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Development and evaluation of hyaluronic acid-based polymeric micelles for targeted delivery of photosensitizer for photodynamic therapy in vitro
Journal of Drug Delivery Science and Technology 48 (2018) 414–421
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Journal of Drug Delivery Science and Technology journal homepage: www.elsevier.com/locate/jddst
Development and evaluation of hyaluronic acid-based polymeric micelles for targeted delivery of photosensitizer for photodynamic therapy in vitro
T
Xiaoling Wanga,b, Junbo Wanga, Junjun Lib, Hongxia Huanga, Xiaoyi Suna,∗, Yuanyuan Lva,∗∗ a b
Department of Pharmacy, Zhejiang University City College, Hangzhou, China Institute of Pharmaceutics, College of Pharmaceutical Sciences, Zhejiang University, Hangzhou, China
A R T I C LE I N FO
A B S T R A C T
Keywords: Drug targeting Photodynamic therapy CD44 Protoporphyrin IX Micelles
To enhance the effectiveness and specificity of photodynamic therapy (PDT) in CD44-overexpressing cancers, the hyaluronic acid-b-poly (d,l-lactide-co-glycolide) copolymer (HA-b-PLGA) was synthesized, and potent photosensitizer protoporphyrin IX (PpIX) loaded polymeric micelles were prepared. The formulation organic solvent, polymer concentration, and feeding weight ratio of PpIX to polymer were optimized. PpIX-loaded HA-b-PLGA micelles (PpIX HA-PLGA micelles) were obtained with a drug loading rate of 2%, encapsulation efficiency of 43%, average diameter of 213 nm and zeta potential of −24 mV. The PpIX HA-PLGA micelles exhibited enhanced phototoxicity toward A549 cells. A cellular uptake study indicated that the micelles were internalized by CD44-mediated endocytosis. Exposure of A549 cell spheroids to HA-PLGA micelles showed deeper penetration and improved phototoxicity compared with nanoparticle and free drug analogues. PpIX HA-PLGA micelles have great potential for CD44-targeting delivery in PDT treatment.
1. Introduction Photodynamic therapy (PDT) is a promising therapeutic treatment for the management of a variety of solid tumors and non-malignant lesions [1]. It is based on administration of a photosensitizer (PS) and light of a specific wavelength. When the PS is exposed to the light, it converts nearby oxygen to highly reactive oxygen species (ROS), which induce an intratumor cytotoxic reaction involving inflammatory innate, adaptive immune reaction culminating in eradication of residual survival cancer cells. Lung cancer, the most common cause of cancer deaths worldwide was one of the first indications for PDT [2]. In the early 1980s, PDT was introduced in human lung cancer therapy [3], and it is now adopted as an alternative treatment for patients for whom surgery is contraindicated [4]. Protoporphyrin IX (PpIX) is a PS that is endogenously produced by the heme biosynthetic pathway. However, low water solubility and aggregation in aqueous solution limit its use. Encapsulation of PpIX in nanocarriers including nanoparticles [5], liposomes [6] and polymeric micelles [7] has been used to overcome these two shortcomings. To enhance the specificity and effectiveness of PDT, long circulation copolymer micelles [8], magnetic field oriented liposomes [9], photothermal triggered nanoparticles [10], and active targeting nanoparticles have been prepared [11,12]. However, until now, there have ∗
no reports of tumor-targeting micelles developed for delivery of PpIX for PDT. Many types of tumors, including non-small cell lung cancer [13], over-express CD44 and RHAMM, which are the receptors for hyaluronic acid (HA), a naturally linear polysaccharide composed of alternating Dglucuronic acid and N-acetyl-D-glucosamine units. The interactions between CD44 and HA have pivotal roles in inflammation and cancer progression. HA is regarded as an excellent tumor-targeting moiety [14,15] because it has good biocompatibility and biodegradability, and is non-toxic and non-immunotoxic [16]. Poly (lactide-co-glycolide) (PLGA) is widely used in the development of various drug delivery systems [17,18] because its application in humans has been approved by the United States Food and Drug Administration. The hydrophobic PLGA segments in the polymeric micelles serve as the inner core for loading hydrophobic drugs, and the HA segments form the hydrophilic shell. The HA chain can be freely extended and directed toward aqueous solution, which may contribute to the longer circulation and greater active targeting efficiency. In this study, PLGA/HA block copolymer was synthesized to encapsulate a photosensitizer for active PDT of CD44 receptor-expressing cells. The physicochemical properties of PpIX HA-PLGA polymeric micelles were evaluated. The uptake in A549 cells and penetration into tumor spheroids was investigated to illustrate enhanced phototoxicity
Corresponding author. Department of Pharmacy, Zhejiang University City College, 55 Huzhou Street, Hangzhou, 310015, China. Corresponding author. Department of Pharmacy, Zhejiang University City College, 55 Huzhou Street, Hangzhou, 310015, China. E-mail addresses: [email protected] (X. Sun), [email protected] (Y. Lv).
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https://doi.org/10.1016/j.jddst.2018.10.018 Received 27 August 2017; Received in revised form 8 October 2018; Accepted 20 October 2018 Available online 21 October 2018 1773-2247/ © 2018 Elsevier B.V. All rights reserved.
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DMF (3:1, v/v) was used as organic solvent, 4 ml water was added. The copolymer concentrations were 1, 1.5, 2, 2.5, 3 mg/ml. Finally, 10 mg polymer with various dose of PpIX (0.2–2.0 mg) was dissolved in DMSO/DMF (3:1, v/v) to investigate the effect of dosage on drug loading and encapsulation rate. For administration in vitro, photosensitizer loaded micelles can be concentrated and collected by centrifugation (15,000 rpm × 20 min). As a comparison, nanoprecipitation method was used to prepare PpIX loaded PLGA nanoparticles (PpIX PLGA-NPs) [5]. Briefly, 125 mg PLGA (Resomer® RG 502H) with 6.25 mg PpIX were dissolved in 2.5 ml DMSO and dripped into 50 ml aqueous solution of the surfactant, Pluronic P-127 (1%, w/v), with agitation on a magnetic stirrer. The nanoparticles were washed and collected by centrifugation. For encapsulation efficiency and drug loading rate determination, micelles and nanoparticles were dissolved in DMSO. The fluorescence intensity was measured at excitation wavelength of 408 nm and emission wavelength of 635 nm. In vitro release was studied using dialysis. PpIX HA-PLGA micelles (20 μg PpIX) were suspended in PBS containing 0.1% Triton X-100 and placed in a dialysis bag, which was immersed into 25 ml PBS (0.1% Triton X-100). Samples were incubated at 37 °C, 100 rpm/min. Dialysate was collected for fluorescence intensity measurement while an equivalent volume of fresh medium was added. PpIX dissolved in DMSO was set as the control. The particle size and zeta potential were determined by dynamic light scattering (ZEN3690, Malvern, UK) and the transmission electron microscope (JEM-1230, JEOL, JP) was utilized for the morphological observation. The samples were diluted with purified water.
compared with PpIX-loaded PLGA nanoparticles. 2. Materials and methods 2.1. Materials Hyaluronic acid (Mw 5763 Da) was purchased from Shandong Freda Biopharmaceutical Co., Ltd. (Shandong, China). PLGA (50:50) polymers, Resomer® RG 502H, coumrain 6, 3-(4,5-Dimethyl-2-thiazolyl)2,5-diphenyl-2H-tetrazolium bromide (MTT), and 2′,7′-dichlorodihydrofluorescin diacetate (H2DCF-DA) were purchased from Sigma-Aldrich (St Louis, MO, USA). Protoporphyrin IX (PpIX), 1,4diaminobutane, 1-(3-dimethylaminopropyl)-3-ethylcarbodiimide hydrochloride (EDC·HCl), N-hydroxysuccinimide (NHS), sodium cyanoborohydride and N,N-Diisopropylethylamine were supplied by Aladdin Reagent (Shanghai, China). Human lung adenocarcinoma cell line A549 was obtained from Shanghai Institute for Biological Sciences, Chinese Academy of Sciences (Shanghai, China). DMEM, trypsin-EDTA and fetal bovine serum (FBS) were supplied by Gibco (BRL, USA). 2.2. Synthesis of hyaluronic acid-b-poly (d,l-lactide-co-glycolide) copolymers Hyaluronic acid-b-poly (d,l-lactide-co-glycolide) copolymers were synthesized by an end to end coupling strategy [19]. Typically, a terminal reductive amination reaction between HA and 1,4-diaminobutane with sodium cyanoborohydride was used to synthesize the amino-functionalized HA. Meanwhile, N-hydroxysuccinmide PLGA (PLGA-NHS) was obtained by adding PLGA-COOH into solution of NHS with EDC·HCl. Finally, the mixture of amino-functionalized HA, PLGANHS and N, N-ciisopropylethylamine was stirred to synthesize HA-bPLGA.
2.5. Internalization and cytotoxicity of PpIX HA-PLGA micelles in 2D cell cultures
The 1H NMR spectra were obtained on a Bruker (Advance DMX500, Switzerland) NMR spectrometer. The molecular weights (Mw) of HA-bPLGA, PLGA-COOH were measured by GPC (Waters, Breeze, USA). The critical micelle concentration (CMC) of HA-b-PLGA was determined by a pyrene fluorescence method. Empty micelles suspension was prepared by the similar procedure as described in “preparation and characterization of PpIX HA-PLGA micelles” in the absence of photosensitizer. The resultant suspension was diluted with deionized water to achieve various HA-b-PLGA concentrations ranged from 1 × 10−5 to 0.1 mg/ml. Pyrene in acetone (6 × 10−5 M) was transferred into a series of volumetric flasks, after which acetone was evaporated. Then, HA-b-PLGA samples were added into each flask, with a final pyrene concentration of 6 × 10−7 M. The solutions were allowed to equilibration for 12 h at 37 °C in a shaker. The fluorescence intensity was measured using a fluorescent spectrometer (RF-5301PC, Shimadzu, JP). The excitation spectrum ranged from 300 to 360 nm with the emission spectrum was fixed at 390 nm. The fluorescence intensities at 334 nm and 336 nm were recorded.
For confocal laser scanning microscopy (CLSM, FLV1000, Olympus, JP) observation of cellular uptake, A549 cells were seeded onto coverslips. Cells were incubated with: (1) free PpIX; (2) PpIX PLGA-NPs; (3) PpIX HA-PLGA micelles; (4) PpIX HA-PLGA micelles with free HA (5 mg/ml) which had an equivalent concentration of PpIX (5 μmol/l) for 2 h at 37 °C. The cells were washed, fixed with 4% paraformaldehyde and stained with DAPI. For intracellular PpIX determination, A549 cells were plated at a density of 105 cell per well. Cells were lysed by 0.15 ml 1% Triton X-100. The samples were measured by the fluorescence spectrophotometer. Phototoxicity of PpIX HA-PLGA micelles was evaluated by MTT assay. Briefly, A549 cells were plated in 96-well plates at a density of 1 × 104 cells per well. Cells were incubated with different concentrations of free PpIX, PpIX PLGA-NPs and PpIX HA-PLGA micelles in serum-free DMEM. After 6 h incubation, cells were washed and exposed to light irradiation for 30 min with a LED bulb (625–630 nm, 18 W, Everfine, Hangzhou, China) to achieve a total light dose of 3 J/cm2. Then, the cells were post-incubated in cell culture medium for 12 h. MTT solution (5 mg/ml) were added and incubated for 4 h at 37 °C. The crystal formed was dissolved using DMSO. Absorbance at 570 nm was measured using a microplate reader (Multiskan GO, Thermo Fisher, USA). Dark cytotoxicity was also determined.
2.4. Preparation and characterization of PpIX HA-PLGA micelles
2.6. ROS analysis
PpIX loaded HA-b-PLGA micelles (PpIX HA-PLGA micelles) were prepared by the solvent-dialysis method. Firstly, organic solvent was optimized. 5 mg HA-b-PLGA with 0.5 mg PpIX was dissolved in 1 ml DMSO, DMF, or DMSO/DMF. Then, 4 ml deionized water was slowly added under stirring for 15 min. Extra organic solvent was removed by dialysis (Spectra/Pro®, MWCO 12–14 kDa) against water for 12 h. Unloaded PpIX was removed though a syringe filter (0.45 μm, Millipore, USA). The effect of copolymer concentration on particle size was determined on blank micelles. 5–15 mg polymer in 1 ml DMSO/
Intracellular ROS production and fluorescence cell imaging was performed by using H2DCF-DA as the indicator. Briefly, 10 mM H2DCFDA stock solution made in DMSO was diluted by serum-free DMEM to yield a 10 μM working solution. After 6 h exposure to PpIX HA-PLGA micelles (5 μmol/l), A549 celles in 6-well plate at a density of 1 × 106 per well were washed with PBS and photoactivation was performed (3 J/cm2). Then, H2DCF-DA working solution was added and cells were incubated at 37 °C for 30 min. Cells were collected and the positive rate was detected by fluorescent-activated cell sorting (FACS, FACSCalibur,
2.3. Characterization of HA-b-PLGA copolymers
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d5). When the copolymer was added to D2O, the proton signals of PLGA were no longer detected (Fig. 1A d) suggesting that HA-b-PLGA selfassembles into a core-shell micelle structure, which shielded the PLGA chain. However, signals for both HA and PLGA were seen (Fig. 1A c) when the copolymer was dissolved in DMOS because the PLGA chain is fully stretched. The molecular weights of PLGA-COOH (19,480 g/mol) and HA-b-PLGA (23,851 g/mol) were measured. The increase in molecular weight measured supported the addition of HA (5763 g/mol). These results indicated that HA was successfully conjugated to PLGA. The amphiphilic macromolecules formed polymeric micelles in aqueous solution with HA forming the outer shell. To clarify the self-assembling behavior of HA-b-PLGA copolymer further, pyrene was employed to determine its critical micelle concentration (CMC). The intensity ratio of I336/I334 vs. logC of HA-b-PLGA was plotted (Fig. 1C) based on the red shift in the excitation spectra resulting from increasing the concentration of copolymer. Linear regression analysis was performed at low and high concentrations and the inflection point, defined as the CMC value, was calculated to be 4 mg/l. This low CMC is critical to maintain the physical stability of the polymeric micelles in macrodilution.
BD, USA) with the FITC channel (excitation = 488 nm). A negative control in the dark was also performed. PpIX solution and PpIX PLGANPs were used as the positive control. For the fluorescence observation, A549 cells were seeded in 24-well plates. The CLSM observation was performed after post-stain with H2DCF-DA. 2.7. Tumor spheroid penetration and viability In order to investigate the penetrability of HA-PLGA micelles, threedimensional spheroids model of A549 cells were employed. The tumor spheroids were grown in a 96-well plate, which was initially coated with 1.5% agarose solution prepared in serum-free DMEM. 2 × 103 A549 cells were plated per well and grown for 4–5 days to aggregate into a single spheroid. Due to the weak fluorescence signal emitted by PpIX, coumarin 6 was used as a probe. Coumarin 6 loaded micelles were prepared in the same way as the optimal PpIX HA-PLGA micelles formulation, except that PpIX was replaced by 60 μg coumarin 6. For the preparation of coumrain 6 loaded PLGA nanoparticles, 0.75 mg coumarin 6 was used. The fluorescence intensity was measured at excitation wavelength of 466 nm and emission wavelength of 504 nm. Free coumarin 6, coumarin 6 loaded PLGA nanoparticles and coumarin 6 loaded HA-PLGA micelles equivalent to 500 ng/ml coumarin 6 were added to tumor spheroids, respectively. After 4 h incubation, the spheroids were washed and fixed with 4% paraformaldehyde for CLSM observation. Z-Stack images were collected. Trypan blue exclusion assay was applied to assess spheriod viability exposed to light irradiation. Spheroids were treated with different concentration of PpIX solution, PpIX PLGA-NPs and PpIX HA-PLGA micelles in serum-free DMEM for 6 h. Then, 50 mW/cm2 irradiation (670 nm, fiber optic-coupled laser system, Changchun New Indusrties Optoelectronics Technology Co., Ltd, China) was applied for 3 min. After 12 h, sets of spheroids (10 per set) were collected and trypsinized with 0.25% trypsin-EDTA. After dispersion cells into 0.4% trypan blue, the number of viable cells and total number of cells were determined on a hemocytometer. The percentage of viable cells was calculated.
3.2. Optimal formulation selection There are 3 main parameters in micelle preparation using the dialysis method; organic solvent, polymer concentration, and feeding weight ratio of drug to polymer. The optimal organic solvent was initially established. DMF and DMSO were chosen because HA-b-PLGA and PpIX were readily dissolved in these solvents and both are miscible with water. In the drug incorporation experiment, the particle size ranged from 145.0 nm to 215.0 nm increased as the proportion of DMSO in the organic phase increased. The same trend was observed for the PDI value (Fig. 2A). The micelles with smaller size were more uniform. Micelles formed when the organic solvent exchanged with water. The solvent viscosity therefore played a critical role in the particle size. DMF, which has a relatively low viscosity exhibited quick phase separation and complete diffusion, which resulted in small polymeric micelles. In terms of drug loading rate, DMF had a negative effect. The PpIX loading rate did not exceed 1% until a greater than 75% volume ratio of DMSO was added to the organic phase (Fig. 2B). The higher loading rate could be explained by the larger size of particles prepared with DMSO [22,23]. The size of micelles is important to their fate in vivo. Smaller particles (< 200 nm) are less prone to RES uptake and can exhibit prolonged circulation times in the blood, and extravasation from leaky capillaries [24]. These particles are therefore prone to accumulation in the tumors and can enhance antitumor efficiency and reduce the toxicity of their payloads. Drug loading rate is considered to be important for evaluating the loading capacity of the preparation. The higher the loading rate achieved, the lower the amount of polymer that must be administrated. Considering this balance of particle size and loading rate, a mixture of DMSO and DMF with the volume ratio fixed at 3:1 as the organic solvent for preparation of PpIX loaded micelles that had relatively small particle size and high drug loading was finally adopted. Next, the copolymer concentration was optimized. The particle size increased with the increase of copolymer concentration. The particle size was greater than 200 nm when the concentration exceeded 2 mg/ ml (Fig. 2C). To achieve a relatively high micelle yield and small particle size, a copolymer concentration of 2 mg/ml was chosen. The relationship between polymer concentration and particle size could also be explained by the viscosity of organic phase determined by copolymer concentration. The retarded diffusion rate and enhanced aggregation of particles in the formation of micelles likely gave rise to the big particle size. A similar trend was also observed by Bei [25] when the dispersion volume was altered. Finally, the optimal feeding weight of PpIX/copolymer was determined. It was found that when the mass of HA-b-PLGA was kept
2.8. Statistical analysis Data were expressed as the mean values of at least three independent determinations ± standard deviation (S.D.). Statistical analysis was done with Student's t-test, one-way ANOVA followed by SNK test (SPSS 18.0 software) with p < 0.05 as the minimal level of significance. 3. Results and discussion 3.1. Characterization of HA-b-PLGA Polymeric micelles formed by PLGA and HA copolymers have recently received significant attention [20,21] owing to their superior stability after bulk dilution in blood circulation, favorable loading capacity for hydrophobic drugs, biodegradability, the specific tumor targeting of HA, and the good biocompatibility of PLGA. In the HA-b-PLGA synthesized according to [19] (Fig. 1B), the HA segment can be freely directed toward the outer environment while the PLGA domain in the core of the micelles directs PpIX incorporation. 1H NMR and GPC were applied to confirm the chemical structure of HA-b-PLGA. Specific peaks attributed to methyl (3H, eCH3), methylene (2H, eOCH2eC]O), and eCHe in PLGA-COOH (Fig. 1A b1-b3) as well as copolymer in DMSO (Fig. 1A c1-c3), were observed at 1.46, 5.25, and 4.91 ppm, respectively. The characteristic peak of the N-acetyl group (3H, eCOCH3) in the HA chain was observed at 1.90 ppm (Fig. 1A a4) with the methylene (2H, eCH2OH) and glucosidic H peaks appearing at 2.85–4.45 ppm (Fig. 1A a5). These characteristic peaks also appeared in the spectrum of HA-b-PLGA dissolved in DMSO (Fig. 1A c4-c5) and H2O (Fig. 1A d4416
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Fig. 1. Characterization of HA-b-PLGA. A. 1H NMR spectra of (a) HA in D2O, (b) PLGA-COOH in DMSO‑d6, (c) HA-b-PLGA in DMSO‑d6, (d) HA-b-PLGA in D2O. B. Chemical structure of HA-b-PLGA. C. The CMC values of HA-b-PLGA evaluated by a pyrene fluorescence method.
solvent. The drug loading rate and encapsulation efficiency of PpIX micelles were 2.1 ± 0.1% and 42.9 ± 2.0%, respectively. Although the drug loading was not high, it was still feasible for PDT application in vivo. PpIX was an effective photosensitizer, so the dosage needed was quite low. 1 μmol/kg and 1.68 μmol/kg PpIX could cause 50% and 70% tumor necrosis in tumor bearing mice, respectively [26]. Moreover, PpIX HA-PLGA micelles were expected to be accumulated in tumor sites, leading to much lower dosage than free PpIX in antitumor activity evaluation. The hydrodynamic size was 213.4 nm with a PDI 0.152 (Fig. 3B). The micelles exhibited a negative charge (−24.3 mV) in the physiological environment (pH 7.4, Fig. 3C), which is attributed to the
constant, increases in the theoretical drug content led to a significant increase in the drug-loading rate—from 0.5 ± 0.03% to 2.3 ± 0.08%—followed by a gradual increase when PpIX/copolymer exceeded 1:10. The encapsulation efficiency first improved and then decreased, showing an inverted bell curve. The results show that PpIX/ copolymer ratio significantly influences the drug loading rate and the encapsulation efficiency. A PpIX/copolymer ratio of 1:20 was chosen because it simultaneously allowed for a high loading rate and high encapsulation efficiency. The optimal organic phase was prepared by dissolving 10 mg of HAb-PLGA and 0.5 mg of PpIX in 1 ml DMSO/DMF (3:1, v/v) mixed
Fig. 2. Influences of formulation factors on PpIX HA-PLGA micelles. A. Influence of organic solvent composition on particle size and PDI. B. Influence of organic solvent composition on drug loading rate. C. Influence of copolymer concentration on particle size. D. Influence of feed weight ratio of PpIX to copolymer on drug loading rate. 417
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Fig. 3. Characterization of PpIX HA-PLGA micelles. A. TEM of PpIX HA-PLGA micelles. B. In vitro release of PpIX HA-PLGA micelles. C. Particle size distribution. D. Zeta potential distribution.
strongly suggested that the enhanced uptake of PpIX HA-PLGA micelles was dependent on the HA-CD44 specific interaction. Interestingly, encapsulation of the photosensitizer into PLGA nanoparticles did not improve the uptake of PpIX. Although PLGA-NPs could modulate PpIX solubility and prevent aggregation, the large particle size and unmodified surface led to the insufficient endocytosis. The intracellular PpIX levels quantitatively analyzed by fluorescence spectrophotometer (Fig. 4B) were in accordance with the results of CLSM. The cell viability was then evaluated by MTT assay. Free PpIX, PpIX PLGA-NPs, and PpIX HA-PLGA micelles showed no dark cytotoxicty in the administrated concentration range, indicating that the micelles are nontoxic when they are not irradiated with red light. All of the groups showed dose-dependent cytotoxicity toward the tumor cells (Fig. 4C) under the same illumination intensity. The phototoxicity of PpIX was greatly enhanced by HA-PLGA encapsulation (p < 0.01). The IC50 was 18 ± 2, 19 ± 1 and 4 ± 0.2 μmol/l for free PpIX, PpIX PLGA-NPs, and PpIX HA-PLGA micelles, respectively. This result suggests that the enhanced phototoxicity might be due to the active cellular uptake mediated by HA-CD44 interaction, which increased the intracellular concentration of PpIX.
negatively charged HA polymer [19,27]. The relatively high zeta potential suggested good physical stability resulting from sufficient electrostatic repulsion. After storage at 4 °C for 3 weeks, the particle size and zeta potential for PpIX HA-PLGA micelles suspension was 219.6 nm with a PDI 0.162 and −23.9 mV, respectively. The micelles were colloidally stable. Fig. 3A shows that the PpIX HA-PLGA micelles were spherical in shape. The diameter of the micelles estimated from TEM micrographs were approximately 150 nm, which is smaller than that obtained from DLS analysis. The reason for this difference is that the micelles measured by DLS were hydrated whereas those measured by TEM were dehydrated. In vitro release of PpIX from HA-PLGA micelles is shown in Fig. 3B. At all indicated time points, the level of PpIX released from HA-PpIX micelles was lower than that in PpIX solution, illustrating a sustained release lasting for 3 days. 3.3. Internalization and phototoxicity of PpIX HA-PLGA micelles Because PpIX is insolubility in both methylene chloride and chloroform, it is challenging to prepare PpIX loaded PLGA nanoparticles by the emulsion-solvent evaporation method, which is widely accepted for hydrophobic drug encapsulation [28]. The nanoprecipitation method adopted here has been reported to result in large particle size and a broad size distribution. The particle size of PpIX loaded PLGA nanoparticles (PpIX PLGA-NPs) was 320.0 nm and drug loading rate was 1.8 ± 0.1%. Both free PpIX and PpIX PLGA-NPs were used as control groups. The internalization of PpIX HA-PLGA micelles was observed by confocal laser scanning microscopy (CLSM) (Fig. 4A). It is noteworthy that the fluorescence in the PpIX HA-PLGA micelles treated group was significantly higher than that of the free PpIX and PpIX PLGA-NP treated groups. After competitive treatment with HA, the fluorescence in the PpIX HA-PLGA + HA treated group was significantly reduced in CD44 over-expressing A549 cells. The results
3.4. ROS production PDT stimulates the overproduction of ROS, which could directly induce apoptosis and necrosis in tumor cells and destroy tumor capillaries and microvasculature by disrupting endothelial cells. A host-defense mechanism involving neutrophils and dendritic cells can also be activated by an acute inflammatory response [29]. Therefore, the generation of ROS was analyzed. The ROS production (Fig. 5) correlated with the MTT results. The greater the decrease in cell viability, the higher the observed ROS level. When combined with the internalization results, these findings suggested that the enhanced intracellular 418
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Fig. 4. Internalization and cytotoxicity of HA-PLGA micelles in 2D cell cultures. A. Intracellular uptake of free PpIX, PpIX PLGA-NPs, PpIX HA-PLGA micelles and PpIX HA-PLGA micelles combined with 5 mg/ml free HA. Images were observed by CLSM. The nucleus was stained with DAPI (blue). PpIX exhibited red fluorescence. B. The quantitative analysis of intracellular PpIX (∗∗p < 0.01). C. Dark and light cytotoxicity of free PpIX, PpIX PLGA-NPs and PpIX HA-PLGA micelles against A549 cells. (For interpretation of the references to colour in this figure legend, the reader is referred to the Web version of this article.)
Fig. 5. Intracellular ROS generation on A549 cells treated with PpIX, PpIX PLGA-NPs and PpIX HA-PLGA micelles after photostimulation. A. Images of A549 cells in which the fluorescent signal developed from the ROS probe. B. The percentage of the ROS positive cells (∗∗p < 0.01). 419
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Fig. 6. Penetration and phototoxicity of HA-PLGA micelles in A549 spheroids. A. Z-Stake confocal images of spheroids incubated with (a) free coumarin 6, (b) coumarin 6 loaded PLGA nanoparticles and (c) coumarin 6 loaded HA-PLGA micelles. B. Phototoxicity of free PpIX, PpIX PLGA-NPs and PpIX HA-PLGA micelles determined by trypan blue exclusion assay. (For interpretation of the references to colour in this figure legend, the reader is referred to the Web version of this article.)
deep within the spheroids, while the free coumarin 6 and coumarin 6 PLGA-NPs exhibited little penetration. To analyze the phototoxicity of PpIX HA-PLGA micelles, the A549 spheroids were incubated with PpIX loaded micelles, PpIX PLGA-NPs, and free PpIX, and were subsequently exposed to light irradiation. Higher dose of light irradiation was applied by using the laser system to achieve phototoxicity in tumor spheroids. The laser power was safe and caused no damage to tumor cells [32]. The numbers of live and dead cells were determined using a trypan blue exclusion assay in which the dead cells were stained blue while live cells excluded the trypan blue dye because of the integrity of their membranes. HA-PLGA micelles showed significant enhancement of the phototoxicity of PpIX at the concentrations both of 5 and 10 μmol/l (Fig. 6B). Little effect was observed when cells were treated with 1 μmol/l PpIX as 100% viability was retained. The results demonstrated that the HA-PLGA micelles could enhance the penetration and increase the phototoxicity of PpIX in A549 spheroids.
delivery of PpIX by HA-PLGA micelles could significantly increase ROS production in A549 cells. 3.5. Tumor spheroid penetration and viability Traditional 2D monolayer cell cultures are not always able to simulate in vivo tumors as extracellular barriers are not reproduced. Spheroids are attractive 3D tumor models, which develop oxygen, nutrient, and energy gradients similar to those found in vivo. They have emerged as a powerful and promising predictive tool for chemotherapeutic and photodynamic therapy evaluation [30]. Thus, the penetration and phototoxicity of PpIX HA-PLGA micelles in A549 spheroids was investigated. CD44 is known to traffic between endosomes and the cell surface. Huang [31] reported that hyaluronan coated nanoparticles exhibited cellular uptake mediated by CD44 and could be transported out of the cells with their cargo. The exported nanoparticles could be taken up by neighboring cells resulting in deeper penetration into tumors in 3D tumor models. Since the fluorescence signal of PpIX was too weak to be detected by CLSM, coumarin 6 was used as a fluorescence probe. The A549 spheroids were incubated with coumarin 6 loaded HAPLGA micelles, coumarin 6 loaded PLGA nanoparticles, and coumarin 6. The uptake of the micelles was significantly higher than those of the PLGA nanoparticles and the solution control (Fig. 6A), correlating well with the 2D cell uptake results. Z-Stack confocal microscopy images of the spheroids after incubation were acquired. The micelles were found
4. Conclusion In this study, preparation and application of HA-PLGA micelles in tumor targeting delivery of PpIX in vitro were demonstrated. Following formulation optimization, PpIX HA-PLGA micelles were obtained by dialysis. The particle size was ∼200 nm, and the drug encapsulation efficiency exceeded 43%. Enhanced phototoxicity towards CD44420
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overexpressing A549 cells was observed in both 2D monolayer cell cultures and 3D tumor spheroids, which is attributed to promoted cell uptake and deeper spheroid penetration. PpIX HA-PLGA micelles are promising candidates for targeted PDT treatment of lung cancer.
cancer, Expet Opin. Drug Deliv. 7 (6) (2010) 681–703. [15] J.M. Wickens, H.O. Alsaab, P. Kesharwani, K. Bhise, M.C. Amin, R.K. Tekade, U. Gupta, A.K. Iyer, Recent advances in hyaluronic acid-decorated nanocarriers for targeted cancer therapy, Drug Discov. Today 22 (4) (2017) 665–680. [16] J.M. Liang, F. Zeng, M. Zhang, Z.Z. Pan, Y.Z. Chen, Y.E. Zeng, Y. Xu, Q. Xu, Y.Z. Huang, Green synthesis of hyaluronic acid-based silver nanoparticles and their enhanced delivery to CD44+ cancer cells, RSC Adv. 5 (54) (2015) 43733–43740. [17] F. Ramazani, W. Chen, C.F. van Nostrum, G. Storm, F. Kiessling, T. Lammers, W.E. Hennink, R.J. Kok, Strategies for encapsulation of small hydrophilic and amphiphilic drugs in PLGA microspheres: state-of-the-art and challenges, Int. J. Pharm. 499 (1–2) (2016) 358–367. [18] Y.N. Cui, M. Zhang, F. Zeng, H.Y. Jin, Q. Xu, Y.Z. Huang, Dual-targeting magnetic PLGA nanoparticles for codelivery of paclitaxel and curcumin for brain tumor therapy, ACS Appl. Mater. Interfaces 8 (47) (2016) 32159–32169. [19] J. Huang, H. Zhang, Y. Yu, Y. Chen, D. Wang, G. Zhang, G. Zhou, J. Liu, Z. Sun, D. Sun, Y. Lu, Y. Zhong, Biodegradable self-assembled nanoparticles of poly(D,Llactide-co-glycolide)/hyaluronic acid block copolymers for target delivery of docetaxel to breast cancer, Biomaterials 35 (1) (2014) 550–566. [20] P. Bhatnagar, A.B. Pant, Y. Shukla, A. Panda, K.C. Gupta, Hyaluronic acid grafted PLGA copolymer nanoparticles enhance the targeted delivery of Bromelain in Ehrlich's Ascites Carcinoma, Eur. J. Pharm. Biopharm. 105 (2016) 176–192. [21] H.K. Park, S.J. Lee, J.S. Oh, S.G. Lee, Y.I. Jeong, H.C. Lee, Smart nanoparticles based on hyaluronic acid for redox-responsive and CD44 receptor-mediated targeting of tumor, Nanoscale Res. Lett. 10 (1) (2015) 981. [22] H.J. Jeon, Y.I. Jeong, M.K. Jang, Y.H. Park, J.W. Nah, Effect of solvent on the preparation of surfactant-free poly(DL-lactide-co-glycolide) nanoparticles and norfloxacin release characteristics, Int. J. Pharm. 207 (1–2) (2000) 99–108. [23] K. Zhang, X. Tang, J. Zhang, W. Lu, X. Lin, Y. Zhang, B. Tian, H. Yang, H. He, PEGPLGA copolymers: their structure and structure-influenced drug delivery applications, J. Contr. Release 183 (2014) 77–86. [24] L.Y. Qiu, Y.H. Bae, Polymer architecture and drug delivery, Pharm. Res. 23 (1) (2006) 1–30. [25] Y.Y. Bei, X.F. Zhou, B.G. You, Z.Q. Yuan, W.L. Chen, P. Xia, Y. Liu, Y. Jin, X.J. Hu, Q.L. Zhu, C.G. Zhang, X.N. Zhang, L. Zhang, Application of the central composite design to optimize the preparation of novel micelles of harmine, Int. J. Nanomed. 8 (2013) 1795–1808. [26] C.L. Conway, I. Walker, A. Bell, D.J. Roberts, S.B. Brown, D.I. Vernon, In vivo and in vitro characterization of a protoporphyrin IX-cyclic RGD peptide conjugate for use in photodynamic therapy, Photochem. Photobiol. Sci. 7 (3) (2008) 290–298. [27] B. Xiao, M.K. Han, E. Viennois, L. Wang, M. Zhang, X. Si, D. Merlin, Hyaluronic acid-functionalized polymeric nanoparticles for colon cancer-targeted combination chemotherapy, Nanoscale 7 (42) (2015) 17745–17755. [28] F. Danhier, E. Ansorena, J.M. Silva, R. Coco, A. Le Breton, V. Préat, PLGA-based nanoparticles: an overview of biomedical applications, J. Contr. Release 161 (2) (2012) 505–522. [29] A.P. Castano, P. Mroz, M.R. Hamblin, Photodynamic therapy and anti-tumour immunity, Nat. Rev. Canc. 6 (7) (2006) 535–545. [30] L.A. Muehlmann, M.C. Rodrigues, J.P. Longo, M.P. Garcia, K.R. Py-Daniel, A.B. Veloso, P.E. de Souza, S.W. da Silva, R.B. Azevedo, Aluminium-phthalocyanine chloride nanoemulsions for anticancer photodynamic therapy: development and in vitro activity against monolayers and spheroids of human mammary adenocarcinoma MCF-7 cells, J. Nanobiotechnol. 13 (2015) 36. [31] M.H. El-Dakdouki, E. Puré, X. Huang, Development of drug loaded nanoparticles for tumor targeting. Part 2: enhancement of tumor penetration through receptor mediated transcytosis in 3D tumor models, Nanoscale 5 (9) (2013) 3904–3911. [32] J. Xue, C. Li, H. Liu, J. Wei, N. Chen, J. Huang, Optimal light dose and drug dosage in the photodynamic treatment using PHOTOCYANINE, Photodiagn. Photodyn. Ther. 8 (3) (2011) 267–274.
Acknowledgments The authors gratefully acknowledge the financial support from the National Natural Science Foundation of China (81402872), Zhejiang Provincial Natural Science Foundation of China (LY17H160002, LY17B020002, LY14C090001). Science and Technology Program of Traditional Chinese Medicine of Zhejiang Province of China (2015ZQ022). References [1] U. Chilakamarthi, L. Giribabu, Photodynamic therapy: past, present and future, Chem. Rec. 17 (8) (2017) 775–802. [2] R. Allison, G. Moghissi, G. Downie, K. Dixon, Photodynamic therapy (PDT) for lung cancer, Photodiagn. Photodyn. Ther. 8 (3) (2011) 231–239. [3] T.G. Sutedja, P.E. Postmus, Photodynamic therapy in lung cancer. A review, J. Photochem. Photobiol. B Biol. 36 (2) (1996) 199–204. [4] K. Moghissi, K. Dixon, J.A. Thorpe, M. Stringer, C. Oxtoby C, Photodynamic therapy (PDT) in early central lung cancer: a treatment option for patients ineligible for surgical resection, Thorax 62 (5) (2007) 391–395. [5] C.L. da Silva, J.O. Del Ciampo, F.C. Rossetti, M.V. Bentley, M.B. Pierre, Improved in vitro and in vivo cutaneous delivery of protoporphyrin IX from PLGA-based nanoparticles, Photochem. Photobiol. 89 (5) (2013) 1176–1184. [6] M. Przybylo, D. Glogocka, J.W. Dobrucki, K. Fraczkowska, H. Podbielska, M. Kopaczynska, T. Borowik, M. Langner, The cellular internalization of liposome encapsulated protoporphyrin IX by HeLa cells, Eur. J. Pharmaceut. Sci. 85 (2016) 39–46. [7] H. Ding, B.D. Sumer, C.W. Kessinger, Y. Dong, G. Huang, D.A. Boothman, J. Gao, Nanoscopic micelle delivery improves the photophysical properties and efficacy of photodynamic therapy of protoporphyrin IX, J. Contr. Release 151 (3) (2011) 271–277. [8] B. Li, E.H. Moriyama, F. Li, M.T. Jarvi, C. Allen, B.C. Wilson, Diblock copolymer micelles deliver hydrophobic protoporphyrin IX for photodynamic therapy, Photochem. Photobiol. 83 (6) (2007) 1505–1512. [9] H. Basoglu, M.D. Bilgin, M.M. Demir, Protoporphyrin IX-loaded magnetoliposomes as a potential drug delivery system for photodynamic therapy: fabrication, characterization and in vitro study, Photodiagn. Photodyn. Ther. 13 (2016) 81–90. [10] Y. Jang, S. Kim, S. Lee, C.M. Yoon, I. Lee, J. Jang, Graphene oxide wrapped SiO2/ TiO2 hollow nanoparticles loaded with photosensitizer for photothermal and photodynamic combination therapy, Chemistry 23 (15) (2017) 3719–3727. [11] W.P. Savarimuthu, P. Gananathan, A.P. Rao, E. Manickam, G. Singaravelu, Protoporphyrin IX-gold nanoparticle conjugates for targeted photodynamic therapy–an in-vitro study, J. Nanosci. Nanotechnol. 15 (8) (2015) 5577–5584. [12] I.T. Teng, Y.J. Chang, L.S. Wang, H.Y. Lu, L.C. Wu, C.M. Yang, C.C. Chiu, C.H. Yang, S.L. Hsu, J.A. Ho, Phospholipid-functionalized mesoporous silica nanocarriers for selective photodynamic therapy of cancer, Biomaterials 34 (30) (2013) 7462–7470. [13] G. Tripodo, A. Trapani, M.L. Torre, G. Giammona, G. Trapani, D. Mandracchia, Hyaluronic acid and its derivatives in drug delivery and imaging: recent advances and challenges, Eur. J. Pharm. Biopharm. 97 (Pt B) (2015) 400–416. [14] D.A. Ossipov, Nanostructured hyaluronic acid-based materials for active delivery to
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Journal of Drug Delivery Science and Technology journal homepage: www.elsevier.com/locate/jddst
Development and evaluation of hyaluronic acid-based polymeric micelles for targeted delivery of photosensitizer for photodynamic therapy in vitro
T
Xiaoling Wanga,b, Junbo Wanga, Junjun Lib, Hongxia Huanga, Xiaoyi Suna,∗, Yuanyuan Lva,∗∗ a b
Department of Pharmacy, Zhejiang University City College, Hangzhou, China Institute of Pharmaceutics, College of Pharmaceutical Sciences, Zhejiang University, Hangzhou, China
A R T I C LE I N FO
A B S T R A C T
Keywords: Drug targeting Photodynamic therapy CD44 Protoporphyrin IX Micelles
To enhance the effectiveness and specificity of photodynamic therapy (PDT) in CD44-overexpressing cancers, the hyaluronic acid-b-poly (d,l-lactide-co-glycolide) copolymer (HA-b-PLGA) was synthesized, and potent photosensitizer protoporphyrin IX (PpIX) loaded polymeric micelles were prepared. The formulation organic solvent, polymer concentration, and feeding weight ratio of PpIX to polymer were optimized. PpIX-loaded HA-b-PLGA micelles (PpIX HA-PLGA micelles) were obtained with a drug loading rate of 2%, encapsulation efficiency of 43%, average diameter of 213 nm and zeta potential of −24 mV. The PpIX HA-PLGA micelles exhibited enhanced phototoxicity toward A549 cells. A cellular uptake study indicated that the micelles were internalized by CD44-mediated endocytosis. Exposure of A549 cell spheroids to HA-PLGA micelles showed deeper penetration and improved phototoxicity compared with nanoparticle and free drug analogues. PpIX HA-PLGA micelles have great potential for CD44-targeting delivery in PDT treatment.
1. Introduction Photodynamic therapy (PDT) is a promising therapeutic treatment for the management of a variety of solid tumors and non-malignant lesions [1]. It is based on administration of a photosensitizer (PS) and light of a specific wavelength. When the PS is exposed to the light, it converts nearby oxygen to highly reactive oxygen species (ROS), which induce an intratumor cytotoxic reaction involving inflammatory innate, adaptive immune reaction culminating in eradication of residual survival cancer cells. Lung cancer, the most common cause of cancer deaths worldwide was one of the first indications for PDT [2]. In the early 1980s, PDT was introduced in human lung cancer therapy [3], and it is now adopted as an alternative treatment for patients for whom surgery is contraindicated [4]. Protoporphyrin IX (PpIX) is a PS that is endogenously produced by the heme biosynthetic pathway. However, low water solubility and aggregation in aqueous solution limit its use. Encapsulation of PpIX in nanocarriers including nanoparticles [5], liposomes [6] and polymeric micelles [7] has been used to overcome these two shortcomings. To enhance the specificity and effectiveness of PDT, long circulation copolymer micelles [8], magnetic field oriented liposomes [9], photothermal triggered nanoparticles [10], and active targeting nanoparticles have been prepared [11,12]. However, until now, there have ∗
no reports of tumor-targeting micelles developed for delivery of PpIX for PDT. Many types of tumors, including non-small cell lung cancer [13], over-express CD44 and RHAMM, which are the receptors for hyaluronic acid (HA), a naturally linear polysaccharide composed of alternating Dglucuronic acid and N-acetyl-D-glucosamine units. The interactions between CD44 and HA have pivotal roles in inflammation and cancer progression. HA is regarded as an excellent tumor-targeting moiety [14,15] because it has good biocompatibility and biodegradability, and is non-toxic and non-immunotoxic [16]. Poly (lactide-co-glycolide) (PLGA) is widely used in the development of various drug delivery systems [17,18] because its application in humans has been approved by the United States Food and Drug Administration. The hydrophobic PLGA segments in the polymeric micelles serve as the inner core for loading hydrophobic drugs, and the HA segments form the hydrophilic shell. The HA chain can be freely extended and directed toward aqueous solution, which may contribute to the longer circulation and greater active targeting efficiency. In this study, PLGA/HA block copolymer was synthesized to encapsulate a photosensitizer for active PDT of CD44 receptor-expressing cells. The physicochemical properties of PpIX HA-PLGA polymeric micelles were evaluated. The uptake in A549 cells and penetration into tumor spheroids was investigated to illustrate enhanced phototoxicity
Corresponding author. Department of Pharmacy, Zhejiang University City College, 55 Huzhou Street, Hangzhou, 310015, China. Corresponding author. Department of Pharmacy, Zhejiang University City College, 55 Huzhou Street, Hangzhou, 310015, China. E-mail addresses: [email protected] (X. Sun), [email protected] (Y. Lv).
∗∗
https://doi.org/10.1016/j.jddst.2018.10.018 Received 27 August 2017; Received in revised form 8 October 2018; Accepted 20 October 2018 Available online 21 October 2018 1773-2247/ © 2018 Elsevier B.V. All rights reserved.
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DMF (3:1, v/v) was used as organic solvent, 4 ml water was added. The copolymer concentrations were 1, 1.5, 2, 2.5, 3 mg/ml. Finally, 10 mg polymer with various dose of PpIX (0.2–2.0 mg) was dissolved in DMSO/DMF (3:1, v/v) to investigate the effect of dosage on drug loading and encapsulation rate. For administration in vitro, photosensitizer loaded micelles can be concentrated and collected by centrifugation (15,000 rpm × 20 min). As a comparison, nanoprecipitation method was used to prepare PpIX loaded PLGA nanoparticles (PpIX PLGA-NPs) [5]. Briefly, 125 mg PLGA (Resomer® RG 502H) with 6.25 mg PpIX were dissolved in 2.5 ml DMSO and dripped into 50 ml aqueous solution of the surfactant, Pluronic P-127 (1%, w/v), with agitation on a magnetic stirrer. The nanoparticles were washed and collected by centrifugation. For encapsulation efficiency and drug loading rate determination, micelles and nanoparticles were dissolved in DMSO. The fluorescence intensity was measured at excitation wavelength of 408 nm and emission wavelength of 635 nm. In vitro release was studied using dialysis. PpIX HA-PLGA micelles (20 μg PpIX) were suspended in PBS containing 0.1% Triton X-100 and placed in a dialysis bag, which was immersed into 25 ml PBS (0.1% Triton X-100). Samples were incubated at 37 °C, 100 rpm/min. Dialysate was collected for fluorescence intensity measurement while an equivalent volume of fresh medium was added. PpIX dissolved in DMSO was set as the control. The particle size and zeta potential were determined by dynamic light scattering (ZEN3690, Malvern, UK) and the transmission electron microscope (JEM-1230, JEOL, JP) was utilized for the morphological observation. The samples were diluted with purified water.
compared with PpIX-loaded PLGA nanoparticles. 2. Materials and methods 2.1. Materials Hyaluronic acid (Mw 5763 Da) was purchased from Shandong Freda Biopharmaceutical Co., Ltd. (Shandong, China). PLGA (50:50) polymers, Resomer® RG 502H, coumrain 6, 3-(4,5-Dimethyl-2-thiazolyl)2,5-diphenyl-2H-tetrazolium bromide (MTT), and 2′,7′-dichlorodihydrofluorescin diacetate (H2DCF-DA) were purchased from Sigma-Aldrich (St Louis, MO, USA). Protoporphyrin IX (PpIX), 1,4diaminobutane, 1-(3-dimethylaminopropyl)-3-ethylcarbodiimide hydrochloride (EDC·HCl), N-hydroxysuccinimide (NHS), sodium cyanoborohydride and N,N-Diisopropylethylamine were supplied by Aladdin Reagent (Shanghai, China). Human lung adenocarcinoma cell line A549 was obtained from Shanghai Institute for Biological Sciences, Chinese Academy of Sciences (Shanghai, China). DMEM, trypsin-EDTA and fetal bovine serum (FBS) were supplied by Gibco (BRL, USA). 2.2. Synthesis of hyaluronic acid-b-poly (d,l-lactide-co-glycolide) copolymers Hyaluronic acid-b-poly (d,l-lactide-co-glycolide) copolymers were synthesized by an end to end coupling strategy [19]. Typically, a terminal reductive amination reaction between HA and 1,4-diaminobutane with sodium cyanoborohydride was used to synthesize the amino-functionalized HA. Meanwhile, N-hydroxysuccinmide PLGA (PLGA-NHS) was obtained by adding PLGA-COOH into solution of NHS with EDC·HCl. Finally, the mixture of amino-functionalized HA, PLGANHS and N, N-ciisopropylethylamine was stirred to synthesize HA-bPLGA.
2.5. Internalization and cytotoxicity of PpIX HA-PLGA micelles in 2D cell cultures
The 1H NMR spectra were obtained on a Bruker (Advance DMX500, Switzerland) NMR spectrometer. The molecular weights (Mw) of HA-bPLGA, PLGA-COOH were measured by GPC (Waters, Breeze, USA). The critical micelle concentration (CMC) of HA-b-PLGA was determined by a pyrene fluorescence method. Empty micelles suspension was prepared by the similar procedure as described in “preparation and characterization of PpIX HA-PLGA micelles” in the absence of photosensitizer. The resultant suspension was diluted with deionized water to achieve various HA-b-PLGA concentrations ranged from 1 × 10−5 to 0.1 mg/ml. Pyrene in acetone (6 × 10−5 M) was transferred into a series of volumetric flasks, after which acetone was evaporated. Then, HA-b-PLGA samples were added into each flask, with a final pyrene concentration of 6 × 10−7 M. The solutions were allowed to equilibration for 12 h at 37 °C in a shaker. The fluorescence intensity was measured using a fluorescent spectrometer (RF-5301PC, Shimadzu, JP). The excitation spectrum ranged from 300 to 360 nm with the emission spectrum was fixed at 390 nm. The fluorescence intensities at 334 nm and 336 nm were recorded.
For confocal laser scanning microscopy (CLSM, FLV1000, Olympus, JP) observation of cellular uptake, A549 cells were seeded onto coverslips. Cells were incubated with: (1) free PpIX; (2) PpIX PLGA-NPs; (3) PpIX HA-PLGA micelles; (4) PpIX HA-PLGA micelles with free HA (5 mg/ml) which had an equivalent concentration of PpIX (5 μmol/l) for 2 h at 37 °C. The cells were washed, fixed with 4% paraformaldehyde and stained with DAPI. For intracellular PpIX determination, A549 cells were plated at a density of 105 cell per well. Cells were lysed by 0.15 ml 1% Triton X-100. The samples were measured by the fluorescence spectrophotometer. Phototoxicity of PpIX HA-PLGA micelles was evaluated by MTT assay. Briefly, A549 cells were plated in 96-well plates at a density of 1 × 104 cells per well. Cells were incubated with different concentrations of free PpIX, PpIX PLGA-NPs and PpIX HA-PLGA micelles in serum-free DMEM. After 6 h incubation, cells were washed and exposed to light irradiation for 30 min with a LED bulb (625–630 nm, 18 W, Everfine, Hangzhou, China) to achieve a total light dose of 3 J/cm2. Then, the cells were post-incubated in cell culture medium for 12 h. MTT solution (5 mg/ml) were added and incubated for 4 h at 37 °C. The crystal formed was dissolved using DMSO. Absorbance at 570 nm was measured using a microplate reader (Multiskan GO, Thermo Fisher, USA). Dark cytotoxicity was also determined.
2.4. Preparation and characterization of PpIX HA-PLGA micelles
2.6. ROS analysis
PpIX loaded HA-b-PLGA micelles (PpIX HA-PLGA micelles) were prepared by the solvent-dialysis method. Firstly, organic solvent was optimized. 5 mg HA-b-PLGA with 0.5 mg PpIX was dissolved in 1 ml DMSO, DMF, or DMSO/DMF. Then, 4 ml deionized water was slowly added under stirring for 15 min. Extra organic solvent was removed by dialysis (Spectra/Pro®, MWCO 12–14 kDa) against water for 12 h. Unloaded PpIX was removed though a syringe filter (0.45 μm, Millipore, USA). The effect of copolymer concentration on particle size was determined on blank micelles. 5–15 mg polymer in 1 ml DMSO/
Intracellular ROS production and fluorescence cell imaging was performed by using H2DCF-DA as the indicator. Briefly, 10 mM H2DCFDA stock solution made in DMSO was diluted by serum-free DMEM to yield a 10 μM working solution. After 6 h exposure to PpIX HA-PLGA micelles (5 μmol/l), A549 celles in 6-well plate at a density of 1 × 106 per well were washed with PBS and photoactivation was performed (3 J/cm2). Then, H2DCF-DA working solution was added and cells were incubated at 37 °C for 30 min. Cells were collected and the positive rate was detected by fluorescent-activated cell sorting (FACS, FACSCalibur,
2.3. Characterization of HA-b-PLGA copolymers
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d5). When the copolymer was added to D2O, the proton signals of PLGA were no longer detected (Fig. 1A d) suggesting that HA-b-PLGA selfassembles into a core-shell micelle structure, which shielded the PLGA chain. However, signals for both HA and PLGA were seen (Fig. 1A c) when the copolymer was dissolved in DMOS because the PLGA chain is fully stretched. The molecular weights of PLGA-COOH (19,480 g/mol) and HA-b-PLGA (23,851 g/mol) were measured. The increase in molecular weight measured supported the addition of HA (5763 g/mol). These results indicated that HA was successfully conjugated to PLGA. The amphiphilic macromolecules formed polymeric micelles in aqueous solution with HA forming the outer shell. To clarify the self-assembling behavior of HA-b-PLGA copolymer further, pyrene was employed to determine its critical micelle concentration (CMC). The intensity ratio of I336/I334 vs. logC of HA-b-PLGA was plotted (Fig. 1C) based on the red shift in the excitation spectra resulting from increasing the concentration of copolymer. Linear regression analysis was performed at low and high concentrations and the inflection point, defined as the CMC value, was calculated to be 4 mg/l. This low CMC is critical to maintain the physical stability of the polymeric micelles in macrodilution.
BD, USA) with the FITC channel (excitation = 488 nm). A negative control in the dark was also performed. PpIX solution and PpIX PLGANPs were used as the positive control. For the fluorescence observation, A549 cells were seeded in 24-well plates. The CLSM observation was performed after post-stain with H2DCF-DA. 2.7. Tumor spheroid penetration and viability In order to investigate the penetrability of HA-PLGA micelles, threedimensional spheroids model of A549 cells were employed. The tumor spheroids were grown in a 96-well plate, which was initially coated with 1.5% agarose solution prepared in serum-free DMEM. 2 × 103 A549 cells were plated per well and grown for 4–5 days to aggregate into a single spheroid. Due to the weak fluorescence signal emitted by PpIX, coumarin 6 was used as a probe. Coumarin 6 loaded micelles were prepared in the same way as the optimal PpIX HA-PLGA micelles formulation, except that PpIX was replaced by 60 μg coumarin 6. For the preparation of coumrain 6 loaded PLGA nanoparticles, 0.75 mg coumarin 6 was used. The fluorescence intensity was measured at excitation wavelength of 466 nm and emission wavelength of 504 nm. Free coumarin 6, coumarin 6 loaded PLGA nanoparticles and coumarin 6 loaded HA-PLGA micelles equivalent to 500 ng/ml coumarin 6 were added to tumor spheroids, respectively. After 4 h incubation, the spheroids were washed and fixed with 4% paraformaldehyde for CLSM observation. Z-Stack images were collected. Trypan blue exclusion assay was applied to assess spheriod viability exposed to light irradiation. Spheroids were treated with different concentration of PpIX solution, PpIX PLGA-NPs and PpIX HA-PLGA micelles in serum-free DMEM for 6 h. Then, 50 mW/cm2 irradiation (670 nm, fiber optic-coupled laser system, Changchun New Indusrties Optoelectronics Technology Co., Ltd, China) was applied for 3 min. After 12 h, sets of spheroids (10 per set) were collected and trypsinized with 0.25% trypsin-EDTA. After dispersion cells into 0.4% trypan blue, the number of viable cells and total number of cells were determined on a hemocytometer. The percentage of viable cells was calculated.
3.2. Optimal formulation selection There are 3 main parameters in micelle preparation using the dialysis method; organic solvent, polymer concentration, and feeding weight ratio of drug to polymer. The optimal organic solvent was initially established. DMF and DMSO were chosen because HA-b-PLGA and PpIX were readily dissolved in these solvents and both are miscible with water. In the drug incorporation experiment, the particle size ranged from 145.0 nm to 215.0 nm increased as the proportion of DMSO in the organic phase increased. The same trend was observed for the PDI value (Fig. 2A). The micelles with smaller size were more uniform. Micelles formed when the organic solvent exchanged with water. The solvent viscosity therefore played a critical role in the particle size. DMF, which has a relatively low viscosity exhibited quick phase separation and complete diffusion, which resulted in small polymeric micelles. In terms of drug loading rate, DMF had a negative effect. The PpIX loading rate did not exceed 1% until a greater than 75% volume ratio of DMSO was added to the organic phase (Fig. 2B). The higher loading rate could be explained by the larger size of particles prepared with DMSO [22,23]. The size of micelles is important to their fate in vivo. Smaller particles (< 200 nm) are less prone to RES uptake and can exhibit prolonged circulation times in the blood, and extravasation from leaky capillaries [24]. These particles are therefore prone to accumulation in the tumors and can enhance antitumor efficiency and reduce the toxicity of their payloads. Drug loading rate is considered to be important for evaluating the loading capacity of the preparation. The higher the loading rate achieved, the lower the amount of polymer that must be administrated. Considering this balance of particle size and loading rate, a mixture of DMSO and DMF with the volume ratio fixed at 3:1 as the organic solvent for preparation of PpIX loaded micelles that had relatively small particle size and high drug loading was finally adopted. Next, the copolymer concentration was optimized. The particle size increased with the increase of copolymer concentration. The particle size was greater than 200 nm when the concentration exceeded 2 mg/ ml (Fig. 2C). To achieve a relatively high micelle yield and small particle size, a copolymer concentration of 2 mg/ml was chosen. The relationship between polymer concentration and particle size could also be explained by the viscosity of organic phase determined by copolymer concentration. The retarded diffusion rate and enhanced aggregation of particles in the formation of micelles likely gave rise to the big particle size. A similar trend was also observed by Bei [25] when the dispersion volume was altered. Finally, the optimal feeding weight of PpIX/copolymer was determined. It was found that when the mass of HA-b-PLGA was kept
2.8. Statistical analysis Data were expressed as the mean values of at least three independent determinations ± standard deviation (S.D.). Statistical analysis was done with Student's t-test, one-way ANOVA followed by SNK test (SPSS 18.0 software) with p < 0.05 as the minimal level of significance. 3. Results and discussion 3.1. Characterization of HA-b-PLGA Polymeric micelles formed by PLGA and HA copolymers have recently received significant attention [20,21] owing to their superior stability after bulk dilution in blood circulation, favorable loading capacity for hydrophobic drugs, biodegradability, the specific tumor targeting of HA, and the good biocompatibility of PLGA. In the HA-b-PLGA synthesized according to [19] (Fig. 1B), the HA segment can be freely directed toward the outer environment while the PLGA domain in the core of the micelles directs PpIX incorporation. 1H NMR and GPC were applied to confirm the chemical structure of HA-b-PLGA. Specific peaks attributed to methyl (3H, eCH3), methylene (2H, eOCH2eC]O), and eCHe in PLGA-COOH (Fig. 1A b1-b3) as well as copolymer in DMSO (Fig. 1A c1-c3), were observed at 1.46, 5.25, and 4.91 ppm, respectively. The characteristic peak of the N-acetyl group (3H, eCOCH3) in the HA chain was observed at 1.90 ppm (Fig. 1A a4) with the methylene (2H, eCH2OH) and glucosidic H peaks appearing at 2.85–4.45 ppm (Fig. 1A a5). These characteristic peaks also appeared in the spectrum of HA-b-PLGA dissolved in DMSO (Fig. 1A c4-c5) and H2O (Fig. 1A d4416
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Fig. 1. Characterization of HA-b-PLGA. A. 1H NMR spectra of (a) HA in D2O, (b) PLGA-COOH in DMSO‑d6, (c) HA-b-PLGA in DMSO‑d6, (d) HA-b-PLGA in D2O. B. Chemical structure of HA-b-PLGA. C. The CMC values of HA-b-PLGA evaluated by a pyrene fluorescence method.
solvent. The drug loading rate and encapsulation efficiency of PpIX micelles were 2.1 ± 0.1% and 42.9 ± 2.0%, respectively. Although the drug loading was not high, it was still feasible for PDT application in vivo. PpIX was an effective photosensitizer, so the dosage needed was quite low. 1 μmol/kg and 1.68 μmol/kg PpIX could cause 50% and 70% tumor necrosis in tumor bearing mice, respectively [26]. Moreover, PpIX HA-PLGA micelles were expected to be accumulated in tumor sites, leading to much lower dosage than free PpIX in antitumor activity evaluation. The hydrodynamic size was 213.4 nm with a PDI 0.152 (Fig. 3B). The micelles exhibited a negative charge (−24.3 mV) in the physiological environment (pH 7.4, Fig. 3C), which is attributed to the
constant, increases in the theoretical drug content led to a significant increase in the drug-loading rate—from 0.5 ± 0.03% to 2.3 ± 0.08%—followed by a gradual increase when PpIX/copolymer exceeded 1:10. The encapsulation efficiency first improved and then decreased, showing an inverted bell curve. The results show that PpIX/ copolymer ratio significantly influences the drug loading rate and the encapsulation efficiency. A PpIX/copolymer ratio of 1:20 was chosen because it simultaneously allowed for a high loading rate and high encapsulation efficiency. The optimal organic phase was prepared by dissolving 10 mg of HAb-PLGA and 0.5 mg of PpIX in 1 ml DMSO/DMF (3:1, v/v) mixed
Fig. 2. Influences of formulation factors on PpIX HA-PLGA micelles. A. Influence of organic solvent composition on particle size and PDI. B. Influence of organic solvent composition on drug loading rate. C. Influence of copolymer concentration on particle size. D. Influence of feed weight ratio of PpIX to copolymer on drug loading rate. 417
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Fig. 3. Characterization of PpIX HA-PLGA micelles. A. TEM of PpIX HA-PLGA micelles. B. In vitro release of PpIX HA-PLGA micelles. C. Particle size distribution. D. Zeta potential distribution.
strongly suggested that the enhanced uptake of PpIX HA-PLGA micelles was dependent on the HA-CD44 specific interaction. Interestingly, encapsulation of the photosensitizer into PLGA nanoparticles did not improve the uptake of PpIX. Although PLGA-NPs could modulate PpIX solubility and prevent aggregation, the large particle size and unmodified surface led to the insufficient endocytosis. The intracellular PpIX levels quantitatively analyzed by fluorescence spectrophotometer (Fig. 4B) were in accordance with the results of CLSM. The cell viability was then evaluated by MTT assay. Free PpIX, PpIX PLGA-NPs, and PpIX HA-PLGA micelles showed no dark cytotoxicty in the administrated concentration range, indicating that the micelles are nontoxic when they are not irradiated with red light. All of the groups showed dose-dependent cytotoxicity toward the tumor cells (Fig. 4C) under the same illumination intensity. The phototoxicity of PpIX was greatly enhanced by HA-PLGA encapsulation (p < 0.01). The IC50 was 18 ± 2, 19 ± 1 and 4 ± 0.2 μmol/l for free PpIX, PpIX PLGA-NPs, and PpIX HA-PLGA micelles, respectively. This result suggests that the enhanced phototoxicity might be due to the active cellular uptake mediated by HA-CD44 interaction, which increased the intracellular concentration of PpIX.
negatively charged HA polymer [19,27]. The relatively high zeta potential suggested good physical stability resulting from sufficient electrostatic repulsion. After storage at 4 °C for 3 weeks, the particle size and zeta potential for PpIX HA-PLGA micelles suspension was 219.6 nm with a PDI 0.162 and −23.9 mV, respectively. The micelles were colloidally stable. Fig. 3A shows that the PpIX HA-PLGA micelles were spherical in shape. The diameter of the micelles estimated from TEM micrographs were approximately 150 nm, which is smaller than that obtained from DLS analysis. The reason for this difference is that the micelles measured by DLS were hydrated whereas those measured by TEM were dehydrated. In vitro release of PpIX from HA-PLGA micelles is shown in Fig. 3B. At all indicated time points, the level of PpIX released from HA-PpIX micelles was lower than that in PpIX solution, illustrating a sustained release lasting for 3 days. 3.3. Internalization and phototoxicity of PpIX HA-PLGA micelles Because PpIX is insolubility in both methylene chloride and chloroform, it is challenging to prepare PpIX loaded PLGA nanoparticles by the emulsion-solvent evaporation method, which is widely accepted for hydrophobic drug encapsulation [28]. The nanoprecipitation method adopted here has been reported to result in large particle size and a broad size distribution. The particle size of PpIX loaded PLGA nanoparticles (PpIX PLGA-NPs) was 320.0 nm and drug loading rate was 1.8 ± 0.1%. Both free PpIX and PpIX PLGA-NPs were used as control groups. The internalization of PpIX HA-PLGA micelles was observed by confocal laser scanning microscopy (CLSM) (Fig. 4A). It is noteworthy that the fluorescence in the PpIX HA-PLGA micelles treated group was significantly higher than that of the free PpIX and PpIX PLGA-NP treated groups. After competitive treatment with HA, the fluorescence in the PpIX HA-PLGA + HA treated group was significantly reduced in CD44 over-expressing A549 cells. The results
3.4. ROS production PDT stimulates the overproduction of ROS, which could directly induce apoptosis and necrosis in tumor cells and destroy tumor capillaries and microvasculature by disrupting endothelial cells. A host-defense mechanism involving neutrophils and dendritic cells can also be activated by an acute inflammatory response [29]. Therefore, the generation of ROS was analyzed. The ROS production (Fig. 5) correlated with the MTT results. The greater the decrease in cell viability, the higher the observed ROS level. When combined with the internalization results, these findings suggested that the enhanced intracellular 418
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Fig. 4. Internalization and cytotoxicity of HA-PLGA micelles in 2D cell cultures. A. Intracellular uptake of free PpIX, PpIX PLGA-NPs, PpIX HA-PLGA micelles and PpIX HA-PLGA micelles combined with 5 mg/ml free HA. Images were observed by CLSM. The nucleus was stained with DAPI (blue). PpIX exhibited red fluorescence. B. The quantitative analysis of intracellular PpIX (∗∗p < 0.01). C. Dark and light cytotoxicity of free PpIX, PpIX PLGA-NPs and PpIX HA-PLGA micelles against A549 cells. (For interpretation of the references to colour in this figure legend, the reader is referred to the Web version of this article.)
Fig. 5. Intracellular ROS generation on A549 cells treated with PpIX, PpIX PLGA-NPs and PpIX HA-PLGA micelles after photostimulation. A. Images of A549 cells in which the fluorescent signal developed from the ROS probe. B. The percentage of the ROS positive cells (∗∗p < 0.01). 419
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Fig. 6. Penetration and phototoxicity of HA-PLGA micelles in A549 spheroids. A. Z-Stake confocal images of spheroids incubated with (a) free coumarin 6, (b) coumarin 6 loaded PLGA nanoparticles and (c) coumarin 6 loaded HA-PLGA micelles. B. Phototoxicity of free PpIX, PpIX PLGA-NPs and PpIX HA-PLGA micelles determined by trypan blue exclusion assay. (For interpretation of the references to colour in this figure legend, the reader is referred to the Web version of this article.)
deep within the spheroids, while the free coumarin 6 and coumarin 6 PLGA-NPs exhibited little penetration. To analyze the phototoxicity of PpIX HA-PLGA micelles, the A549 spheroids were incubated with PpIX loaded micelles, PpIX PLGA-NPs, and free PpIX, and were subsequently exposed to light irradiation. Higher dose of light irradiation was applied by using the laser system to achieve phototoxicity in tumor spheroids. The laser power was safe and caused no damage to tumor cells [32]. The numbers of live and dead cells were determined using a trypan blue exclusion assay in which the dead cells were stained blue while live cells excluded the trypan blue dye because of the integrity of their membranes. HA-PLGA micelles showed significant enhancement of the phototoxicity of PpIX at the concentrations both of 5 and 10 μmol/l (Fig. 6B). Little effect was observed when cells were treated with 1 μmol/l PpIX as 100% viability was retained. The results demonstrated that the HA-PLGA micelles could enhance the penetration and increase the phototoxicity of PpIX in A549 spheroids.
delivery of PpIX by HA-PLGA micelles could significantly increase ROS production in A549 cells. 3.5. Tumor spheroid penetration and viability Traditional 2D monolayer cell cultures are not always able to simulate in vivo tumors as extracellular barriers are not reproduced. Spheroids are attractive 3D tumor models, which develop oxygen, nutrient, and energy gradients similar to those found in vivo. They have emerged as a powerful and promising predictive tool for chemotherapeutic and photodynamic therapy evaluation [30]. Thus, the penetration and phototoxicity of PpIX HA-PLGA micelles in A549 spheroids was investigated. CD44 is known to traffic between endosomes and the cell surface. Huang [31] reported that hyaluronan coated nanoparticles exhibited cellular uptake mediated by CD44 and could be transported out of the cells with their cargo. The exported nanoparticles could be taken up by neighboring cells resulting in deeper penetration into tumors in 3D tumor models. Since the fluorescence signal of PpIX was too weak to be detected by CLSM, coumarin 6 was used as a fluorescence probe. The A549 spheroids were incubated with coumarin 6 loaded HAPLGA micelles, coumarin 6 loaded PLGA nanoparticles, and coumarin 6. The uptake of the micelles was significantly higher than those of the PLGA nanoparticles and the solution control (Fig. 6A), correlating well with the 2D cell uptake results. Z-Stack confocal microscopy images of the spheroids after incubation were acquired. The micelles were found
4. Conclusion In this study, preparation and application of HA-PLGA micelles in tumor targeting delivery of PpIX in vitro were demonstrated. Following formulation optimization, PpIX HA-PLGA micelles were obtained by dialysis. The particle size was ∼200 nm, and the drug encapsulation efficiency exceeded 43%. Enhanced phototoxicity towards CD44420
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overexpressing A549 cells was observed in both 2D monolayer cell cultures and 3D tumor spheroids, which is attributed to promoted cell uptake and deeper spheroid penetration. PpIX HA-PLGA micelles are promising candidates for targeted PDT treatment of lung cancer.
cancer, Expet Opin. Drug Deliv. 7 (6) (2010) 681–703. [15] J.M. Wickens, H.O. Alsaab, P. Kesharwani, K. Bhise, M.C. Amin, R.K. Tekade, U. Gupta, A.K. Iyer, Recent advances in hyaluronic acid-decorated nanocarriers for targeted cancer therapy, Drug Discov. Today 22 (4) (2017) 665–680. [16] J.M. Liang, F. Zeng, M. Zhang, Z.Z. Pan, Y.Z. Chen, Y.E. Zeng, Y. Xu, Q. Xu, Y.Z. Huang, Green synthesis of hyaluronic acid-based silver nanoparticles and their enhanced delivery to CD44+ cancer cells, RSC Adv. 5 (54) (2015) 43733–43740. [17] F. Ramazani, W. Chen, C.F. van Nostrum, G. Storm, F. Kiessling, T. Lammers, W.E. Hennink, R.J. Kok, Strategies for encapsulation of small hydrophilic and amphiphilic drugs in PLGA microspheres: state-of-the-art and challenges, Int. J. Pharm. 499 (1–2) (2016) 358–367. [18] Y.N. Cui, M. Zhang, F. Zeng, H.Y. Jin, Q. Xu, Y.Z. Huang, Dual-targeting magnetic PLGA nanoparticles for codelivery of paclitaxel and curcumin for brain tumor therapy, ACS Appl. Mater. Interfaces 8 (47) (2016) 32159–32169. [19] J. Huang, H. Zhang, Y. Yu, Y. Chen, D. Wang, G. Zhang, G. Zhou, J. Liu, Z. Sun, D. Sun, Y. Lu, Y. Zhong, Biodegradable self-assembled nanoparticles of poly(D,Llactide-co-glycolide)/hyaluronic acid block copolymers for target delivery of docetaxel to breast cancer, Biomaterials 35 (1) (2014) 550–566. [20] P. Bhatnagar, A.B. Pant, Y. Shukla, A. Panda, K.C. Gupta, Hyaluronic acid grafted PLGA copolymer nanoparticles enhance the targeted delivery of Bromelain in Ehrlich's Ascites Carcinoma, Eur. J. Pharm. Biopharm. 105 (2016) 176–192. [21] H.K. Park, S.J. Lee, J.S. Oh, S.G. Lee, Y.I. Jeong, H.C. Lee, Smart nanoparticles based on hyaluronic acid for redox-responsive and CD44 receptor-mediated targeting of tumor, Nanoscale Res. Lett. 10 (1) (2015) 981. [22] H.J. Jeon, Y.I. Jeong, M.K. Jang, Y.H. Park, J.W. Nah, Effect of solvent on the preparation of surfactant-free poly(DL-lactide-co-glycolide) nanoparticles and norfloxacin release characteristics, Int. J. Pharm. 207 (1–2) (2000) 99–108. [23] K. Zhang, X. Tang, J. Zhang, W. Lu, X. Lin, Y. Zhang, B. Tian, H. Yang, H. He, PEGPLGA copolymers: their structure and structure-influenced drug delivery applications, J. Contr. Release 183 (2014) 77–86. [24] L.Y. Qiu, Y.H. Bae, Polymer architecture and drug delivery, Pharm. Res. 23 (1) (2006) 1–30. [25] Y.Y. Bei, X.F. Zhou, B.G. You, Z.Q. Yuan, W.L. Chen, P. Xia, Y. Liu, Y. Jin, X.J. Hu, Q.L. Zhu, C.G. Zhang, X.N. Zhang, L. Zhang, Application of the central composite design to optimize the preparation of novel micelles of harmine, Int. J. Nanomed. 8 (2013) 1795–1808. [26] C.L. Conway, I. Walker, A. Bell, D.J. Roberts, S.B. Brown, D.I. Vernon, In vivo and in vitro characterization of a protoporphyrin IX-cyclic RGD peptide conjugate for use in photodynamic therapy, Photochem. Photobiol. Sci. 7 (3) (2008) 290–298. [27] B. Xiao, M.K. Han, E. Viennois, L. Wang, M. Zhang, X. Si, D. Merlin, Hyaluronic acid-functionalized polymeric nanoparticles for colon cancer-targeted combination chemotherapy, Nanoscale 7 (42) (2015) 17745–17755. [28] F. Danhier, E. Ansorena, J.M. Silva, R. Coco, A. Le Breton, V. Préat, PLGA-based nanoparticles: an overview of biomedical applications, J. Contr. Release 161 (2) (2012) 505–522. [29] A.P. Castano, P. Mroz, M.R. Hamblin, Photodynamic therapy and anti-tumour immunity, Nat. Rev. Canc. 6 (7) (2006) 535–545. [30] L.A. Muehlmann, M.C. Rodrigues, J.P. Longo, M.P. Garcia, K.R. Py-Daniel, A.B. Veloso, P.E. de Souza, S.W. da Silva, R.B. Azevedo, Aluminium-phthalocyanine chloride nanoemulsions for anticancer photodynamic therapy: development and in vitro activity against monolayers and spheroids of human mammary adenocarcinoma MCF-7 cells, J. Nanobiotechnol. 13 (2015) 36. [31] M.H. El-Dakdouki, E. Puré, X. Huang, Development of drug loaded nanoparticles for tumor targeting. Part 2: enhancement of tumor penetration through receptor mediated transcytosis in 3D tumor models, Nanoscale 5 (9) (2013) 3904–3911. [32] J. Xue, C. Li, H. Liu, J. Wei, N. Chen, J. Huang, Optimal light dose and drug dosage in the photodynamic treatment using PHOTOCYANINE, Photodiagn. Photodyn. Ther. 8 (3) (2011) 267–274.
Acknowledgments The authors gratefully acknowledge the financial support from the National Natural Science Foundation of China (81402872), Zhejiang Provincial Natural Science Foundation of China (LY17H160002, LY17B020002, LY14C090001). Science and Technology Program of Traditional Chinese Medicine of Zhejiang Province of China (2015ZQ022). References [1] U. Chilakamarthi, L. Giribabu, Photodynamic therapy: past, present and future, Chem. Rec. 17 (8) (2017) 775–802. [2] R. Allison, G. Moghissi, G. Downie, K. Dixon, Photodynamic therapy (PDT) for lung cancer, Photodiagn. Photodyn. Ther. 8 (3) (2011) 231–239. [3] T.G. Sutedja, P.E. Postmus, Photodynamic therapy in lung cancer. A review, J. Photochem. Photobiol. B Biol. 36 (2) (1996) 199–204. [4] K. Moghissi, K. Dixon, J.A. Thorpe, M. Stringer, C. Oxtoby C, Photodynamic therapy (PDT) in early central lung cancer: a treatment option for patients ineligible for surgical resection, Thorax 62 (5) (2007) 391–395. [5] C.L. da Silva, J.O. Del Ciampo, F.C. Rossetti, M.V. Bentley, M.B. Pierre, Improved in vitro and in vivo cutaneous delivery of protoporphyrin IX from PLGA-based nanoparticles, Photochem. Photobiol. 89 (5) (2013) 1176–1184. [6] M. Przybylo, D. Glogocka, J.W. Dobrucki, K. Fraczkowska, H. Podbielska, M. Kopaczynska, T. Borowik, M. Langner, The cellular internalization of liposome encapsulated protoporphyrin IX by HeLa cells, Eur. J. Pharmaceut. Sci. 85 (2016) 39–46. [7] H. Ding, B.D. Sumer, C.W. Kessinger, Y. Dong, G. Huang, D.A. Boothman, J. Gao, Nanoscopic micelle delivery improves the photophysical properties and efficacy of photodynamic therapy of protoporphyrin IX, J. Contr. Release 151 (3) (2011) 271–277. [8] B. Li, E.H. Moriyama, F. Li, M.T. Jarvi, C. Allen, B.C. Wilson, Diblock copolymer micelles deliver hydrophobic protoporphyrin IX for photodynamic therapy, Photochem. Photobiol. 83 (6) (2007) 1505–1512. [9] H. Basoglu, M.D. Bilgin, M.M. Demir, Protoporphyrin IX-loaded magnetoliposomes as a potential drug delivery system for photodynamic therapy: fabrication, characterization and in vitro study, Photodiagn. Photodyn. Ther. 13 (2016) 81–90. [10] Y. Jang, S. Kim, S. Lee, C.M. Yoon, I. Lee, J. Jang, Graphene oxide wrapped SiO2/ TiO2 hollow nanoparticles loaded with photosensitizer for photothermal and photodynamic combination therapy, Chemistry 23 (15) (2017) 3719–3727. [11] W.P. Savarimuthu, P. Gananathan, A.P. Rao, E. Manickam, G. Singaravelu, Protoporphyrin IX-gold nanoparticle conjugates for targeted photodynamic therapy–an in-vitro study, J. Nanosci. Nanotechnol. 15 (8) (2015) 5577–5584. [12] I.T. Teng, Y.J. Chang, L.S. Wang, H.Y. Lu, L.C. Wu, C.M. Yang, C.C. Chiu, C.H. Yang, S.L. Hsu, J.A. Ho, Phospholipid-functionalized mesoporous silica nanocarriers for selective photodynamic therapy of cancer, Biomaterials 34 (30) (2013) 7462–7470. [13] G. Tripodo, A. Trapani, M.L. Torre, G. Giammona, G. Trapani, D. Mandracchia, Hyaluronic acid and its derivatives in drug delivery and imaging: recent advances and challenges, Eur. J. Pharm. Biopharm. 97 (Pt B) (2015) 400–416. [14] D.A. Ossipov, Nanostructured hyaluronic acid-based materials for active delivery to
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