Transferrin-conjugated magnetic silica PLGA nanoparticles loaded with doxorubicin and paclitaxel for brain glioma treatment

Biomaterials 34 (2013) 8511e8520

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Transferrin-conjugated magnetic silica PLGA nanoparticles loaded with doxorubicin and paclitaxel for brain glioma treatment Yanna Cui a, b, Qingxing Xu b, Pierce Kah-Hoe Chow c, d, e, Deping Wang a, *, Chi-Hwa Wang b, * a

School of Materials Science and Engineering, Tongji University, 4800 Caoan Road, Shanghai 201804, PR China Department of Chemical and Biomolecular Engineering, National University of Singapore, 4 Engineering Drive 4, Singapore 117576, Singapore Department of General Surgery, Singapore General Hospital, Outram Road, Block 6, Level 7, Singapore 169608, Singapore d Office of Clinical Sciences, Duke-NUS Graduate Medical School, 8 College Road, Singapore 169857, Singapore e Department of Surgical Oncology, National Cancer Centre, 11 Hospital Drive, Singapore 169610, Singapore b c

a r t i c l e i n f o

a b s t r a c t

Article history: Received 26 June 2013 Accepted 21 July 2013 Available online 6 August 2013

The effective treatment of malignant brain glioma is hindered by the poor transport across the blood ebrain barrier (BBB) and the low penetration across the blood-tumor barrier (BTB). In this study, transferrin-conjugated magnetic silica PLGA nanoparticles (MNP-MSN-PLGA-Tf NPs) were formulated to overcome these barriers. These NPs were loaded with doxorubicin (DOX) and paclitaxel (PTX), and their anti-proliferative effect was evaluated in vitro and in vivo. The in vitro cytotoxicity of drug-loaded NPs was evaluated in U-87 cells. The delivery and the subsequent cellular uptake of drug-loaded NPs could be enhanced by the presence of magnetic field and the usage of Tf as targeting ligand, respectively. In particular, cells treated with DOX-PTX-NPs-Tf with magnetic field showed the highest cytotoxicity as compared to those treated with DOX-PTX-NPs-Tf, DOX-PTX-NPs, DOX-PTX-NPs-Tf with free Tf. The in vivo therapeutic efficacy of drug-loaded NPs was evaluated in intracranial U-87 MG-luc2 xenograft of BALB/c nude mice. In particular, the DOX-PTX-NPs-Tf treatment exhibited the strongest anti-glioma activity as compared to the PTX-NPs-Tf, DOX-NPs-Tf or DOX-PTX-NPs treatment. Mice did not show acute toxicity after administrating with blank MNP-MSN-PLGA-Tf NPs. Overall, MNP-MSN-PLGA-Tf NPs are promising carriers for the delivery of dual drugs for effective treatment of brain glioma. Ó 2013 Elsevier Ltd. All rights reserved.

Keywords: Magnetic silica PLGA nanoparticles Transferrin Doxorubicin Paclitaxel Brain glioma

1. Introduction Malignant brain glioma is the most lethal and aggressive type of cancer in oncology, with a survival time of 12e15 months and a 5year survival rate of less than 5% [1,2]. With the development of chemotherapeutic drugs, chemotherapy has been considered to be effective in inhibiting the growth of brain tumor. However, due to the existence of bloodebrain barrier (BBB), which separates blood from the cerebral tissue and prevents drug penetration into the central nervous system [3e5], and blood-tumor barrier (BTB), which hampers drug accumulation and uptake in tumor, the chemotherapeutic efficacy is greatly reduced. Due to the poor transport efficiency, high drug dosage is often necessary to reach therapeutic concentration within the tumor, and this poses severe toxicity to the normal tissues. Thus, there is a pressing need to * Corresponding authors. E-mail addresses: [email protected] (C.-H. Wang).

(D. Wang), [email protected]

0142-9612/$ e see front matter Ó 2013 Elsevier Ltd. All rights reserved. http://dx.doi.org/10.1016/j.biomaterials.2013.07.075

overcome these challenges and improve brain cancer treatment. Recently, the modification of drug delivery carriers with targeting ligands can improve transport efficiency across BBB, target tumor cells by recognizing their specific over-expressed receptor and trigger receptor-mediated endocytosis [6e9]. Although many targeting ligands have been developed to improve drug delivery [10,11], the targeting and infiltrating of glioma cells still remain a challenge. Transferrin (Tf) has been widely applied to enhance cellular uptake of drug-loaded nanoparticles (NPs) since Tf-receptors (Tf-Rs) are overexpressed in brain capillary endothelium and glioma cells [12,13]. Designing Tf-conjugated drug delivery systems can improve drug transport across the BBB and enhance drug accumulation in the tumor cells. Such drug delivery systems include not only Tf-drug conjugates, but also Tfmodified nanocarriers involving liposomes [14], polymersomes [15], and micelles [4,6]. Compared with these nanocarriers, inorganiceorganic hybrid NPs could offer great advantages because the favorable properties of organic and inorganic components could be combined in a single drug carrier.

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Mesoporous silica NPs (MSNs) are ideal drug delivery systems due to their high pore volume, large specific surface area and biocompatibility [16,17]. Superparamagnetic iron oxide (or magnetic nanoparticles (MNPs)) can be loaded into silica to form uniform coreeshell structured composite, consisting of a magnetic core and a silica shell. These NPs could improve transport to a specific location and enhance cellular uptake in the presence of a magnetic field [18e20]. Poly(D,L-lactic-co-glycolic acid) (PLGA) has been widely used in drug delivery systems due to its biocompatibility and biodegradability [21]. However, studies on magnetic silica PLGAbased carriers loaded with hydrophilic and hydrophobic drugs are limited. In our previous studies, PLGA was loaded with MSNs, and Tf was conjugated on the surface of MSN-PLGA NPs to develop a targeting drug delivery system for the treatment of cancer [22,23]. In this study, Tf-conjugated MNP-MSN-PLGA NPs (MNP-MSN-PLGA-Tf NPs) were prepared to deliver doxorubicin (DOX) and paclitaxel (PTX) simultaneously for brain glioma treatment. Both DOX and PTX are common chemotherapeutic drugs as they have excellent anti-tumor efficacy against various solid tumors [24,25]. DOX can bind with DNA by intercalation and induce apoptosis in various cancer cells. PTX is a mitotic inhibitor [26], which interferes with the normal breakdown of microtubules during cell division and has been licensed to treat ovary, breast, lung, head and neck cancers, and Kaposi’s sarcoma [27]. Simultaneous delivery of these two drugs is a promising strategy to improve the efficacy of cancer treatment. However, due to multidrug resistance and poor targeting efficiency, the administration of DOX and/or PTX as free drug(s) has produced poor clinical outcomes. The aim of this study is to synthesize MNP-MSN-PLGA-Tf NPs that could deliver DOX and PTX across BBB and target brain tumor. The core phase of MNP-MSNs was used to load DOX while the shell phase of PLGA was used to load PTX. Tf was conjugated on the surface of MNP-MSN-PLGA NPs to enhance the transport efficiency across BBB and target the brain glioma cells. The cytotoxicity and cellular uptake of the drug-loaded NPs were evaluated in vitro in the presence and absence of a magnetic field. The therapeutic efficacy of the drug-loaded NPs was also evaluated in vivo in U-87 tumor-bearing BALB/c nude mice. 2. Materials and methods 2.1. Materials Iron (III) chloride hexahydrate (FeCl3∙6H2O, 98%), sodium sulfite (Na2SO3,
98%), ammonium hydroxide (NH3∙H2O, 28e30%), tetramethyl orthosilicate (TMOS, 98%), (3aminopropyl)triethoxysilane (APTES,
98%), 1,3,5-triisopropylbenzene (TIPB, 96%), decane (
99%), human holo-transferrin (
98%), zinc sulfate monohydrate (ZnSO4$H2O), poly(vinyl alcohol) (PVA, MW ¼ 30,000e70,000 Da), dichloromethane (DCM), coumarin-6 (CM), hydrochloric acid (HCl), N-hydroxysuccinimide (NHS, 98%), 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC) hydrochloride, oleic acid (OA,
99%), acetonitrile (ACN,
99.5%) and ethylene glycol (EG, anhydrous,
98%) were obtained from SigmaeAldrich Corp. (St. Louis, MO). Hexadecyltrimethylammonium bromide (HTAB,
99%) was purchased from Acros Organics (Geel, Belgium). Poly(D,Llactic-co-glycolic acid) (PLGA 50:50, inherent viscosity ¼ 0.24e0.54 dL/g in hexafluoroisopropanol) was purchased from Lactel Absorbable Polymers (Pelham, AL). Doxorubicin (DOX), in the form of hydrochloride salt, was obtained from Boryung, Inc. (Seoul, Korea). Paclitaxel (PTX) was obtained as a gift from Bristol-Myers Squibb Company (New York, NY). CellTiter 96Ò Non-Radioactive Cell Proliferation Assay was purchased from Promega Corp. (Madison, WI). Formaldehyde solution (16%, methanolfree) was purchased from Thermo Fisher Scientific, Inc. (Rockford, IL). Phosphatebuffered saline (PBS, pH ¼ 7.4) was purchased from Mediatech, Inc. (Manassas, VA). 2.2. Preparation of magnetic nanoparticles (MNPs) MNPs were synthesized by semi-reduction co-precipitation method. The FeCl3 solution (8 ml, 1 M) was diluted in 100 ml of deionized water, stirred and protected under nitrogen at room temperature. The Na2SO3 solution (1.5 ml, 1 M) was added to the above solution, and the mixture was stirred. Subsequently, NH3∙H2O solution (5 ml, 28e30%) was added, and the mixture was stirred at 70e80  C for 30 min. In order to modify the surface of MNPs with OA, 0.33 ml of OA was added dropwise to

the above mixture with stirring during the 30-min period. Here, a short chain length of fatty acid was coated on the surface of NPs to produce sterically stabilized MNPs. The mixture was further incubated at 80  C for 50 min. The mixture was then heated to 110  C in order to evaporate excess ammonium hydroxide and water. Lastly, the pH of the mixture was adjusted by adding HCl in order to precipitate the MNPs. The oleic acid modified MNPs (OA-MNPs) were separated from the solution by using magnetic force, washed repeatedly with acetone and deionized water to remove excess OA, and lyophilized to obtain the powder [28]. 2.3. Preparation of DOX- and PTX-loaded magnetic silica PLGA nanoparticles (DOXPTX-NPs) MNP-MSNs were synthesized based on a revised method [29e31]. Briefly, 15 mg of OA-MNPs were dissolved in 1 ml of chloroform and added with HTAB solution (10 ml, 0.41 M). The mixture was stirred at 70  C in order to evaporate chloroform and obtain a homogeneous microemulsion. Subsequently, 35 ml of water and 5 ml of EG were added into the above mixture. After 15 min, NH3∙H2O solution (0.7 ml, 28e 30%) was added. To obtain a large pore size, decane and TIPB of different ratios were added into the mixture by stirring for 2 h. Finally, TMOS was added into the mixture under nitrogen protection at 60  C and stirred for another 2 h. The resulting product was washed and lyophilized. To remove the organic surfactant, the above product was dissolved and stirred in a solution containing 160 mg of ammonium nitrate in 60 ml of ethanol at 60  C for 30 min. The product was washed with ethanol and water, centrifuged and dried. To load DOX into the mesopores of MNP-MSNs, 10 mg of MNP-MSNs was first suspended in a 1:1 (v/v) methanol/H2O mixture. The mixture was then added with 15 mg of DOX and stirred under dark condition for 24 h. The DOX-loaded MNP-MSNs were centrifuged and washed three times with PBS. The surface of MNP-MSNs was later functionalized with amino groups [32]. Briefly, MNP-MSNs were immersed in absolute ethanol. The mixture was added with APTES and stirred at room temperature for 12 h. The modified NPs (MNP-MSNNH2 NPs) were then isolated and purified by centrifugation to remove excess APTES. DOX- and PTX-loaded magnetic silica PLGA NPs (DOX-PTX-NPs) were prepared by a double emulsion solid-in-oil-in-water solvent evaporation method [33,34]. The emulsification was performed using a microtip probe sonicator (VC 505, Sonics & Materials, Inc., Newtown, CT) at 25% amplitude. To load PTX in PLGA, 20 mg of PTX and 100 mg of PLGA were dissolved in 1 ml of DCM. The solution was added with 15 mg of DOX-MNPMSN-NH2 NPs, and the mixture was sonicated for 1 min. PVA (2 ml, 2% (w/v) in water) was added, and the mixture was sonicated for another 1 min. PVA (15 ml, 2% (w/v) in water) was added, and the mixture was stirred for 4 h in order to ensure complete evaporation of the organic solvent. The samples were washed and lyophilized. 2.4. Preparation of Tf-conjugated drug-loaded magnetic silica PLGA nanoparticles (DOX-PTX-NPs-Tf) Tf was conjugated on the surface of DOX-PTX-NPs by a two-step EDC/NHS activation and surface-grafting method [35,36]. Briefly, 10 mg of DOX-PTX-NPs were dispersed in 5 ml of PBS with EDC (250 ml, 1 mg/ml) and NHS (250 ml, 1 mg/ml), and the mixture was stirred at room temperature for 4 h to prepare amine-reactive esters from the carboxylic acid terminated PLGA. The NPs were dispersed in 2 ml of PBS, and Tf (100 ml, 1 mg/ml) was added dropwise to the mixture. The mixture was stirred at room temperature for 2 h and incubated at 4  C overnight. The samples were washed and lyophilized. 2.5. X-ray photoelectron spectroscopy (XPS) The surface chemistry of MNP-MSN-PLGA NPs and MNP-MSN-PLGA-Tf NPs was analyzed using XPS (AXIS Ultra DLD, Kratos Analytical Ltd., Manchester, UK) equipped with an Al Ka source (hn ¼ 1486.71 eV). The angle between sample surface and detector was kept at 90 . The survey spectrum ranging from 0 to 1100 eV was acquired. All binding energies were referenced to the C 1s (CeC bond) at 284.50 eV. 2.6. Drug loading and release The encapsulation efficiency (EE) and loading content (LC) of DOX and PTX in the NPs were determined by using the following equations. EE of DOXð%Þ ¼

initial amount of DOX used  amount of DOX in supernatant  100% initial amount of DOX used (1)

initial amount of DOX used  amount of DOX in supernatant  100% amount of NPs used (2) amount of PTX loaded into NPs EE of PTXð%Þ ¼  100% (3) initial amount of PTX used

LC of DOXð%Þ ¼

LC of PTXð%Þ ¼

amount of PTX loaded into NPs  100% amount of NPs used

(4)

Y. Cui et al. / Biomaterials 34 (2013) 8511e8520 The drug release was performed by suspending w10 mg of DOX-PTX-NPs in 5 ml of PBS with 0.05% (w/v) TweenÒ 80 in triplicate, and shaken at 37  C and 150 rpm. At specific time intervals, the supernatant was removed, and fresh PBS was replaced. The concentration of DOX was determined by measuring its absorbance using a spectrophotometer at a wavelength of 481 nm. The concentration of PTX was determined by high performance liquid chromatography (HPLC).

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SingHealth Institutional Animal Care and Use Committee (IACUC). Experimental practices were conducted in accordance with the National Advisory Committee for Laboratory Animal Research (NACLAR) guidelines in facilities licensed by the Agrifood and Veterinary Authority (AVA) of Singapore. Male BALB/c nude mice used in the experiments were 6e8 weeks old and weighed 20e25 g. All mice were housed in ventilated cages with free access to sterilized food and water, and acclimatized for 3e5 days prior to the experiments.

2.7. Cell culture and maintenance U-87 MG-luc2 cells were obtained from Caliper Life Sciences, Inc. (Hopkinton, MA). The cells were grown in Dulbecco’s modified Eagle’s medium (DMEM) (Gibco, Life Technologies Corp., Carlsbad, CA) with 10% fetal bovine serum (FBS) (HyClone, Thermo Fisher Scientific, Inc., Logan, UT) and 1% penicillin-streptomycin (PS) (PANBiotech GmbH, Aidenbach, Germany) in a humidified incubator under the conditions of 37  C and 5% CO2. Mouse brain endothelial cells, bEND.3, were obtained from American Type Culture Collection (ATCC Number: CRL-2299Ô). Similarly, the cells were grown in DMEM with 10% FBS and 1% PS in the humidified incubator at 37  C and 5% CO2. 2.7.1. In vitro cytotoxicity of drug-loaded NPs The cytotoxic effects of blank MNP-MSN-PLGA-Tf NPs, DOX-PTX-NPs, DOX-PTXNPs-Tf, DOX-PTX-NPs-Tf with free Tf, DOX-PTX-NPs-Tf with magnetic field, free DOX and free Taxol were determined using an MTT cell viability assay. Briefly, U-87 cells were seeded onto 96-well plates at a density of 5000 cells per well. On the following day, the cells were treated with different concentrations of blank MNP-MSN-PLGA-Tf NPs, DOX-PTX-NPs, DOX-PTX-NPs-Tf, DOX-PTX-NPs-Tf with free Tf, DOX-PTX-NPs-Tf with magnetic field, free DOX and free Taxol. For DOX-PTX-NPs-Tf with free Tf, free Tf was added together with the drug-loaded NPs to the cells. For DOX-PTX-NPs-Tf with magnetic field, the cells were treated with drug-loaded NPs and cultivated in the presence of a 0.1 T neodymiumeironeboron (NdFeB) permanent magnet (Master Magnetics, Inc., Castle Rock, CO) placed externally under each well. After 48 or 96 h of incubation, the cells were washed with PBS and replaced with fresh growth medium. The cell viability was then determined using a CellTiter 96Ò Non-Radioactive Cell Proliferation Assay (Promega Corp., Madison, WI). In particular, 15 ml of dye solution was added to 100 ml of medium in each well, and the cells were further incubated for 4 h. Following that, 100 ml of solubilization solution was added to each well, and the plates were gently shaken at room temperature for 24 h. The samples were then assayed with a spectrophotometer at 570 nm with a reference wavelength of 630 nm. The data were expressed as mean and standard deviation of six replicates. For bEND.3 cells, the method to evaluate the cytotoxic effect of blank MNP-MSN-PLGA-Tf NPs at different concentrations was the same as described above. 2.7.2. Cellular uptake of drug-loaded NPs For qualitative study using laser scanning confocal microscope, U-87 cells were seeded onto 24-well plates with glass cover slips at a density of 50,000 cells per well, incubated for 24 h, and treated with DOX-CM-NPs, DOX-CM-NPs-Tf, DOX-CM-NPs-Tf with free Tf and DOX-CM-NPs-Tf with magnetic field at a concentration of 200 mg/ml for 4 h. The cells were then washed with PBS and fixed using 4% formaldehyde at room temperature for 15 min. Subsequently, the cells were washed with PBS for three times and stained with Hoechst dye (1 mg/ml) for 30 min. The cells were washed with PBS for three times before the cover slips were mounted onto microscope slides and visualized using a Fluoview FV1000 laser scanning confocal microscope (Olympus Corp., Tokyo, Japan). Confocal images were captured using Olympus Fluoview software. For quantitative study using flow cytometry, U-87 cells were seeded onto 12well plates at a density of 100,000 cells per well, incubated for 24 h, and treated with CM-NPs, CM-NPs-Tf, CM-NPs-Tf with free Tf and CM-NPs-Tf with magnetic field at a concentration of 100 mg/ml for 4 h. Cells without any treatment were used as the control. The cells were then washed with PBS and harvested. The cellular uptake of CM was measured by flow cytometry analysis using FL-6 log filter for collection of fluorescence intensity. The transport of drug-loaded NPs across BBB model in vitro was also examined. The BBB model was established using a method previously described [37]. Briefly, a 2% gelatin solution was used to pre-coat the transwell inserts for 1 h. The bEND.3 cells were seeded onto the transwell inserts at a density of 50,000 cells per well and incubated for 7 days. The growth medium was changed every 2 days. After 7 days, the transendothelial electrical resistance (TEER) value of the BBB model was measured using EVOM & EVOMX Epithelial Voltohmmeters (World Precision Instruments, Inc., Sarasota, FL). The TEER value had to be over 250 U cm2 to be used for the experiments. The transwell inserts with the bEND.3 cells were then transferred to a 12-well plate with U-87 cells that was cultured for 24 h. Subsequently, the cells were subjected to CM-NPs, CM-NPs-Tf and CM-NPs-Tf with magnetic field treatments for 4 h. The U-87 cells were then washed with PBS and harvested. The cellular uptake of CM was measured by flow cytometry analysis as described above. 2.8. Animal care and maintenance All animal experiments were performed at the SingHealth Experimental Medicine Centre (SEMC), Singapore General Hospital (SGH), with approval from the

2.8.1. In vivo therapeutic efficacy of drug-loaded NPs BALB/c nude mice were anesthetized via intraperitoneal injection of ketamine (50 mg/kg) and diazepam (5 mg/kg), and immobilized on a stereotactic head frame (Thermo Fisher Scientific, Inc., Rockford, IL). A midline incision was made on the scalp and a burr hole was created 2.5 mm lateral to the sagittal sinus at the midpoint between bregma and lambda [22]. U-87 MG-luc2 cells (3  105 cells in 10 ml medium) were injected to a depth of 2.5 mm using a hypodermic needle at a flow rate of 1.0 ml/min, and the wound was closed by suture. After tumor cell inoculation, the non-invasive bioluminescence imaging (BLI) was performed using an IVISÒ animal imaging system (Xenogen, Alameda, CA) equipped with LivingImageÔ software (Xenogen) to monitor the intracranial tumor growth every five days. D-Luciferin potassium salt, the substrate for firefly luciferase, was dissolved in DPBS (without Ca2þ and Mg2þ) at a concentration of 15 mg/ml and filtered through a 0.2 mm filter before use. Mice were injected with luciferin at a dose of 150 mg/kg body weight. During image acquisition, isoflurane anesthesia in oxygen was delivered via a nose cone system. An integration time of 1 min was used for luminescent image acquisition. The signal intensity was quantified as the flux of all detected photon counts within a constant area of the region of interest (ROI) using the LivingImageÔ software. Peak signal time was obtained after 15 min of luciferin administration. The tumor-bearing mice were randomly separated into five groups, mainly i) saline or blank MNP-MSN-PLGA-Tf NPs, ii) DOX-NPs-Tf, iii) PTX-NPs-Tf, iv) DOX-PTXNPs, and v) DOX-PTX-NPs-Tf. Each group had five mice. After 20 days post tumor cell inoculation, the mice were injected with 100 ml of NPs via the tail vein, at a dose of 50 mg/kg for every three days with a total of four doses. BLI was performed 5, 10, 15 and 20 days after the start of treatment to monitor the intracranial tumor growth. After 20 days post-treatment, the mice were sacrificed for hematoxylin and eosin (H&E) staining. Brains were harvested and sliced into 3-mm coronal sections using a microtome, and fixed in 4% formaldehyde for 48 h. H&E staining was performed, and the cells were visualized under an optical microscope. 2.8.2. In vivo biodistribution of drug-loaded NPs The tumor-bearing mice were randomly separated into three groups, mainly i) blank MNP-MSN-PLGA-Tf NPs, ii) CM-NPs, and iii) CM-NPs-Tf. After 20 days post tumor cell inoculation, the mice were injected with 100 ml of NPs via the tail vein, at a dose of 100 mg/kg body weight. After 2 h, the in vivo distribution of NPs in the mice was visualized using the IVISÒ animal imaging system.

3. Results 3.1. Characterization of MNP-MSN-PLGA-Tf NPs TEM images show that MNP-MSN-PLGA-Tf NPs were spherical with a size of w150 nm (Fig. 1A and B). The MNPs loaded in mesoporous silica were localized at the center of the PLGA particles. After coating with PLGA and conjugating with Tf, the mesoporous structure of silica was not obvious. The surfaces of MNP-MSN-PLGA NPs with and without Tf conjugation were analyzed by XPS (Fig. 1C). In comparison to MNP-MSN-PLGA NPs, the XPS spectrum of MNP-MSN-PLGA-Tf NPs showed an obvious peak at w397 eV, which may be attributed to the 1s orbital of nitrogen atom from Tf. This indicates that Tf has been successfully conjugated on the surface of MNP-MSN-PLGA-Tf NPs. Zeta potentials of MNP-MSNs, MNP-MSN-NH2 NPs, MNP-MSNPLGA NPs and MNP-MSN-PLGA-Tf NPs were 9.3  0.8, 38.2  0.9, 34.3  1.1 and 18.1  0.5 mV, respectively. The negatively-charged surface of MNP-MSNs facilitated the loading of positively-charged DOX. After conjugating with Tf, the zeta potential of MNP-MSN-PLGA-Tf NPs remained negative when compared to that of MNP-MSN-PLGA NPs. It has been demonstrated that the anionic surface of drug delivery system would provide improved blood compatibility as compared to the cationic carrier [38]. In order to improve drug loading capacity in the mesopores, the swelling agents, decane and TIPB, were used to expand mesopores during the preparation of MNP-MSNs. Fig. 2A shows the nitrogen

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suitable in targeting drug-loaded MNP-MSN-PLGA-Tf NPs to a specific tumor site. 3.2. Encapsulation efficiency and in vitro release of DOX and PTX An ideal drug carrier should have a high drug loading capacity and encapsulation efficiency. Here, DOX and PTX were loaded into MNP-MSN-PLGA-Tf NPs. For PTX, it was effectively encapsulated into the NPs, achieving a drug loading of 136.4  2.1 mg/mg of NPs and an encapsulation efficiency of 89.2  1.1%. For DOX, the drug loading and encapsulation efficiency were 31.3  2.8 mg/mg of NPs and 22.0  0.8%, respectively. The controlled and sustained release of drugs from a carrier system is an important property for its clinical application in the biomedicine for cancer treatment. The drug release of DOX and PTX from MNP-MSN-PLGA-Tf NPs was performed in PBS with 0.05% (w/ v) TweenÒ 80 (Fig. 3). During the first 24 h, w19% of DOX and 57% of PTX were released. After 24 h, PTX was released at a slower rate. The release rate of DOX was also slower due to the presence of the external PLGA coating that effectively delayed its release from the mesopores of silica. The sustained release of DOX and PTX is highly beneficial in enhancing long-term anticancer efficiency and improving drug accumulation at the targeted site. 3.3. In vitro cytotoxicity of drug-loaded NPs

Fig. 1. (A) and (B) TEM images of MNP-MSN-PLGA-Tf NPs; (C) XPS spectra of MNPMSN-PLGA NPs and MNP-MSN-PLGA-Tf NPs.

gas adsorptionedesorption isotherms and pore size distribution for MNP-MSN-NH2 NPs. The BarretteJoynereHalenda (BJH) method was used to calculate the pore size distribution from the experimental isotherms after removing the structure-directing HTAB template. The adsorptionedesorption isotherms were considered as Type IV. The BrunauereEmmetteTeller (BET) specific surface area and pore volume of the NPs were w687.5 m2/g and 1.2 cm3/g, respectively. The peak pore diameter was w3.6 nm. Fig. 2B shows the magnetic property of the NPs using vibrating sample magnetometer (VSM). Both MNP-MSN-PLGA NPs and MNPMSN-PLGA-Tf NPs have superparamagnetism property, showing no remnant magnetization and hysteresis. Saturation magnetism (Ms) decreased only slightly after Tf conjugation. In particular, the Ms value of MNP-MSN-PLGA-Tf NPs was w4.5 emu/g, and it was higher than other reported result of 2 emu/g [18]. This property could be

MTT cell viability assay was performed to examine the cytotoxicity of drug-loaded NPs in U-87 cells under various conditions (Fig. 4A and B). Overall, the cytotoxicity of drug-loaded NPs depended on drug concentration and treatment period. Based on a 48-h treatment period (Fig. 4A), the IC50 values of the drug combination for DOX-PTX-NPs-Tf with free Tf, DOX-PTX-NPs, DOX-PTXNPs-Tf and DOX-PTX-NPs-Tf with magnetic field were 9.79, 8.89, 1.03 and 0.52 mg/ml, respectively. The IC50 of the drug combination could potentially be lowered by the presence of Tf-conjugation (w8.6-fold) and by the presence of magnetic field (w2.0-fold) to achieve a comparable growth inhibition of cells. Based on a 96-h treatment period (Fig. 4B), the IC50 values of the drug combination for DOX-PTX-NPs-Tf with free Tf, DOX-PTX-NPs, DOX-PTX-NPsTf and DOX-PTX-NPs-Tf with magnetic field were 2.85, 0.28, 0.13 and 0.045 mg/ml, respectively. With a longer treatment period, improvement in the reduction of the IC50 of the drug combination was also observed by the presence of Tf-conjugation (w2.2-fold) and by the presence of magnetic field (w2.9-fold). These results indicate that the delivery and the subsequent cellular uptake of drug-loaded NPs could be enhanced by the presence of magnetic field and the usage of Tf as targeting ligand, respectively, and hence,

Fig. 2. (A) Nitrogen gas adsorptionedesorption isotherms for MNP-MSN-NH2 NPs. Inset: Pore size distribution obtained from adsorption measurements (V ¼ pore volume and D ¼ pore size); (B) Magnetization curves for MNP-MSN-PLGA NPs and MNP-MSN-PLGA-Tf NPs as a function of applied magnetic field measured at 300 K.

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was achieved in bEND.3 and U-87 cells based on a concentration of 100 mg/ml and a treatment period of 96 h (Fig. 4D). 3.4. Cellular uptake of drug-loaded NPs

Fig. 3. In vitro release profiles of DOX and PTX from MNP-MSN-PLGA-Tf NPs.

leading to greater cytotoxicity of the combined drug treatment. However, when free Tf was added with the NPs, the cytotoxicity was reduced, possibly due to the competition between free Tf and Tf-conjugated NPs to bind with the Tf-Rs. Cells treated by DOX-PTXNPs-Tf with magnetic field achieved the highest chemotherapeutic efficiency, with w7.1- and 1.8-fold reduction in the IC50 of the drug combination after 96 h when compared to those treated by free DOX and free Taxol, respectively (Fig. 4C). The blank NPs did not show severe cytotoxicity in which cell viability of more than 80%

The uptake of drug-loaded NPs in U-87 cells was examined by laser scanning confocal microscopy (Fig. 5). Here, PTX was replaced by CM in order to visualize the drug distribution within the cells. Cells treated by DOX-CM-NPs-Tf showed higher intracellular accumulation of CM and DOX than those treated by DOX-CM-NPs. However, the drug distribution was uneven throughout the cells. Cells treated by DOX-CM-NPs-Tf with free Tf showed lower intracellular accumulation of CM and similar degree of DOX distribution as compared to those treated by DOX-CM-NPs-Tf without free Tf. Cells treated by DOX-CM-NPs-Tf with magnetic field showed higher intracellular accumulation of CM and DOX than those treated by DOX-CM-NPs-Tf without magnetic field. The uptake of drug-loaded NPs in U-87 cells was also quantified by flow cytometry (Fig. 6A). The geometric mean fluorescence intensities of CM in cells treated with CM-NPs, CM-NPs-Tf with free Tf, CM-NPs-Tf and CM-NPs-Tf with magnetic field were determined to be w388, 578, 775 and 1013, respectively (Fig. 6B). Cells treated with CM-NPs-Tf in the presence of a magnetic field showed the highest accumulation of CM among the four experimental groups. Meanwhile, a BBB model was established in vitro to assess whether drug-loaded NPs could be transported across the BBB (Fig. 6A). The geometric mean fluorescence intensities of CM in cells treated with

Fig. 4. Viability of U-87 cells after treatment with drug-loaded NPs for (A) 48 h, and (B) 96 h; (C) Viability of U-87 cells after treatment with free drugs for 48 and 96 h; (D) Viability of U-87 and bEND.3 cells after treatment with blank NPs for 48 and 96 h.

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Fig. 5. Confocal images of U-87 cells after treatment with drug-loaded NPs under various conditions at a concentration of 200 mg/ml for 4 h. The cell nuclei were stained by Hoechst dye and indicated in blue. CM and DOX were indicated in green and red, respectively. Scale bar ¼ 20 mm. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)

CM-NPs, CM-NPs-Tf and CM-NPs-Tf with magnetic field were w8, 13 and 33, respectively (Fig. 6B). In this case, the fluorescence intensities were lower than those in cells treated in the absence of the BBB. Though BBB was present, cells treated with CM-NPs-Tf in the presence of a magnetic field still showed the highest accumulation of CM among the three experimental groups. 3.5. In vivo biodistribution of drug-loaded NPs To evaluate in vivo BBB penetration and Tf-targeting efficacy of MNP-MSN-PLGA-Tf NPs, the biodistributions of two formulations, CM-NPs and CM-NPs-Tf, were examined in tumor-bearing mice (Fig. 7A). Blank NPs were used for the control group. A noticeable fluorescence signal at the brain tumor site was observed after CMNPs-Tf injection as compared to after CM-NPs injection. In particular, the fluorescence intensities of the circled regions in Fig. 7A were 0, 5.9  1010 and 1.3  1011 after injection of blank NPs, CMNPs and CM-NPs-Tf, respectively. Results indicate that Tf is an effective targeting ligand that could facilitate transport of NPs across BBB and improve drug accumulation in the brain tumor. This observation is consistent with other Tf-conjugated NPs [1,6].

3.6. Body weight and histological examination It is critical to evaluate the in vivo toxicity of the NPs used for drug delivery. The fluctuation in animal body weight is recognized as a useful indicator to assess in vivo toxicity of drug delivery systems. Here, the mice body weight was recorded for a period of 20 days after the start of treatment, and mice injected with saline were treated as the control group (Fig. 7B). Mice administrated with saline showed a steadily increasing body weight. Similar to the control group, mice administered with blank MNP-MSN-PLGA-Tf NPs exhibited no decline in body weight, indicating the nontoxicity of the NPs. Mice administered with DOX-PTX-NPs-Tf showed no serious body weight loss and deaths, indicating the good health condition of the mice in the drug treatment group. The histological examination of the tumor slices harvested from mice after 20 days post-treatment was performed to further evaluate the in vivo toxicity of the NPs (Fig. 7C). The DOX-PTX-NPs-Tf treatment showed the presence of apoptotic cells in the tumor tissue due to high drug accumulation possibly by Tf-mediated effect. In contrast, the blank MNP-MSN-PLGA-Tf NPs treatment showed normal and non-necrotic cells, similar to the saline

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Fig. 6. Flow cytometric analysis of U-87 cells after treatment with drug-loaded NPs under various conditions at a concentration of 200 mg/ml for 4 h. CM intensity is denoted as FL-6 Log.

treatment. Results indicate that the MNP-MSN-PLGA NPs are an ideal drug delivery system with no significant toxicity observed in vivo. The non-toxicity of the NPs could be attributed to the inertness of silica, and the biocompatibility and biodegradability of PLGA. Besides, iron oxide can be degraded by hydrolytic enzymes in lysosomes [39,40]. Thus, MNP-MSN-PLGA-Tf NPs could serve as useful drug delivery systems.

3.7. In vivo therapeutic efficacy of drug-loaded NPs To examine the therapeutic efficacy of the drug-loaded NPs in vivo, the intracranial tumor size was measured every five days by the IVISÒ animal imaging system. Fig. 8 shows the representative BLI of intracranial tumor before start of treatment (i.e., day 0) and after treatment (day 20). By day 20, it was apparent that the tumor

Fig. 7. (A) Fluorescence images showing biodistribution of NPs in tumor-bearing mice after 2 h tail vein injection of MNP-MSN-PLGA-Tf NPs (control), CM-NPs and CM-NPs-Tf. The blue circle indicates the location of the brain tumor; (B) Variation of mice body weight as a function of post-treatment time; (C) Histological examination of tumor slices harvested from mice after 20 days post-treatment. The blue circle represents the necrotic regions in the tumor interior. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)

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Fig. 8. Representative BLI of intracranial U-87 MG-luc2 xenograft in mice ((a)e(e)) before start of treatment and ((f)e(j)) after 20 days post-treatment. ((a) and (f)) Saline, ((b) and (g)) PTX-NPs-Tf, ((c) and (h)) DOX-NPs-Tf, ((d) and (i)) DOX-PTX-NPs, and ((e) and (j)) DOX-PTX-NPs-Tf. Luciferase expressing tumor cells are shown in the brain with a redeblue color bar, where the red and blue indicate the highest and lowest bioluminescence intensity, respectively. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)

growth was effectively inhibited by the DOX-PTX-NPs-Tf treatment as compared to the saline, PTX-NPs-Tf, DOX-NPs-Tf or DOX-PTXNPs treatments. Due to the different initial tumor size at day 0, the bioluminescence intensities of tumor at various time points were normalized with respect to the initial value at day 0 (Fig. 9). In comparison to the saline treatment, the DOX-PTX-NPs, DOX-NPs-Tf, PTX-NPs-Tf and DOX-PTX-NPs-Tf treatments resulted in w1.3-, 1.7-,

2.2- and 47.5-fold reduction in tumor size, respectively. It was clear that the DOX-PTX-NPs-Tf treatment exhibited the strongest inhibition of tumor growth as compared to other treatments. Both DOX-NPs-Tf and PTX-NPs-Tf treatments showed only minor inhibitory effect on the tumor size. In contrast, with the absence of Tf as the targeting ligand, the DOX-PTX-NPs treatment showed the least inhibitory effect on the tumor size. 4. Discussion

Fig. 9. Normalized bioluminescence intensity of intracranial U-87 MG-luc2 xenograft in mice as a function of post-treatment time.

The efficacy of chemotherapeutic agents in treating malignant brain glioma is poor due to low drug penetration and accumulation within the tumor. The aim of this study is to improve the chemotherapeutic efficacy in brain glioma treatment by using inorganicorganic composite NPs with a targeting ligand as the drug delivery system. In particular, Tf-conjugated magnetic silica PLGA NPs (MNP-MSN-PLGA-Tf NPs) were synthesized to deliver DOX and PTX (Fig. 10A). These drug-loaded NPs (DOX-PTX-NPs-Tf) could enhance transport across BBB and target glioma cells for effective cancer treatment (Fig. 10B). In order to prepare DOX-PTX-NPs-Tf, MNP-MSNs were first prepared to load DOX into the mesopores of silica. Subsequently, the DOX-loaded MNP-MSNs were coated with PTX-loaded PLGA on the surface via double emulsion solid-in-oil-in-water solvent evaporation method. The surface of DOX-PTX-NPs was then

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Fig. 10. (A) Preparation of DOX-PTX-NPs-Tf; (B) The transport of DOX-PTX-NPs-Tf across BBB and the targeting of glioma cells for cancer treatment.

modified with Tf via amide bond formation through a two-step EDC/NHS activation and surface-grafting method. These coree shell structured NPs could encapsulate hydrophilic and hydrophobic drugs, allow sustained drug release and achieve synergistic therapeutic effects. MSN-PLGA NPs could provide compatible cell micro-environment due to their ability to counteract the acidification of the environment by PLGA degradation products [41] via buffering with degradation products of silica particles. Besides, the incorporation of MNPs in MSN-PLGA NPs could provide magneticmediated delivery of drugs to the tumor site. However, the clinical application of MNPs in brain is limited [42], partly because of Food and Drug Administration (FDA) restrictions on the magnetic field strength subjected to humans. The XPS spectrum of MNP-MSN-PLGA-Tf NPs indicated the successful conjugation of Tf on the surface of the NPs (Fig. 1). In fact, only a small degree of Tf conjugation on the drug carrier is required to enhance endocytosis [43]. The superparamagnetism property of MNP-MSN-PLGA NPs before and after Tf conjugation was maintained, indicating their suitability for magnetic-mediated drug delivery (Fig. 2). The drug release profiles showed that DOX and PTX were released quickly during the first 24 h, followed by slow release over the next 4 days (Fig. 3).

For in vitro experiments, the delivery and the subsequent cellular uptake of DOX-PTX-NPs-Tf could be enhanced by the presence of magnetic field and the usage of Tf as targeting ligand, respectively. Here, the DOX-PTX-NPs-Tf with magnetic field treatment resulted in the greatest cytotoxicity as compared to the DOXPTX-NPs-Tf, DOX-PTX-NPs or DOX-PTX-NPs-Tf with free Tf (Fig. 4A and B). This may be attributed to the high intracellular drug accumulation as evident from confocal images (Fig. 5) and flow cytometric analysis (Fig. 6). By having the BBB model in vitro, the uptake of NPs was also enhanced by the presence of magnetic field and the usage of Tf as targeting ligand (Fig. 6). For in vivo experiments, the DOX-PTX-NPs-Tf treatment resulted in the strongest inhibition of tumor growth, followed by PTX-NPs-Tf, DOX-NPs-Tf and DOX-PTXNPs (Figs. 8 and 9). The DOX-PTX-NPs treatment showed the least inhibitory effect on the tumor size, indicating the importance of Tf as the targeting ligand for the delivery of drugs to the brain tumor. For in vitro experiments, cells treated with blank MNP-MSNPLGA-Tf NPs showed no severe cytotoxicity even though a high concentration of NPs (100 mg/ml) and a long treatment duration (96 h) were used (Fig. 4D). Moreover, mice administrated with blank MNP-MSN-PLGA-Tf NPs did not exhibit a decline in body weight, a profile similar to the saline treatment (Fig. 7B). Besides, the

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histological examination of tumor slices harvested from mice that were administrated with blank MNP-MSN-PLGA-Tf NPs showed normal and non-necrotic cells, an observation similar to the saline treatment (Fig. 7C). These results indicate that MNP-MSN-PLGA-Tf NPs are promising materials for use in drug delivery systems. Overall, it is clear that the transport efficiency across BBB and the cellular uptake of drugs are critical factors that affect the chemotherapeutic efficacy on brain glioma treatment, and these challenges can be overcome by the use of MNP-MSN-PLGA-Tf NPs. 5. Conclusions In this study, Tf-conjugated magnetic silica PLGA NPs (MNPMSN-PLGA-Tf NPs) loaded with DOX and PTX (DOX-PTX-NPs-Tf) were successfully prepared for effective brain glioma treatment. The delivery and the subsequent cellular uptake of DOX-PTX-NPs-Tf could be enhanced by the presence of magnetic field and the usage of Tf as the targeting ligand, respectively, leading to the greatest cytotoxicity as observed in vitro. Moreover, the DOX-PTX-NPs-Tf treatment exhibited the strongest inhibition of tumor growth as observed in vivo. Due to the biocompatibility of MNP-MSN-PLGA-Tf NPs, they serve as a suitable drug delivery system that could be applied for the treatment of brain glioma. Acknowledgments The authors acknowledge the funding support from National Medical Research Council (NMRC, Singapore) (NMRC/EDG/0067/ 2009), National University of Singapore (NUS) (R279-000-257-731), and Shanghai Municipal Science and Technology Commission Foundation Research Project (12JC1408500). We thank Professor Meng Cheong Wong (Department of Pharmacology, NUS) for his advice on the design of in vivo experiments. Yanna Cui also acknowledges Chinese Scholarship Council (CSC, State-Sponsored Graduate Scholarship Program for Building High-Level Universities) for supporting her Ph.D. program during her exchange in NUS. References [1] Zhang P, Hu L, Yin Q, Feng L, Li Y. Transferrin-modified c[RGDfK]-paclitaxel loaded hybrid micelle for sequential blood-brain barrier penetration and glioma targeting therapy. Mol Pharm 2012;9:1590e8. [2] Mangiola A, Anile C, Pompucci A, Capone G, Rigante L, De Bonis P. Glioblastoma therapy: going beyond Hercules Columns. Expert Rev Neurother 2010;10:507e14. [3] Xin H, Sha X, Jiang X, Zhang W, Chen L, Fang X. Anti-glioblastoma efficacy and safety of paclitaxel-loading angiopep-conjugated dual targeting PEG-PCL nanoparticles. Biomaterials 2012;33:8167e76. [4] Ding H, Inoue S, Ljubimov AV, Patil R, Portilla-Arias J, Hu J, et al. Inhibition of brain tumor growth by intravenous poly (b-L-malic acid) nanobioconjugate with pH-dependent drug release. Proc Natl Acad Sci U S A 2010;107:18143e8. [5] Zhan C, Wei X, Qian J, Feng L, Zhu J, Lu W. Co-delivery of TRAIL gene enhances the anti-glioblastoma effect of paclitaxel in vitro and in vivo. J Control Release 2012;160:630e6. [6] Yue J, Liu S, Wang R, Hu X, Xie Z, Huang Y, et al. Transferrin-conjugated micelles: enhanced accumulation and antitumor effect for transferrin-receptoroverexpressing cancer models. Mol Pharm 2012;9:1919e31. [7] Ojima I. Modern molecular approaches to drug design and discovery. Acc Chem Res 2008;41:2e3. [8] Thorpe PE. Vascular targeting agents as cancer therapeutics. Clin Cancer Res 2004;10:415e27. [9] Parren PW, van de Winkel JG. An integrated science-based approach to drug development. Curr Opin Immunol 2008;20:426e30. [10] Pang Z, Feng L, Hua R, Chen J, Gao H, Pan S, et al. Lactoferrin-conjugated biodegradable polymersome holding doxorubicin and tetrandrine for chemotherapy of glioma rats. Mol Pharm 2010;7:1995e2005. [11] Xin H, Jiang X, Gu J, Sha X, Chen L, Law K, et al. Angiopep-conjugated poly(ethylene glycol)-co-poly(ε-caprolactone) nanoparticles as dual-targeting drug delivery system for brain glioma. Biomaterials 2011;32:4293e305. [12] Jones AR, Shusta EV. Blood-brain barrier transport of therapeutics via receptor-mediation. Pharm Res 2007;24:1759e71. [13] Martell LA, Agrawal A, Ross DA, Muraszko KM. Efficacy of transferrin receptortargeted immunotoxins in brain tumor cell lines and pediatric brain tumors. Cancer Res 1993;53:1348e53.

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