Folate-conjugated and pH-responsive polymeric micelles for target-cell-specific anticancer drug delivery

Accepted Manuscript Folate-conjugated and pH-responsive Polymeric Micelles for Target-Cell-Specific Anticancer Drug Delivery Jiao Guan, Zun-Qiang Zhou, Mao-Hua Chen, Hui-Yan Li, Da-Nian Tong, Jun Yang, Jing Yao, Zheng-Yun Zhang PII: DOI: Reference:

S1742-7061(17)30455-5 http://dx.doi.org/10.1016/j.actbio.2017.07.018 ACTBIO 4985

To appear in:

Acta Biomaterialia

Received Date: Revised Date: Accepted Date:

4 March 2017 3 July 2017 12 July 2017

Please cite this article as: Guan, J., Zhou, Z-Q., Chen, M-H., Li, H-Y., Tong, D-N., Yang, J., Yao, J., Zhang, Z-Y., Folate-conjugated and pH-responsive Polymeric Micelles for Target-Cell-Specific Anticancer Drug Delivery, Acta Biomaterialia (2017), doi: http://dx.doi.org/10.1016/j.actbio.2017.07.018

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1

Folate-conjugated and pH-responsive Polymeric Micelles for Target-Cell-Specific Anticancer Drug Delivery

Jiao Guan&,1, Zun-Qiang Zhou&,1, Mao-Hua Chen2 , Hui-Yan Li3, Da-Nian Tong1, Jun Yang1, Jing Yao1 , Zheng-Yun Zhang* ,1

1 Department of Surgery, Shanghai Jiao Tong University Affiliated Sixth People's Hospital, No. 600 Yishan Road,Shanghai, 200233,China 2 Key Laboratory of Advanced Technologies of Materials, Ministry of Education, Southwest Jiaotong University, Chengdu 610031, China 3 Key Laboratory of Medical Molecular Virology, Ministry of Education and Health, Shanghai Medical College, Fudan University, No.138 Yi xue yuan Road, Shanghai, 200032, China

& These authors contributed equally to this work.

*Correspondence to: Dr. Zheng-Yun Zhang, Department of Surgery, Shanghai Jiao Tong University Affiliated Sixth People's Hospital, No. 600, Yishan Road, 200233, Shanghai, China. Tel: +86-21-64369181; Fax: +86-21-64701361; E-mail:[email protected]

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2

Abstract: In this study, we developed a folate (FA)-conjugated and pH-responsive active targeting micellar system for anti-cancer drug delivery. In this system, FA was attached to the terminal of the hydrophilic segment of poly(lactic acid)–poly(L-lysine) (PLA–PLL), and PLL was modified by a citric acid group. The FA receptor-mediated active targeting and electrostatic interaction between micelles and cell membrane due to a negative-to-positive charge reversal was combined in one micellar anti-cancer drug delivery system to enhance the tumour targeting and cellular internalisation of micelles. In vitro and in vivo anti-cancer studies demonstrated that the doxorubicin-loaded, FA-conjugated and pH-responsive polymeric micelles possess an enhanced and effective cancer efficiency. Keywords: FA receptor-mediated active targeting; charge reversal; pH-sensitive; micelle; drug delivery; cancer therapy 1. Introduction At present, systemic chemotherapy accompanied with surgical resection or radiotherapy remains the most commonly utilised therapeutic strategy for cancer inhibition. However, most anti-cancer drugs (e.g. doxorubicin (DOX) and cisplatin) that can effectively induce cancer cell death present tremendous side effects mainly because of their lack of targeting ability [1–2]. Nano-carriers have attracted wide attention as a promising drug delivery system in cancer chemotherapy because of their capacity to carry drugs preferentially into tumour tissue [3–4] through the well-known enhanced permeation and retention (EPR) effect [4–5]. However, the route from the injection site to the target cells is blocked by numerous barriers, such as mucosal barriers and nonspecific uptake, which limit the effectiveness of nano-drug delivery systems [6]. 2

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To overcome these barriers, two aspects should be considered simultaneously. One is to prolong the blood circulation of nano-carriers, and the other is to enhance their targeting ability. To prolong blood circulation, polyethylene glycol (PEG) is often used for the surface modification of these nano-carriers; this process renders the nano-carriers invisible from the mucosal barriers and prolongs their circulating time in vivo, thereby improving the accumulation in tumour tissues [7–9]. However, PEGylation may also impede the cancer cell uptake of nano-carriers and lead to insufficient anti-tumor efficacy [10–11]. Even if nano-carriers are located at the target cells, their internalisation by cancer cells is difficult. Thus, efforts should be exerted to improve the targeting ability of nano-carriers and enhance the internalisation of micelles by cancer cells. At present, nano-carriers are modified by various targeting ligands with receptors overexpressed on the surface of cancer cells through the so-called active targeting [12]. For example, folate (FA), which possesses high binding affinity to the FA receptor on the cancer cell surface, has been widely employed to enhance the cellular uptake of micelles. [13–15] Positively charged nano-carriers (including micelles) that can be more readily uptaken by cancer cells owing to the negative charge of cell membranes have also been employed [16–17]. However, positively charged nano-carriers are easy to be cleared by the reticulo-endothelial system (RES) and present a strong interaction with serum components, resulting in severe aggregation and a short blood circulation half-life [18]. Meanwhile, negatively charged nano-carriers overcome this problem to a certain degree, thereby prolonging blood circulation [19]; however, these nano-carriers are difficult to be internalised. Therefore, a negative-to-positive charged reversal micelle surface, i.e., one that maintains a negative charge in blood circulation and becomes 3

4

positively charged upon entering the tumour tissue, could improve selectivity for tumour cells and increase uptake chance [20-23]. Furthermore, even if the nanocarriers pass through the barriers exist in the blood, arrive in tumor tissue and internalized by tumor cells, they will still encounter intracellular compartments (e.g. early endosomes, late endosomes, and lysosomes). Thus, the escape of nanocarriers or drug from the intracellular compartments should be an urgent task to realize the efficient intracellular drug release. Considering the acidic environment of endo/lysosomes (pH 4~5) [24], pH-responsive delivery systems would be a good choice to achieve the goal, for example nanocarriers, which can induce a “proton sponge effect” to absorb protons in the endo/lysosomes, would be more easier to escape from the endo/lysosomes. Because the “proton sponge effect” can increases the osmotic pressure inside the endo/lysosomes which can result in plasma membrane disruption [36]. Taken together, to achieve the goal of efficient intracellular drug release, the pH-responsiveness should occur at least twice in the delivery of nanocarriers from extracellular environment to intracellular compartments. However, most of these systems are only responsive to a single pH variation, either in the tumor extracellular environment or in the intracellular compartments. Based on those thoughts, in this study, we combined FA- mediated active targeting and a negative-to-positive charge reversal in the same micellar drug delivery system as shown in Scheme 1. Relative to other nano-carriers, polymeric micelles with distinct core/shell architecture self-assembled from amphiphilic co-polymers have been extensively explored [25–26] because of their advantages [27–29]. These advantages include high stability ascribed to low critical micelle concentration (CMC), improved solubility of water-insoluble drugs, prolonged blood circulation and ease of functionalisation. In the micellar system that we 4

5

designed, poly(lactic acid) (PLA), an FDA-approved biodegradable material, was employed as the hydrophobic segment, and poly(L-lysine) (PLL), a cationic polymer widely applied in gene delivery vectors [30–31], was used as the hydrophilic segment. FA was conjugated onto the terminal of PLL segment. A negative-to-positive charge reversal was acquired by integrating a citric acid (CA) group onto the side NH2 group of PLL. This group would turn the positively charged polymer into a negatively charged polymer. Additionally, the amide bond between CA and NH2 is sensitive to pH and can be cleaved in acidic environments. Thus, a negative-to-positive charge reversal of micelles could be achieved in acidic tumour microenvironments, thereby increasing their probability to bind onto the negatively charged membrane of tumour cells. Furthermore, PLL exhibits a proton sponge effect in acidic solution because of the large amount of unsaturated amide bonds, which may help accelerate the escape of micelles from the endo/lysosomes to achieve an effective intracellular drug release. 2. Materials and methods 2.1. Materials Lactide, N6-[(phenylmethoxy)carbonyl]-L-lysine ((Z)LL), 3-amino-1-propanol

(AP), CA,

succinic acid (SA), FA, trifluoroacetic acid (TFA) and triphosgene (BTC) were purchased from Sigma–Aldrich (Shanghai, China). Doxorubicin (DOX), SnCl2, dicyclohexylcarbodiimide (DCC), N-hydroxysuccinimide (NHS), (4-dimethylaminopyridine) DMAP and di-tert-butyl dicarbonate (Boc2 O) were obtained from Aladdin (Shanghai, China). All other chemicals were acquired from commercial suppliers and used as received. 2.2. Cells and animals The human cervical cancer cell line HeLa, human liver hepatocellular carcinoma cell line 5

6

HepG2 and lung epithelial cancer cell line A549 (FA receptor positive) were obtained from Cell Bank of Chinese Academy of Sciences and Sichuan University (China). Cells were cultured in RPMI 1640 (Hyclone) supplemented with 10% foetal bovine serum (FBS, Gibco) at 37 °C in an atmosphere of 5% CO2 and under humidified conditions. Male BALB/c nude mice (20 ± 2 g, 5–6 weeks of age) were approved for use by the Institutional Animal Care and Use Committee of Shanghai Jiao Tong University and fed under 25 °C and 55% humidity. All animal experiments were performed in compliance with the guidelines of National Institutes of Health guide for the care and use of Laboratory animals (NIH Publications No. 8023, revised 1978). HeLa tumour models were established via the subcutaneous injection of 100 µL HeLa cell suspension containing 1×106 cells at the left flank of BALB/c nude mice. The tumour volume was calculated based on the longest (a) and shortest (b) diameters by following equation: volume = ab2/2. 2.3. Synthesis procedures The target product was synthesised as previously described [32] with slight modifications as shown in the following steps. 2.3.1 Synthesis of tert-butyl-N-(3-hydroxypropyl) carbamate AP (75 mg, 1 mmol) and triethylamine (1 mmol) dissolved in 10 mL of dichloromethane (DCM) were cooled to 0℃, and followed by dropwise addition of Boc2O (240 mg, 1.1 mmol) for 1 h with continuous stirring.[33] The mixture was kept stirred at room temperature for another 4h, diluted with 20 ml 5% potassium hydrogen sulphate aqueous solution and deionized water, extracted with ice ethyl ether (2×30ml) . The organic layers were desiccated by MgSO4. After filtration and evaporation, the residue was purified by silica gel column chromatography using an 6

7

eluent gradient from 0/5 (v/v) to 1/5 (v/v) ethyl acetate / hexane to obtain the product. Purified tert-butyl-N-(3-hydroxypropyl) carbamate was present as colourless viscous oil and was over 98% pure as determined by high-performance liquid chromatography. Yield: 85%. 2.3.2 Preparation of Boc–NH–PLA Tert-butyl-N-(3-hydroxypropyl) carbamate (87.5 mg, 0.5 mmol), L-lactide (2.6 g, 18 mmol) and SnCl2 (0.0275 g) (as a 0.5 g/mL solution in dry DCM) were added into a 100 mL oven-dried round-bottom flask equipped with a stopcock, vacuum purged with Ar for three times and then degassed under a high vacuum for 4 h to remove the residual DCM at 40 °C. Afterwards, the device was heated to 140 °C to initiate ring-opening polymerisation, which lasted for 6 h. After the device was cooled, the products were dissolved in DCM, precipitated in excess cold ethanol and then vacuum dried for 24 h. Thus, white Boc–NH–PLA powders were obtained in 83% yield. 2.3.3 Synthesis of amine-terminated PLA (NH2-PLA) Boc–NH–PLA (2.6 g, 0.5 mmol) in distilled DCM (20 mL) was added in an oven-dried flask, followed by the addition of a large excess amount of anhydrous TFA (3 mL). The solution was kept stirred for approximately 1 h at 0 °C. After all solvents were evaporated, the product was redissolved in DCM, washed with 5% NaHCO3 aqueous solution and water and finally dried over MgSO4. After filtration and precipitation in chilled methanol, the product amine-terminated PLA (NH2–PLA) was vacuum dried at room temperature. Yield: 80%. 2.3.4 Synthesis of P(Z)LL–PLA Prior to the synthesis of P(Z)LL–PLA, ω-benzyloxycarbonyl-L-lysine-N-carboxyanhydride ((Z)LL–NCA) was prepared as previously described [34]. In brief, (Z)LL (4.2 g, 15 mmol) was dissolved in tetrahydrofuran (THF, 40 mL) and reacted with BTC at 40 °C for 4 h to obtain 7

8

(Z)LL–NCA. Then, (Z)LL–NCA (0.918 g, 3 mmol) in 10 mL of dimethylformamide (DMF) solution was added into the solution of NH2–PLA (0.5 g, 0.1 mmol) dissolved in 10 mL of DMF and then reacted for 72 h at 35 °C under stirring. After purification of the products by ethyl ether, light-coloured powders of P(Z)LL–PLA were obtained in 81% yield. The polymerisation degrees of P(Z)LL and PLA in P(Z)LL–PLA were finally determined to be 31 and 68 by 1 H NMR. 2.3.5 Synthesis FA–P(Z)LL–PLA The co-polymer FA–P(Z)LL–PLA was synthesised as follows: P(Z)LL–PLA (0.64 g, 0.05 mmol) was dissolved in 15 mL of dimethylsulphoxide (DMSO) with FA (26.5 mg, 0.06 mmol), DCC (15 mg, 0.072 mmol) and NHS (8.3 mg, 0.072 mmol) under stirring at 40 °C for 12 h and then purified with ethyl ether to obtain FA–P(Z)LL–PLA. Yield: 87%. 2.3.6 Synthesis of FA–PLL–PLA and PLL–PLA FA–P(Z)LL–PLA (0.52 g, 0.04 mmol) was dissolved in a 20 mL mixture solution of TFA (15 mL) and 33% HBr/HAc solution (5 mL) under stirring for 2 h at 0 °C under N2 protection and then purified with ethyl ether to obtain FA–PLL–PLA in 75% yield. Similar procedures were carried out to prepare PLL–PLA, in which P(Z)LL–PLA (0.128 g, 0.01 mmol) was used, and the resultant yield was 80%. 2.3.7 Synthesis of FA–PLL(CA)–PLA, FA–PLL(SA)–PLA and PLL(CA)–PLA FA–PLL(CA)–PLA was synthesised as follows: FA–PLL–PLA (0.196 g, 0.002 mmol) was dissolved in 5 mL of DMSO with CA (156 mg, 1 mmol) and DMAP (24 mg, 0.2 mmol) with stirring for 12 h and then purified with ethyl ether. Yield: 90%. FA–PLL(SA)–PLA was synthesised as follows: FA–PLL–PLA (0.196 g, 0.002 mmol) was 8

9

dissolved in 5 mL of DMSO with SA (100 mg, 1 mmol) and DMAP (24 mg, 0.2 mmol) with stirring for 12 h and then purified with ethyl ether. Yield: 89%. PLL(CA)–PLA was synthesised as follows: PLL–PLA (0.188 g, 0.002 mmol) was dissolved in 15 mL of DMSO with CA (156 mg, 1 mmol) and DMAP (24 mg, 0.2 mmol) under stirring for 12 h. After purification, PLL(CA)–PLA was obtained in 86% yield. 2.3.6. Characterisation of co-polymers 1

H nuclear magnetic resonance (1H NMR) spectra were obtained on a Bruker AM 300

apparatus to confirm the structures of the prepared co-polymers. CDCl3 was used as the solvent, and tetramethylsilane (TMS) was used as the internal reference. The weight-average molecular weight (Mw) and Polymer Dispersion Index(PDI) were measured by gel permeation chromatography (Waters 1515, USA). DMF/LiBr was used as the eluent at a flow rate of 1 mL/min at 40 °C, and polymethyl methacrylate was used as the reference. 2.4. Micelle formation and characterisation The

blank

or

DOX-loaded

micelles

from

the

co-polymers,

PLL(CA)–PLA,

FA–PLL(CA)–PLA or FA–PLL(SA)–PLA, were prepared using the solvent evaporation method. In brief, to prepare blank micelles, a 10 mL block co-polymer solution in THF was added dropwise into 10 mL of deionised water under stirring, followed by evaporation of THF at room temperature. Finally, the blank micelle solution was obtained. The DOX-loaded micelles were prepared using the same procedure, except a certain amount of DOX (DOX:co-polymer = 1: 10, w/w) was dissolved in THF together with the co-polymers. All processes were performed in the dark. Then, the prepared drug-loaded micelle solution was pipetted into a dialysis bag (MWCO 1000) against deionised water to remove unloaded drugs. For storage, micelle solution samples 9

10

were lyophilised and then stored under dry conditions. For characterisation, a 1 mg/mL micelle solution was filtered using a 0.22 µm syringe filter in advance. The size distributions of different micelles were determined by dynamic light scattering (DLS, Nano-ZS90, Malvern Zetasizer) measurements. Micellar morphologies were observed and imaged by transmission electron microscopy (TEM, HT7700, Hitachi). The DOX loading content and encapsulation efficiency were determined using a ultraviolet (UV)-vis spectrophotometer (U2910, Hitachi). Briefly, the calibration curve of DOX in dimethyl sulphoxide (DMSO) have been established in advance, in which the UV absorption changes of different amount of DOX in DMSO at the excitation wavelength of 488 nm have been determined. Then the preweighed freeze-dried sample was re-dissolved in DMSO, and the DOX concentration in the solution can be determined according to the absorbance of DOX at 488 nm and the standard curve. Furtherly, the drug loading content (LC) and encapsulation efficiency (EE) can be calculated according to formulas as below: LC=(Weight loaded drug/Weight loaded drug and polymer)*100% EE=(Weight loaded drug/Weight drug in feed)*100% 2.5. pH-responsive behaviours and in vitro DOX release The

pH-responsive

behaviours

of

PLL(CA)–PLA,

FA–PLL(CA)–PLA

and

FA–PLL(SA)–PLA micelles were determined by monitoring the zeta potentials of different micelles under diverse conditions with pH values ranging from 7.4 to 4.5 by using a Malvern Zetasizer Nano-ZS90 instrument. The in vitro release behaviours of DOX from different micelles (PLL(CA)–PLA, FA–PLL(CA)–PLA or FA–PLL(SA)–PLA) were investigated. Different DOX-loaded micelle 10

11

solutions 1 mL(1 mg/mL) were pipetted into a dialysis bag (MWCO 1000), which was immersed into a tube containing 30 mL of phosphate buffered saline (PBS, pH 7.4 or pH 6.5). The tubes were placed on a shaking bed at 100 rpm and 37 °C. At a preset time point, 1 mL of the release medium was collected to determine the released amount on a fluorescence spectrophotometer (F7000, Hitachi), and 1 mL of fresh PBS or ABS was supplemented into the tube. The release behaviour of 30 µg/mL DOX solution in a dialysis bag was also investigated. 2.6. Cytotoxicity assay Alamar Blue assay was employed to investigate the cytotoxicity of blank micelles from PLL(CA)–PLA, FA–PLL(CA)–PLA or FA–PLL(SA) –PLA with cancer cells (HepG2, HeLa and A549). Cells (2×104 cells/well) were seeded in a 48-well plate and then cultured for 24 h. Subsequently, different blank micelle solutions were added into the culture medium (micelle concentration ranging from 50 µg/mL to 800 µg/mL) and then co-cultured with cells. After 24 h, the medium was removed, and the cells were washed three times with PBS. Then, 300 µL of Alamar Blue solution (10% Alamar Blue, 10% FBS and 80% media 199 (Gibco), v/v) was added and incubated for 24 h. Finally, the Alamar Blue solution was determined on a microplate spectrophotometer (FlexStation 3, MolecularDevices) 2.7. Investigation of in vitro anti-tumour effects Alamar Blue assay was also used to investigate the anti-cancer activities of free DOX and DOX-loaded micelles (PLL(CA)–PLA, FA–PLL(CA)–PLA or FA–PLL(SA)–PLA) with HeLa cells at both pH 7.4 and 6.5, respectively. In brief, cells (2×104 cells/well) were seeded onto a 48-well plate and then cultured for 24 h. Then, DOX and different DOX-loaded micelles were added into the culture medium (pH 7.4 or pH 6.5) for incubation for another 24 h. The DOX dose 11

12

ranged from 0.025 µg/mL to 10 µg/mL. Alamar Blue assay was conducted as described above. The morphologies of HeLa cells (seeded into a six-well plate at a density of 5×104 cells/well) after incubation with different samples for 4 h at both pH 7.4 and 6.5 were directly viewed using confocal laser scanning microscopy (CLSM) to demonstrate the intracellular drug release of the micelles. An equivalent DOX dose (5 µg/mL) was applied. Cells without any treatments served as control. After incubation, the cells were fixed by with 2.5% glutaraldehyde for 60 min, and cell nuclei were stained with DAPI (blue) for approximately 10 min, followed by rinsing with PBS for three times. Then, the cells of each group were imaged via CLSM (FV1000, Olympus, Japan). Cell apoptosis was quantified by flow cytometry to further confirm the anti-tumour effect of free DOX and different DOX-loaded micelles at both pH 7.4 and 6.5. In brief, HeLa cells (1 × 105 cells/well) were seeded in a six-well cell culture plate and then cultured for 24 h. Then, the cells were treated with different DOX-loaded micelles at an equivalent DOX dose of 10 µg/mL. After 48 h of exposure, cells were obtained, and their apoptosis rates were measured using an Annexin V–FITC/propidium iodine (PI) apoptosis detection kit (Beijing 4A Biotech Co., Beijing, China) in accordance with the manufacturer’s protocols by flow cytometry (Accrue C6, BD Biosciences). Cells in a culture medium of pH 7.4 or 6.5 without any treatment were used as control groups. 2.8. In vivo biodistribution Saline, fluorescent probe DIR and DIR-loaded micelles (PLL(CA)–PLA, FA–PLL(CA)–PLA or FA–PLL(SA)–PLA) were administrated to HeLa tumour-bearing BALB/c nude mice via the lateral tail vein at an equivalent DIR dose of 5 mg/kg. Then, at preset times (0.5, 2, 4 and 8 h), the mice were anaesthetised and imaged by an in vivo imaging system (Spectrum CT, Perkin-Elmer). After being imaged at 8 h, the mice were sacrificed, and the organs and tumours were excised for 12

13

ex vivo imaging. A 740 nm excitation filter was used, and the emission fluorescence from 780 nm to 800 nm was collected. 2.9. Investigation of in vivo anti-tumour effects At approximately 10 days after the subcutaneous inoculation of HeLa cells, the tumours in the mice reached a volume of around 50 mm3 (labelled as day 0). Then, 5–6 week tumour-bearing BALB/c nude mice were randomly grouped (n = 7). Different samples, including saline, free DOX and

DOX-loaded

micelles

(PLL(CA)–PLA/DOX,

FA–PLL(CA)–PLA/DOX

or

FA–PLL(SA)–PLA/DOX) were injected into the mice from each group with a DOX dose of 5 mg/kg via the lateral tail vein. Intravenous administration was carried out every 3 days for four times since day 0. Tumour size (the longest (a) and shortest (b) diameters) and the weight of each mice were measured every 3 days since day 0. Furthermore, at 21 days after the first administration, mice from each group were randomly selected and sacrificed. The tumour tissues of the mice were prepared into 5 µm-thick tumour sections, stained with haematoxylin/eosin (H&E) and then immunostained with a rabbit polyclonal antibody for Ki-67 or caspase-3. The H&E- and caspase-3- or Ki-67-stained tumour sections were observed and imaged by optical microscopy. 2.10. Statistical analysis Data were displayed as mean values ± standard deviation. The statistical significance of the data was determined by single factorial ANOVA. 3. Results and discussion 3.1. Synthesis and characterisation of co-polymers The FA-attached and pH-responsive co-polymer FA–PLL(CA)–PLA and the co-polymers 13

14

PLL(CA)–PLA and FA–PLL(SA) –PLA were synthesised as shown in Figures 1, FiguresS1 and Figures S2. 1H NMR spectral analyses (Figure 2A and Figures S3–S6 in the Supporting Information) were carried out to verify the composition and structures of the synthesised co-polymers. The characteristic peaks for each product can be found as labelled in the spectra, suggesting the successful synthesis of the co-polymers. The molecular weights of the different co-polymers are shown in Figure S7 and Table 1. 3.2. Micelle formation and characterisation A solvent

evaporation

method

was

performed

to

form stable

micelles

from

FA–PLL(CA)–PLA, PLL(CA)–PLA and FA–PLL(SA)–PLA. A series of measurements was carried out to confirm the successful formation of these micelles (Table 2). First, the CMC values of different micelles (1.86 mg/L for FA–PLL(CA)–PLA, 2.29 mg/L for PLL(CA)–PLA and 2.04 mg/L for FA–PLL(SA)–PLA) were calculated from the plot of I339/I333 ratio over co-polymer concentration (Figure S8A) by using pyrene as a probe. Low CMC values can ensure that the micelle remains stable under an extremely dilute aqueous condition (such as blood) until the target site is reached. The size distribution (Figure 2B and Figure S8B) of these micelles was measured by DLS. The unimodal peak was observed for all types of micelles, suggesting that all micelles possess a uniform size distribution. The average diameters of the micelles were 125 nm (FA–PLL(CA)–PLA), 148 nm (PLL(CA)–PLA) and 132 nm (FA–PLL(SA)–PLA). TEM images of the micelles (Figure 2C and Figure S8C) revealed their presence as uniform spheres with smaller diameters compared with those found in the DLS results. This difference can be attributed to the shrinkage of the mPEG shell due to the dehydration of micelles during sample preparation for TEM. These results demonstrate that stable micelles were successfully self-assembled from the 14

15

prepared co-polymers. The characteristics of different DOX-loaded micelles, including the DOX loading content and encapsulation efficiency were also determined and the data can be found in Table 3. The morphology and size distribution of each micelles are shown in Figure S9A. The morphologies of DOX-loaded micelles maintained spherical shapes. To further evaluate the stability of micelles, the size change of different DOX-loaded micelles in normal physiological conditions within 24 h have been determined as shown in Figure S9B We can find that all the DOX-loaded

micelles

(PLL(CA)-PLA/DOX,

FA-PLL(SA)-PLA/DOX

and

FA-PLL(CA)-PLA/DOX) have shown slight changes in average particle sizes, suggesting these micelles would be stable during blood circulation. 3.3. pH-responsive behaviours and in vitro DOX release The pH-responsive behaviours of the different micelles were investigated by monitoring the changes in their zeta potentials under diverse conditions with pH ranging from 7.4 to 4.5. As shown in Figure 2D, the average zeta potentials of the micelles in a weakly alkaline environment (pH

7.4)

were

−19.1

(FA–PLL(CA)–PLA),

−20.8

(PLL(CA)–PLA)

and

−17.6mV

(FA–PLL(SA)–PLA). These values indicate that all micelles possess negative charges because of the existence of a carboxyl group in CA and SA. Meanwhile, the CA group-containing micelles in weakly acidic aqueous solutions (pH 6.5) started to carry positive charges (+16.4 mV for PLL(CA)–PLA and +15.5 mV for FA–PLL(CA)–PLA). Furthermore, the value of the positive zeta charges increased with decreasing pH. Meanwhile, the SA group-containing micelles (FA–PLL(SA)–PLA) remained negatively charged under conditions with diverse pH values. To explain, at pH 6.5, the amide bonds formed between lysine and CA in PLL(CA)–PLA and FA–PLL(CA)–PLA were cleaved, thereby exposing the positively charged amino groups. 15

16

However, FA–PLL(SA) –PLA is non-sensitive to the acidic pH, explaining the absence of a ‘charge reversal’ [7]. These results indicate that these CA group-containing micelles (PLL(CA)–PLA and FA–PLL(CA)–PLA) are extremely sensitive to pH changes, thus enabling a successful and rapid negative-to-positive charge reversal in the acidic tumour microenvironment and increasing the opportunities for binding onto the negatively charged tumour cell membrane. To further investigate the controlled release behaviour of the pH-sensitive micelles, we determined the cumulative DOX release profiles from free DOX in dialysis of the PLL(CA)–PLA/DOX, FA–PLL(SA)–PLA/DOX and FA–PLL(CA)–PLA/DOX micelles at pH 7.4 (Figure 2E) and pH 6.5 (Figure 2F). Evidently, free DOX at pH 7.4 and pH 6.5 showed a burst release, in which over 90% of the drug was released at both pH 7.4 and 6.5 within only 8 h. However, DOX encapsulated in the micelles was slowly released. Only 27% (pH 7.4) and 43.7% (pH 6.5) of DOX were released from the PLL(CA)–PLA/DOX micelles within 8 h, and the cumulative release reached 42.8% (pH 7.4) and 66.4% (pH 6.5) within 72 h. These results suggest that encapsulating the drug in the polymer micelles allows a sustainable drug release. Notably, the release behaviours of FA–PLL(CA)–PLA/DOX at both pH 7.4 and 6.5 were virtually the same as that of PLL(CA)-PLA/DOX, indicating that the attachment of FA did not affect the DOX release process. Furthermore, DOX released from FA–PLL(SA)–PLA/DOX presented a cumulative release of approximately 21% (pH 7.4) and 23% (pH 6.5) within 8 h and 34.9% (pH 7.4) and 40.7%

(pH

6.5)

within

72

h.

These

results

indicate

that

relative

to

that

in

FA–PLL(CA)–PLA/DOX, the DOX release rate in FA–PLL(SA)–PLA/DOX decreased and did not respond to the variation in pH level of the release medium.

16

17

3.4. Cytotoxicity assay To investigate the biocompatibility of the prepared co-polymers, we cultured HepG2, HeLa and A549 cells with blank micelles from the FA–PLL(CA)–PLA, PLL(CA)–PLA and FA–PLL(SA)–PLA co-polymers at concentrations ranging from 50 µg/mL to 800 µg/mL. As shown in Figures 3A–3C), a cell viability of more than 85% can be found in all groups for all micelle concentrations, suggesting that the prepared co-polymers are non-cytotoxic at the tested concentrations. 3.5. In vitro anti-tumour activity study HeLa cells were further employed for incubation with different DOX formulations (e.g. free DOX, PLL(CA)–PLA/DOX, FA–PLL(SA)–PLA/DOX and FA–PLL(CA)–PLA/DOX) for 24 h to investigate their anti-tumour effects at diverse DOX concentrations and pH 7.4 and 6.5. The cell viabilities of different groups under various DOX doses (0.025–10 µg/mL) are shown in Figures 3D and 3E). FA–PLL(CA)–PLA/DOX was the most effective treatment among all the micellar groups, possessing a half-maximal inhibitory concentration (IC50) of 2.4 µ g/mL (pH 7.4) and 0.8 µg/mL (pH 6.5), which were considerably lower than that of PLL(CA)–PLA/DOX (IC50: 8.8 µg/mL at pH 7.4 and 6.2 µg/mL at pH 6.5). This result indicates that the conjugation of FA helped increase the anti-tumour efficiency of DOX-loaded micelles, which can be attributed to the enhancement of internalisation by tumour cells via FA receptor-mediated endocytosis [35]. Meanwhile, the IC50 values of FA–PLL(SA)–PLA/DOX were measured as 4.2 µg/mL at pH 7.4 and 3.4 µg/mL at pH 6.5, respectively. This result indicates that FA–PLL(SA)–PLA/DOX is evidently less efficient than FA–PLL(CA)–PLA/DOX, especially under acidic conditions. Thus, the replacement of CA groups in the co-polymer by SA groups decreased the anti-tumour 17

18

efficiency. In other words, the CA groups grafted to the poly lysine segment via pH-sensitive amide bond can help enhance the anti-tumour effect. The amide bond-linked CA groups could be ‘cut off’ from the poly lysine segment at pH 6.5, resulting in a negative-to-positive charge reversal. Therefore, considering the conjugated FA, the internalisation of the FA––PLL(CA)–PLA/DOX micelles can be enhanced through electrostatic interaction from the negative-to-positive charge reversal in response to the acidic environment and FA receptor-mediated endocytosis. Furthermore, the large amount of amide bonds in the backbone of the poly lysine segment may induce a ‘proton sponge effect’. This effect may help facilitate the escape of FA–PLL(CA)–PLA/DOX or the other micelles from the acidic endo/lysosomes through their ability to absorb protons in the endo/lysosomes as a second-stage pH response. This phenomenon increases the osmotic pressure inside the endo/lysosomes and sequential plasma membrane disruption [36]. In addition, the IC50 values of free DOX were 1.6 µg/mL (pH 7.4) and 1.5 µg/mL (pH 6.5), which are higher than that of FA–PLL(CA)–PLA/DOX at pH 6.5 but lower than that of FA–PLL(CA)–PLA/DOX at pH 7.4. This result could be due to the ability of free DOX to directly diffuse into the tumour cells, whereas the internalisation of DOX-loaded micelles via endocytosis is time consuming. However, the internalisation of the FA–PLL(CA)–PLA/DOX micelles under acidic conditions was enhanced both through an FA receptor-mediated endocytosis and electrostatic interaction via the negative-to-positive charge reversal as described above. CLSM observation was further applied to demonstrate whether the DOX encapsulated in the micelles can be efficiently released to the nucleus after internalisation of the micelles by HeLa cells. Before observation, HeLa cells were incubated with different DOX formulations at pH 6.5 and 7.4 for 3 h. Figure 4 shows that the HeLa cells treated with the FA–PLL(CA)–PLA micelles 18

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presented the strongest red fluorescence in the nuclei at both pH 6.5 and 7.4, suggesting the successful nuclear delivery of DOX from the micelles. In addition, although the cells treated with the FA–PLL(SA)–PLA micelles displayed a relatively stronger fluorescence than the cells in the PLL(CA)–PLA group, the DOX fluorescence was mainly distributed in the cytoplasm; in particular, few DOX was found at pH 7.4. These findings indicate that although the FA–PLL(SA)–PLA micelles can be internalised more by cancer cells through FA targeting, the intracellular drug release remained limited to intracellular trafficking because of insensitivity to acidic environments. Overall, the results show that the delivery of DOX into cancer cells, especially to the nucleus, can be efficiently enhanced by the FA–PLL(CA)–PLA micelles, which are capable of FA receptor targeting, negative-to-positive charge reversal and second-stage pH response. The apoptosis of HeLa cells induced by free DOX and DOX-loaded micelles (PLL(CA)–PLA, FA–PLL(SA)–PLA and FA–PLL(CA)–PLA) after 48 h incubation at both pH 6.5 and 7.4 was also evaluated (Figure 5). The total apoptotic HeLa cell populations induced by the FA–PLL(CA)–PLA micelles were 93.7% (pH 6.5) and 70.3% (pH 7.4), which are higher than those of the DOX (80.5% (pH 6.5) and 78.1% (pH 7.4)), FA–PLL(SA)–PLA (60.8% (pH 6.5) and 52.4% (pH 7.4)) and PLL(CA)–PLA (39.86% (pH 6.5) and 32.4% (pH 7.4)) groups. These data agree with the above results and demonstrate again that the FA–PLL(CA)–PLA micelles can efficiently deliver the anti-cancer drug to cancer cells and achieve an excellent anti-cancer effect. 3.6. In vivo biodistribution To directly visualise the in vivo distribution of micelles, we administered different DIR-labelled micelles, including PLL(CA)–PLA, FA–PLL(SA)–PLA and FA–PLL(CA)–PLA, 19

20

into HeLa tumour-bearing BALB/c nude mice via the lateral tail vein. Then, the mice were anaesthetised and imaged by an in vivo imaging system at preset times (0.5, 2, 4 and 8 h). The in vivo fluorescence images in Figure 6A show that the intensity of DIR fluorescence at the tumour sites for all micelles increased with observation time post-injection. This result can be attributed to the EPR effect in tumour tissues because of their tortuous and leaky vasculatures. At 8 h post-administration, the FA–PLL(CA)–PLA group possessed the highest fluorescence intensity. The 3D images (Figure 6B) of in vivo fluorescence further confirm that the accumulation of DIR-labelled micelles was considerably higher in the tumour of FA–PLL(CA)–PLA group than that in the other groups. Furthermore, the representative ex vivo fluorescence images (Figure 6C) and the corresponding fluorescent intensity of the DIR in the various organs such as tumor, brain, liver, heart, spleen, kidneys and lungs (Figure S10) obtained by Spectrum CT, show that the tumours from the DIR group lacked fluorescence, whereas all other micelles-treated groups, presented stronger fluorescence in tumours. Furthermore, the FA–PLL(CA)–PLA group showed the strongest fluorescence which is about 20 times of the DIR group, 6 times of the PLL-(CA)-PLA group and 2 times of the FA-PLL(SA)-PLA group. Although DIR fluorescence can be found in normal tissues (e.g. liver, spleen and lung) from all groups, the FA–PLL(CA)–PLA group presented relatively lower fluorescence in the lung. Overall, these results demonstrate that all micelles from the prepared polymers can be successfully accumulated at the tumour sites. Moreover, the FA–PLL(CA)–PLA micelles presented better accumulation than the other micelles because of the enhanced active targeting effect via the combination of FA receptor targeting and electrostatic interaction between cells and micelles resulting from the charge reversal in response to the acidic tumour microenvironment. 20

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3.7. In vivo anti-tumour activity study To further evaluate the in vivo anti-tumour effect of DOX-loaded micelles, we injected various formulations (saline (as control), free DOX and DOX-loaded micelles (PLL(CA)–PLA, FA–PLL(SA)–PLA and FA–PLL(CA)–PLA)) into HeLa tumour-bearing BALB/c nude mice. Figure 7 illustrates the anti-tumour effect and toxicity of these formulations. As illustrated in Figures 7A and 7C, although tumours from all groups increased in size over time, different growth rates of tumours from different groups were observed. The tumour volume of the control groups increased considerably faster than those of the treatment groups. Among these treatment groups, the DOX group was the most ineffective compared with the micellar groups because of rapid drug clearance by the RES [37–38]. In addition, an overaccumulation in non-target tissue [39] causes the loss of a large amount of DOX, thereby reducing the anti-cancer effect. Meanwhile, encapsulation of DOX into the micelles can solve this problem to a certain degree owing to the protection of micelles and the enhanced accumulation at the tumour site conferred by the EPR effect. However, certain barriers, such as mucosal barriers and non-specific uptake, were still observed. The tumour volume of the micelles without FA conjugation was larger than those of the FA-conjugated micelles. This result may be due to the fact that FA modification can increase the anti-tumour effect of micelles through the active targeting ability to overexpressed FA receptors on the surface of tumour cells and enhanced the internalisation of drug-loaded micelles by tumour cells. In addition, the proton sponge effect of the poly lysine segment accelerated the micellar escape from the endo/lysosomes in the cells and intracellular drug release. Moreover, the FA–PLL(CA)–PLA group possessed the slowest tumour growth rate, suggesting that the CA-modified micelles featuring a negative-to-positive charge reversal can further enhance the 21

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active targeting ability of micelles and increase the internalisation opportunities of tumour cells. Histological (H&E) and immunohistochemical (caspase-3 and Ki-67,a marker of cell proliferation) analyses were applied to further confirm the anti-tumour effect (Figure 8). The fewest tumour cells and the highest level of cell apoptosis (nucleus lysis, appearance of vacuoles and destruction of membrane integrity) were visualised by H&E staining of the FA–PLL(CA)–PLA/DOX group. This finding agreed with the caspase-3 expression level, as shown in the caspase-3-stained immunohistochemical images. In these images, the activity of caspase-3 was higher in the FA–PLL(CA)–PLA/DOX group than in the other groups, suggesting that the treatment could induce caspase-dependent mitochondria-mediated apoptosis. In addition, the Ki-67 expression level of the FA–PLL(CA)–PLA/DOX group was significantly lower than that of the other groups, implying that the tumour cell proliferation was suppressed and thus its anti-tumour effect was promoted. In other words, these results support the excellent anti-cancer effect of the active targeting and pH-sensitive micelles (FA–PLL(CA)–PLA/DOX). Meanwhile, changes in survival rate (Figure 7B) showed that the FA-conjugated micelles, especially the FA-conjugated pH-responsive micelles (FA–PLL(CA)–PLA/DOX), possessed a higher survival rate and therefore offered better safety to normal tissues compared with free DOX and common micelles. 4. Conclusion We developed an active targeting and pH-sensitive micellar system for anti-cancer drug delivery based on an FA-conjugated PLA–PLL co-polymer with a pH-sensitive amide bond between CA and the PLL segment. These DOX-loaded micelles possess excellent anti-cancer effect and improved safety to normal tissues owing to the enhanced active targeting and cellular 22

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internalisation via the combination of FA-receptor mediated targeting and the electrostatic interaction from the negative-to-positive charge reversal in response to the acidic tumour microenvironment. Furthermore, the micelles can escape from the endo/lysosomes of tumour cells because of the proton sponge effect caused by the unsaturated amide bond of the PLL segment, thereby ensuring the nuclear delivery of the anti-cancer drug and the anti-cancer effect. Thus, the prepared active targeting and pH-sensitive micelles are promising tools for effective cancer therapy. Acknowledgments This work was partly supported by grants from the National Natural Science Foundation of China (81502014)，and Medical Guiding Project of Shanghai Municipal Science and Education Commission (14411960700).

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Figure Caption

Scheme 1. Illustration of the folate-conjugated pH-responsive polymeric micelles based on FA–PLL(CA)–PLA for receptor-mediated endocytosis and pH-triggered release.

Figure 1. Synthetic route of FA–PLL(CA)–PLA.

Figure 2. A) 1H NMR spectra of FA-PLL(CA)-PLA. B) The size distribution of micelles fabricated from FA-PLL(CA)-PLA copolymers measured by DLS. C) The corresponding morphologies of these micelles viewed by TEM. D) Variations in zeta potentials of the PLL(CA)–PLA, FA–PLL(SA)–PLA and FA–PLL(CA)–PLA micelles under conditions with pH values ranging from 7.4 to 4.5, suggesting the different pH sensitivities of these micelles. E) and F) In vitro DOX release profiles of different DOX formulations, including free DOX in dialysis, PLL(CA)–PLA/DOX, FA–PLL(SA)–PLA/DOX and FA–PLL(CA)–PLA/DOX in (B) PBS (pH 7.4) or (C) PBS (pH 6.5)

Figure 3. Viabilities of A) HepG2, B) HeLa and C) A549 cells treated with different concentrations of blank micelles prepared from PLL(CA)–PLA, FA–PLL(SA)–PLA and FA–PLL(CA)–PLA after 24 h (n = 3), respectively. E) and F) Viability of HeLa cells treated with different DOX formulations at DOX doses ranging from 0.025 µg/mL to 10 µg/mL for 24 h, including free DOX and DOX-loaded micelles (PLL(CA)–PLA, FA–PLL(SA)–PLA and FA–PLL(CA)–PLA)under the conditions of E) pH 7.4 and F) pH 6.5, respectively. 29

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Figure 4. A) and B) CLSM images of HeLa cells incubated with different DOX formulations at both pH 6.5 and 7.4 for 4 h, including free DOX and DOX-loaded micelles (PLL(CA)–PLA, FA–PLL(SA)–PLA and FA–PLL(CA)–PLA. DOX-equivalent dose: 5 µg/mL.

Figure 5. Cell apoptosis of HeLa cells induced by free DOX and DOX-loaded micelles (PLL(CA)–PLA, FA–PLL(SA)–PLA and FA–PLL(CA)–PLA after 24 h of incubation at both pH 6.5 and 7.4. Cells were stained with annexin V–FITC and PI for flow cytometry. DOX-equivalent dose: 5 µg/mL.

Figure 6. (A) In vivo fluorescence biodistribution of different micelles (PLL(CA)–PLA, FA––PLL(SA)–PLA and FA–PLL(CA)–PLA) labelled by fluorescent probe DIR in HeLa tumour-bearing BALB/c nude mice at different times post-injection. Tumour sites were enclosed by red circles. (B) 3D images of in vivo fluorescence biodistribution. Tumour sites were indicated by red arrows. (C) Ex vivo fluorescence imaging of the tissues and organs excised from sacrificial mice at 8 h post-injection.

Figure 7. In vivo anti-tumour effects and survival ratio of tumour-bearing BALB/c nude mice after treatment with different formulations of saline (as control), free DOX, PLL(CA)–PLA/DOX, FA–PLL(SA)–PLA/DOX and FA–PLL(CA)–PLA/DOX at an equivalent DOX dose of 5 mg/kg. (A) Variations in tumour volume; the tumor size of the saline control, DOX, PLL(CA)–PLA/DOX, FA–PLL(SA)–PLA/DOX and FA–PLL(CA)–PLA/DOX micelles were 1321 ± 107.5 mm3, 909 ± 30

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97.5mm3, 673 ± 93 mm3, 473 ± 88.4 mm3 and 309 ± 83.5mm3, respectively after the various treatment (*p<0.05)。(B)Survival ratio of tumour-bearing BABL/c nude mice after treatment. The median survival time of the saline control, DOX, PLL(CA)–PLA/DOX, FA–PLL(SA)–PLA/DOX and FA–PLL(CA)–PLA/DOX micelles were 16.5d, 20d, 25d, 29d and37d,respectively after the various treatment (p<0.05) .(C) Photos of HeLa solid tumours excised from different treatment groups at 21 day. Scale bars = 1 cm.

Figure 8. Histological analyses of tumour sections from different groups at XX days post-first treatment, including: (A) H&E, (B) caspase-3 protein and (C) Ki-67 protein analysis. Scale bars = 200 µm.

Table 1. The Mw and corresponding PDI values of different copolymers determined by GPC (DMF/LiBr as the eluent) Sample

Mw

PDI

PLA

4962

1.34

P(Z)LL-PLA

12857

1.52

PLL-PLA

9470

1.58

PLL(CA)-PLA

12071

1.63

FA-PLL(CA)-PLA

15394

1.49

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Table 2. The characteristics of different blank micelles Average Size a

Zeta Potential a

CMC b

(mV)

(mg/L)

PDI a

Sample (nm) PLL(CA)-PLA

148±2.4

0.229±0.04

-20.8±1.39

2.29±0.06

FA-PLL(SA)-PLA

132±3.7

0.186±0.03

-17.6±2.01

2.04±0.03

FA-PLL(CA)-PLA

125±4.1

0.174±0.01

-19.1±1.57

1.86±0.02

a

Determined by DLS.

b

Determined by Fluorescence spectrophotometer

Table 3. The characteristics of different DOX-loaded micelles Loading

Encapsulation

Content b

Efficiency b

(%)

(%)

Size a PDI a

Sample (nm)

PLL(CA)-PLA/DOX

154±5.2

0.308±0.05

6.39±0.78

56.3±1.87

FA-PLL(SA)-PLA/DOX

138±4.7

0.216±0.03

5.82±1.03

54.7±2.06

FA-PLL(CA)-PLA/DOX

131±6.1

0.184±0.02

6.07±0.95

59.2±1.41

a

Determined by DLS.

b

Determined by UV-vis

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Statement of Significance Negatively charged nano-carriers prolonged anti-cancer drugs’ blood circulation. However it is difficult to be internalised. Therefore, a negative-to-positive charged micelle surface could improve selectivity for tumour cells and increase uptake chance. In this study, we developed a folate (FA)-conjugated and pH-responsive active targeting micellar system for anti-cancer drug delivery.The FA receptor-mediated active targeting and electrostatic interaction between micelles and cell membrane due to a negative-to-positive charge reversal was combined in one micellar anti-cancer drug delivery system to enhance the tumour targeting and cellular internalisation of micelles. In vitro and in vivo anti-cancer studies demonstrated that the doxorubicin-loaded, FA-conjugated and pH-responsive polymeric micelles possess an enhanced and effective cancer efficiency.

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