Colloids and Surfaces B: Biointerfaces 123 (2014) 486–492

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Controlled co-delivery nanocarriers based on mixed micelles formed from cyclodextrin-conjugated and cross-linked copolymers Shuai Li 1 , Qin He 1 , Tianchan Chen, Wei Wu, Kening Lang, Zhong-Ming Li ∗ , Jianshu Li ∗ College of Polymer Science and Engineering, Sichuan University, Chengdu 610065, China

a r t i c l e

i n f o

Article history: Received 9 June 2014 Received in revised form 2 September 2014 Accepted 23 September 2014 Available online 30 September 2014 Keywords: Co-delivery nanocarrier Mixed micelles Host–guest interaction Core-stabilized Cyclodextrin

a b s t r a c t The combination of multiple drugs within a single nanocarrier can provide significant advantages for disease therapy and it is desirable to introduce a second drug based on host–guest interaction in these co-delivery systems. In this study, a core-stabilized mixed micellar system consisting of ␤cyclodextrin-conjugated poly(lactic acid)-b-poly(ethylene glycol) (␤-CD-PLA-mPEG) and DL-Thioctic acid (TA) terminated PLA-mPEG (TA-PLA-mPEG) was developed for the co-delivery of DOX and fluorescein isothiocyanate labeled adamantane (FA). DOX can be loaded within the hydrophobic segment of PLA and FA may form stable complexation with ␤-CD in the core. The mixed micelles (MM) are based on well-accepted medical materials and can be easily cross-linked by adding 1,4-dithio-d,l-threitol (DTT), which can enhance the stability of the system. Drug-loaded MM system was characterized in terms of particle size, morphology, drug loading and in vitro release profile. Cytotoxicity test showed that blank MM alone showed negligible cytotoxicity whereas the drug-loaded MM remained relatively high cytotoxicity for HeLa cancer cells. Confocal laser scanning microscopy (CLSM) demonstrated that the MM could efficiently deliver and release DOX and FA in the same tumor cells to effectively improve drugs’ bioavailability. These results suggested that the core-stabilized MM are highly promising for intracellular co-delivery of multiple drugs. © 2014 Elsevier B.V. All rights reserved.

1. Introduction Drug delivery system has attracted considerable interest because it can significantly improve the solubility of poorly watersoluble drug, carry it into target sites, reach for sustained/controlled release profiles, prolong circulation time and improve biodistribution/bioavailability [1,2]. However, most of existing drug delivery systems are applied to load and delivery a single drug, which does not satisfy the demands of combination therapy, and far from perfect due to their limited bioavailability, toxicity and other side effects. Therefore, it is necessary to develop co-delivery system, which can carry two or more drugs, for superior profiles such as controlled release sequence, timing, doses and duration of each drug, to achieve an effective combination therapy [3,4]. For example, Sengupta et al. reported a new drug carrier named as ‘nanocell’ for solid tumor treatment, in which doxorubicin (DOX) was conjugated to the poly(lactic-co-glycolic) acid (PLGA)-based nanoparticle

∗ Corresponding authors. Tel.: +86 28 85466755; fax: +86 28 85405402. E-mail addresses: [email protected] (Z.-M. Li), jianshu [email protected], [email protected] (J. Li). 1 These two authors are equally contributed to this work. http://dx.doi.org/10.1016/j.colsurfb.2014.09.049 0927-7765/© 2014 Elsevier B.V. All rights reserved.

while an anti-angiogenesis agent (combretastatin-A4) was trapped within the outer lipid envelope. This ‘nanocell’ enabled a temporal release model of the two drugs: the outer envelope first released the combretastatin-A4 and caused a vascular shutdown; the inner nanoparticle could be trapped inside the solid tumor and then released DOX to kill the tumor cells [5]. Also, we note that Celator Pharmaceuticals has carried out Phase II/III clinical trials with CombiPlex® , which indicates the potential of drug combination therapy. As a well-studied drug delivery system, polymeric micelle consists of hydrophobic core and hydrophilic shell. The micelles are formed by macromolecular amphiphiles when the concentration is above the critical micelle concentration (CMC). They have been well investigated for drug co-delivery systems with improved properties [6–13]. Shin and coworkers reported the multi-agent loaded micelles based on poly(ethylene glycol)-blockpoly(D,L-lactic acid) (PEG-b-PLA), which were utilized to delivery several drugs, e.g. paclitaxel (PTX), etoposide (ETO), docetaxel (DCTX) and 17-allylamino-17-demethoxygeldanamycin (17-AAG). Combination of PTX/17-AAG, ETO/17-AAG, DCTX/17-AAG and PTX/ETO/17-AAG were all solubilized at the level of mg/mL and stable for 24 h in the micelles. The presence of 17-AAG could help to keep the stability of 2- or 3-drug combination PEG-b-PLA

S. Li et al. / Colloids and Surfaces B: Biointerfaces 123 (2014) 486–492

micelles [7]. Other strategies, such as forming micelle-drug conjugates via complexation or electrostatic interactions [8–12], have also been used to load drugs or bioactive agents in co-delivery systems. For instance, Lee et al. reported that cationic micellar nanoparticle could be employed as carriers to co-delivery PTX and Herceptin through different interaction mechanisms. This system achieved targeted delivery of PTX to human epidermal growth factor receptor-2 (HER2/neu)-overexpressing human breast cancer cells, and enhanced cytotoxicity through synergistic activities as compared with single-drug treatment [12]. Although previous researches have made great progress in this field, there are very few reports about introducing the second drug via the inclusion interaction between host and guest molecules in co-delivery systems. Cyclodextrins (CDs) are wellknown host molecules consisting of six to eight glucose units. The internal hydrophobic cavities in the CDs facilitate the inclusion of various guest molecules [14–16], thus CDs could be desirable functional units for co-delivery systems. In our previous work, we synthesized a well-defined amphiphilic copolymer, i.e., ␤cyclodextrin-conjugated poly(lactic acid)-b-poly(ethylene glycol) (␤-CD-PLA-mPEG), which could self-assemble in aqueous solution to form micelles with high drug loading efficiency and controlled release profile, especially in the control of the initial burst release of a single drug (indomethacin, IND) [17]. Since the cavity of ␤-CD may encapsulate a second drug in addition to one drug incorporated in the hydrophobic core of micelles, the micelles assembled by ␤-CD-PLA-mPEG copolymer could be a promising co-delivery system to achieve differential release profiles. Another requirement for micellar delivery system is that it should have enough stability to avoid the drug leakage during circulation. Since micelles self-assembled from single copolymer often lack multiple functionalities due to the limited number of building blocks, mixed micelles (MM) can provide an alternative approach to enhance the stability of micellar system. MM formed from two or more dissimilar block copolymers are efficient to introduce considerable functional units without the need of complicated synthetic steps [18–20], and can also achieve significant improvements, e.g., lower CMC [21], higher drug loading efficiency [22], controlled particle size [23], enhanced stability and more efficient tumor inhibition [24,25]. Recently, several groups have reported the enhanced stabilization of polymeric micelles via forming cross-linked cores [26–31], thus the introduction of cross-linking units into one of the copolymers of mixed micelles could be a facile strategy to stabilize the co-delivery system. Herein, we design a core-stabilized mixed micellar system, which is formed by ␤-CD-PLA-mPEG and DL-Thioctic acid (TA) terminated PLA-mPEG (TA-PLA-mPEG), for the co-delivery of DOX and fluorescein isothiocyanate labeled adamantane (FA), as shown in Scheme 1. DOX is loaded within the hydrophobic segment of PLA and FA may form stable complexation with ␤-CD in the core. TA is produced naturally in the human body and could be widely used as an antioxidant drug for treating diseases such as HIV and diabetes [27,32]. Thus the MM are based on well-accepted

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medical materials. The MM can be easily cross-linked by adding 1,4-dithio-D,L-threitol (DTT) relative to the lipoyl groups at the end of TA-PLA-mPEG copolymer, which can improve the stability of the system. Drug-loaded MM were characterized in terms of particle size, morphology, drug loading and in vitro release profile. Cytotoxicity test showed that blank MM alone showed negligible cytotoxicity whereas the drug-loaded MM remained relatively high cytotoxicity for HeLa cancer cells. Furthermore, confocal laser scanning microscopy (CLSM) demonstrated that the MM system could efficiently deliver and release DOX and FA in the same tumor cells to effectively improve drugs’ bioavailability. These results suggested that the core-stabilized MM system are highly promising for intracellular co-delivery of multiple drugs. 2. Experimental details 2.1. Materials PLA-mPEG and ␤-CD-PLA-mPEG were synthesized in our previous work [17]. DL-Thioctic acid (TA) was purchased from Acros (Acros Organics, Geel, Belgium). 1,4-dithio-D,L-threitol (DTT) was purchassed from Merck (Darmstadt, Germany). 4-dimethylamino pyridine (DMAP, 99%) was provided by Kelong Chemical Reagent Plant (Chengdu, China). 1-(3-Dimethylaminopropyl)3-ethylcarbodiimide hydrochloride (EDC) and fluorescein isothiocyanate were purchased from Aladdin (Shanghai, China). 1-adamantanamine was purchased from Adamas (Shanghai, china). FA was prepared following reported procedure in literature [33]. Doxorubicin hydrochloride (DOX HCl) was obtained from Melone biotechnology Co., Ltd. (Dalian, China) and desalted with triethylamine and dimethyl sulfoxide. N,N-dimethylformamide (DMF) was purchased from Bodi Chemical Holding Co., Ltd. (Tianjin, China) and used after drying. 2.2. Synthesis of TA-PLA-mPEG PLA-mPEG (Mn = 4000, 2 g, 0.5 mmol) was dissolved in dry DMF (15 mL). Under a nitrogen atmosphere, EDC (0.19 g, 1 mmol), DMAP (0.1125 g, 0.75 mmol), and TA (0.15 g, 0.75 mmol) were added in order. After stirring for 15 min, the reaction bottle was sealed and placed into an oil bath thermostatted at 25 ◦ C for 2 days. The obtained solution was diluted with 15 mL water and dialyzed against water by a dialysis tube (MWCO 3500 Da) for 3 days. The obtained copolymer solution was lyophilized. Products were characterized by 1 H NMR (400 MHz) (AVII-400; Bruker, Karlsruche, Germany). 2.3. MM formation and determination of CMC ␤-CD-PLA-mPEG and TA-PLA-mPEG were dissolved in DMF at three different mass ratios (8:2, 5:5, 2:8) and stirred overnight. Then a certain amount of water was dropwise added to the solution under stirring. The resultant solution was dialyzed against distilled

Scheme 1. Illustration of mixed micelles based on ␤-CD-PLA-mPEG and TA-PLA-mPEG for the co-delivery of DOX and FA.

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water (MWCO 1000 Da) over 2 days and then lyophilized. Their CMCs were determined by fluorescence probe technique using pyrene as the fluorescence probe. The concentration of pyrene in micelle solution was 6 × 10−7 mol/L and the concentration of mixed copolymers was varied from 0.0001 to 1 mg/mL. The fluorescent spectra were recorded using Shimadzu RF-5301PC luminescence spectrometer at room temperature. The measured emission wavelength was set at 395 nm with the excitation wavelength of 330 nm and band widths of 3 nm. The ratios of the excitation spectra’s fluorescent intensities at 337.0 nm and 335.0 nm (I337.0 /I335.0 ) were calculated and plotted against the logarithm of polymer concentration.

(150 mm × 4.6 mm, 5.0 ␮m particles). A degassed mixture of acetonitrile and water (32:68, v/v, adjusted to pH 2.4 by phosphoric acid) was used as the mobile phase and pumped at a flow rate of 1.0 mL/min. The column temperature was maintained at 37 ◦ C. The detection was at 256 nm for DOX and 234 nm for FA. For determination of drug loading amount, an aliquot of lyophilized MM-DOX-FA was dissolved in acetonitrile and analyzed with RP-HPLC. The loading efficiency (LE) was calculated according to the equation LE = [WDrug ]/[WPolymer + WDrug ] × 100%, and the entrapment efficiency (EE) according to the equation EE = [WDrug ]/[Wadded ] × 100%, where WDrug , WPolymer and Wadded refer to the weights of drug encapsulated by micelles, copolymers in formulations and added drug, respectively.

2.4. Preparation of drug-loaded MM (MM-DOX-FA) 2.8. In vitro drug release The preparation procedure of drug-loaded mixed micelles was similar to that of blank mixed micelles. Typically, 30 mg copolymers (mixtures with different mass ratios between ␤-CD-PLA-mPEG and TA-PLA-mPEG) and 12 mg drugs (FA:DOX = 1:1, wt/wt) were dissolved in 5 mL DMF. Then 15 mL distilled water was added dropwise into the solution under vigorous stirring over 45 min. Subsequently, under a N2 atmosphere, DTT solution in distilled water (10 mol% relative to the amount of lipoyl units) was added to the resultant solution. After being stirred for another 30 min, the polymer solution was dialyzed against distilled water (MWCO 1000 Da) over 3 days, during which micelles underwent a cross-linking reaction by oxygen in air. The dialysis medium was refreshed for six times. Finally, the obtained solution was filtered through a 450 nm filter to remove large aggregates (about 5–10% loss of polymer micelle formulations) and then lyophilized for further experiments. The whole procedure was performed in the dark.

The release profiles of drugs from mixed micelles were studied using a dialysis bag (MWCO 1000 Da) in two different media, i.e., PBS (10 mM, pH 7.4) and the same PBS with 10 mM DTT. In this part, DTT was used as the mimic of glutathione (GSH) which acted as a disulfide-reducing agent in the cytosol of cells. 10 mg freeze-dried drug-loaded micelles were suspended in 10 mL of release medium and transferred into the dialysis bag. The release experiment was initiated by placing the end-sealed dialysis bag into 90 mL of release medium at 37 ◦ C with continuous shaking at 100 rpm. At selected time intervals, 5 mL of external release medium was taken out and replenished with an equal volume of fresh medium. The amount of drugs released was determined by RP-HPLC as described above. The accumulative percentage of the released drugs were calculated as a function of time. 2.9. Cell viability assays and intracellular drug delivery

2.5. Mean particle size distribution and zeta-potential measurements The mean particle size distribution and zeta-potential of MM-DOX-FA were determined by dynamic light scattering (DLS, Zetasizer Nano ZS90, Malvern) with the sample in distilled water (pH 7.4) at a concentration of 1 mg/mL. The size of cross-linked MM was also investigated by DLS measurements in different solvents and different MM concentrations. All measurements were carried out at 37 ◦ C with a 90◦ scattering angle. The micelle aqueous solution was subjected to ultrasonic treatment and passed through 0.45 ␮m pore-sized filters before being transferred into sample cuvettes. Each sample was measured three times. 2.6. Surface morphology The morphologies of MM-DOX-FA were observed by transmission electron microscope (TEM) (Hitachi H-600 transmission electron microscope, JOEL Ltd., Japan) with an accelerating voltage of 75 kV. To prepare the samples for TEM measurements, a drop of sample solution (1 mg/mL) was deposited on a carbon-coated copper grid (200 meshes), then excess of micelle solution was removed with filter paper. After being negatively stained with 1 wt/v% phosphotungstic acid for 60 s and a thorough air-drying, the samples were observed in the electron microscope. 2.7. Drug-loading and entrapment efficiency The content of drugs loaded in mixed micelles was quantified by reversed-phase high-performance liquid chromatography (RP-HPLC) and obtained on the basis of the standard curves. The HPLC system was equipped with a LC-15C HPLC pump, SPD-15C UV/Vis detector (Shimadzu, Japan) and reversed phase C18 column

The relative cytotoxicities of blank and drug-loaded mixed polymeric micelles against HeLa cells were evaluated in vitro by standard 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide (MTT) assay. Live tumor cells without any treatment were used as control. In brief, the cells were cultured with 10% heatinactivated fetal bovine serum (FBS), 100 units/mL of penicillin and 100 ␮g/mL of streptomycin in Dulbecco’s modified eagle medium (DMEM) at 37 ◦ C with an atmosphere of 5% CO2 and 95% relative humidity. Growing cells (1 × 104 cells/well) were seeded on 96well micro plate (Nunc Co., Wiesbaden, Germany) and incubated in 100 ␮L of DMEM/well for 24 h. Then fresh culture media containing serial dilutions of blank MM or MM-DOX-FA were used to treat the cells for 24 h. After that, cells were incubated with 10 ␮L of MTT stock solution in PBS (5 mg/mL) each well for 4 h. Then the medium was removed and the produced formazan crystals were solubilized in DMSO (100 ␮L/well). The absorbance was measured at 570 nm by a micro plate reader (Spectra Plus, Tecan, Zurich, Switzerland). The relative cell viability (%) was calculated using the following equation: Cell viability =

Atest × 100% Acontrol

where Atest and Acontrol are the mean absorbance value of tested group and control group (without MM or MM-DOX-FA), respectively. The co-delivery of DOX and FA was observed by a TCRS SP5 confocal laser scanning microscope (CLSM, Leica, Germany). HeLa cells were seeded in 6-well cell culture plates with a clean cover slip (24 mm × 24 mm) placed into each well and incubated at 37 ◦ C with an atmosphere of 5% CO2 . After cells in culture media were grown to ∼70% confluency (∼2 × 105 cells/well), the culture media were replaced by fresh media containing free DOX, FA and MM-DOX-FA (␤-CD-PLA-mPEG/TA-PLA-mPEG = 5:5, filtered with

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0.45 ␮m syringe driven filters before use). After 4 h, the culture media were removed and the cells were washed three times with PBS (pH 7.4). Then, the cells were fixed with 4 w/v% formaldehyde at 4 ◦ C for 20 min. After that, cell nuclei were stained with Hoechst 33,342 (5 ␮g/mL) for 10 min. Finally, the cover slips were placed on a glass microscope slide and the prepared samples were subjected for fluorescence imaging by CLSM. DOX was excited at 352 nm with emission at 455 nm, and 494/518, 345/478 nm for FA and Hoechst 33,342, respectively.

Table 1 The loading efficiency (LE) and entrapment efficiency (EE) of drug-loaded micelles.

3. Results and discussion

co-delivery system may significantly affect the release profiles of these drugs. In this study, MM-DOX-FA were prepared by adding water dropwise to the mixture of drugs and copolymers in DMF under stirring, followed by dialysis against distilled water. The yield of MM-DOX-FA after dialysis was about 80% (w/w). In this formulation, DOX was a common chemotherapeutic drug with excellent anti-tumor efficiency against various solid tumors. It was mainly loaded within the micellar core of PLA by hydrophobic entrapment. FA was prepared by bonding 1-adamantanamine together with fluorescein isothiocyanate via a thiourea linkage (Fig. S4). Since FA could form quite stable inclusion complex with ␤-CD [33], it was selected as model drugs (and also a model molecular probe) to be loaded mainly by ␤-CD units in the central core of the MM system. Then, the mixed micelles could be easily cross-linked by adding DTT relative to the lipoyl groups at the end of TA-PLA-mPEG copolymer, which could improve the stability of the system. Theoretically, based on the stoichiometry of FA:␤-CD = 1:1 inclusion complex, the maximum loading amounts of FA in these three samples (8:2, 5:5, 2:8) were calculated as 7.16, 4.64 and 1.97%, respectively (according to the calculation in supporting information). However, as shown in Table 1, the actual loading amounts of FA (7.2, 5.8 and 4.3%, respectively) were higher than the theoretical values. It indicated that some excessive FA could be loaded within the hydrophobic PLA segment of the MM after the hydrophobic cores of ␤-CD had been occupied. Also, we could see the differences between them became more obvious with the decrease of ␤-CD-PLA-mPEG ratio in the MM system. On the other hand, the loading amounts of DOX became higher as the mass ratio of ␤-CDPLA-mPEG decrease. These phenomena could be attributed to the different encapsulating manner of FA and DOX within the micellar core and also the changes of the hydrophobicity of the micellar core. As explained above, the decrement of the ratio of ␤-CD-PLAmPEG enhanced the hydrophobicity of the inner core, resulting in an increase of the hydrophobic entrapment to DOX, even to FA (excessive portion not in ␤-CD). Eventually, there were more DOX and FA being loaded within the inner cores of micelles through hydrophobic entrapment. The size of micelles is an important parameter which may significantly influence the circulation time and organ distribution of micelles. It has been reported that micelles at the nanoscale could not only avoid renal excretion and bypass filtration by interendothelial cell slits in the spleen, but also enhance drug accumulation within tumors due to the relatively high permeability of the vascular endothelium of tumors [34,35]. DLS measurements showed that these MM-DOX-FA were in the size range from 130 to 190 nm with relatively low polydispersities and negative surface charge (Fig. 1 and Table 2). As for in vivo applications, the negative surface charge could reduce plasma protein bioadhesion and facilitate the stealth property of micelles itself [36]. Also, from DLS results, we found that the sizes of micelles became smaller with the decrease of the ␤-CD-PLA-mPEG ratio in the MM system. It might be because the decrement of relatively hydrophilic ␤-CD in the hydrophobic micellar core made the mixed copolymers form a more compact structure. In addition, DLS also revealed the existence of crosslinked structure and it endowed micelles with certain stability in diluted condition (Fig. S5 and Fig. S6). The morphologies of

3.1. Synthesis and characterization of TA-PLA-mPEG TA-PLA-mPEG was synthesized by the reaction between the carboxyl group of TA and the hydroxyl group of PLA-mPEG in the presence of EDC and DMAP. The successful synthesis was confirmed by 1 H NMR spectroscopy using CDCl3 as the solvent. 1 H NMR spectrum (Fig. S1) displayed signals at ı 3.1–3.2 assigned to the methylene protons next to the disulfide bond on the lipoyl moiety. Meanwhile, the signal at ı 3.38 ascribed to the O CH3 group was one of the characteristic peaks of PEG-PLA. By comparing the intensities of signals at ı 3.1–3.2 and 3.38, it could be estimated that each block copolymer chain was conjugated with approximately one lipoyl functional group. 3.2. MM formation and CMC characterization As a type of drug carrier, micelles can effectively enhance the aqueous solubility and bioavailability of various hydrophobic drugs. Due to their excellent characters that different types of copolymers can be concentrated in a single polymeric micellar system, MM systems have been proved to be ideal candidates for functional drug carriers. However, the stability of MM systems in diluted condition is a main concern in this field because they are formed from two or more kinds of different polymers. CMC is an important parameter indicative of the ability of micelles to retain drug molecules, especially when being diluted in vivo. In this study, these MM systems were prepared by a dialysis method which was often used to form micelles from amphiphilic copolymers. 1 H NMR (Fig. S2) indicated the existence of core–shell structure with the PEG block as the hydrophilic outer shell and the PLA block as the hydrophobic inner core, which was a typical micellar structure. Fluorescence probe technique with pyrene as the fluorescence probe was used to study the CMCs of mixed copolymers. The CMC values of ␤-CD-PLA-mPEG/TA-PLA-mPEG at different mass ratios were obtained by plotting the ratio of I337.0 /I335.0 of the emission spectra profile vs. the logarithm of polymer mass concentration as shown in Fig. S3. The CMC values for three samples (␤-CD-PLA-mPEG/TAPLA-mPEG = 8:2, 5:5, 2:8) were as low as 0.02547, 0.01832 and 0.00824 mg/mL, respectively. It is noteworthy that the decrement of the ratio of ␤-CD-PLA-mPEG lowers the CMC of mixed micelles. It may be because ␤-CD exhibits certain hydrophilicity due to a large number of hydroxyl groups on its exterior. The decrement of the ratio of ␤-CD-PLA-mPEG can improve the hydrophobicity of ␤CD-PLA-mPEG/TA-PLA-mPEG system, which causes the decrease of CMC values. 3.3. Preparation and characterization of drug-loaded mixed micelles (MM-DOX-FA) Co-delivery systems can synergistically combine different drugs with different physiochemical properties, thus they have been proposed to solve problems associated with current treatments, such as undesirable side effects, low bioavailability and drug resistance. It is well-known that the loading strategy of different drugs in the

CD/TA (w/w)a

8:2 5:5 2:8 a

DOX

FA

LE (%)

EE (%)

LE (%)

EE (%)

8.1 9.2 10.4

47.5 53.9 60.9

7.2 5.8 4.3

42.3 33.8 25.2

Mass ratio of ␤-CD-PLA-mPEG to TA-PLA-mPEG.

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Fig. 1. DLS characterizations of drug loaded ␤-CD-PLA-mPEG/TA-PLA-mPEG mixed micelles (8:2, 5:5, 2:8, w/w) with a concentration of 1 mg/mL.

Fig. 3. Release profiles of drug-loaded mixed micelles (5:5) in the absence or presence of 10 mM DTT at pH 7.4 and 37 ◦ C.

Table 2 Characteristics of drug-loaded micelles. CD/TA (w/w)a

Particle size (nm)b

PDIb

Zeta potential (mv)b

8:2 5:5 2:8

186.8 175.9 132.8

0.142 0.153 0.178

−12.6 −16.4 −13.8

Mass ratio of ␤-CD-PLA-mPEG to TA-PLA-mPEG. Determined by Zetasizer Nano ZS90 at a concentration of 1 mg/mL in PBS (10 mM, pH 7.4) at 37 ◦ C. a

b

drug-loaded mixed micelles were examined by TEM as shown in Fig. 2. As can be seen, these micelles had spherical morphologies with a narrow size dispersity. Compared with the values determined by DLS, the sizes shown in the TEM were smaller which could be attributed to the shrinkage of the micelles during the drying process of sample preparation. 3.4. In vitro drug release The MM-DOX-FA with a mixed ratio of 5:5 was chosen as an example to investigate the drug release profiles in the presence or absence of DTT. As shown in Fig. 3, the release profile of FA in the presence of DTT is almost the same as that without DTT. It could be explained that the inclusion interaction between FA and ␤-CD was not significantly influenced by the presence of DTT. In contrast, the release of DOX from micelles was affected by the presence of DTT in the release medium. The presence of DTT accelerated the release of DOX. This was most likely due to the open of the

Fig. 4. Cytotoxicity determined by the MTT assay of blank and drug-loaded mixed micelles (5:5) against HeLa cells after incubation for 24 h. Data are presented as the average ± standard deviation (n = 6).

cross-linked structure located at the core of micelles, which not only rendered micelles better stability but also formed a physical barrier to suppress DOX release from micelles in the transport process. In addition, the release of FA was slower than that of DOX, regardless of whether DTT was added or not. This difference in release behavior between FA and DOX could be mainly attributed to their different loading mechanism. DOX is loaded within the

Fig. 2. TEM images of drug-loaded ␤-CD-PLA-mPEG/TA-PLA-mPEG mixed micelles.

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Fig. 5. Confocal images of HeLa cells incubated with free DOX, free FA and drug-loaded mixed micelles (5:5) for 4 h. Blue: Hoechst 33,342; green: FA; red: DOX (scale bar: 20 ␮m). (For interpretation of the references to color in figure legend, the reader is referred to the web version of the article.)

hydrophobic segment of PLA by common hydrophobic entrapment, while FA was mainly hosted by the CD cavities (also part in the hydrophobic core) which resulted in a retarded release profile [37].

intracellular drug accumulation [41]. These results strongly suggested that MM-DOX-FA could efficiently deliver and release DOX and FA in the same tumor cells to effectively improve the bioavailability of drugs.

3.5. Cell viability assays and intracellular drug delivery 4. Conclusions MTT assay was used to determine the cytotoxicity of blank and drug-loaded MM (5:5) with different concentrations in HeLa cancer cells. As shown in Fig. 4, the MTT assay showed that blank mixed micelles alone showed negligible cytotoxicity with ≥90% cell viability even at a micellar concentration of up to 100 ␮g/mL, which was in accordance with the well-recognized biocompatibility of these materials used. Nevertheless, MM-DOX-FA resulted in a concentration-dependent increase in cytotoxicity with the increase of the concentration from 0.78125 to 100 ␮g/mL. These results demonstrate that the mixed micellar system has a potential for two drugs loading and delivery into cancer cells to kill cancer cells. Then, CLSM was used to evaluate the intracellular accumulation of both drugs in HeLa cells after being treated with mixed micellar co-delivery systems. Hoechst 33,342 was used to identify cell nuclei [38]. Fig. 5 shows the fluorescence microscopy photographs of HeLa cells treated with free DOX, free FA, and MM-DOX-FA (5:5). After a 4 h incubation, the cells treated with free DOX remained blue with little red fluorescence appearing in the nucleus. Previous studies have demonstrated that free DOX could diffuse across the cellar membrane into tumor cells in a very short time but were rapidly extruded by P-glycoprotein (P-gp) expressed on the cell membrane [39,40]. P-gp, as a ATP-binding cassette transporter, contributes mainly to the drug resistance of tumor cells. The cells treated with free FA did not show obvious green fluorescence since free FA could hardly be internalized by tumor cells. In comparison, the treatment with MM-DOX-FA led to higher DOX and FA accumulation within HeLa cells, and significant overlap of fluorescence signals from them at the same incubation time of 4 h. As a type of nanocarrier, MM-DOX-FA could enter tumor cells via endocytosis, which might bypass multidrug resistance mechanisms involving cell-surface protein pumps (e.g., P-gp) and increase the

In summary, we successfully prepared a core-stabilized mixed micellar system by ␤-CD-PLA-mPEG and TA-PLA-mPEG as a novel nanocarrier for the co-delivery of DOX and FA. DOX is loaded within the hydrophobic segment of PLA and FA can form stable complexation with ␤-CD in the core. The MM system can be easily cross-linked by adding DTT and enhance the stability of the system. The drug-loaded MM systems exhibited controlled release profiles due to their well-designed structure. The blank MM alone showed negligible cytotoxicity while the drug-loaded MM remained relatively high cytotoxicity for HeLa cancer cells. CLSM investigation demonstrated that the MM could efficiently deliver and release DOX and FA in the same tumor cells and effectively improve bioavailability of drugs. These core-stabilized MM show great potential as nanocarriers for intracellular co-delivery of multiple drugs. Acknowledgements Financial support from the National Natural Science Foundation of China (51322303, 51121001), Program for Changjiang Scholars and Innovative Research Team in University (IRT1163) and Science & Technology Foundation of Sichuan Province (2012JQ0009) are gratefully acknowledged. Appendix A. Supplementary data Supplementary data associated with this article can be found, in the online version, at http://dx.doi.org/10.1016/j.colsurfb. 2014.09.049.

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Controlled co-delivery nanocarriers based on mixed micelles formed from cyclodextrin-conjugated and cross-linked copolymers.

The combination of multiple drugs within a single nanocarrier can provide significant advantages for disease therapy and it is desirable to introduce ...
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