Noncovalent Surface Locking of Mesoporous Silica ... - ACS Publications

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Article

Noncovalent Surface Locking of Mesoporous Silica Nanoparticles for Exceptionally High Hydrophobic Drug Loading and Enhanced Colloidal Stability L. Palanikumar,1 Ho Young Kim,1 Joon Yong Oh,1 Ajesh P. Thomas,1 Eun Seong Choi,1 M. T. Jeena,1 Sang Hoon Joo,1,2 and Ja-Hyoung Ryu1,* 1

Department of Chemistry, School of Natural Science and 2Department of Chemical

Engineering, School of Energy and Chemical Engineering, Ulsan National Institutes of Science and Technology (UNIST), Ulsan, 689-798, Korea *To whom correspondence should be addressed. E-mail: [email protected] Abstract: Advances in water-insoluble drug delivery systems are limited by selective delivery, loading capacity, and colloidal and encapsulation stability. We have developed a simple and robust hydrophobic-drug delivery platform with different types of hydrophobic chemotherapeutic agents using a non-covalent gatekeeper’s technique with mesoporous silica nanoparticles (MSNs). The unmodified pores offer a large volume of drug loading capacity, and the loaded drug is stably encapsulated until it enters the cancer cells owing to the noncovalently bound polymer-gatekeeper. In the presence of polymer-gatekeepers, the drugloaded mesoporous silica nanoparticles showed enhanced colloidal stability. The simplicity of drug encapsulation allows any combination of small chemotherapeutics to be coencapsulated and thus produce synergetic therapeutic effects. The disulfide moiety facilitates decoration of the nanoparticles with cysteine containing ligands through thiol-disulfide chemistry under mild conditions. To show the versatility of drug targeting to cancer cells, we decorated the surface of the shell-crosslinked nanoparticles with 2 types of peptide ligands, SP94 and RGD. The nanocarriers reported here can release encapsulated drugs inside the reducing microenvironment of cancer cells via degradation of the polymer shell, leading to cell death.

Keywords: mesoporous silica nanoparticle, hydrophobic drug delivery, polymer-gatekeepers, drug-loading capacity, colloidal stability 1

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INTRODUCTION Many small therapeutic molecules used as anticancer agents are limited by their poor solubility in aqueous solutions.1 Methods of solubilizing and delivering such drugs to specific cancer cells with nanoscale carriers, including spherical micelles, cylindrical micelles, hollow capsules, nanogels, and inorganic nanoparticles, have been extensively explored.2 However, development of a versatile drug delivery platform for noncovalent encapsulation of various kinds of hydrophobic drug molecules without leakage, with a large loading capacity, and with high colloidal stability, has been challenging. Among drug nanocarriers, polymeric assembling systems have been considered the most promising hydrophobic drug carriers owing to their unique core-shell architecture.3 The hydrophobic core is used to improve the solubility of hydrophobic drugs, whereas the hydrophilic shell provides colloidal stability in aqueous media.2a, 3d, 4 In addition, various molecular designs allow incorporation of desired functions into drug delivery systems, such as on-demand control of drug release profiles using stimuli-sensitive chemical functional groups and high cancer selectivity by addition of targeting ligands.2c,

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incorporation of hydrophobic drug molecules into the nanocarrier core requires optimization on a case-by-case basis by varying the drug molecules, because this process primarily depends on the interaction between the core materials and the hydrophobic molecules.3b, 4-6 Most reported polymeric assemblies showed a moderate loading capacity (10–20 wt%, drugto-carrier weight ratio) for particular drug molecules which have low hydrophobicity, such as doxorubicin (Dox), curcumin (Cur), and tamoxifen (TMX). However, for highly hydrophobic drug molecules, including camptothecin (CPT) and paclitaxel (PTX), polymeric assemblies show low loading capacity (typically less than 10 wt%). Mesoporous silica nanoparticles (MSNs) have received wide attention in nanomedicine research and have been used as drug delivery carriers, gene transfection agents, 2


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and cell markers, due to a large surface area, tunable pore size, and an easily functionalizable surface.7 Recently, several researchers have reported utilization of MSNs as a reservoir for hydrophobic drugs via physical adsorption of drug molecules onto the surface of the mesoporous structure.8 However, drug-loaded MSNs showed poor colloidal stability in aqueous solutions and uncontrolled drug release due to their inherent leaky nature, whereas unmodified MSNs showed potential toxicity at a high concentration.9 Increasing the colloidal stability of drug-loaded MSNs requires complicated chemical modification of the MSN surface, which can limit loading efficiency due to reduced surface area or affect drug release.10 While hydrophilic drug delivery systems using mechanized MSNs have been extensively investigated, hydrophobic drug delivery systems utilizing MSNs have not been widely studied due to the difficulties mentioned above. In this work, we report the development of a simple, facile, and versatile drug delivery platform utilizing MSNs with polymer-gatekeepers, which is capable of encapsulating a variety of hydrophobic drug molecules with exceptionally high loading capacity and enhanced colloidal stability (Scheme 1). We hypothesized that non-covalent post-modification of the surface of the hydrophobic drug-loaded MSN with a selfcrosslinkable hydrophilic polymer would increase the colloidal stability, leading to effective dispersion into physiological media and preventing drug leakage by blocking the gate to the particle core. Because the polymer at the shell is degradable in response to a specific stimulus, we are able to finely control the time and place at which the drug is released from the carrier. We have previously reported the use of a non-covalent polymer-gatekeeper in MSNs, allowing entrapment of large amounts of hydrophilic drug molecules in the pores of pristine MSNs, as well as release of drug molecules inside cancer cells by degradation of the disulfide functional groups in the polymer shells.11 In the present investigation, we applied this strategy to develop a versatile hydrophobic drug delivery system with a large payload and enhanced 3


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colloidal stability. Consequently, various drug molecules can be easily encapsulated in the MSNs without consideration of changes in drug loading capacity specific to each drug. MSNs have a very large surface area, allowing physical adsorption of large amounts of hydrophobic drug molecules.9 Moreover, the drug loading capacity of MSNs can be increased by using high drug concentrations in organic solutions during the drug feeding process or repeating the drug adsorption procedure.9, 12 The polymer-gatekeeper strategy, using a self-crosslinkable random copolymer containing pyridine disulfide hydrochloride (PDS) and polyethylene glycol (PEG) as side chains, could be applied to wrap the hydrophobic drug-loaded MSNs through the weak electrostatic interaction between the MSNs and the polymer, followed by crosslinking of the polymer shell through a disulfide exchange reaction.11 We assessed the versatility of MSNs as a hydrophobic drug delivery system by encapsulating clinically important hydrophobic drugs Dox, PTX, CPT, TMX, and Cur. Owing to the simplicity of hydrophobic drug encapsulation, co-delivery of combinations of these multifunctional agents, such as that of Dox, CPT, and PTX, or of Dox, CPT, and Cur, in one nanoparticle could achieve enhanced therapeutic effects. In addition, the surface of the polymer-capped MSNs can be functionalized with a target ligand to increase selectivity for cancer cells. We utilized 2 types of peptide ligands, cyclic RGD and SP94, to selectively target cancer cells.

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A

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PEG-PDS In-situ crosslinking Enhanced Colloidal & encapsulation stability

High hydrophobic drug loading

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PEG-b-PLA

Hydrophobic drugs

Aggregated

Scheme 1. Hydrophobic drug encapsulation with polymer-gatekeeper MSN (PMSN) or a block copolymer. Schematic diagram for A) preparation of a large mass of hydrophobic drugloaded MSNs and installation of non-covalent polymer-gatekeepers to show enhanced colloidal and encapsulation stability, and B) comparison of hydrophobic drug encapsulation in block copolymer micelles at low and high drug feeding concentrations. At a low feeding concentration, the block copolymer self-assembles into micelles, while at a high feeding rate the block copolymer tends to aggregate.

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EXPERIMENTAL METHODS Reagent. All materials and reagents were obtained from Sigma-Aldrich (Yongin, Korea) and Tokyo Chemical Industries (TCI) (Tokyo, Japan) unless otherwise specified and were used as received without any purification. CPT and PTX were obtained from Ontario Chemicals Inc. (Ontario, Canada). Doxorubicin (Dox) was obtained from Acorn PharmaTech (Redwood City, CA, USA). Tamoxifen citrate was obtained from Medchemexpress LLC (Princeton, NJ, USA). All cell culture reagents were obtained Invitrogen, Gibco, and Life Technologies (Seoul, Korea). Preparation of CPT drug-loaded MSNs and size analysis. MSNs (5.0 mg) were dispersed well in 1 mL of drug solution (0.5 mg, 1.5 mg, 2.5 mg, 5.0 mg, 7.5 mg, or 10.0 mg in DMSO) and stirred for a period of 3 h at room temperature. After the stipulated time periods, the drug-loaded nanoparticles were centrifuged and the supernatant was collected. The supernatant sample was used to measure drug loading by collecting UV-Vis absorption spectra using a molar absorption coefficient of 42,282 M-1cm-1 with λmax = 365 nm.28 Drug loading was calculated using the following equation: Drug loading capacity (%) = Mass of the drug in MSN/Mass of MSN × 100 Entrapment efficiency (%) = Mass of the loaded drug/Initial mass of the drug × 100 For sequential drug loading studies (cycle experiments), 5.0 mg of MSN was dispersed well in 1 mL of drug solution (5.0 mg/mL) and stirred for a period of 3 h at room temperature (for 1 cycle). Next, the drug-loaded nanoparticles were collected by centrifugation, vacuum-dried, and redispersed in 5.0 mg of drug solution. The drug-loading process was repeated, and the size of the drug-loaded MSNs was measured by DLS in an aqueous suspension at 1 mg/mL of MSN after each cycle. Preparation of CPT-loaded polymeric micelles. The block copolymer (10 mg) and hydrophobic drug CPT (0.1 mg, 0.5 mg, or 1.0 mg) were dissolved in 1 mL DMSO. The 6


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resulting solution was stirred vigorously for a period of 24 h and filtered using a 0.45-µm PTFE filter to remove any undissolved components. The polymer solution was dialyzed against water for 24 h to induce micelle formation, with frequent replacement of water. Due to observation of an off-white solid precipitate, the micelle solutions were passed through 0.45-µm filters to remove any free CPT and precipitate, which were then lyophilized and redissolved in water or DMSO for characterization.6a CPT drug loading and entrapment efficiency, as a percentage of weight, were determined using UV-Vis spectroscopy and calculated using the molar absorption coefficient (42,282 M-1cm-1 at λmax = 365 nm). Drug loading capacity (%) = Mass of CPT in micelles/Mass of Polymer × 100 Entrapment efficiency (%) = Mass of CPT in micelles/Initial Mass of CPT × 100 Synthesis of hydrophobic drug-loaded MSNs and wrapping of the MSN surface with PEG-PDS copolymer. MSNs (5 mg) were dispersed in 1 mL of drug solution (dichloromethane for Dox, DMSO for CPT, acetonitrile for PTX, acetone for Cur, and methanol for TMX) and stirred for 4 h at room temperature. The drug-loaded nanoparticles were collected by centrifugation and washed 3 times with DMSO to remove the unabsorbed drug on the particle surface, and the supernatant and washed solution were subjected to UVVis spectroscopy and high-performance liquid chromatography (HPLC) to measure drug loading capacity and entrapment efficiency.29 The drug-loaded nanoparticles were vacuum-dried, redispersed in PEG-PDS copolymer solution (15 mg/mL), and stirred for 3 h. To crosslink the surface-wrapped polymer, a partial amount of DTT (10 to 50 mol% against the PDS group) was added and the resulting solution was stirred for 3 h at room temperature. Drug-loaded PMSNs were collected by centrifugation and washed extensively with pH 7.4 phosphate buffer solution and distilled water. The crosslinking density was analyzed by checking the release of the byproduct pyridothione using the known molar extinction coefficient of pyridiothione (8.08 × 103 M-1 7


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cm-1 at 343 nm).2b, 2c Co-encapsulation of 3 hydrophobic drug molecules in PMSNs. Three anticancer drugs (CPT, PTX, and Dox; 5.0 mg each) were dissolved in 1 mL of DMSO.15c Next, 5 mg of MCM-41 MSNs were suspended in the CPT/PTX/Dox solution, dispersed by sonication, and stirred continuously for 3 h at room temperature. The drug-loaded MSNs were collected by centrifugation, washed with DMSO, and vacuum dried. In detail, the drug loaded nanoparticles were vacuum dried, once again washed with DI water, centrifuged and freeze dried in order to remove any DMSO/water molecules available in the nanoparticles. The polymer gatekeeper was installed using the procedure described above. Drug loading parameters of the PMSNs were calculated using UV-Vis spectroscopy and HPLC. Drug release profile. The release kinetics of the drug-loaded nanoparticles in phosphatebuffered saline (pH 7.4) at room temperature were analyzed at excitation wavelengths of 480 nm, 430 nm, and 360 nm for Dox, Cur, and CPT, respectively, at several time points using a Shimadzu RFPC 5430 Spectrofluorometer. The drug release profiles of TMX and PTX were analyzed using HPLC with methanol or acetonitrile and water. In order to measure triggered drug release profiles, 5 mM of GSH was added to the nanoparticle solution at 4 h and the release profile was analyzed using a fluorometer and HPLC. Ligand decoration. Approximately 1 mg of CPT-loaded PMSNs (CPT-PMSN) was dispersed well in 1 mL of DI water. To this solution, 1 mg of cyclo(RGDfC) peptide ligand (RGD) was added, and the resulting solution was stirred at 200 rpm for 24 h at room temperature. For SP94, 3 mg of CPT-PMSNs was dispersed well in 1 mL of DI water containing 3 mg of SP94 (AcCGGSFSIIHTPILPL) peptide ligand, and the resulting solution was stirred for 24 h at room temperature. Ligand decoration was confirmed by the release of the byproduct pyridiothione via UV-Vis spectroscopy. Cell viability assays.

KB cells were cultured in sterile 96-well Nunc (Thermo Fisher 8


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Scientific Inc.) microtiter plate at a seeding density of 5 x 103 cells/well and allowed to settle for 24 h under incubation at 37 °C and 5% CO2 in RPMI 1640 medium. In order to check cell viability, the cells were then treated with different concentrations of plain MSN (0.01, 0.05, 0.10. 0.25, 0.50, 1.00 and 2.00 mg/mL), PMSN (0.01, 0.05, 0.10. 0.25, 0.50, 1.00 and 2.00 mg/mL), drug loaded PMSN (concentrations of 0.01, 0.05, 0.10. 0.25, 0.50, 1.00 and 2.00 µg/mL of drugs). Cell viability was measured at 48 h using the alamar blue assay, with each data point measured in triplicate. Fluorescence measurements were made using the plate reader (Tecan Infinite Series, Germany) by setting the excitation wavelength at 565 nm and monitoring emission at 590 nm on the 96 well plates. Similarly the cell viability analysis for RGD-PMSN and SP94-PMSN were analyzed in KB cells and HepG2 cells (cells were washed and replaced with fresh medium after 1 h of incubation) by checking the viability after 48 h using the alamar blue assay. In-situ confocal microscopy for cellular internalization. KB cells were seeded in one well glass cover slips at a seeding density of 2 x 105 cells/well. After 24 h, cells were treated with CPT loaded RGD-PMSN nanoparticles(10 µg/mL of CPT). Similarly, HepG2 cells were seeded at 2 x 105 cells/well in one well glass cover slips. After 24 h, cells were treated with CPT-loaded SP94-PMSN nanoparticles(10 µg/mL of CPT). The cellular uptake was monitored in the coverglass (Lab Tek II glass chamber coverglass, Thermo Fisher Scientific Inc) were monitored periodically using Carl Zeiss LSM 700 microscope connected to CO2 incubator. Flow-cytometry analysis. To analyze the cellular intake non-ligand and ligand decorated nanoparticles, KB and HepG2 cells were seeded at 5 x 103 cells in 24 well plates. After 24 h, cells were treated with non-ligand and ligand decorated CPT-PMSN for 3 h. Then the cellular uptake of CPT was analyzed using FACS (FACS caliber, BD Bioscience). Combination index analysis. In order to check the effect of drug combination effect in KB 9


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cells, the cell viability of three different drugs loaded (PTX, CPT and Dox) PMSNs nanoparticles and single drug loaded-PMSNs was compared. In order to check this, KB cells were cultured in sterile 96-well Nunc microtiter plate at a seeding density of 5 x 103 cells/well and they were allowed to settle for 24 h under incubation at 37 °C and 5% CO2 in RPMI 1640 medium. Cells were incubated with different concentrations of single drug loaded-PMSN and three drugs loaded-PMSN for a period of 24 h, and replaced with fresh medium. After 24h, cell viability were analyzed using the alamar blue assay. Combination index (CI) analysis were calculated using the previously reported Chou-Talay principle; CI = ICx(D1) /ICx(Dm1) + ICx(D2) /ICx(Dm2) + ICx(D3)/ICx (Dm3) where D1, D2 and D3 are the concentrations of drug 1, drug 2 and drug 3 that in combination to produce certain cytotoxicity (e.g., 50% inhibition of cells), while Dm1, Dm2 and Dm3 are the concentrations of three drugs at which single drugs alone have the same effect when administered alone.15c, 23

RESULTS AND DISCUSSION Synthesis of MSNs (MCM41 and MCM48) and CPT encapsulation. Ordered MCM-41 and MCM-48 type MSNs were prepared according to a previously reported procedure.13 Scanning electron microscopy (SEM) and transmission electron microscopy (TEM) observations showed well-defined spherical nanoparticles with an average diameter of approximately 130 nm and 190 nm for MCM-41 and MCM-48 MSNs, respectively (Figure S1). Magnified TEM images and X-ray diffraction (XRD) analysis confirmed the well-ordered hexagonal mesostructure of the MCM-41 MSNs and the highly ordered 3-D cubic Ia3d mesostructure of the MCM-48 MSNs. Nitrogen adsorptiondesorption isotherm measurements showed that the MCM-48 MSNs had a larger surface area (1241 m2/g) and total pore volume (0.96 cm3/g) than those of the MCM-41 MSNs (789 m2/g and 0.79 cm3/g, respectively). 10


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We investigated the loading capacity of hydrophobic drug CPT in the unmodified MSNs via physical entrapment (Scheme 1). The MSNs were immersed in organic solutions containing CPT at a range of concentrations for several hours. The excess drug solution and drug-loaded MSNs were separated using centrifugation and the MSN pellets were dried in a vacuum. The drug loading capacity (DLC, weight ratio between drug and MSNs) can be increased by increasing the drug concentration (from 0.5 mg/mL to 10.0 mg/mL, the maximum solubility of CPT in dimethyl sulfoxide (DMSO)) of the feeding solution when the concentration of the nanoparticles is fixed (5 mg/mL). The loading capacities are summarized in Table 1. The DLC of the MCM-41 MSNs ranged from 9 wt% to 120 wt%. The MCM-48 MSNs had a loading capacity much larger than that of the MCM-41 MSNs (up to 130 wt%), implying that the surface area of MSNs was directly related to loading capacity, because the primary drug encapsulation mechanism is physical adsorption of the drug molecule on the MSN surface by non-covalent interactions, including hydrogen bonding and van der Waals interactions.9 It should be noted that CPT is a very hydrophobic drug molecule that generally shows very low loading capacity in polymer assembling systems. Although the encapsulation efficiency (weight ratio of the loaded-drug and the fed-drug) was low, the process was costeffective, because the remaining drug solution in the supernatant could be reused after centrifugation without further purification. Silica nanoparticles containing different amounts of CPT were resuspended in deionized water by sonication, and dynamic light scattering (DLS) measurement showed no change in particle size, indicating that most of the drug molecules were located in the appropriate place inside the mesoporous nanoparticles (Figure 1a and 1c). Another strategy to increase the loading capacity of drug carriers is repeating the process of immersing, centrifuging, and drying the particles. When we repeated these steps 5 times using the same mass ratio of the drug and MSNs (5 mg/mL each), we found that the drug loading capacity was continuously increased up to 291 wt% and 339 wt% for MCM-41 11


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and MCM-48 MSNs, respectively (Table S1). However, DLS showed that drug aggregates were formed from the 4th cycle, indicating that 3 repetitions of the drug loading procedure is the optimal method for maximizing drug loading capacity (141 wt% and 200 wt% for MCM41 and MCM-48 MSNs, respectively) without affecting nanoparticle size (Figure 1b and 1d).

Table 1. Drug loading capacity and entrapment efficiency of CPT in MCM-41 and MCM-48 nanoparticles Drug conc (Feeding)

CPT DLC wt% (EE %)* MCM-41

MCM-48

0.5 mg/mL

9 wt% (90%)

9 wt% (90%)

1.5 mg/mL

15 wt% (50%)

17 wt% (57%)

2.5 mg/mL

33 wt% (66%)

40 wt% (80%)

5.0 mg/mL

62 wt% (62%)

86 wt% (86%)

7.5 mg/mL

88 wt% (59%)

105 wt% (70%)

10.0 mg/mL

120 wt% (60%)

130 wt% (65%)

* Drug loading capacity (DLC) (wt%) = mass of drug loaded in nanoparticles/mass of nanoparticles. Entrapment efficiency (EE) (%) = mass of drug loaded in nanoparticles/initial mass of drug.

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100

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Figure 1. Hydrodynamic particle size distribution analysis for A) MCM-41 and C) MCM-48 MSNs with different CPT loading capacity, and B) MCM-41 and D) MCM-48 MSNs during the cyclic loading experiments. The loading capacity is expressed in parentheses.

The suspension of drug-loaded MSNs showed short-term colloidal stability. When the suspension was stored without motion, the MSNs began to aggregate in less than 30 min, (confirmed by DLS, Figure 2a), probably due to increased hydrophobicity conferred by the drug molecules adsorbed onto the surface of the MSNs. Furthermore, the drug molecules were continuously released from the MSNs into the aqueous solution (Figure 2b). These results indicate that pristine MSNs are not an ideal drug delivery vehicle, because an a ideal nanocarrier should be stable under physiological conditions and should not release its payload until it reaches the target tumor cell to prevent undesired side effects in healthy cells. To solve the issues of limited colloidal and encapsulation stability, we applied the polymer-gatekeeper strategy, which was previously developed for the production of hydrophilic drug delivery platforms.11 Briefly, the surface of the MSNs was coated with a self-crosslinkable copolymer 13


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using the electrostatic interaction between the positively charged polymers and the negatively charged silica surface, while subsequent crosslinking of the polymer shell generated stable colloids (Figure S3b and S3c). The polymer-wrapped MSNs (PMSNs) produced by this process were assessed by TEM, and changes in size and surface charge were evaluated via DLS and zeta potential measurement, respectively (Figure 2c, d). The polymer coating provided long-term colloidal stability. As shown in Figure 2e, the size of the PMSNs was not changed during 72 h in phosphate buffer (pH 7.4), sodium acetate buffer (pH 4.0), or serum containing solution (Figure S3), and no precipitation was observed for 3 months. This high colloidal stability may be conferred by the copolymer, which contains polyethylene glycol units for water solubility and serum stability, while the polymer-gatekeeper blocked undesired drug leakage from the MSNs.

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80 60 40 20 0

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Figure 2. A) Hydrodynamic particle size distribution to assess the colloidal stability of CPTloaded MCM-41 MSNs after 30 min. The inset image of precipitation/flocculation indicates low colloidal stability. B) Drug release profile for CPT from MSNs under normal physiological conditions (PBS at pH 7.4). C) Hydrodynamic particle size analysis for MSNs and PMSNs and inset shows a TEM image of PMSNs. D) Zeta potential analysis for MSNs and PMSNs in an aqueous suspension. E) Hydrodynamic particles size of PMSNs in RPMI 1640 medium with 10% fetal bovine serum. F) Drug release pattern of PMSNs under normal physiological conditions (PBS at pH 7.4).

We hypothesized that the method we used would not be significantly affected by the hydrophobicity of the encapsulated drug molecules, because, in contrast to block co-polymer 15


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micelles, encapsulation of hydrophobic drug molecules inside MSNs is performed independently of the installation of the polymer-gatekeeper. Generally, encapsulation of hydrophobic drug molecules in block copolymer micelle systems occurs during the selfassembly process. To investigate this difference, we compared the degree of encapsulation of strongly hydrophobic drug CPT in PMSNs and block copolymer micelles. We synthesized a poly(ethylene glycol)-b-poly(lactic acid) (PEG-b-PLA) block copolymer that consisted of 5000 Mw PEG and 5000 Mw PLA. PEG-b-PLA self-assembled into micelles with an average diameter of 240 nm in an aqueous solution (Figure S3a). When 1 wt% of CPT was added to the block copolymer micelle solution, most of the drug molecules in the feeding solution were loaded into the micelles with 90% entrapment efficiency (Table 2). However, when we increased the feeding amount to 5 wt% and 10 wt%, we observed precipitation from the solution and larger aggregates in DLS, although the amount of encapsulated CPT was increased (Figure 3a and 3b). This precipitation and aggregation might have been caused by inter-particular aggregation due to surface-bound hydrophobic drug molecules or fast selfaggregation of CPT molecules due to strong π-π intermolecular interaction of its planar aromatic ring structure, because of the weak interaction between the block copolymer core and CPT. These results indicate that block copolymer micelles can tend to precipitate upon the increase in concentration of drug feeding and it may lead to loss of polymer concentration upon the filtration of polymeric micelle. In contrast, PMSNs showed no precipitation or changes in size using the same drug feeding procedures, as well as reasonable loading capacity and long-term colloidal stability (Figure 3c and 3d). The stability of block copolymer micelle systems could be affected by the hydrophobicity and amount of the loaded drug molecule, because these characteristics may alter the hydrophilic-lipophilic balance during the block copolymer assembly process. However, the polymer-gatekeeper strategy allows non-covalent modification of the MSN surface with the hydrophilic polymer after the 16


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drug loading procedure. Drug-loaded MSNs can be well-dispersed in aqueous solutions for short time periods, without formation of large flocculated aggregates, by vigorous stirring or sonication. The surface modification of drug-loaded MSNs conferred enhanced colloidal stability regardless of the hydrophobicity of encapsulated drug molecules, because hydrophobic drug molecules were physically adsorbed inside the mesoporous structure.

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10

5

0 10

100

1000

10000

Diameter (nm)

Figure 3. Images of varying feeding concentration (1 wt%, 5 wt%, and 10 wt%) of the CPTloaded block copolymer (A) and CPT-PMSNs (C). Precipitation/aggregation of the CPTloaded block copolymer at high feeding concentrations. Hydrodynamic particle size analysis for CPT-micelles (B) and CPT-PMSNs (D).

Table 2. DLC and EE for CPT loading in micelles and PMSNs

* Drug loading capacity = mass of CPT loaded in micelles/mass of micelles (final yield, after filtration); DLC, drug loading capacity; EE, entrapment efficiency. 17


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Versatility for intracellular hydrophobic drug delivery A significant merit of independent drug encapsulation in the rigid mesoporous core and installation of the polymer-gatekeeper is that any hydrophobic drug molecules can be entrapped inside the mesoporous silica core with high loading capacity inside the pore (< 2–3 nm). To investigate the versatility of drug encapsulation in MSNs, we investigated the loading capacity of 5 different drug molecules, Dox, PTX, CPT, TMX, and Cur, which are currently used in the clinic. Very high drug loading capacities of 106 wt%, 108 wt%, 108 wt%, 108 wt%, and 104 wt% were observed for Dox, PTX, CPT, Cur, and TMX, respectively, when the ratio of the drug and MCM-41 was around 2:1 by weight (Table 3 and Figure 4). It should be noted that loading capacity was not significantly affected by the type of hydrophobic drug molecule incorporated into the MSNs. To investigate the encapsulation stability of MSNs for hydrophobic drugs in the presence and absence of a polymer shell crosslinked by adding a deficient amount (19 to 50 mol% against the PDS group in the PEG-PDS polymer) of DTT over 24 h, we examined the release profile of different drugs loaded into MSNs and PMSNs in PBS. The PMSNs with the crosslinked shell (30% crosslinking density) showed no leakage of drug molecules for all 5 drugs at the high loading percentages of 64% for Dox, 65 wt% for PTX, 66 wt% for CPT, 58 wt% for TMX, and 66 wt% for Cur (Figure 5), while the non-crosslinked and non-polymerwrapped MSNs showed a burst release profile pattern (Figure S4). The PMSNs showed good colloidal and encapsulation stability under normal physiological conditions, indicating that the PMSNs provided stable encapsulation of large amounts of hydrophobic drug molecules without premature leakage. The disulfide functionalities of the crosslinked shell can be cleaved via thioldisulfide exchange reactions in the presence of GSH, a small peptide that may be abundant in cancer cells. To evaluate the effect of GSH on the PMSNs, we added 5 mM GSH after a 18


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period of 4 h to the PMSN, after which a burst release pattern was noted, indicating that GSH reduced disulfide crosslinking on the particle surface. These results confirm that drug molecules can be released by external stimuli inside target cancer cells, including the presence of GSH, which can cleave disulfide links and trigger drug release. Table 3. DLC, DLE and EE for Dox-, PTX-, CPT-, Cur-, and TMX-loaded PMSNs Drug conc (Feeding)

Dox DLC wt%, (DLE wt%, EE %)

PTX DLC wt% (DLE wt%, EE %)

CPT DLC wt% (DLE wt%, EE %)

Cur DLC wt% (DLE wt%, EE %)

TMX DLC wt% (DLE wt%, EE %)

0.5 mg/mL

6% (6%, 60%)

6% (6%, 60%)

4 % (4%, 40%)

4% (4%, 40%)

6% (6%, 60%)

1.5 mg/mL

17% (15%, 57%)

13% (11%, 43%)

15% (13%, 50%)

15% (13%, 51%)

16% (14%, 53%)

2.5 mg/mL

30% (23%, 60%)

26% (21%, 52%)

28% (22%, 55%)

26% (21%, 53%)

30% (23%, 60%)

5.0 mg/mL

64% (39%, 64%)

65% (39%, 65%)

66% (40%, 66%)

66% (40%, 66%)

58% (37%, 58%)

7.5 mg/mL

80% (44%, 53%)

80% (44%, 53%)

80% (44%, 53%)

86% (46%, 57%)

84% (46%, 56%)

10.0 mg/mL

106% (51%, 53%)

108% (52%, 54%)

108% (52%, 54%)

108% (52%, 54%)

104% (51%, 52%)

Drug loading capacity (DLC) = mass of drug loaded in nanoparticles / mass of nanoparticles Drug loading efficiency (DLE) = mass of drug loaded in nanoparticles / mass of drug loaded nanoparticles

A

Doxorubicin (Dox)

Paclitaxel (PTX)

Camptothecin (CPT)

B

Curcumin (Cur)

Tamoxifen (TMX)

C TMX

Dox

CPT

Cur

PTX

TMX

Dox

CPT CPT

Cur Cur

PTX PTX

Figure 4. A) Investigated hydrophobic drugs. B,C) Images of hydrophobic drug-loaded PMSNs (at 66 wt% loading capacity; 2.5 mg/mL of the drug-loaded PMSNs dispersed in DI water) under visible light and UV-light. Precipitation and aggregation were not observed after storage for more than 1 month.

19


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B

Drug Release (%)

100 80 60 40 Crosslinked

5mM GSH

20

*

*

**

0 Ctrl 0.001 0.01 0.10 0.25 0.50 1.00 2.00

Concentration of Dox (µg/mL)

D

80

*

60 40 5mM GSH

Crosslinked

20

0 2

80

* **

40

** ***

20

***

0

4 6 8 10 12 14 16 18 20 22 24

Ctrl 0.001 0.01 0.1 0.25 0.5

F

Cell Viability (%)

*

60 40 Crosslinked

0 0 2

4 6

2

CPT-PMSN Free-CPT

100

80

5mM GSH

1

Concentration of PTX (µg/mL)

100

20

*

60

Time (h)

Drug Release (%)

PTX-PMSN Free-PTX

100

Cell Viability (%)

Drug Release (%)

100

80

*

*

60

**

40

** **

20

***

0

8 10 12 14 16 18 20 22 24

Ctrl 0.001 0.01 0.10 0.25 0.50 1.00 2.00

Concentration of CPT (µg/mL)

Time (h)

H

100

TMX-PMSN Free-TMX

100

*

80 60 40 5 mM GSH Crosslinked

20

Cell Viability (%)

Drug Release (%)

**

20

0

*

80

*

60

**

40 ***

20 0

0

Ctrl 0.1 0.25 0.5

0 2 4 6 8 10 12 14 16 18 20 22 24

Time (h)

J

100

1

2

5

10

Concentration (µg/mL) Cur-PMSN Free-Cur

100

80

*

60 40

5 mM GSH

20

Crosslinked

0 0

2 4 6

8 10 12 14 16 18 20 22 24

Time (h)

Cell Viability (%)

I

*

40

Time (h)

G

*

60

0 2 4 6 8 10 12 14 16 18 20 22 24

E

*

80

0

C

Dox-PMSN Free-Dox

100

Cell Viability (%)

A

Drug Release (%)

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

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80

* * **

60

**

40

***

20

***

0 Ctrl 0.1 0.25 0.5

1

2

5

10

Concentration (µg/mL)

Figure 5. Drug release patterns of Dox (A), PTX (C), CPT (E), TMX (G), and Cur (I) from 20


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PMSNs with and without addition of 5 mM GSH after 4 h. In-vitro cell viability analysis after 48 h incubation with Dox- (B), PTX- (D), CPT- (F), TMX- (H), or Cur-loaded (J) PMSNs or free drugs in KB cells using alamar blue staining. * P < 0.05, ** P < 0.01, *** P < 0.001 compared to control, analyzed using Students t-test.

To assess the biocompatibility of the developed nanocarriers, cytotoxicity was evaluated using alamar blue cell viability assays in KB cells (human nasopharyngeal carcinoma cells) incubated with different concentrations of MSNs and PMSNs (0.1–2 mg/mL) for 24 h. KB cells are capable of nanoparticle uptake and are thus suitable for testing cell viability using pristine MSNs and PMSNs. No significant cell death was noted for PMSNs at concentrations lower than 1 mg/mL, suggesting that PMSNs are a biocompatible nanocarrier (Figure S5). The neutral outer PEG layers of the PMSNs confer protein resistance, prolong circulation time, and improve biocompatibility.14 The viability of KB cells exposed to free drugs and drug-loaded PMSNs was evaluated for a period of 48 h using the alamar blue assay. The cytotoxicity of the drug-loaded PMSNs was little lower to that of free drugs at equal concentrations (Figure 6b, 6d, 6f, 6hm, and 6j), as evidenced by the diffusion of encapsulated drugs through nanopores of silica particles at slow rate upon the addition of GSH. These results clearly demonstrate that the drug-loaded PMSNs and the respective free drugs have equivalent capacities to kill cancer cells.

Versatile co-delivery of several hydrophobic drugs by PMSNs Despite remarkable progress in chemotherapy, substantial low antitumor efficacy and poor inhibition of drug-resistant cancer cells are urgent and unsolved problems.15 Utilization of single agents in treatment often fails to completely cure patients due to the development of drug-resistant cancer cells.15c In the effort to overcome such issues and enhance the therapeutic efficacy of chemotherapeutic drugs, co-delivery of multifunctional agents to 21


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achieve synergistic effects has received considerable recent interest.16 Antitumor drugs have diverse inhibitory mechanisms. Dox can generate free radicals and intercalate into DNA;17 CPT can deregulate/degrade DNA topoisomerase I;17b PTX can act as a microtubule inhibitor in cell plasma, which affects cell fate during mitosis by disrupting spindle assembly, chromosome aggregation, and cell division;18 Cur suppresses NF-κB activation by inhibiting IκKB activity;19 and TMX binds inhibits cell growth via estrogen receptor binding.20 In recent years, co-delivery of multifunctional agents to cancer cells by various nanosized carriers has become a research focus;21 however, such delivery may require chemical modifications and case-specific optimization for different drugs, which may impair encapsulation. Our approach can be applied easily to combination treatments due to the simplicity of hydrophobic encapsulation of a wide range of drug molecules. As a proof of concept, we prepared PMSNs with 2 combinations of 3 drugs: (i) Dox-CPT-PTX and (ii) Dox-CPT-Cur. We were able to load Dox, PTX, and CPT in PMSNs at loading capacities of 24 wt%, 23 wt%, and 25 wt% (72 wt% total), respectively, when the ratio of MSN to each drug was 1:2. Similarly, the loading capacities of CPT, Cur, and Dox were 21 wt%, 21 wt%, and 19 wt% (61 wt% total), respectively. In comparison with other types of carriers, the drug loading capacities of the PMSNs were exceptionally high. The anticancer efficacy of each drug combination was assessed by measuring cellular uptake and cell viability in KB cells. As shown in Figure 6a and 6d, co-localized fluorescence associated with Dox, CPT, and Cur was observed inside the cells after 4 h of incubation, indicating that all of these drug molecules were localized in a single nanoparticle. The cell viability analysis revealed significant inhibition of cell growth within 12 h of incubation, which was significantly different from the effects of each drug alone (Figure 6b and 6e). The difference in cytotoxicity of the single drugs and the drug combinations can be explained by the multiple actions of the drugs in the combinations, allowing the treatment to overcome cellular 22


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resistance.22 To further confirm the synergistic effects of the combinations, we determined the combination index (CI) using the Chou-Talay principle, in which CI is calculated from the dose-effect profiles of a given drug combination after 24 h of incubation, and CI values lower than, equal to, and higher than 1 represent synergism, additivity, and antagonism, respectively.23 Plotting of the CI values against the drug effect levels (ICx) provides information regarding therapeutic efficiency.15c CI vs. ICx plots were determined for both tested drug combinations in KB cells (Figure 6c and 6f), and the results for both combinations indicated synergistic effects. These results show that PMSNs containing drug combinations can increase intracellular concentrations of drug inside cancer cells and thus enhance their cytotoxic effects. The enhanced cytotoxic effects of PMSNs containing drug combinations indicate that they can be utilized to solve issues involving drug resistance, encapsulation of different drugs in a single carrier, short term treatment techniques, and encapsulation stability. Similarly we have also compared the cytotoxicity of dual drug combination using the same procedure (Table S2 and Figure S7). The loading capacities of the dual drugs were in the ratio of 1:1 for the three different combinations in PMSN. The cytotoxicity analysis in KB cells revealed that, dual drug combinations also show a significant cytotoxic response than that of single ones. This combination drug loading can be useful for designing specific and drug resistant cancer treatment.15c

23


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A

CPT

BF

B

100

Cell Viability (%)

Dox PTX

80

CPT

60 Combination

40 20 0 Ctrl 0.001 0.01 0.1 0.25 0.5

Merged

C

2

Combination Index (CI)

Dox

Additive = 1

1

Synergism < 1

0

Cur

E

20

30

40

50

60

70

Inhibitory Concentrationx 100

Cell Viability (%)

CPT

2

Antagonism > 1

10

D

1


Cur Dox

80

CPT

60 Combination

40 20 0 Ctrl 0.001 0.01 0.1 0.25 0.5

1

2


Dox

Merged

F

Combination Index (CI)

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

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2

Antagonism > 1

Additive = 1

1

Synergism < 1

0 10

20

30

40

50

60

70

Inhibitory Concentrationx

Figure 6. A) Confocal microscopy images of Dox-PTX-CPT combination PMSNs showing fluorescence of Dox and CPT inside KB cells after 4 h of incubation (BF: bright field image). B) In vitro cell viability analysis for Dox-PTX-CPT combination PMSNs, Dox, CPT, and PTX in KB cells. C) Combination index of Dox-PTX-CPT combination PMSNs in KB cells. D) Confocal microscopy images of Dox-Cur-CPT combination PMSNs in KB cells after 4 h of incubation. E) In vitro cell viability analysis for Dox, Cur, CPT, and Dox-Cur-CPT combination PMSNs in KB cells. F) Combination index of Dox-Cur-CPT combination PMSNs in KB cells. Scale bar: 20 µm. Note: Combination index analysis based on ChouTalay principle: CI < 1, synergism; CI = 1, additive effect; and CI > 1, antagonism. 24


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Target-ligand decoration and cellular uptake investigation. The surfaces of PMSNs can be decorated with ligands to allow targeting of cancer cells. Disulfide crosslinking of PDS functionalities provides enhanced non-covalent encapsulation stability to overcome the leaky nature of the carrier, and the remaining PDS functionalities in the polymer shell can be used to conjugate thiol containing molecules to improve cellular uptake efficiency. Thus, PDS functionalities provide an opportunity to tailor the surface with ligands that facilitate selective and rapid internalization. Cysteine containing ligands can be easily decorated onto the surface of PMSNs via disulfide exchange reactions between PDS units of the polymer shell and thiol groups of the ligand molecules under mild conditions and without any further chemical modification of PMSNs surface. Therefore, we used 2 different ligands that contain cysteine: cRGDfC and SP94 (Figure 7a). The RGD (arginine-glycine-aspartic acid) peptide is an effective targeting agent that can be recognized by αvβ3 integrin receptors, which are overexpressed in several human cancers.23 SP94 is a novel peptide that binds specifically to hepatocellular carcinoma cells.24 SP94-PMSNs and RGD-PMSNs were prepared by simple mixing of PMSN with peptide ligands in aqueous suspensions, and ligand attachment was assessed by measuring the appearance of pyridothione, a byproduct of disulfide exchange reactions between thiol groups of peptide ligands and PDS functional groups on PMSNs (Figure S4). Motivated by the ligand-mediated cellular uptake, we examined the viability of KB cells after treatment with CPT-loaded RGD-PMSNs. RGD-PMSNs produced significantly greater inhibition of cell viability than that of PMSNs (Figure 7b). To further confirm ligand-mediated internalization, we measured the fluorescence signal of CPT from RGD-PMSNs and PMSNs inside KB cells using a confocal microscope after 60 min of treatment. As shown in Figure 7c, the RGDPMSNs showed higher fluorescence intensity than the PMSNs. As shown in the flow cytometry (FACS) histogram, the ligand decorated and non-ligand-decorated carriers showed 25


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significant differences in uptake (Figure 7d). Endocytosis is an important mechanism of cellular uptake for several drug carriers.25 Clathrin-mediated endocytosis, caveolin-mediated endocytosis, and macropinocytosis are the 3 major endocytic pathways.26 The localization of RGD-PMSN in KB cells was assessed with endocytosis inhibitors methyl-β-cyclodextrin (MβCD) (inhibiting caveolin-dependent endocytosis), sucrose (inhibiting clathrin-mediated endocytosis), and amiloride (inhibiting macropinocytosis).25 The fluorescence signal (blue) of CPT in the cells was observed by confocal imaging (Figure 7e) and FACS (Figure 7f). Fluorescence in the cells pretreated with MβCD and amiloride was decreased up to 17% and 35%, respectively, in comparison with the control cells. However, the fluorescence intensity of sucrose-pretreated cells was more than 90% of that of the control cells. These results suggest that caveolin-mediated endocytosis and macropinocytosis are predominantly involved in cellular uptake of RGDPMSN. To show the versatility of the ligand decoration, we used another cysteine containing peptide ligand, SP94 peptide, which targets hepatocellular carcinoma cells (HepG2). The SP94 peptide-decorated PMSN specifically targeted HePG2 cells and showed increased cytotoxicity (Figure 7g). FACS analysis showed a significant difference in cytotoxicity between the PMSNs and SP94-PMSNs (Figure 7i). These results suggest that liganddecorated PMSNs can be rapidly internalized inside cancer cells and subsequently release drug molecules that kill the cancer cells, and this strategy could be used to build targeted drug delivery systems using simple thiol containing ligands.27

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B

A

120

RGD-PMSN PMSN

Cell Viability (%)

100

** ** ***

60

***

40

***

20 0 Control 0.01

RGD-PMSN

BF

Merged

PMSN

BF

Merged

MβCD

BF

Merged

0.1

0.25

0.5

1

2

Concentration of CPT (µg/mL)

D

% of Max

C

*

80

Cyclo (RGDf C) or SP94 (AcCGGSFSIIHTPILP-L)

Fluorescence

F

Relative Uptake Efficiency(%)

E

100

AMI

Merged

BF

80 60 40 20 0 Ctrl

MβCD

AMI

SUC

Endocytosis Inhibitors SUC

G

Merged

BF

SP94-PMSN PMSN

100

Cell Viability (%)

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

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H

*

80

** **

60

** ***

40 20 0 Control0.001 0.01

0.1

0.25

0.5

1

2


Figure 7. A) Schematic presentation of PMSN ligand decoration. B) In vitro cell viability analysis for CPT-loaded PMSNs and RGD-PMSNs (washed and replaced with fresh medium after 1 h of incubation) at 48 h in KB cells. C) Cellular uptake studies using confocal 27


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microscope imaging of RGD-PMSNs and PMSNs in KB cells after 1 h of incubation. D) Flow cytometry analysis for PMSNs and RGD-PMSNs after 3 h of incubation in KB cells (red, PMSNs; blue, RGD-PMSNs). E) Confocal microscopy images of RGD-PMSNs in the presence of MβCD (methyl-β-cyclodextrin), SUC (sucrose), or AMI (amiloride). F) Flow cytometry analysis to assess cellular uptake in the presence of endocytosis inhibitors. G) In vitro cell viability analysis for SP94-PMSNs and PMSNs in SP94-positive HepG2 cells (cells were washed and replaced with fresh medium after 1 h of incubation) at 48 h of incubation. H) Flow cytometry analysis for SP94-PMSNs and PMSNs after 3 h incubation with HepG2 cells. * P < 0.05, ** P < 0.01, *** P < 0.001 compared to control cells, analyzed using Student’s t-test.

CONCLUSIONS In summary, the polymer gate keeper strategy was successfully adopted to develop targeted drug delivery carriers for hydrophobic anticancer drugs. The higher drug loading capacity obtained in PMSNs was due to the stable polymer wrapping. Also, the enhanced colloidal, encapsulation stability of PMSN and controlled release of drugs proved the utility of the design. Further, the versatility of the hydrophobic drug carrier platform was demonstrated with a series of drugs such as Dox, PTX, CPT, TMX and Cur. All drug molecules were simply encapsulated in PMSNs with high loading capacity and without loosing its colloidal stability. Cancer targeting ligands (SR94 and cyclic RGD) were integrated into the surface of PMSNs to direct into the specific site and enhance the therapeutic effect. In vitro investigations with multidrug encapsulated nanocarriers showed enhanced cytotoxicity, underscore the ability to hold and deliver the hydrophobic anticancer drugs in great control. Taken together, the design strategy outlined here could have broad implications in a variety of bioengineering area including therapeutics, diagnostics, and nanotheanostics.

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Supporting Information Available: Experimental detail for synthesis of PEG-PDS copolymer and block copolymer preparation. Results of crosslinking density, drug release profile for different drug loaded nanoparticles, in-vitro cytotoxicity analysis of nanoparticles in KB cells. This material is available free of charge via the Internet at http://pubs.acs.org.

Corresponding Authors [email protected] (J.-H.R.)

Authors Contributions The manuscript was written through contributions of all authors.

Conflict of Interest: The authors declare no competing financial interest.

Acknowledgment. This work was supported by the Basic Science Research Program (2013R1A1A2061694, 2014R1A1A1002642, 2011-35B-C00024) through the National Research Foundation of Korea. We thank Prof. S. Kang for kind providing KB and HepG2 cell lines, and thank W. M. Shin for helping drug release experiments.

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Zeng, Y. F.; Li, P. Z.; Li, J. H.; Liu, J. J.; Zhao, Y. L. Intracellular redox-activated anticancer drug delivery by functionalized hollow mesoporous silica nanoreservoirs with tumor specificity. Biomaterials 2014, 35, 7951-7962. (e) Wang, Y. J.; Yan, Y.; Cui, J. W.; Hosta-Rigau, L.; Heath, J. K.; Nice, E. C.; Caruso, F. Encapsulation of Water-Insoluble Drugs in Polymer Capsules Prepared Using Mesoporous Silica Templates for Intracellular Drug Delivery. Adv. Mater. 2010, 22, 4293-4297. (f) Chen, Y.; Gao, Y.; Chen, H.; Zeng, D.; Li, Y.; Zheng, Y.; Li, F.; Ji, X.; Wang, X.; Chen, F.; He, Q.; Zhang, L.; Shi, J. Engineering Inorganic Nanoemulsions/Nanoliposomes by Fluoride-Silica Chemistry for Efficient Delivery/Co-Delivery of Hydrophobic Agents. Adv. Funct. Mater. 2012, 22, 1586-1597. 3. (a) Yang, X. L.; Luo, Y. L.; Xu, F.; Chen, Y. S. Thermosensitive mPEG-b-PA-g-PNIPAM comb block copolymer micelles: effect of hydrophilic chain length and camptothecin release behavior. Pharm. Res. 2014, 31, 291-304. (b) Oh, K. T.; Bronich, T. K.; Kabanov, A. V. Micellar formulations for drug delivery based on mixtures of hydrophobic and hydrophilic Pluronic block copolymers. J. Control. Release 2004, 94, 411-422. (c) Kawano, K.; Watanabe, M.; Yamamoto, T.; Yokoyama, M.; Opanasopit, P.; Okano, T.; Maitani, Y. Enhanced antitumor effect of camptothecin loaded in long-circulating polymeric micelles. J. Control. Release 2006, 112, 329-332. 4. McRae Page, S.; Martorella, M.; Parelkar, S.; Kosif, I.; Emrick, T. Disulfide Cross-Linked Phosphorylcholine Micelles for Triggered Release of Camptothecin. Mol. Pharmaceut. 2013, 10, 2684-2692. 5. Yen, H.-C.; Cabral, H.; Mi, P.; Toh, K.; Matsumoto, Y.; Liu, X.; Koori, H.; Kim, A.; Miyazaki, K.; Miura, Y.; Nishiyama, N.; Kataoka, K. Light-Induced Cytosolic Activation of Reduction-Sensitive Camptothecin-Loaded Polymeric Micelles for Spatiotemporally Controlled in Vivo Chemotherapy. ACS Nano 2014, 8, 11591-11602. 6. (a) Page, S. M.; Martorella, M.; Parelkar, S.; Kosif, I.; Emrick, T. Disulfide Cross-Linked Phosphorylcholine Micelles for Triggered Release of Camptothecin. Mol. Pharmaceut. 2013, 10, 2684-2692. (b) Lee, S. Y.; Tyler, J. Y.; Kim, S.; Park, K.; Cheng, J. X. FRET Imaging Reveals Different Cellular Entry Routes of Self-Assembled and Disulfide Bonded Polymeric Micelles. Mol. Pharmaceut. 2013, 10, 3497-3506. 7. (a) Lu, J.; Liong, M.; Zink, J. I.; Tamanoi, F. Mesoporous Silica Nanoparticles as a Delivery System for Hydrophobic Anticancer Drugs. Small 2007, 3, 1341-1346. (b) Argyo, C.; Weiss, V.; Bräuchle, C.; Bein, T. Multifunctional Mesoporous Silica Nanoparticles as a 30


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