM images of MSPMSRGD at pH 7.4 and 6.5. (d) The particle size of SSPM, SSPMRGD, MSPM and MSPMRGD at pH 7.4. (e) Zeta potentials of SSPM, SSPMRGD, MSPM, MSPMRGD and MSPMSRGD under different pH conditions. Error bars denote the standard deviations (s.d.) over ten times of measurements with the same micelle suspension.

For the better study of the uniqueness of the targeting-ligand-switchable MSPMSRGD, similarly micelles with different surfaces were prepared and employed as comparative groups

(Figure

2a, Table

mixed-shell-polymeric-micelle

S1

in Supporting Information), including the

without

RGD

(MSPM),

the

non-switchable

mixed-shell-polymeric-micelle with RGD at the terminal of PEG (MSPMRGD), the single-shell-polymeric-micelle (SSPM) and the single-shell-polymeric-micelle with RGD at the terminal of PEG (SSPMRGD). DLS results demonstrated that these four types of micelles had similar sizes (~90 nm at pH 7.4) (Figure 2d), which were consistent with the results determined by TEM (Figure S4).

Zeta potential analyses were used to evaluate the surface charge of all the micelles (Figure 2e). All micelles were negatively charged at pH 7.4. In contrast, the mixed-shell-polymeric micelles, i.e., MSPM, MSPMRGD and MSPMSRGD, exhibited significant change in surface charge as the change of environmental pH to 6.5, while the zeta potential of the single shell polymeric micelles (SSPM, and SSPMRGD) remained negative under the same pH. Obviously, the introduction of PAE to the shell endows the mixed-shell-polymeric micelles with the surface-charge conversion capability, which was attributed to the protonation/deprotonation of PAE.

3.2. Reversible and Rapid Hiding/Exposing of Targeting Ligand. Besides the 13



charge conversion, changing in the environmental pH also leads to the phase transition due to the change in the hydrophobicity of the PAE chains (Figure S5), resulting in the collapse or stretch of the PAE domains on the micelle surface. As depicted in Figure 3a, since RGD is conjugated covalently on the terminal of PAE, the phase transition leads to the hiding/exposing of the RGD on the surface of MSPMSRGD when pH is high than 7.4 or lower than 6.5 respectively. Since both the phase transition and charge conversion was caused by deprotonation/protonation of PAE, the switching process of MSPMSRGD is fast and fully reversible. This is confirmed by measuring the zeta potentials of MSPMSRGD when immersing the micelles alternatively in phosphate buffer at pH 7.4 and 6.5 for several cycles. As shown in Figure 3b, the zeta potential of MSPMSRGD was +4.31 mV at pH 6.5, and then deceased to -8.52 mV when pH was tuned from 6.5 to 7.4. By adjusting the pH back to 6.5, the micelles again represented a positively charged surface with a very similar zeta potential (+3.74 mV) to that of their original state, indicating the high reversibility of the switching processes. Importantly, such switching process of MSPMSRGD could be finished in less than 2 min (the shortest time that we could achieved for finishing the pH adjustments and zeta potential measurements), suggesting a rapid responsiveness of MSPMSRGD to the changes in environmental pH. The results of the next cycle was shown in the supporting information.

To further validate the reversible phase transition of PAE on the surface of micelles, the structures of MSPMSRGD at different pH conditions were analyzed using 1H NMR. As shown in the 1H NMR spectra (Figure 3c), the characteristic peaks of PAE around 14


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2.5-3.0 ppm can be observed when the pH was adjusted below 6.5, but disappeared when the pH was adjusted to 7.4 due to the low mobility of the hydrogen of PAE when it is in insoluble state. This observation indicated that RGD-PAE segments in MSPMSRGD was insoluble at pH 7.4 and soluble at pH 6.5, confirming the reversible phase transition of PAE in MSPMSRGD.

Figure 3. Reversible and rapid hiding/exposing of targeting ligand. (a) Mechanism of the switching process of surface ligands based on the phase transition of PAE. (b) Charge conversion within two minutes of MSPMSRGD with alternative pH variation from 6.5 to 7.4. Error bars denote the standard deviations over ten times of measurements with the same micelle suspension. (c) 1H NMR spectra of MSPMSRGD under different pH conditions. (d,e) Inverted fluorescent microscopy image of HepG2 cells after incubation with MSPM/DOX, MSMRGD/DOX and MSPMSRGD/DOX at pH 7.4 (d) and pH 6.5 (e) for 2h. The cell nuclei were stained by DAPI (blue). (f) The flow cytometric analysis of MSPM/DOX, MSMRGD/DOX and MSPMSRGD/DOX at pH 7.4 and 6.5. The data were analyzed using Student's t-test, *P < 0.05.

The capability of MSPMSRGD to control the exposure of the RGD on the micelle’s surface allows the temporary deactivation of the targeting capability when circulating in the bloodstream (pH 7.4), and reactivation of the targeting capability when reaching tumor environment (pH 6.5). This is confirmed by comparing the cellular 15



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uptake of the micelles under different pH. HepG2 cells, which over-express RGD receptors32-34,

were

incubated

with

doxorubicin

(DOX)-loaded

MSPMSRGD

(MSPMSRGD/DOX) at different pH 6.5 and 7.4, respectively. For better comparison, DOX-loaded MSPM and MSPMRGD (MSPM/DOX and MSPMRGD/DOX) were also employed as negative and positive controls. After 2h incubation, cells were washed and the uptake were observed using fluorescent microscopy. Figure 3d and 3e compared the uptake efficiency of the micelles with the same DOX concentration at pH 7.4 and pH 6.5, respectively. As expected, cells treated with MSPMRGD/DOX presented much stronger fluorescence than those treated with MSPM/DOX, which was similar to the cellular uptake of single shell polymeric micelle as shown in Figure S6. Since the uptake of micelles was achieved through RGD-receptor mediated endocytosis, the uptake efficiency of both MSPMRGD/DOX and MSPM/DOX did not change significantly at different pH. In contrast, the uptake efficiency of MSPMSRGD/DOX varied

significantly, which indicated that the RGD on

MSPMSRGD/DOX was deactivated at pH 7.4 and reactivated at pH 6.5. Further quantification of the cell uptake using flow cytometry analysis (Figure 3f) revealed that the mean fluorescence intensity of cells incubated with MSPMSRGD showed no significant difference with that of the negative control, and significantly lower (-58%) than that of the positive control (*P < 0.05 was considered significant), indicating that the cells cannot recognize and interact with the RGD on MSPMSRGD when the RGD hiding inside the shell at pH 7.4. At pH 6.5, the mean fluorescence intensity of the cells treated with MSPMSRGD exhibited no significant difference with the positive 16


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control and 40% higher than the negative control ((*P < 0.05 was considered significant), which suggested that the RGD was exposed the surface and enhanced the uptake of MSPMSRGD. These quantitative analyses agreed with the results observed by fluorescent microscopy, confirming that MSPMSRGD can hide the target ligands at a physiological pH, as well as expose them on the surface in acidic condition such as tumor microenvironment. In addition, The cytotoxicity of MSPMRGD/DOX and MSPMSRGD/DOX were evaluated on HepG2 cells at different pH values (7.4 and 6.5) by 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyl-tetrazolium bromide (MTT) assay as shown in Figure S7. The increased cytotoxicity at pH 6.5 was observed at each DOX concentration

compared

with

that at

pH

7.4

for

MSPMSRGD/DOX

and

MSPMRGD/DOX, resulting from the charge conversion of the mixed shell micelles at pH 6.5.

3.3. In vivo Blood Clearance of Polymeric Micelles. Since the capability of hiding the targeting ligands of MSPMSRGD decrease the recognition and uptake by cells effectively at pH 7.4, we expect MSPMSRGD could achieve a prolonged circulation in vivo compared to the micelles without this capability. To evaluate the circulating capability of the micelles, five types of cyanine5 (Cy5) labeled micelles (Cy5-SSPM, Cy5-SSPMRGD, Cy5-MSPM, Cy5-MSPMRGD and Cy5-MSPMSRGD) were injected into healthy rats respectively, following by measuring the residual blood fluorescence over 24 h. The residual amount of micelles in blood is expressed as the percentage of the injected dose (%ID). Figure 4a compared the clearance profiles of SSPM, SSPMRGD and MSPM. According to the results, SSPMRGD showed the fastest clearance from the 17



bloodstream with a circulating half-life of 2.1±0.1 h. In contrast, non-targeted SSPM and MSPM showed a much slower clearance with the half-lives of 5.5±0.86 h and 10.4±1.7 h, respectively. The shortest circulation time of SSPMRGD in blood suggested that the RGD on the surface accelerated the immune clearance. In addition, MSPM exhibited the longer circulation time than that of SSPM because of the micro-phase separation on the micelle surface as reported in our previous work29,35.

Figure 4. Blood clearance and tumor accumulation efficacy of polymeric micelles. (a) Blood clearance profiles of SSPM, SSPMRGD and MSPM. Error bars indicate s.d (n=3). (b) Blood RGD

clearance curves of MSPM, MSPM

and MSPMSRGD. Error bars indicate s.d (n=3). (c) In 18


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vivo fluorescence imaging of the HepG2 tumor-bearing nude mice at 1, 6 and 24 h after intravenous injection of Cy5 labeled micelles. Arrow indicate the site of tumor. (d) Ex vivo fluorescence imaging of the tumor and normal tissues harvested from the HepG2 tumor-bearing nude mice at 1 h post-injection. The numeric label for each organ is as follows: 1, heart; 2, liver; 3, spleen; 4, lung; 5, kidney; and 6, tumor. (e) Quantitative analysis of fluorescent signals from tissues and tumor from the ex vivo fluorescence image. Error bar indicated s.d. (n = 3). The data were analyzed using Student's t-test, *P < 0.05.

Furthermore, the blood clearance rates of MSPM, MSPMRGD and MSPMSRGD were compared and summarized in Figure 4b. As a result, the circulation half-life of MSPMSRGD (9.8±0.7 h) is identical to that of MSPM (10.4±1.7 h), implying the two types of micelles presented similar surface structure in the blood circulation. In contrast, MSPMRGD showed the shortest circulation half-life (5.9±0.2 h) compared with those of MSPM and MSPMSRGD. Obviously, the only structural difference is that the RGD peptides of MSPMRGD are exposed on the micelle surface all the time, while MSPMSRGD can hide the RGD from exposing during the circulation. Therefore, the capability of hiding targeting ligands during circulating in the bloodstream makes MSPMSRGD an effective strategy for the construction of targeted drug delivery system without compromise the circulation time.

3.4. Tumor Accumulation Efficacy and Biodistribution. Besides the stability during circulating in blood, effective accumulating and binding to the target sites is also crucial for a targeted delivery system. To evaluate the capability of tumor accumulation, Cy5-MSPM, Cy5-MSPMRGD and Cy5-MSPMSRGD were injected into HepG2 tumor-bearing nude mice with the same dose of Cy5. A non-invasive near-infrared optical imaging technique was employed to investigate the accumulation of micelles in the tumors. As shown in Figure 4c, the fluorescence intensity at tumor 19



site was increasing along with time after the injection, indicating that all three types of micelles were capable to accumulate in tumor. Particularly, MSPMSRGD presented the strongest fluorescence signal at the tumor area (at each time point of 1 h, 6 h and 24 h), suggesting the highest accumulation efficacy in tumor among the three types of micelles. For a better quantification of the biodistribution of micelles, major organs (1, heart; 2, liver; 3, spleen; 4, lung; 5, kidney) of the animals were harvested and imaged at 1 h post injection. According to the results (Figure 4d, 4e, S8 and S9), MSPMSRGD showed the highest level of tumor accumulation (23% higher than MSPMRGD, and 33% higher than MSPM at 1h post injection), indicating that the conjugated RGD on MSPMSRGD exposed on the micelle surface and remained capable to bind to the tumor cells when entering the tumor microenvironment. Moreover, quantification of the biodistribution also indicated that the accumulation of MSPMSRGD in immune organs was identical to that of non-target MSPM, and significant lower than that of MSPMRGD (50% lower in liver and 20% lower in kidney). These results collectively confirmed the effectiveness of the switchable structure of MSPMSRGD in hiding and exposing surface ligands in response to the different microenvironments. By hiding the RGD peptides during the circulation, immune clearance against the RGD peptides can be largely inhibited, resulting in a higher portion of MSPMSRGD reaching to tumor microenvironment and subsequently exposing the RGD peptides to identify and binding to the cancer cells. With a higher targeting efficiency, MSPMSRGD is expected to deliver drugs and accumulate them in tumor in a highly specific and efficient manner, resulting an optimal antitumor effect. 20


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3.5. Antitumor Activities. DOX-loaded MSPMRGD/DOX and MSPMSRGD/DOX were prepared and employed for the evaluation of the antitumor efficacy. The sizes of the micelles were presented in Figure 5a, with average diameters of 92±2 nm and 88±3 nm for MSPMRGD/DOX and MSPMSRGD/DOX, respectively. In addition, the drug loading content (DLC) of MSPMSRGD was 12.5±2.8 wt%, showing no significant difference with MSPMRGD (11.3±2.1 wt%). The release profiles of DOX from MSPMRGD/DOX and MSPMSRGD/DOX were investigated under different pH (5.0, 6.5 and 7.4) (Figure 5b). Similar release kinetics were observed from both MSPMSRGD and MSPMRGD, with less than 10% of DOX released over 48 h incubation at pH 7.4, and 35% of DOX released at pH 5.0. Such acid-promoted drug release is essential to facilitate an effective intracellular drug release, which is important for the drugs to realize their anti-tumor functions.

21



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Figure 5. Antitumor activities. (a) The size and drug loading content of MSPMRGD/DOX and MSPMSRGD/DOX. (b) In vitro drug release profiles of MSPMRGD/DOX and MSPMSRGD/DOX under different pH conditions (7.4, 6.5 and 5.0) for 48 h. Error bars indicate s.d. (n = 3). (c) HepG2 tumor growth curves after injection with saline, free DOX, MSPMRGD/DOX and MSPMSRGD/DOX at a dose of 5 mg DOX/kg body weight. The arrows show the time of intravenous injection. Representative images (inset) of the tumors harvested from the mice after treating with the DOX formulations at Day 21. 1, free DOX; 2, free DOX; 3, MSPMRGD/DOX; 4, MSPMSRGD/DOX. Error bars indicate s.d. (n = 5). *P < 0.05, **P < 0.01 (two-tailed Student’s t-test). (d) Relative body weight during treatment for 21 days. Error bars indicate s.d. (n = 5). (e) Histological observation of the tumor tissues after treatment. The tumor sections were stained with hematoxylin and eosin.

The

antitumor

efficacy

was

evaluated

by

systemic

administration

of

MSPMSRGD/DOX, MSPMRGD/DOX, and free DOX into tumor-bearing nude mice 22


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respectively, following by monitoring the tumor volumes and the body weights of the animals. A dose of 5 mg DOX/kg body weight was administered into each mouse four times every two days. As showed in Figure 5c, all mice treated with the DOX formulations showed certain degrees of inhibition in tumor growth over 21 days. Both micelle-based formulations including MSPMRGD/DOX and MSPMSRGD/DOX inhibited the tumor growth more effectively than free DOX due to the higher efficiency of DOX in tumor accumulation. Particularly, the most effective inhibition of the tumor growth was observed from the mice treated with MSPMSRGD/DOX, which agrees with the results from the biodistribution studies. Further evaluation of the cell apoptosis in the tumor tissue using the hematoxylin and eosin (H&E) staining examination (Figure 5e) indicated that the administration of MSPMSRGD triggered the highest level of tumor apoptosis, which again confirmed the highest anti-tumor efficacy of MSPMSRGD among all the formulations. In addition, the micelle-based formulations reduced the undesired distribution of DOX throughout the body, leading to the relief in the side-effects caused by the non-specific toxicity of DOX. This was confirmed by monitoring the body weights of the mice over the treatments, indicating that the body weights of the mice treated with MSPMRGD/DOX and MSPMSRGD/DOX changed slightly during the treatment, while significant weight loss was observed from the mice treated with free DOX as shown in Figure 5d.

4. CONCLUSIONS In summary, we have demonstrated a targeted drug delivery system based on ligand-switchable micelles with enhanced tumor targeting efficiency. Such micelle is 23



capable to hiding the targeting ligands during circulating in the bloodstream, and expose them back on the surface when approaching tumor microenvironments. Compared to traditional targeted nanocarriers that usually have poor circulating capability due to undesired immune recognition of the targeting ligands conjugated statically on their surfaces, the micelle-based system reported here has overcome this problem with its ligand-switching structure, leading to an effective inhibition of the immune clearance of the micelle during the circulation without compromising the tumor-targeting capability. With the reduction of plasma clearance, a higher portion of the micelles achieved to reach and accumulate in the tumor sites, bind and internalize into the cancer cells, and release the drugs, leading to a higher overall drug delivery efficiency and antitumor efficacy than those of traditional targeted systems. Moreover, such ligand-switching process is achieved by the specially designed shell structure of the micelle and compatible to a large variety of tumor-targeting ligands, providing a general approach to develop new cancer nanotherapeutics with better targeting and antitumor efficacy.

ASSOCIATED CONTENT Supporting Information The Supporting Information is available free of charge on the ACS Publications website. The detailed synthetic routes and characterization of block copolymers, TEM images of micelles with different formulations. AUTHOR INFORMATION 24


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Corresponding Author *E-mail: [email protected] (L.S.). *E-mail: [email protected] (J.L.). *E-mail: [email protected] (Y.L.). Notes The authors declare no competing financial interest. ACKNOWLEDGMENTS This work was financially supported by the National Natural Science Foundation of China (51390483, 91527306) and Thousand Talents Program for Young Professionals. REFERENCES 1.

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Ligand-Switchable Micellar Nanocarriers for Prolonging Circulation

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