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On-demand combinational delivery of curcumin and doxorubicin via a pH-labile micellar nanocarrier
Accepted Manuscript Title: On-demand combinational delivery of curcumin and doxorubicin via a pH-labile micellar nanocarrier Author: Haoyu Li Man Li Chao Chen Aiping Fan Deling Kong Zheng Wang Yanjun Zhao PII: DOI: Reference:
S0378-5173(15)30220-9 http://dx.doi.org/doi:10.1016/j.ijpharm.2015.09.022 IJP 15207
To appear in:
International Journal of Pharmaceutics
Received date: Revised date: Accepted date:
3-8-2015 1-9-2015 12-9-2015
Please cite this article as: Li, Haoyu, Li, Man, Chen, Chao, Fan, Aiping, Kong, Deling, Wang, Zheng, Zhao, Yanjun, On-demand combinational delivery of curcumin and doxorubicin via a pH-labile micellar nanocarrier.International Journal of Pharmaceutics http://dx.doi.org/10.1016/j.ijpharm.2015.09.022 This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting proof before it is published in its final form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.
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On-demand combinational delivery of curcumin and doxorubicin via
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a pH-labile micellar nanocarrier
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Haoyu Li, a Man Li,a Chao Chen, a Aiping Fan, a Deling Kong,b,c Zheng Wang*a,c and Yanjun Zhao*a,c
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a
School of Pharmaceutical Science & Technology, Tianjin Key Laboratory for Modern Drug Delivery & High Efficiency, and Collaborative Innovation Center of Chemical Science and Engineering (Tianjin), Tianjin University, Tianjin 300072, People’s Republic of China
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b
Key Laboratory of Bioactive Materials, Ministry of Education, College of Life Science, and Collaborative Innovation Center of Chemical Science and Engineering (Tianjin), Nankai University, Tianjin 300071, People’s Republic of China
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c
State Key Laboratory of Medicinal Chemical Biology, Nankai University, Tianjin 300071, People’s Republic of China
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To whom correspondence should be addressed.
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Dr. Yanjun Zhao or Prof. Zheng Wang
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School of Pharmaceutical Science & Technology, Tianjin University
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92 Weijin Road, Nankai District, Tianjin 300072，China
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Tel: +86-22-2740 7882, Fax: +86-22-2740 4018;
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Email: [email protected]; [email protected]
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Graphical abstract
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Highlights
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Curcumin and doxorubicin were co-loaded in a pH-responsive micellar nanocarrier.
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Both agents exhibited an accelerated drug release under acid conditions.
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Such system would enhance the therapeutic efficacy without delay of action onset.
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ABSTRACT
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The combinational delivery of doxorubicin and curcumin in a physically loaded nanocarrier offers
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the benefits of enhanced therapeutic efficacy and reduced adverse effects, but this strategy often
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suffers from the slow drug release followed by delayed onset of pharmacological action. This work
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reported the hydrazone-linked polymer-curcumin conjugate micelles containing physically loaded
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doxorubicin to address this problem; the ester-linked conjugate micelles were produced as the
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control. The pH-labile spherical micelles were less than 100 nm with a loading at 9.3 ± 0.5% (w/w,
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Curcumin) and 2.5 ± 0.1(w/w, Doxorubicin). Both agents were released at a faster rate in the
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pH-labile micelles compared to the control. The confocal laser scanning microscopy revealed a
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time-dependent co-localization of both agents in HepG2 cells. The IC50 of pH-labile conjugate
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micelles without doxorubicin in HepG2 cells was 27.7 ± 5.3 (µM), whereas the co-loaded micelles
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was lowered to 10.8 ± 3.4 (µM) (Cur-equivalent dose). The combination index calculation
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demonstrated a synergistic action of both agents in the co-loading nanocarrier. The current work
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provided an efficient nanocarrier system to achieve rapid on-demand drug release without onset
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delay of therapeutic action, which might add value to the clinical translation of the combinational
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delivery systems.
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Key words: Micelle; combinational delivery; curcumin; doxorubicin; stimuli-responsive.
1. INTRODUCTION
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Cancer treatment via nanoscale delivery systems has been promising with several commercial
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success e.g. Doxil® and Abraxane®. The benefits of cancer nanomedicine include the permeability
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and retention (EPR) effect, sufficient loading of hydrophobic agents without organic vehicle,
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transcytosis of drugs across tigh epithelial and endothelial barriers, optimized pharmacokinetics and
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biodistribution, and integrated delivery with imaging (Peer et al., 2007; Farokhzad and Langer,
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2009). Nanoparticulate codelivery of multiple active agents for anticancer therapy shows additional
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advantage of synergistic pharmacological action, decrease of effective dose of each agent, reduction
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of adverse effects, and reversal of multidrug resistance (MDR). (Parhi et al., 2012; Markovsky et al.,
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2014).
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Doxorubicin (Dox) is a potent wide-spectrum cytotoxic agent, but it has been surffering from
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the notorious carditoxicity as well as the MDR problem. One popular strategy to address this issue
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is the co-encapsulation of Dox with a pleiotropic agent, curcumin (Cur) in nanocarriers (Sadzuka et
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al., 2012; Yallapu et al., 2012; Li et al., 2015). In such combination Cur played two roles and acted
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as both an anticancer agent and a chemosensitizer for MDR supression. A variety of nanocarriers
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have been reported for co-encapsulation of Dox and Cur, including liposomes, lipid nanoparitcles,
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and polymeric nanoparticles (Misra and Sahoo, 2011; Duan et al., 2012; Abouzeid et al., 2013;
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Wang et al., 2013; Barui et al., 2014; Zhao et al., 2015).
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However, Dox and Cur co-loaded non-covalently in a single nanocarrier has to experience a
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multi-step process prior to the onset of pharmacological action. Firstly, it is compulsory for the
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water to enter the particle, followed by drug dissolution. Then, the drug molecule diffuses out off
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the particles and reaches the target site. During these steps, the drug release is often very slow,
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which presents an obstacle of rapid inception of therapeutic action (Zhao et al., 2009; Meng et al.,
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2012). It was anticipated that the generation of stimuli-responsive nanocarriers would speed up the
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cargo release leading to a fast initiation of therapeutic effect (Mura et al., 2013; Cao et al., 2015).
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Taking advantage of the acid microenvironment of endosome and lysosome inside cells (Such et al.,
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2015), the aim of this study was to design acid-labile on-demand polymeric micelles for
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coadministration of Dox and Cur for achieving rapid drug release upon cellular internalization.
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Biodegradable methoxy poly(ethylene glycol)-poly(lactic acid) (mPEG-PLA) was selected as
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the amphiphilic model polymer that was covalently linked to hydrophobic Cur via a pH-responsive
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hydrazone linker. Such polymer-curcumin conjugate maintained the amphiphilicity of mPEG-PLA
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and could self-assemble into micellar nanocarriers displaying higher physical stability (Wang et al.,
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2015). Hydrophobic Dox was physically encapsulated within such conjugate micelles (Fig. 1). As
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the negative control with a low degree of pH-sensitivity, ester-linked polymer-curcumin conjugate
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was also generated with Dox being non-covalently loaded.
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2. EXPERIMENTAL
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2.1 Materials:
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Curcumin, pyridine, N, N-Dimethylformamide (DMF), sodium dodecyl sulfate and citric acid were
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obtained from Guangfu Fine Chemical Research Institute (Tianjin, China). Doxorubicin
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hydrochloride (Dox·HCl) was purchased from huafeng lianbo Technology Co., Ltd. (Beijing,
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China). Levulinic acid, 4-Dimethylaminopyridine (DMAP), stannous octoate, glutaric anhydride,
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pyrene and hydrazine were purchased from Jingchun reagent Co., Ltd. (Shanghai, China).
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1-[3-(dimethylamino)propyl]-3-ethylcarbodiimide was from Medpep Co., Ltd. (Shanghai, China).
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Monomethoxy poly(ethylene glycol) (mPEG, 2000Da) and 4’,6-diamidino-2-phenylindole (DAPI)
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were purchased from Sigma-Aldrich (Beijing, China). D,L-lactic acid was from Daigang
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Biomaterial Co., Ltd. (Jinnan, China). 4-Nitrophenyl chloroformate was obtained from Energy
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Chemical Co., Ltd. (Shanghai, China). Zn and fetal bovine serum were purchased from Jiangtian
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Chemical Technology Co., Ltd. (Tianjin, China). Human liver cancer (HepG2) cells were sourced
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from Institute of Biomedical Engineering (Chinese Academy of Medical Sciences & Peking Union
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Medical College). Cell Counting Kit-8 (CCK-8) was purchased from Dojindo Laboratories
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(Shanghai, China). DMEM was purchased from HyClone Inc. (Logan City, Utah, USA).
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2.2 Synthesis of mPEG-PLA-Hyd-Cur
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The synthesis of curcumin derivative (Cur-L), mPEG-PLA, and mPEG-PLA-NHNH2 used a
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previous published method (Wang et al., 2015). Cur-L (0.4 g, 0.858 mmol) and
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mPEG-PLA-NHNH2 (2.0 g, 0.572 mmol) were mixed in 15 mL DMF with light protection at room
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temperature. After 48 h, the reactants were further dialyzed against diluted ammonium solution
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using a dialysis tube (molecular weight cut-off/MWCO: 3,500 Da). Afterwards, the solution was
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centrifuged (12,000 rpm/min, 5 min) prior to supernatant lyophilization. 1H NMR (600 MHz,
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CDCl3 ), δ [ppm]: 1.55 (m, -CH3 PLA repeating unit); 2.22 (s, -CH3 spacer); 2.88 (s, -CH2 spacer);
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3.38 (s, -CH3 PEG end group); 3.64 (m, -CH2 PEG repeating unit); 5.15 (m, -CH PLA repeating
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unit); 5.80 (d, 2H); 6.46-6.49 (q, 2H); 6.92-7.04 (m, 6H); 7.56-7.89 (q, 2H). The synthesis of
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ester-linked control conjugate (mPEG-PLA-Cur) utilized our previously published method (Yang et
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al., 2012).
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2.3 Conjugate self-assembly and doxorubicin loading
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The self-assembly of amphiphilic conjugates utilized a typical dialysis method (Yang et al., 2012).
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The conjugate, mPEG-PLA-Hyd-Cur (100 mg, 0.024 mmol) or mPEG-PLA-Cur (100 mg, 0.025
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mmol) was dissolved in 2 mL dimethyl sulfoxide (DMSO) prior to dialysis. In terms of loading,
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Dox·HCl (20 mg, 0.034 mmol) was first dissolved in 2 mL DMSO, followed by triethylamine
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(TEA) (14.4 μL, 0.10 mmol) supplementation. The solution was stirred in dark at room temperature
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for 24 h. Afterwards, either mPEG-PLA-Hyd-Cur (100 mg, 0.024 mmol) or mPEG-PLA-Cur (100
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mg, 0.025 mmol) was added to the above Dox solution; the excess Dox was removed via dialysis
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and the solution was filtered through a 0.45 µm PVDF syringe filter and then freeze-dried ready for
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use.
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2.4 Micelle characterization
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The determination of critical micelle concentration (CMC) of mPEG-PLA-Hyd-Cur conjugate
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micelles employed a classic fluorescence method with pyrene as the probe (Cao et al., 2015). A
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Fluorolog®3-21 fluorescence spectrometer (HORIBA Jobin Yvon, France) was used for the analysis.
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The hydrodynamic diameter and zeta potential of micelle samples were analyzed by a Malvern
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Zetasizer Nano ZS. The micelle morphology was observed by a JEM-100 CXII transmission
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electron microscope (JEOL, Japan). The quantitative analysis of Cur (Cur-L) and Dox was
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performed using a gradient HPLC (high performance liquid chromatography) elution method
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(Table 1). The Agilent 1100 HPLC coupled with a UV detector was used for the separation with a
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corresponding detection wavelength of 499 nm (Dox), 421 nm (Cur), and 411 nm (Cur-L). The
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injection volume was 20 μL with a flow rate of 1.0 mL/min at constant elution temperature (30oC).
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2.5 In vitro drug release
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Upon hydrolysis of hydrazone and ester bond, Cur-L and Cur was released from the conjugate
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micelles, respectively (Scheme 1). A diffusion-cell based method was used to monitor drug release
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in vitro (Gao et al., 2015). Briefly, 15.0 mg mPEG-PLA-Hyd-Cur/Dox and 18.0 mg
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mPEG-PLA-Cur/Dox was dissolved into 2 mL PBS buffer separately as the donor. The receiver
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fluid was PBS buffer (pH 7.4, 6.0, and 5.0) containing 5% (w/v) SDS. A regenerated cellulose
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membrane (MWCO: 3,500 Da) was used to separate the donor and receptor compartment. The
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release experiment was maintained at 37oC (n = 3). At pre-determined time points, around 0.5 mL
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receiver fluid was taken out for HPLC analysis with the same volume receiver fluid being
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supplemented. The cumulative amount of drug was plotted against time to get the release profile.
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2.6 Cytotoxicity and cellular uptake
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HepG2 cells were selected for the cytotoxicity and cellular uptake study. The cell viability
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assessment used a typical CCK-8 method (Cao et al., 2015). The cells were cultured in DMEM
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medium containing 10% fetal bovine serum and 100 U/mL penicillin-streptomycin in a humidified
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atmosphere with 5% carbon dioxide at 37oC. The cells were incubated with micelle samples as well
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as control (Cur, Cur-L, and Dox) for 24 h prior to viability evaluation and IC50 determination.
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Regarding the internalization study, the cells were maintained in a 35-mm plate at a density of 1 x
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104 cells/well containing 1 mL DMEM medium. After 24 h’s culturing in regular conditions, the
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medium was discarded, followed by cell washing with PBS (0.5 mL) twice. Then the two types of
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co-loaded micelle samples (200 µg/mL) in fresh DMEM medium (0.5 mL) were added. At different
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time points (2 h, 6 h, and 12 h) the drug-containing medium was removed and the cells were
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washed with PBS three times, followed by paraformaldehyde (4%) fixation for 20 min, PBS
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washing, and DAPI staining (300 µL, 1 µg/mL) for 5 min. In the end, the cells were washed again
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with PBS, and supplemented with DMEM medium (1 mL) ready for image-taking. The excitation
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wavelength was 405 nm (DAPI), 488 nm (Cur/Cur-L), and 559 nm (Dox), respectively.
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RESULTS AND DISCUSSION
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The hydrodynamic size of mPEG-PLA-Hyd-Cur conjugate micelles was slightly smaller
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compared to that of Cur and Dox co-loaded mPEG-PLA-Hyd-Cur/Dox micelles; both were within
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100 nm (Fig. 2). Such difference was mainly due to the presence of Dox inside the hydrophobic
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micelle core. This is a typical phenomenon for micellar encapsulation that was often observed in
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previous investigations (Torchilin et al., 2007). The transmission electron microscopy (TEM)
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analysis revealed a consistant nanoscale core diameter and a spherical morphology of both micelles
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(Fig. 3). One reason to design polymer-Cur conjugate other than polymer-Dox conjugate was
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mainly because Cur exhibited higher hydrophobicity (Log P = 3.2) compared to Dox (Log P = 1.3).
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The existance of a more hydrophobic moiety in the amphiphilic polymer-drug conjugate usually
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enhanced the driving force of conjugate self-assembly, resulting in a more stable micellar nanoarrier
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(Yang et al., 2012). This concurred well with the critical micelle concentration data (Fig. 4). The
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mPEG-PLA-Hyd-Cur conjugate micelles showed a CMC of 7.7 ± 0.1 (µg/mL) that was three times
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lower compared to the control (mPEG-PLA) micelles (Yang et al., 2012), indicating an enhanced
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nanocarrier stability. Another reason for generating polymer-Cur conjugate was due to the
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pleitropic acitivity of curucmin, for which the polymeric curcumin conjugate micellar nanosystem
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might act as an universal platform for co-deliver various therapeutic agents. There was no
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significant difference between the zeta potentials of both micelles (p > 0.05, t-test).
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The HPLC assay demonstrated that the drug loading in acid-lablie conjugate micelles was 9.3 ±
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0.5% (w/w, Cur) and 2.5 ± 0.1(w/w, Dox), respectively. In contrast, the loading in ester-linked
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control conjugate micelles was 3.6% (w/w, Cur, via 1H-NMR) and 2.0 ± 0.1 (w/w, Dox),
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correspondingly. The decreased Cur loading was primarily a consequence of difficulty in
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conjugating the phenolic hydroxy groups of Cur with the polymer. The selection of hydrazone
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linker is because of not only its pH-dependent hydrolysis, but also the high conjugation efficiency
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with satisfactory cargo loading, ambient reaction conditions and short reaction time compared to
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other linkers like acetal (Gu et al., 2013). However, there was no dramatic discrepancy in terms of
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Dox loading. In both cases, the Dox loading was not very high. This was probably due to the
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relativley low hydrophobicity of Dox. Since the hydrophobic interaction between the carrier and the
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drug was the main powder to hold the cargo within micelles, Dox displayed relatively limited
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affinity or cohesive forces with the carrier, leading to a loading of less than 3% (w/w).
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The drug release in vitro was carried out in a static Franz-type diffusion cell under sink
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conditions at different pH values (7.4, 6.0, and 5.0) (Fig. 5). A gradient HPLC elution method was
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developed for quantifying Cur and Dox at the same time. As expected, the curcumin release from
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hydrazone-linked conjugate micelles was strongly correlated to the pH of release medium. This was
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obviously due to the pH-dependent hydrolysis behavior of hydrazone. The lower the medium pH,
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the higher the hydrolysis/drug release rate (Gao et al., 2015). In contrast, although the hydrolysis
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rate of ester bond in lower pHs (e.g. 5.0 and 6.0) was elevated compared to that in pH 7.4, its
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overall pH-sensitity could not compete with hydrazone. In addition, it should be noted that for ease
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of conjugation, Cur was modified with levulinic acid to form a derivative (Cur-L) prior to
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connecting with the polymer (Scheme 1). Upon hydrolysis, the released curucmin from pH-labile
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micelles was the modified form (i.e. Cur-L), other than the parent form of curcumin (Cur). In terms
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of Dox release, the assessment via the Higuchi model revealed that the steady-state flux from
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mPEG-PLA-Hyd-Cur/Dox was significantly higher compared to that of ester-linked micelles at
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lower pH (5.0 and 6.0) (p < 0.05, t-test). However, at neutral condition (pH 7.4), the Dox release
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rate from both nanocarriers were not statistically different (p > 0.05, t-test). This trend
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commensurated with the pH-dependent hydrolysis of hydrazone (Wang et al., 2005). Since the Dox
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release was mainly controlled by the diffusion process, the nanocarrier stability and integrity would
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play a key role in this process. For acid-responsive conjugate micelles, the lower pH enabled a more
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rapid Cur-L release and hence decreased carrier stability, which explained well the more rapid Dox
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release at low pHs. The Dox release from ester-linked control micelles also showed similar trend,
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that was consistent with the Cur release profile. The in vitro release study clearly demonstrated
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pH-labile co-loading micelles could speed up the liberation of both Cur and Dox under acid
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conditions (e.g. endosomes and lysosomes).
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The cellular uptake and distribution of Dox and Cur was investigated by confocal laser
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scanning microscopy (Fig. 6). As Dox and Cur exhibited intrinsic green and red fluorescence,
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respectively, no additional probe labelling was required for image-taking. At 2 h, both agents
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primarily located outside the nuclei; whereas their nuclear entry at 6 h was well-defined. This can
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be verified by the fact that the nuclei-indicating blue color at 2 h (in the last column) was still
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distinct, but get less evident at 6 h. Both control nanocarrier (mPEG-PLA-Cur/Dox) and acid-labile
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micelles displayed similar color distribution pattern. As the same nanocarrier dosing was applied for
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both samples and the loading of Cur and Dox in the control micelles was lower than that in
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pH-labile micelles, the drug mass that were internalized by cells was less, which was reflected by
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the relatively lower color intensity for the control sample.
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The overlap of green and red color produces yellow color, which can be used for
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co-localization analysis of both active agents (French et al., 2008). At 12 h, both control and
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acid-labile micelles exhibited a decreasing degree of co-localization, which signified a higher extent
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of drug release. This was consistent with the in vitro drug release data and the release curve got
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plateaued post 12 h (Fig. 5). Theoretically, co-loaded nanocarrier without drug release would
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geneate a complete co-localization since both drugs resided in the same place. Practically, some
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active agents would be liberated from the carrier in the culture medium prior to cellular entry. In
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addition, the
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even fully liberated drug confined in the endosomes/lysomomes would give a similar
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co-localization information to that without any release (Wang et al., 2014). Another disadvantage of
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such analysis was the location of Cur (Cur-L) and its polymeric conjugate could not be
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distingguished as both the conjugate and the small molecules showed green fluorescence.
barrier of endosomal escape also hinders the true assessment of drug location since
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The derivative Cur-L was less potent than Cur, which signified that the phenolic hydroxyl
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group was vital for anticancer activity (Fig. 7). The IC50 of mPEG-PLA-Hyd-Cur conjugate
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micelles in HepG2 cells was 27.7 ± 5.3 (µM), whereas the co-loaded mPEG-PLA-Hyd-Cur/Dox
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micelles was lowered to 10.8 ± 3.4 (µM) (Cur-equivalent dose). In terms of the ester-linked control
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(mPEG-PLA-Cur), the obtained IC50 was 68.8 ± 8.0 (µM), which agreed well with the in vitro
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release results since the ester hydrolysis was a much slower process. This also indicated that Cur
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had to be released first pior to taking therapeutic action and the polymer-drug conjugate unimer
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showed negligible anticancer capability.
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The classic Chou-Talalay method was used to calculate the combination index (CI) for
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co-loading synergism evaluation (Chou 2010); CI = (DCur)/(IC50)Cur + (DDox)/(IC50)Dox, where
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(IC50)Cur and (IC50)Dox represent the IC50 of Cur and Dox when used alone, and DCur and DDox are the
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corresponding Cur and Dox doses to produce the same effect in combination. The CI theorem offers
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a quantitative assessment of drug combinations for antagomism (CI > 1), additive effect (CI =1),
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and synergism (CI < 1). In the current study, the calculated CI was 0.9, which demonstrated the
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synergism of Dox and Cur co-delivered in a pH-responsive micellar nanocarrier in spite of the huge
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potency gap between Cur (Cur-L) and Dox. Previous work on non-responsive nanocarriers for
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co-administration of Cur and Dox proved that such combination could induce apotosis enven in a
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lower concentration compared to either drug alone (Misra and Sahoo, 2011). The CI result in the
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current work concurred well with these reports.
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CONCLUSION
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In summary, the acid-responsive Dox and Cur co-delivery conjugate micelles could be an efficient
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nanosytem to achieve rapid on-demand drug release without onset delay of therapeutic action. The
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co-administration of both agents was also beneficial in terms of reducing the adverse effects of Dox,
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particularly cardiotoxicity. In addition, the pH-resposnive polymer-Cur conjugate micelles could
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also provide a versatile platform for delivering other hydrophobic anticancer agents. However, the
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current covalent and non-covalent dual loading mode could not precisely tune the drug content and
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their ratio in the carrier. This could be addressed via the co-assembly of polymer-Cur and
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polymer-Dox conjugates.
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ACKNOWLEDGEMENT
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We gratefully acknowledge the funding support from National Basic Research Program of China
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(2015CB856500), National Natural Science Foundation of China (2134068), the open fund of the
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State Key Laboratory of Medicinal Chemical Biology (Nankai Univeristy) (201503007), and
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Tianjin
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(14JCZDJC38400).
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and biology for drug carrier design. Adv. Mater. 27, 2278-2297.
309
Torchilin, V.P., 2007. Micellar nanocarriers: pharmaceutical perspectives. Pharm. Res. 24, 1-16.
310
Wang, B., Shen, Y., Zhang, Q., Li, Y., Luo, M., Liu, Z., Li, Y., Qian, Z., Gao, X., Shi, H., 2013.
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Codelivery of curcumin and doxorubicin by MPEG-PCL results in improved efficacy of
312
systemically administered chemotherapy in mice with lung cancer. Int. J. Nanomedicine 8,
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3521-3531.
314
Wang, Z., Chen, C., Liu, R., Fan, A., Kong, D., Zhao, Y., 2014. Two birds with one stone:
315
dendrimer surface engineering enables tunable periphery hydrophobicity and rapid endosomal
316
escape. Chem. Commun. 50, 14025-14028.
317
Wang, Z., Chen, C., Zhang, Q., Gao, M., Zhang, J., Kong, D., Zhao, Y., 2015. Tuning the
318
architecture of polymeric conjugate to mediate intracellular delivery of pleiotropic curcumin. Eur. J.
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Pharm. Biopharm. 90, 53-62.
320
Yallapu, M.M., Jaggi, M., Chauhan, S.C., 2012. Curcumin nanoformulations: a future
321
nanomedicine for cancer. Drug Discov. Today 17, 71-80.
322
Yang, R., Zhang, S., Kong, D, Gao, X., Zhao, Y., Wang, Z., 2012. Biodegradable
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polymer-curcumin conjugate micelles enhance the loading and delivery of low-potency curcumin.
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Pharm. Res. 29, 3512-3525.
325
Zhao, X., Chen, Q., Li, Y., Tang, H., Liu, W., Yang, X., 2015. Doxorubicin and curcumin
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co-delivery by lipid nanoparticles for enhanced treatment of diethylnitrosamine-induced
327
hepatocellular carcinoma in mice. Eur. J. Pharm. Biopharm. 93, 27-36.
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Zhao, Y., Moddaresi, M., Jones, S.A., Brown, M.B., 2009. A dynamic topical hydrofluoroalkane
329
foam to induce nanoparticle modification and drug release in situ. Eur. J. Pharm. Biopharm. 72,
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521-528.
331
LEGEND
332
Scheme 1. The chemical structure of two polymer-curcumin conjugates and their released drug
333
upon acid-catalyzed hydrolysis: mPEG-PLA-Hyd-Cur (top) and mPEG-PLA-Cur (bottom).
334
Figure 1. The chemical structure of mPEG-PLA-Hyd-Cur conjugate and illustration of conjugate
335
self-assembly, doxorubicin encapsulation in micelles, and acid-responsive cargo release. mPEG,
336
PLA, Hyd, Cur, and Dox represent methoxy poly(ethylene glycol), poly(lactic acid), hydrazone,
337
curcumin, and doxorubicin respectively.
338
Figure 2. Hydrodynamic size and zeta potential of two types of acid-responsive micelles:
339
mPEG-PLA-Hyd-Cur and mPEG-PLA-Hyd-Cur/Dox (n = 3).
340
Figure 3. TEM analysis of two types of conjugate micelles. (A and C): mPEG-PLA-Hyd-Cur; (B
341
and D): mPEG-PLA-Cur.
342
Figure 4. Critical micelle concentration (CMC) determination of mPEG-PLA-Hyd-Cur conjugate
343
micelles using pyrene probe (n = 3).
344
Figure 5. Cumulative drug release from ester- and hydrazone-linked conjugate micelles: (A)
345
Curcumin derivative (Cur-L) and (B) Doxorubicin release profile from mPEG-PLA-Hyd-Cur
346
micelles; (C) Curcumin and (D) Doxorubicin release profile from the mPEG-PLA-Cur control
347
micelles. The release study was performed under sink conditions at 37oC (n = 3).
348
Figure 6. The intracellular uptake and distribution of doxorubicin (Dox) and curcumin derivative
349
(Cur-L) in HepG2 cells by confocal laser scanning microscopy. Both agents were delivered in an
350
acid-responsive micellar nanocarrier (mPEG-PLA-Hyd-Cur/Dox; top) and an ester-linked control
351
micelle (mPEG-PLA-Cur/Dox; bottom). The green, red, and blue fluorescence represents Cur-L,
352
Dox, and nuclei, respectively. The fourth column was produced by merging the first and second
353
column; the last column was obtained by merging the first three columns. Scale bar: 10 µm.
354
Figure 7. The half maximal inhibitory concentration (IC50) of doxorubicin (Dox), curcumin (Cur),
355
levulinic acid-modified curcumin derivative (Cur-L), hydrazone-linked mPEG-PLA-Hyd-Cur
356
conjugate micelles, and the corresponding co-loaded micelles (mPEG-PLA-Hyd-Cur/Dox)
357
expressed as curcumin equivalent dose; HepG2 cells were employed for all samples (n ≥ 4).
358
Table 1. HPLC gradient elution conditions for Cur (Cur-L) and Dox quantification. Time (min)
359 360 361 362
Flow Rate (mL/min)
Chanel (A)
Chanel (B)
(water, %, v/v)
(Acetonitrile, %, v/v)
0
1.0
50
50
6
1.0
50
50
10
1.0
0
100
13
1.0
0
100
15
1.0
50
50
363
Fig. 1
364 365 366 367 368 369
370 371
Fig. 2
372
Fig. 3
373 374
375 376 377
Fig. 4
378 379
380
Fig. 5
381 382 383
384
Fig. 6
385
386 387 388
389 390 391
Fig. 7
S0378-5173(15)30220-9 http://dx.doi.org/doi:10.1016/j.ijpharm.2015.09.022 IJP 15207
To appear in:
International Journal of Pharmaceutics
Received date: Revised date: Accepted date:
3-8-2015 1-9-2015 12-9-2015
Please cite this article as: Li, Haoyu, Li, Man, Chen, Chao, Fan, Aiping, Kong, Deling, Wang, Zheng, Zhao, Yanjun, On-demand combinational delivery of curcumin and doxorubicin via a pH-labile micellar nanocarrier.International Journal of Pharmaceutics http://dx.doi.org/10.1016/j.ijpharm.2015.09.022 This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting proof before it is published in its final form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.
1
2
On-demand combinational delivery of curcumin and doxorubicin via
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a pH-labile micellar nanocarrier
4 5
Haoyu Li, a Man Li,a Chao Chen, a Aiping Fan, a Deling Kong,b,c Zheng Wang*a,c and Yanjun Zhao*a,c
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a
School of Pharmaceutical Science & Technology, Tianjin Key Laboratory for Modern Drug Delivery & High Efficiency, and Collaborative Innovation Center of Chemical Science and Engineering (Tianjin), Tianjin University, Tianjin 300072, People’s Republic of China
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b
Key Laboratory of Bioactive Materials, Ministry of Education, College of Life Science, and Collaborative Innovation Center of Chemical Science and Engineering (Tianjin), Nankai University, Tianjin 300071, People’s Republic of China
13 14
c
State Key Laboratory of Medicinal Chemical Biology, Nankai University, Tianjin 300071, People’s Republic of China
15 16 17
To whom correspondence should be addressed.
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Dr. Yanjun Zhao or Prof. Zheng Wang
19
School of Pharmaceutical Science & Technology, Tianjin University
20
92 Weijin Road, Nankai District, Tianjin 300072，China
21
Tel: +86-22-2740 7882, Fax: +86-22-2740 4018;
22
Email: [email protected]; [email protected]
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24
25
Graphical abstract
26
Highlights
27
Curcumin and doxorubicin were co-loaded in a pH-responsive micellar nanocarrier.
28
Both agents exhibited an accelerated drug release under acid conditions.
29
Such system would enhance the therapeutic efficacy without delay of action onset.
30
31
ABSTRACT
32
The combinational delivery of doxorubicin and curcumin in a physically loaded nanocarrier offers
33
the benefits of enhanced therapeutic efficacy and reduced adverse effects, but this strategy often
34
suffers from the slow drug release followed by delayed onset of pharmacological action. This work
35
reported the hydrazone-linked polymer-curcumin conjugate micelles containing physically loaded
36
doxorubicin to address this problem; the ester-linked conjugate micelles were produced as the
37
control. The pH-labile spherical micelles were less than 100 nm with a loading at 9.3 ± 0.5% (w/w,
38
Curcumin) and 2.5 ± 0.1(w/w, Doxorubicin). Both agents were released at a faster rate in the
39
pH-labile micelles compared to the control. The confocal laser scanning microscopy revealed a
40
time-dependent co-localization of both agents in HepG2 cells. The IC50 of pH-labile conjugate
41
micelles without doxorubicin in HepG2 cells was 27.7 ± 5.3 (µM), whereas the co-loaded micelles
42
was lowered to 10.8 ± 3.4 (µM) (Cur-equivalent dose). The combination index calculation
43
demonstrated a synergistic action of both agents in the co-loading nanocarrier. The current work
44
provided an efficient nanocarrier system to achieve rapid on-demand drug release without onset
45
delay of therapeutic action, which might add value to the clinical translation of the combinational
46
delivery systems.
47
48
49
Key words: Micelle; combinational delivery; curcumin; doxorubicin; stimuli-responsive.
1. INTRODUCTION
50
Cancer treatment via nanoscale delivery systems has been promising with several commercial
51
success e.g. Doxil® and Abraxane®. The benefits of cancer nanomedicine include the permeability
52
and retention (EPR) effect, sufficient loading of hydrophobic agents without organic vehicle,
53
transcytosis of drugs across tigh epithelial and endothelial barriers, optimized pharmacokinetics and
54
biodistribution, and integrated delivery with imaging (Peer et al., 2007; Farokhzad and Langer,
55
2009). Nanoparticulate codelivery of multiple active agents for anticancer therapy shows additional
56
advantage of synergistic pharmacological action, decrease of effective dose of each agent, reduction
57
of adverse effects, and reversal of multidrug resistance (MDR). (Parhi et al., 2012; Markovsky et al.,
58
2014).
59
Doxorubicin (Dox) is a potent wide-spectrum cytotoxic agent, but it has been surffering from
60
the notorious carditoxicity as well as the MDR problem. One popular strategy to address this issue
61
is the co-encapsulation of Dox with a pleiotropic agent, curcumin (Cur) in nanocarriers (Sadzuka et
62
al., 2012; Yallapu et al., 2012; Li et al., 2015). In such combination Cur played two roles and acted
63
as both an anticancer agent and a chemosensitizer for MDR supression. A variety of nanocarriers
64
have been reported for co-encapsulation of Dox and Cur, including liposomes, lipid nanoparitcles,
65
and polymeric nanoparticles (Misra and Sahoo, 2011; Duan et al., 2012; Abouzeid et al., 2013;
66
Wang et al., 2013; Barui et al., 2014; Zhao et al., 2015).
67
However, Dox and Cur co-loaded non-covalently in a single nanocarrier has to experience a
68
multi-step process prior to the onset of pharmacological action. Firstly, it is compulsory for the
69
water to enter the particle, followed by drug dissolution. Then, the drug molecule diffuses out off
70
the particles and reaches the target site. During these steps, the drug release is often very slow,
71
which presents an obstacle of rapid inception of therapeutic action (Zhao et al., 2009; Meng et al.,
72
2012). It was anticipated that the generation of stimuli-responsive nanocarriers would speed up the
73
cargo release leading to a fast initiation of therapeutic effect (Mura et al., 2013; Cao et al., 2015).
74
Taking advantage of the acid microenvironment of endosome and lysosome inside cells (Such et al.,
75
2015), the aim of this study was to design acid-labile on-demand polymeric micelles for
76
coadministration of Dox and Cur for achieving rapid drug release upon cellular internalization.
77
Biodegradable methoxy poly(ethylene glycol)-poly(lactic acid) (mPEG-PLA) was selected as
78
the amphiphilic model polymer that was covalently linked to hydrophobic Cur via a pH-responsive
79
hydrazone linker. Such polymer-curcumin conjugate maintained the amphiphilicity of mPEG-PLA
80
and could self-assemble into micellar nanocarriers displaying higher physical stability (Wang et al.,
81
2015). Hydrophobic Dox was physically encapsulated within such conjugate micelles (Fig. 1). As
82
the negative control with a low degree of pH-sensitivity, ester-linked polymer-curcumin conjugate
83
was also generated with Dox being non-covalently loaded.
84
2. EXPERIMENTAL
85
2.1 Materials:
86
Curcumin, pyridine, N, N-Dimethylformamide (DMF), sodium dodecyl sulfate and citric acid were
87
obtained from Guangfu Fine Chemical Research Institute (Tianjin, China). Doxorubicin
88
hydrochloride (Dox·HCl) was purchased from huafeng lianbo Technology Co., Ltd. (Beijing,
89
China). Levulinic acid, 4-Dimethylaminopyridine (DMAP), stannous octoate, glutaric anhydride,
90
pyrene and hydrazine were purchased from Jingchun reagent Co., Ltd. (Shanghai, China).
91
1-[3-(dimethylamino)propyl]-3-ethylcarbodiimide was from Medpep Co., Ltd. (Shanghai, China).
92
Monomethoxy poly(ethylene glycol) (mPEG, 2000Da) and 4’,6-diamidino-2-phenylindole (DAPI)
93
were purchased from Sigma-Aldrich (Beijing, China). D,L-lactic acid was from Daigang
94
Biomaterial Co., Ltd. (Jinnan, China). 4-Nitrophenyl chloroformate was obtained from Energy
95
Chemical Co., Ltd. (Shanghai, China). Zn and fetal bovine serum were purchased from Jiangtian
96
Chemical Technology Co., Ltd. (Tianjin, China). Human liver cancer (HepG2) cells were sourced
97
from Institute of Biomedical Engineering (Chinese Academy of Medical Sciences & Peking Union
98
Medical College). Cell Counting Kit-8 (CCK-8) was purchased from Dojindo Laboratories
99
(Shanghai, China). DMEM was purchased from HyClone Inc. (Logan City, Utah, USA).
100
2.2 Synthesis of mPEG-PLA-Hyd-Cur
101
The synthesis of curcumin derivative (Cur-L), mPEG-PLA, and mPEG-PLA-NHNH2 used a
102
previous published method (Wang et al., 2015). Cur-L (0.4 g, 0.858 mmol) and
103
mPEG-PLA-NHNH2 (2.0 g, 0.572 mmol) were mixed in 15 mL DMF with light protection at room
104
temperature. After 48 h, the reactants were further dialyzed against diluted ammonium solution
105
using a dialysis tube (molecular weight cut-off/MWCO: 3,500 Da). Afterwards, the solution was
106
centrifuged (12,000 rpm/min, 5 min) prior to supernatant lyophilization. 1H NMR (600 MHz,
107
CDCl3 ), δ [ppm]: 1.55 (m, -CH3 PLA repeating unit); 2.22 (s, -CH3 spacer); 2.88 (s, -CH2 spacer);
108
3.38 (s, -CH3 PEG end group); 3.64 (m, -CH2 PEG repeating unit); 5.15 (m, -CH PLA repeating
109
unit); 5.80 (d, 2H); 6.46-6.49 (q, 2H); 6.92-7.04 (m, 6H); 7.56-7.89 (q, 2H). The synthesis of
110
ester-linked control conjugate (mPEG-PLA-Cur) utilized our previously published method (Yang et
111
al., 2012).
112
2.3 Conjugate self-assembly and doxorubicin loading
113
The self-assembly of amphiphilic conjugates utilized a typical dialysis method (Yang et al., 2012).
114
The conjugate, mPEG-PLA-Hyd-Cur (100 mg, 0.024 mmol) or mPEG-PLA-Cur (100 mg, 0.025
115
mmol) was dissolved in 2 mL dimethyl sulfoxide (DMSO) prior to dialysis. In terms of loading,
116
Dox·HCl (20 mg, 0.034 mmol) was first dissolved in 2 mL DMSO, followed by triethylamine
117
(TEA) (14.4 μL, 0.10 mmol) supplementation. The solution was stirred in dark at room temperature
118
for 24 h. Afterwards, either mPEG-PLA-Hyd-Cur (100 mg, 0.024 mmol) or mPEG-PLA-Cur (100
119
mg, 0.025 mmol) was added to the above Dox solution; the excess Dox was removed via dialysis
120
and the solution was filtered through a 0.45 µm PVDF syringe filter and then freeze-dried ready for
121
use.
122
2.4 Micelle characterization
123
The determination of critical micelle concentration (CMC) of mPEG-PLA-Hyd-Cur conjugate
124
micelles employed a classic fluorescence method with pyrene as the probe (Cao et al., 2015). A
125
Fluorolog®3-21 fluorescence spectrometer (HORIBA Jobin Yvon, France) was used for the analysis.
126
The hydrodynamic diameter and zeta potential of micelle samples were analyzed by a Malvern
127
Zetasizer Nano ZS. The micelle morphology was observed by a JEM-100 CXII transmission
128
electron microscope (JEOL, Japan). The quantitative analysis of Cur (Cur-L) and Dox was
129
performed using a gradient HPLC (high performance liquid chromatography) elution method
130
(Table 1). The Agilent 1100 HPLC coupled with a UV detector was used for the separation with a
131
corresponding detection wavelength of 499 nm (Dox), 421 nm (Cur), and 411 nm (Cur-L). The
132
injection volume was 20 μL with a flow rate of 1.0 mL/min at constant elution temperature (30oC).
133
2.5 In vitro drug release
134
Upon hydrolysis of hydrazone and ester bond, Cur-L and Cur was released from the conjugate
135
micelles, respectively (Scheme 1). A diffusion-cell based method was used to monitor drug release
136
in vitro (Gao et al., 2015). Briefly, 15.0 mg mPEG-PLA-Hyd-Cur/Dox and 18.0 mg
137
mPEG-PLA-Cur/Dox was dissolved into 2 mL PBS buffer separately as the donor. The receiver
138
fluid was PBS buffer (pH 7.4, 6.0, and 5.0) containing 5% (w/v) SDS. A regenerated cellulose
139
membrane (MWCO: 3,500 Da) was used to separate the donor and receptor compartment. The
140
release experiment was maintained at 37oC (n = 3). At pre-determined time points, around 0.5 mL
141
receiver fluid was taken out for HPLC analysis with the same volume receiver fluid being
142
supplemented. The cumulative amount of drug was plotted against time to get the release profile.
143
2.6 Cytotoxicity and cellular uptake
144
HepG2 cells were selected for the cytotoxicity and cellular uptake study. The cell viability
145
assessment used a typical CCK-8 method (Cao et al., 2015). The cells were cultured in DMEM
146
medium containing 10% fetal bovine serum and 100 U/mL penicillin-streptomycin in a humidified
147
atmosphere with 5% carbon dioxide at 37oC. The cells were incubated with micelle samples as well
148
as control (Cur, Cur-L, and Dox) for 24 h prior to viability evaluation and IC50 determination.
149
Regarding the internalization study, the cells were maintained in a 35-mm plate at a density of 1 x
150
104 cells/well containing 1 mL DMEM medium. After 24 h’s culturing in regular conditions, the
151
medium was discarded, followed by cell washing with PBS (0.5 mL) twice. Then the two types of
152
co-loaded micelle samples (200 µg/mL) in fresh DMEM medium (0.5 mL) were added. At different
153
time points (2 h, 6 h, and 12 h) the drug-containing medium was removed and the cells were
154
washed with PBS three times, followed by paraformaldehyde (4%) fixation for 20 min, PBS
155
washing, and DAPI staining (300 µL, 1 µg/mL) for 5 min. In the end, the cells were washed again
156
with PBS, and supplemented with DMEM medium (1 mL) ready for image-taking. The excitation
157
wavelength was 405 nm (DAPI), 488 nm (Cur/Cur-L), and 559 nm (Dox), respectively.
158
RESULTS AND DISCUSSION
159
The hydrodynamic size of mPEG-PLA-Hyd-Cur conjugate micelles was slightly smaller
160
compared to that of Cur and Dox co-loaded mPEG-PLA-Hyd-Cur/Dox micelles; both were within
161
100 nm (Fig. 2). Such difference was mainly due to the presence of Dox inside the hydrophobic
162
micelle core. This is a typical phenomenon for micellar encapsulation that was often observed in
163
previous investigations (Torchilin et al., 2007). The transmission electron microscopy (TEM)
164
analysis revealed a consistant nanoscale core diameter and a spherical morphology of both micelles
165
(Fig. 3). One reason to design polymer-Cur conjugate other than polymer-Dox conjugate was
166
mainly because Cur exhibited higher hydrophobicity (Log P = 3.2) compared to Dox (Log P = 1.3).
167
The existance of a more hydrophobic moiety in the amphiphilic polymer-drug conjugate usually
168
enhanced the driving force of conjugate self-assembly, resulting in a more stable micellar nanoarrier
169
(Yang et al., 2012). This concurred well with the critical micelle concentration data (Fig. 4). The
170
mPEG-PLA-Hyd-Cur conjugate micelles showed a CMC of 7.7 ± 0.1 (µg/mL) that was three times
171
lower compared to the control (mPEG-PLA) micelles (Yang et al., 2012), indicating an enhanced
172
nanocarrier stability. Another reason for generating polymer-Cur conjugate was due to the
173
pleitropic acitivity of curucmin, for which the polymeric curcumin conjugate micellar nanosystem
174
might act as an universal platform for co-deliver various therapeutic agents. There was no
175
significant difference between the zeta potentials of both micelles (p > 0.05, t-test).
176
The HPLC assay demonstrated that the drug loading in acid-lablie conjugate micelles was 9.3 ±
177
0.5% (w/w, Cur) and 2.5 ± 0.1(w/w, Dox), respectively. In contrast, the loading in ester-linked
178
control conjugate micelles was 3.6% (w/w, Cur, via 1H-NMR) and 2.0 ± 0.1 (w/w, Dox),
179
correspondingly. The decreased Cur loading was primarily a consequence of difficulty in
180
conjugating the phenolic hydroxy groups of Cur with the polymer. The selection of hydrazone
181
linker is because of not only its pH-dependent hydrolysis, but also the high conjugation efficiency
182
with satisfactory cargo loading, ambient reaction conditions and short reaction time compared to
183
other linkers like acetal (Gu et al., 2013). However, there was no dramatic discrepancy in terms of
184
Dox loading. In both cases, the Dox loading was not very high. This was probably due to the
185
relativley low hydrophobicity of Dox. Since the hydrophobic interaction between the carrier and the
186
drug was the main powder to hold the cargo within micelles, Dox displayed relatively limited
187
affinity or cohesive forces with the carrier, leading to a loading of less than 3% (w/w).
188
The drug release in vitro was carried out in a static Franz-type diffusion cell under sink
189
conditions at different pH values (7.4, 6.0, and 5.0) (Fig. 5). A gradient HPLC elution method was
190
developed for quantifying Cur and Dox at the same time. As expected, the curcumin release from
191
hydrazone-linked conjugate micelles was strongly correlated to the pH of release medium. This was
192
obviously due to the pH-dependent hydrolysis behavior of hydrazone. The lower the medium pH,
193
the higher the hydrolysis/drug release rate (Gao et al., 2015). In contrast, although the hydrolysis
194
rate of ester bond in lower pHs (e.g. 5.0 and 6.0) was elevated compared to that in pH 7.4, its
195
overall pH-sensitity could not compete with hydrazone. In addition, it should be noted that for ease
196
of conjugation, Cur was modified with levulinic acid to form a derivative (Cur-L) prior to
197
connecting with the polymer (Scheme 1). Upon hydrolysis, the released curucmin from pH-labile
198
micelles was the modified form (i.e. Cur-L), other than the parent form of curcumin (Cur). In terms
199
of Dox release, the assessment via the Higuchi model revealed that the steady-state flux from
200
mPEG-PLA-Hyd-Cur/Dox was significantly higher compared to that of ester-linked micelles at
201
lower pH (5.0 and 6.0) (p < 0.05, t-test). However, at neutral condition (pH 7.4), the Dox release
202
rate from both nanocarriers were not statistically different (p > 0.05, t-test). This trend
203
commensurated with the pH-dependent hydrolysis of hydrazone (Wang et al., 2005). Since the Dox
204
release was mainly controlled by the diffusion process, the nanocarrier stability and integrity would
205
play a key role in this process. For acid-responsive conjugate micelles, the lower pH enabled a more
206
rapid Cur-L release and hence decreased carrier stability, which explained well the more rapid Dox
207
release at low pHs. The Dox release from ester-linked control micelles also showed similar trend,
208
that was consistent with the Cur release profile. The in vitro release study clearly demonstrated
209
pH-labile co-loading micelles could speed up the liberation of both Cur and Dox under acid
210
conditions (e.g. endosomes and lysosomes).
211
The cellular uptake and distribution of Dox and Cur was investigated by confocal laser
212
scanning microscopy (Fig. 6). As Dox and Cur exhibited intrinsic green and red fluorescence,
213
respectively, no additional probe labelling was required for image-taking. At 2 h, both agents
214
primarily located outside the nuclei; whereas their nuclear entry at 6 h was well-defined. This can
215
be verified by the fact that the nuclei-indicating blue color at 2 h (in the last column) was still
216
distinct, but get less evident at 6 h. Both control nanocarrier (mPEG-PLA-Cur/Dox) and acid-labile
217
micelles displayed similar color distribution pattern. As the same nanocarrier dosing was applied for
218
both samples and the loading of Cur and Dox in the control micelles was lower than that in
219
pH-labile micelles, the drug mass that were internalized by cells was less, which was reflected by
220
the relatively lower color intensity for the control sample.
221
The overlap of green and red color produces yellow color, which can be used for
222
co-localization analysis of both active agents (French et al., 2008). At 12 h, both control and
223
acid-labile micelles exhibited a decreasing degree of co-localization, which signified a higher extent
224
of drug release. This was consistent with the in vitro drug release data and the release curve got
225
plateaued post 12 h (Fig. 5). Theoretically, co-loaded nanocarrier without drug release would
226
geneate a complete co-localization since both drugs resided in the same place. Practically, some
227
active agents would be liberated from the carrier in the culture medium prior to cellular entry. In
228
addition, the
229
even fully liberated drug confined in the endosomes/lysomomes would give a similar
230
co-localization information to that without any release (Wang et al., 2014). Another disadvantage of
231
such analysis was the location of Cur (Cur-L) and its polymeric conjugate could not be
232
distingguished as both the conjugate and the small molecules showed green fluorescence.
barrier of endosomal escape also hinders the true assessment of drug location since
233
The derivative Cur-L was less potent than Cur, which signified that the phenolic hydroxyl
234
group was vital for anticancer activity (Fig. 7). The IC50 of mPEG-PLA-Hyd-Cur conjugate
235
micelles in HepG2 cells was 27.7 ± 5.3 (µM), whereas the co-loaded mPEG-PLA-Hyd-Cur/Dox
236
micelles was lowered to 10.8 ± 3.4 (µM) (Cur-equivalent dose). In terms of the ester-linked control
237
(mPEG-PLA-Cur), the obtained IC50 was 68.8 ± 8.0 (µM), which agreed well with the in vitro
238
release results since the ester hydrolysis was a much slower process. This also indicated that Cur
239
had to be released first pior to taking therapeutic action and the polymer-drug conjugate unimer
240
showed negligible anticancer capability.
241
The classic Chou-Talalay method was used to calculate the combination index (CI) for
242
co-loading synergism evaluation (Chou 2010); CI = (DCur)/(IC50)Cur + (DDox)/(IC50)Dox, where
243
(IC50)Cur and (IC50)Dox represent the IC50 of Cur and Dox when used alone, and DCur and DDox are the
244
corresponding Cur and Dox doses to produce the same effect in combination. The CI theorem offers
245
a quantitative assessment of drug combinations for antagomism (CI > 1), additive effect (CI =1),
246
and synergism (CI < 1). In the current study, the calculated CI was 0.9, which demonstrated the
247
synergism of Dox and Cur co-delivered in a pH-responsive micellar nanocarrier in spite of the huge
248
potency gap between Cur (Cur-L) and Dox. Previous work on non-responsive nanocarriers for
249
co-administration of Cur and Dox proved that such combination could induce apotosis enven in a
250
lower concentration compared to either drug alone (Misra and Sahoo, 2011). The CI result in the
251
current work concurred well with these reports.
252
CONCLUSION
253
In summary, the acid-responsive Dox and Cur co-delivery conjugate micelles could be an efficient
254
nanosytem to achieve rapid on-demand drug release without onset delay of therapeutic action. The
255
co-administration of both agents was also beneficial in terms of reducing the adverse effects of Dox,
256
particularly cardiotoxicity. In addition, the pH-resposnive polymer-Cur conjugate micelles could
257
also provide a versatile platform for delivering other hydrophobic anticancer agents. However, the
258
current covalent and non-covalent dual loading mode could not precisely tune the drug content and
259
their ratio in the carrier. This could be addressed via the co-assembly of polymer-Cur and
260
polymer-Dox conjugates.
261
ACKNOWLEDGEMENT
262
We gratefully acknowledge the funding support from National Basic Research Program of China
263
(2015CB856500), National Natural Science Foundation of China (2134068), the open fund of the
264
State Key Laboratory of Medicinal Chemical Biology (Nankai Univeristy) (201503007), and
265
Tianjin
266
(14JCZDJC38400).
267
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LEGEND
332
Scheme 1. The chemical structure of two polymer-curcumin conjugates and their released drug
333
upon acid-catalyzed hydrolysis: mPEG-PLA-Hyd-Cur (top) and mPEG-PLA-Cur (bottom).
334
Figure 1. The chemical structure of mPEG-PLA-Hyd-Cur conjugate and illustration of conjugate
335
self-assembly, doxorubicin encapsulation in micelles, and acid-responsive cargo release. mPEG,
336
PLA, Hyd, Cur, and Dox represent methoxy poly(ethylene glycol), poly(lactic acid), hydrazone,
337
curcumin, and doxorubicin respectively.
338
Figure 2. Hydrodynamic size and zeta potential of two types of acid-responsive micelles:
339
mPEG-PLA-Hyd-Cur and mPEG-PLA-Hyd-Cur/Dox (n = 3).
340
Figure 3. TEM analysis of two types of conjugate micelles. (A and C): mPEG-PLA-Hyd-Cur; (B
341
and D): mPEG-PLA-Cur.
342
Figure 4. Critical micelle concentration (CMC) determination of mPEG-PLA-Hyd-Cur conjugate
343
micelles using pyrene probe (n = 3).
344
Figure 5. Cumulative drug release from ester- and hydrazone-linked conjugate micelles: (A)
345
Curcumin derivative (Cur-L) and (B) Doxorubicin release profile from mPEG-PLA-Hyd-Cur
346
micelles; (C) Curcumin and (D) Doxorubicin release profile from the mPEG-PLA-Cur control
347
micelles. The release study was performed under sink conditions at 37oC (n = 3).
348
Figure 6. The intracellular uptake and distribution of doxorubicin (Dox) and curcumin derivative
349
(Cur-L) in HepG2 cells by confocal laser scanning microscopy. Both agents were delivered in an
350
acid-responsive micellar nanocarrier (mPEG-PLA-Hyd-Cur/Dox; top) and an ester-linked control
351
micelle (mPEG-PLA-Cur/Dox; bottom). The green, red, and blue fluorescence represents Cur-L,
352
Dox, and nuclei, respectively. The fourth column was produced by merging the first and second
353
column; the last column was obtained by merging the first three columns. Scale bar: 10 µm.
354
Figure 7. The half maximal inhibitory concentration (IC50) of doxorubicin (Dox), curcumin (Cur),
355
levulinic acid-modified curcumin derivative (Cur-L), hydrazone-linked mPEG-PLA-Hyd-Cur
356
conjugate micelles, and the corresponding co-loaded micelles (mPEG-PLA-Hyd-Cur/Dox)
357
expressed as curcumin equivalent dose; HepG2 cells were employed for all samples (n ≥ 4).
358
Table 1. HPLC gradient elution conditions for Cur (Cur-L) and Dox quantification. Time (min)
359 360 361 362
Flow Rate (mL/min)
Chanel (A)
Chanel (B)
(water, %, v/v)
(Acetonitrile, %, v/v)
0
1.0
50
50
6
1.0
50
50
10
1.0
0
100
13
1.0
0
100
15
1.0
50
50
363
Fig. 1
364 365 366 367 368 369
370 371
Fig. 2
372
Fig. 3
373 374
375 376 377
Fig. 4
378 379
380
Fig. 5
381 382 383
384
Fig. 6
385
386 387 388
389 390 391
Fig. 7