A prominent anchoring effect on the kinetic control of drug release from mesoporous silica nanoparticles (MSNs)

Accepted Manuscript A Prominent Anchoring Effect on the Kinetic Control of Drug Release from Mesoporous Silica Nanoparticles (MSNs) Anh-Vy Tran, Sang-Wha Lee PII: DOI: Reference:

S0021-9797(17)31103-7 https://doi.org/10.1016/j.jcis.2017.09.072 YJCIS 22824

To appear in:

Journal of Colloid and Interface Science

Received Date: Revised Date: Accepted Date:

26 June 2017 15 September 2017 20 September 2017

Please cite this article as: A-V. Tran, S-W. Lee, A Prominent Anchoring Effect on the Kinetic Control of Drug Release from Mesoporous Silica Nanoparticles (MSNs), Journal of Colloid and Interface Science (2017), doi: https://doi.org/10.1016/j.jcis.2017.09.072

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A Prominent Anchoring Effect on the Kinetic Control of Drug Release from Mesoporous Silica Nanoparticles (MSNs)

Anh-Vy Tran, Sang-Wha Lee*

Department of Chemical and Biochemical Engineering, Gachon University, San 65, Bokjeong-Dong, Sujeong-Gu, Seongnam City, 461-701, South Korea

Abstract This work demonstrated kinetically controlled release of model drugs (ibuprofen, FITC) from welltailored mesoporous silica nanoparticles (MSNs) depending on the surface charges and molecular sizes of the drugs. The molecular interactions between entrapped drugs and the pore walls of MSNs controlled the release of the drugs through the pore channels of MSNs. Also, polydopamine (PDA) layer-coated MSNs (MSNs@PDA) was quite effective to retard the release of large FITC, in contrast to a slight retardation effect on relatively small Ibuprofen. Of all things, FITC (Fluorescein isothiocyanate)-labeled APTMS (3aminopropyltrimethoxysilane) (APTMS-FITC conjugates) grafted onto the MSNs generate a pinch-effect on the pore channel (so-called a prominent anchoring effect), which was highly effective in trapping (or blocking) drug molecules at the pore mouth of the MSNs. The anchored APTMS-FITC conjugates provided not only tortuous pathways to the diffusing molecules, but also sustained release of the ibuprofen over a long period of time (~7 days). The fast release kinetics was predicted by an exponential equation based on Fick’s law, while the slow release kinetics was predicted by Higuchi model.

Keywords: Anchoring effect, FITC-APTMS conjugates, Polydopamine, Release kinetics, Controlled release, MSNs

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1. Introduction Over the past few decades, the nanovehicles used for therapeutic drug delivery mostly include liposomes [1-3], micelles [4-9], dendrimers [10-13], nanoparticles [14-18], nanofibers [19-21], and inorganic materials [22-25]. Such nanoscale carriers are intended to deliver therapeutic agents at a cellular level, such as in the cell membrane, cytoplasm, and nucleus via a facile endocytosis [26-29]. However, the most nanovehicles have some shortcomings because of inefficient cellular uptake, endosomal escape, and premature release that frequently occurs before they reach to the target area [30-33]. Recently, nanostructured inorganic materials have opened up a new and exciting pathway for an effective design of controlled drug delivery system. The encapsulation of anticancer drugs within well-tailored nanovehicles can enhance the effectiveness of conventional chemotherapy by reducing side effects. Among the various nanomaterials currently under investigation, mesoporous silica nanoparticles (MSNs) have attracted significant interests as a prospective delivery vehicle of therapeutic drugs because of cost-effective fabrication, non-toxic nature, large surface area, as well as plenty of silanol groups. The textural characteristics of MSNs can increase the loading of anti-cancer drugs in pore tunnels. Furthermore, the silanol-containing surface is easily modified with specific functional ligands which can enhance pharmacokinetics of the therapeutic agents [28, 34-37]. In particular, the pore tunnels that encapsulate the drug cargo can be sealed with various gatekeepers [32, 3841], which offers an opportunity to design controlled drug delivery system with stimuli-responsive drug release. The self-gated strategy of mesoporous nanocarrier shows prospective option for self-controlled delivery which does not require complex capping agents and also possesses switchable drug release behavior [42]. The outer surface of MSNs can be functionalized with switchable gatekeepers that are sensitive to certain external stimuli, such as pH-sensitive dynamic bond, light, magnetic fields, temperature, enabling on-command delivery of therapeutic agents [43-46]. However, it is not easy to control the release of various drugs (different sizes and surface charges) from the MSNs with a uniform pore size. 2

Multifunctional MSNs have highlighted the potential of nanoscale carriers for controlled release of therapeutic agents at desired sites [47-52]. The main factors that influence drug release from MSNs are considered to be its pore size relative to drug molecules and chemical functionalization of the pore walls [53]. The pore diameter of mesoporous materials generally determines the drug release rate, e.g., a reduction of the pore size delays the release of a drug [54, 55]. Also, the pore surface can interact with drug molecules and consequently prolong their retention in the highly porous structure [56-64]. In order to ensure optimal drug retention and controlled release, it is highly desirable to have optimally designed MSNs which can adjust the binding strength of drug molecules and the pore size of MSNs. MSNs can be decorated with molecular or polymer moieties on the external surface to make them more controllable in the delivery process [65, 66]. In particular, mussel-inspired polydopamine (PDA) can be used as a universal intermediate to anchor functional molecules onto various drug delivery substrates [67-70] Furthermore, pH-sensitive PDA can block drugs in MSNs at neutral pH and release drugs at lower pH [71, 72]. In this work, we prepared the MSNs with the modified inner pore walls and exterior surface in order to improve drug loading, storage, and controlled release. We systematically investigated the characteristics and release kinetics of our MSNs-drug systems in terms of surface charge and molecular size of the model drugs (ibuprofen and FITC). Surprisingly, FITC-labeled APTMS (APTMS-FITC) conjugates produced a pinch-effect on the pore channel (so-called a pronounced anchoring effect), which was highly effective in trapping (or blocking) the drug molecules at the pore mouth of MSNs with an PDA coating layer, consequently leading to the highly sustained release of drug molecules. Appropriate diffusion models were applied to analyze the release kinetics of various MSNs-drug systems. The chemical strategy used in this work is very useful in preparing mesoporous drug vehicles with the prominent “anchoring effect” which can kinetically control the release of various drug molecules with different sizes and surface charges.

2. Experimental Methods 3

2.1. Chemical Materials Tetraethyl orthosilicate (TEOS, 99%), cetyltrimethylammonium bromide (CTAB, 99%), ammonium fluoride (NH4F, 99.99%), dopamine hydrochloride (DA, powder), 3-aminopropyltrimethoxysilane (APTMS, 97%), ibuprofen (Ibu, 98%), and fluorescein isothiocyanate (FITC, 90%) were purchased from Sigma-Aldrich (Seoul, South Korea). Phosphate-buffered saline (PBS) was purchased from Bioneer (South Korea). Ethanol and water (HPLC grade) were used as received without further purification. All other chemicals were of the highest commercially available quality and used as received. Glassware was cleaned by an acidic solution of HNO3:HCl (3:1 v/v) and rinsed with deionized (DI) water several times.

2.2. Synthesis of mesoporous silica nanoparticles (MSNs) Mesoporous silica nanoparticles (MSNs) were synthesized by cationic surfactant of cetyltrimethyl ammonium bromide (CTAB) via a sol-gel reaction with tetraethyl orthosilicate (TEOS) [71, 73, 74]. CTAB

(0.15 g) and NH4F (0.4 g) were dissolved in 100 mL of HPLC water and heated to 75 °C under vigorous stirring. When the solution color turned into clear, aliquots (1.52.5 mL) of TEOS were added drop wise and mixed to achieve a milky white solution at 75 °C for 8 h. Then, the collected white solid products were centrifuged (10,000 rpm, 15 min) to separate the precipitates from the sol-gel solution, washed several times with DI water and ethanol, and dried at 60 °C in a vacuum oven for 10 h. To remove the surfactant template (CTAB), the white precipitates were dispersed in ethanol (150 mL) containing hydrochloric acid (2.5 mL, 37%) and refluxed at 75 °C for 24 h. This procedure was repeated twice to ensure the complete removal of the surfactant. The final products were centrifuged, washed several times with DI water and dried in a vacuum oven at 60 °C for 8 h.

2.3. Surface modification of MSNs

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A titer of FITC (2.5 mg) was dissolved in absolute ethanol (3 mL), and an aliquot of APTMS (100 μL) was mixed with FITC solution. The mixed solution kept in the dark with vigorous stirring for 6 h to yield the APTMS-FITC complex. For surface modification of MSNs, CTAB (0.15 g) and NH4F (0.4 g) were continuously dissolved in 100 mL of water with stirring at 70 °C. After the color of mixed solution turned into white, TEOS (2.5 mL) was added drop wise and the APTMS-FITC solution was gradually added into the sol-gel solution. The procedure was performed in the dark with constant stirring at 70 °C. After 24 h, the yellow solution was centrifuged and washed with water and ethanol several times. The solid product was dispersed in water and CTAB was removed by the addition of hydrochloric acid in an ethanolic solvent. The yellow solid products, i.e., MSNs grafted with APTMS-FITC conjugates (MSNs@APTMSFITC), were centrifuged, dried in vacuum oven at 60 °C for 6 h, and stored in the dark before further characterization.

2.4. Drug loading The dried MSNs (30 mg) were added into ethanol (10 mL) containing ibuprofen (Ibu), and the particles were dispersed by sonication. The mixed solution was stirred under constant stirring at 800 rpm for additional 24 h to load the maximum amounts of Ibu. Subsequently, the Ibu-loaded MSNs was centrifuged and washed with water several times in order to remove the Ibu that was attached on the surface of MSNs. Ibu-loaded MSNs (MSNs-Ibu) were dried in a vacuum oven at 40 °C. The same loading procedure was conducted to load Ibu into the MSNs@APTMS-FITC, but was performed in the dark at room temperature. FITC-loaded MSNs (MSNs-FITC) were prepared using ethanol solvent in the dark. MSNs (20 mg) in ethanol (5 mL) were dispersed by sonication. FITC (15 mg) in acetone (5 mL) was added in the MSNs solution for 24 h. The resulting MSNs-FITC were collected by centrifugation and dried in a vacuum oven at 40 °C.

2.5. Polydopamine (PDA) coating 5

Ibu-loaded MSNs (MSNs-Ibu) was coated with a thin PDA layer via the oxidative self-polymerization of dopamine at a neutral pH. 30 mg of MSNs-Ibu was dispersed in tris(hydroxymethyl) aminomethaneHCl buffer (Tris-HCl) (10 mL, pH 8.5) and an aliquot of dopamine (10 mg) was added afterwards. The resulting black solution was stirred at room temperature for 3 h in the dark. The PDA-coated MSNs-Ibu (MSNs-Ibu@PDA) were collected by centrifugation and washed several times with water to remove unpolymerized dopamine. The separated product was dried in a vacuum oven at 60 °C for 4 h and stored in the dark until it was subjected to a further characterization. The PDA-coated MSNs-FITC (MSNsFITC@PDA) was prepared by the same procedure.

2.6. Drug release test Weakly acidic ibuprofen molecules contain the polar carboxyl group (COOH), but the presence of non-polar alkyl groups and benzene ring significantly reduces the solubility of the ibuprofen in aqueous solution [75]. Thus, ibuprofen (Ibu) is slightly soluble in water about 0.11 mg/ml at 25 oC. However, the solubility of Ibu is significantly increased to 3 and 60 mg/ml in phosphate buffer (pH7.4) and ethanol, respectively [76, 77]. The solubility of FITC is 1 mg/ml in PBS (pH 7.4) and 20 mg/ml in ethanol [78, 79]. Drug-loaded MSNs (MSNs-Ibu, MSNs-FITC) was added into 15 ml of PBS under vigorously stirring at 37±0.5 °C that is a physiological body temperature. To estimate the release amounts of the drugs, standard curves were obtained over specific concentration ranges, which followed the linear behavior of Beer-Lambert law. The detailed procedures are described in the supporting information (including Fig. S1 and Fig. S2). During an in-vitro release test, an aliquot was periodically sampled from the solution containing drug-loaded MSNs in order to monitor the release kinetics by UV-vis spectroscopy (NanoDrop; NanoDrop Technologies, Wilmington, Delaware, USA). The in-vitro release tests of other samples (MSNs-Ibu@PDA, MSNs-FITC@PDA, MSNs-Ibu@APTMS-FITC@PDA, and MSNs-FITC@APTMSFITC@PDA) were also carried out under otherwise identical conditions.

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2.7. Evaluation of drug loss by the surface modification process Even though the ibuprofen is completely released from the MSNs in PBS solution, the released amounts should be lower than the initial loading of ibuprofen, due probably to the loss of the drug during the surface modification process (such as washing, coating, and drying steps). In order to confirm any loss of drug during the modification process, we conducted additional release tests of ibuprofen in pure ethanol, using MSNs-Ibu, MSNs-Ibu@PDA, MSNs-Ibu@APTMS-FITC, and MSNs-Ibu@APTMS-FITC@PDA. Ibuprofen loading into the MSNs was carried out in ethanol solvent according to the section 2.4. The absorbance changes of the solution before and after ibuprofen loading were shown in Fig. S3(b and c), and finally converted into loading amounts of ibuprofen based on the standard curve calibrated at 264 nm. The loading amounts of ibuprofen can be calculated by subtracting the residual amounts in the solution (after loading experiment) from the initial amounts of ibuprofen in the solvent (before loading experiment). After then, the release test of the samples was carried out in ethanol solvent by replacing the solvent two times during the test, in order to release the drug completely from the MSNs. It is plausible to assume the complete release of ibuprofen from the MSNs because the solubility of ibuprofen in ethanol is 60 mg/ml that is much larger than that of PBS solution.

a. In case of MSNs-Ibu and MSNs-Ibu@PDA (based on 40 mg of MSNs), the release test of the samples was carried out in 5 ml of ethanol for 3 days. After then, the samples were separated by centrifugation, followed by the successive release of the sample in fresh ethanol (5 ml) for 7 days. b. In case of MSNs-Ibu@APTMS-FITC and MSNs-Ibu@APTMS-FITC@PDA (based on 40 mg of MSNs@APTMS-FITC), the release test of the samples was carried out in 35 ml of ethanol for 3 days. After then, they were separated by centrifugation, followed by the successive release of the sample in fresh ethanol (20 ml) for 7 days.

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The loading amounts of ibuprofen and release amounts of ibuprofen in ethanol (including release amounts of ibuprofen in PBS) are summarized in Table S1. The loss of ibuprofen during the modification process is estimated about 1020% of the initial loading amounts. As a result, the final release fraction of ibuprofen in PBS solution almost reached to the saturation value based on the loading amounts subtracted by the drug loss during the modification process.

3. Results and discussion 3.1 Physicochemical properties of as-synthesized MSNs As previously described in the experimental section, MSNs were synthesized using a sol-gel method (Stöber method) by dropping TEOS into the mixture of CTAB and NH4F, followed by the removal of CTAB under reflux with the mixture of ethanol and HCl [71, 80, 81]. The removal of CTAB surfactants can prevent the aggregation of MSNs [82]. The surfactants were dissolved easily in hot ethanol, and the extracted surfactants were simultaneously detached from the pore walls [83]. APTMS reacts with FITC and forms a stable fluorescent conjugate (APTMS-FITC complex), i.e., FITC is covalently linked to APTMS via a thiourea linkage. The fluorescent conjugate is finally grafted onto the pore walls of MSNs. According to Fig. S4, the synthetic procedures of MSNs followed two different pathways of (a) and (b) depending on the absence and presence of APTMS-FITC conjugates, respectively. Model drugs (ibuprofen or FITC) are loaded into the MSNs and MSNs@APTMS-FITC and subsequently coated with polydopamine (PDA) layer. Finally, in-vitro release test of using the drug-loaded MSNs was carried out in PBS. The particle size distribution was analyzed by dynamic light scattering (DLS) method. The average particle size of the MSNs was 94  13 nm, which increased to 97  25 nm after grafting with APTMSFITC conjugates, and further increased to 99  9 nm after coating with the PDA layer (Fig. S5). The surface charge of as-synthesized particles was measured by a zeta potential meter. The zeta-potential of the 8

MSNs was -19.06 mV, which increased to -3.25 mV after grafting with APTMS-FITC conjugates, and further increased to +28.02 mV after coating with the PDA layer (see Table S2). The N2 adsorption/desorption isotherms and pore size distribution by Barrett-Joyner-Halenda (BJH) analysis are shown in Fig. 1. The adsorption/desorption isotherms are close to type II isotherms [84]. At the initial part of the adsorption isotherm, the adsorption capacity increases sharply, which is a clear indication of the abundant nanopores [85]. After the pore surface of MSNs was grafted with APTMS-FITC conjugates, the BET surface area of MSNs@APTMS-FITC increased five-fold, i.e., from 293 m2/g to 1156 m2/g. The pore volumes also increased from 0.70 cm3/g to 1.67 cm3/g, but the average pore size decreased from 10.3 nm to 5.6 nm due to the presence of APTMS-FITC conjugates. The APTMS-FITC conjugates are expected to provide a fluorescence and confer a positive surface charge on MSNs which can facilitate the attachment and internalization of negatively-charged drugs [86]. Fig. S6 shows the nitrogen adsorption/desorption isotherms and average pore size of MSNs before the removal of CTAB. After CTAB was removed, the BET surface area of MSNs increased from 145 m2/g to 293 m2/g. The pore volume of MSNs increased from 0.46 cm3/g to 0.70 cm3/g, and the average pore size of MSNs was decreased from 12.8 nm to 10.3 nm. 1200

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Fig. 1. Nitrogen adsorption/desorption isotherms and pore size distribution by BJH analysis: (a) CTAB-removed MSNs and (b) CTAB-removed MSNs@APTMS-FITC.

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As-synthesized MSNs were characterized by SEM and TEM instruments as shown in Fig. 2. The SEM images of MSNs showed a spherical shape with a uniform size distribution (Fig. 2a). In comparison to the pristine MSNs, the MSNs@APTMS-FITC, i.e., MSNs grafted with APTMS-FITC conjugates, exhibited rougher surface morphology (Fig. 2b). When the MSNs were grafted with APTMS-FITC conjugates and subsequently coated with a PDA layer, the resulting MSNs@APTMS-FITC@PDA exhibited less uniformity in both the spherical shape and particle size (Fig. 2c). High-resolution of TEM images are shown in the insets of Figs. 2a and 2b that clearly show the two-dimensional pore arrangement of MSNs. The insets in Figs. 2c and 2d also show the distinct presence of the PDA coating layer (~3 nm) on the periphery of MSNs.

(a)

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Fig. 2. SEM and TEM (inset) images of (a) MSNs, (b) MSNs@APTMS-FITC, (c) MSNs@PDA, and (d)

Fig. 4 SEM & TEM (insets) images of (a) MSNs, (b) MSNs@APTMS-FITC, (c) MSNs@PDA MSNs@APTMS-FITC@PDA. and (d) MSNs@APTMS-FITC@PDA.

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The small angle x-ray diffraction patterns of MSNs and MSNs@APTMS-FITC exhibited wellresolved peaks which were indexed as (100) reflections associated with hexagonal symmetry, indicating a good order of pores in mesoporous silica nanoparticles [87]. On the other hand, the diffraction peaks such as (110) and (200) reflections were not shown in Fig. 3 [88]. When compared with the MSNs, the MSNs@APTMS-FITC showed the decreased intensity of (100) diffraction peak with a slight shift toward higher 2θ values, due probably to the incorporated organic molecules as disordered/ordered species [89]. After the MSNs was grafted with APTMS-FITC conjugates, an average pore diameter of the MSNs was decreased from 2.39 nm to 1.44 nm, and the thickness of pore wall was increased from 3.26 nm to 3.92 nm (see Table S3).

MS Ns MS Ns @ A P T MS -F IT C

In te n s ity (a .u .)

(100)

1

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The successive surface modification of MSNs was confirmed by the corresponding FT-IR spectra. Fig. 4a shows the FT-IR spectra of MSNs, in which the broad bands at approximately 3685 cm−1 and 995 cm−1 are attributed to O-H stretching vibrations of silanol groups and absorbed water, respectively. Moreover, the peaks at 1064 cm−1 and 797 cm−1 are attributed to the asymmetric and symmetric stretching vibrations of the Si-O-Si bonds, respectively. After grafting with APTMS-FITC conjugates, the resulting MSNs@APTMS-FITC exhibited new bands at 642 cm−1 and 1681 cm−1, which were assigned as C-S and CO vibrations. In addition, the peaks at 2395 cm−1 and 3413 cm−1 are assigned as C-N-C and NH groups of APTMS-FITC conjugates, all of which indicate the successful functionalization of MSNs by APTMSFITC conjugates (Fig. 4b). Upon further coating with the polydopamine (PDA) layer, new peaks appeared at approximately 3673 cm−1 corresponding to the OH and NH bond vibrations in PDA. New doublet peaks also appeared at approximately 2950 cm−1, which were attributed to the C-H asymmetric vibration of CH2 units in PDA. The peaks at 1212, 1370, and 1543 cm−1 also confirmed the presence of C-C, C-N, and C=C groups of the benzene ring (Fig. 4c). According to the magnified inset of Fig. 4b, a small peak at 2950 cm−1 attributed to CH2 unit of APTMS was overlapped with relatively strong N-H band of MSNs@APTMS-FITC. In practice, a very small amount of APTMS was used to modify the surface of MSNs so that they did not completely cover the surface of MSNs. On the other hand, polydopamine coating showed the full coverage over the outer surface of MSNs as shown in Fig. 2d. As a result, the FTIR spectrum of MSNs@APTMS-FITC@PDA displayed more distinct CH2 peak with strong intensity, whereas the CH2 peak of APTMS was barely observed in the MSNs@APTMS-FITC sample.

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O -H S i-O -S i

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through three Si-O bonding sites which consequently reduce the number of OH groups on the surface. It is plausible to assume that the anchored conjugates cover the pore surface and influences the transport of drug molecules. Firstly, drug molecules confront many obstacles in MSNs@APTMS-FITC@PDA due to the brush-like anchored conjugates which render them to follow a convoluted pathway during the diffusion process, resulting in a prolonged release of drug molecules.

high density of functional groups at

the pore mouth covered with PDA coating layer can lead to a highly effective pore closure to the releasing molecules. To be conclusive, the APTMS-FITC conjugates can possibly trap (or block) drug molecules at 13

the pore outlet with the assistance of PDA layer. In addition,

conjugates can provide a

fluorescent functionality to the MSNs. According to Fig. S7, an emission peak of photoluminescence (PL) was observed at 513 nm from the MSNs@APTMS-FITC, but the MSNs did not show any emission peak at that wavelength. Thus, the PL spectra at 513 nm confirmed the successful grafting of FITC onto the pore surface.

Fig. 5. Schematic snapshots illustrating the MSNs grafted with APTMS-FITC conjugates, followed by the PDA coating layer and its release mechanism of the drug.

3.2 Kinetic study of in-vitro drug release As described in the experimental section and supporting information (Figs. S1, S2, S3), the releasing amounts of the drug were measured periodically to monitor the release kinetics of MSNs-drug systems. According to Table S1, our MSNs-drug systems showed the almost saturation of drug release after a longperiod of time. Furthermore, the PBS solution provided the sufficient solubility to the released drugs. Thus, it is plausible to investigate the release kinetics based on the cumulative release fraction. Firstly, the release kinetics of MSNs-Ibu and MSNs-Ibu@PDA were measured in PBS solution at pH5 and pH7.4, respectively. At both pH conditions, there was an initial burst of ibuprofen (Ibu) release within 14

5 h, followed by a slower release of Ibu that almost leveled off after a long-period of time (as seen from Fig. S8). In order words, Ibu-loaded MSNs exhibited an initial burst of Ibu followed by a slow release at the longer period [73, 90, 91]. The MSNs-Ibu@PDA released ~0.111 mg of ibuprofen per 1.0 g of MSNs at 37 oC. The release amount of MSNs-Ibu@PDA was slightly lower than that of MSNs-Ibu. The initial burst of release was mostly attributed to the rapid dissolution of ibuprofen located near the surface of MSNs. The prolonged release stems from the entrapped ibuprofen inside the porous channel. The release kinetics of ibuprofen from the MSNs was analyzed by an exponential equation based on Noyes-Whitney equation and applying the Fick’s law, so-called Fickian exponential model (or Fickian model): (1) Where Qt and Q∞ are the amounts of the drug in the porous matrix of spherical shape at time t and equilibrium time t=∞, respectively; kF is a diffusion rate constant independent of the concentration, but contains information about the solvent accessibility to the substrate and the diffusion coefficient through mesoporous channels [92, 93]. According to Figs. 6a and 6c, the release kinetics of MSNs-Ibu and MSNs-Ibu@PDA shows an approximate compliance to the Fickian model. As summarized in Table S4, the Fickian model fit gave kF values of 0.29, 0.26, and 0.27 corresponding to MSNs-Ibu (at pH7.4), MSNs-Ibu@PDA (at pH5), and MSNs-Ibu@PDA (at pH7.4), respectively. The diffusion rate constant, kF, of MSNs-Ibu is higher than those of MSNs-Ibu@PDA at pH5.0 and pH7.4. These differences indicated that the diffusion rate of ibuprofen was somewhat slowed down by the PDA coating layer. Furthermore, the slight difference in the release rates between pH5 and pH7.4 was attributed to the differential changes of hydrogen-bonding interactions depending on the pH values. The negative charge of COOH groups in ibuprofen is increased at higher pH7.4, which leads to the decrease of hydrogen-bonding interactions between ibuprofen and the

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pore walls of MSNs. As a result, ibuprofen tends to release more easily from the pore channels of MSNs at higher pH7.4.

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Fig. 6. (a), (c) Model fits of ibuprofen release from MSNs-Ibu and MSNs-Ibu@PDA at pH5.0 and pH7.4 in PBS at 37 oC by the Fickian exponential model versus square root time; (b), (d) Model fits of ibuprofen release from MSNsIbu and MSNs-Ibu@PDA at pH5.0 and pH7.4 in PBS at 37 oC by the K-P model versus cumulative times.

The release profile was also analyzed by a power law equation based on the Korsmeyer-Peppas (K-P) model which was usually utilized to find out the mechanism of drug release: Qt

Q

 kRt n

(2)

Where Qt/Q∞ is a released fraction of drug at time t, kR is the relaxation rate constant, and n is the release exponent. In the K-P model, the value of n characterizes the release mechanism [92, 93]. In the case of 16

spherical matrix, n = 0.43 corresponds to Fickian diffusion, 0.43 < n < 0.85 to non-Fickian transport, and n = 0.85 to Case II (relaxation transport) [94, 95]. According to Figs. 6b and 6d, the K-P model was applied to fit the first 60% of the normalized release because the model was not valid at the high fraction of released drug [96]. The curve fit by K-P model with two adjustable parameters produced the following fitted results: i) kR = 36.8, n=0.22 for MSNs-Ibu at pH7.4, ii) kR = 43.6, n=0.12 for MSNs-Ibu@PDA at pH5.0, and iii) kR = 37, n=0.2 for MSNs-Ibu@PDA at pH7.4. All the release exponents were fitted less than 0.43, indicating the deviation of release kinetics from the classical Fickian diffusion behavior. The MSNs-Ibu system exhibited a sort of pseudo-Fickian behavior showing a shorter linearity (less than 60%) in the normalized release fraction over square root of time [97-99]. Table S4 summarized the parameter values for the release kinetics of asprepared samples by the Fickian and K-P models. Secondly, the release kinetics of MSNs-FITC and MSNs- FITC@PDA were investigated as a function of release time. According to Figs. 7a and 7b, MSNs-FITC showed a rapid release of FITC from the MSNs within a day, whereas MSNs-FITC@PDA showed a substantially retarded release and finally reached to the saturation after a long-period of time (~7 days) [52]. In case of MSNs-FITC, the initial burst of FITC approached 80% release within the first 5 h, and then the release fraction of FITC increased up to ~95% after 20 h. The rapid release kinetics of MSNs-FITC was predicted by the exponential equation based on Fick’s law. As shown in Fig. 7c, the release kinetics of MSNs-Ibu complied well with the Fickian model with k=1.02. On the other hand, the release kinetics of MSNs-FITC@PDA exhibited the very slow release of FITC, which was analyzed by the Higuchi and K-P models, respectively. The square-root-time dependence of release kinetics can be described by Higuchi model [93, 100]: Qt  kH t1/2 , kH   2DS ( A  0.5S )

1/2

(3)

Where, Qt is the released amount of drug in time t and kH is the release rate constant which contains information about the diffusivity and solubility of the drug, the porosity, and drug contents in the particles. According to Fig. 7c, the release kinetics of MSNs-FITC@PDA showed the approximate linearity of 17

fractional release over the square-root-time. The Higuchi model provided a good compliance with the release kinetics of MSNs-FITC@PDA with the kH of 19.90. When the K-P model was applied to MSNsFITC@PDA (Fig. 7d), n value was fitted as 0.68, indicating the non-Fickian behavior that represent the combined relaxation kinetics and Fickian diffusion. 0.030

(a)

0.0020

0.020 0.015 0.010 0.005

MSNs-FITC MSNs-FITC@PDA

0.000

0

20

40

60

80

100

120

Released FITC (g)

Absorbance

0.025

(b)

0.0015

0.0010

0.0005

MSNs-FITC MSNs-FITC@PDA

0.0000

140

0

20

40

Time (hour) 100

100

(c)

80

80

100

120

140

(d)

80

Fickian model

Qt/Q

Qt/Q

60

Time (hour)

60

Higuchi model 1/2 Qt/Q=kHt

40 20

60

K-P model n Qt/Q=kRt

40 20

MSNs-FITC MSNs-FITC@PDA

0 0

2

4

6

time

8

10

MSNs-FITC@PDA

0

12

0

20

40

60

80

100

120

140

Time (hour)

Fig. 7. (a) Release profiles of FITC absorbance from MSNs-FITC and MSNs-FITC@PDA in PBS at 37 oC, (b) the correlation between released amounts of FITC and testing times, c) the model fits of released amounts versus square root time by the Fickian and Higuchi models, respectively and (d) the model fits of release amounts versus cumulative times by the K-P model.

The release rate of FITC from the MSNs-FITC was much faster than that from the MSNs-FITC@PDA, clearly indicating that PDA coating layer played the critical role for the highly retardation of FITC release [101]. For instance, inter-associated FITC molecules can increase their size more than the molecular dimension of FITC (~1.4 nm  0.9 nm) [102-104], which therefore required much longer times in their 18

transport to pass through the PDA coating layer. On the other hand, the release process of MSNs-FITC was completed within 24 h due to the absence of the blocking PDA layer. Furthermore, the associated FITC molecules experience a considerable reduction in their functional groups which can interact with the OH groups on the pore surface, consequently leading to the faster release of FITC from the MSNs due to the weakened molecular interactions. In summary, the fast release kinetics of MSNs-FITC was well predicted by the Fickian model, and the slow release kinetics of MSNs-FITC@PDA was well predicted by the Higuchi model. The PDA coating layer was very effective to control the release rate of large FITC molecules, in contrast with a slight retardation effect on the transport of small ibuprofen (~1.0 nm  0.6 nm) [105]. Finally, the release kinetics of MSNs-Ibu@APTMS-FITC and MSNs-Ibu@APTMS-FITC@PDA were analyzed by the three diffusion models, as shown in Fig. 8. According to the kinetic analysis of MSNsIbu@APTMS-FITC, Figs. 8a and 8b show a two-step release of ibuprofen with an initial rapid release followed by an asymptotic approach to the release saturation after a long-period of time. The initial burst within the first 3 h comprised ~85% release of ibuprofen, which gradually increased to ~90% after 20 h, and gradually approached to the saturation after 100 h (Fig. 8c). According to the Fickian model, the kF value of MSNs-Ibu@APTMS-FITC was fitted as 0.73 which was fairly larger than that of MSNs-Ibu (see Table S4). The reason may be that the APTMS-FITC conjugates significantly reduced the number of OH groups on the pore surface, leading to the faster release of the ibuprofen due to the weakened interactions between the pore surface and the drug. In the meantime, the release kinetics of Ibu from MSNsIbu@APTMS-FITC exhibited the gradual approach to the saturation at the longer period of time probably because the anchored conjugates acted as brush-like hurdles for the diffusing molecule to pass through the porous channels of MSNs. The release kinetics of MSNs-Ibu@APTMS-FITC@PDA was also analyzed by the Higuchi and K-P models, respectively. As seen from Fig. 8c, the release kinetics of MSNs-Ibu@APTMS-FITC@PDA exhibited an approximate compliance with the Higuchi model, showing the linear plot of cumulative 19

release over the square-root-time. The release rate constant, kH, of MSNs-Ibu@APTMS-FITC@PDA was fitted as 19.67 (converted into 11.88 mg/g h1/2). The converted value is in the low range of the reported values for various MSNs-ibuprofen systems (kH=10~61 mg/g h1/2) [90]. The anchored FITC-APTMS conjugates provided not only tortuous pathways to the diffusing molecules, but also sustained release of ibuprofen from the PDA-coated MSNs. The release kinetics of MSNs-Ibu@APTMS-FITC@PDA was also analyzed by the K-P model as shown in Fig. 8d. The release exponent, n, of MSNs-Ibu@APTMSFITC@PDA was fitted as 0.62, indicative of non-Fickian diffusion behavior. 0.08

(a)

Absorbance

0.5 0.4 0.3 0.2 0.1

MSNs-Ibu@APTMS-FITC MSNs-Ibu@APTMS-FITC@PDA

0.0 0

20

40

60

80

100

120

Released Ibuprofen (g)

0.6

0.07

(b)

0.06 0.05 0.04 0.03 0.02 0.01

MSNs-Ibu@APTMS-FITC MSNs-Ibu@APTMS-FITC@PDA

0.00 0

140

20

40

100

100

(c)

80

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(d)

80

Fickian model

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Huguchi model 1/2 Qt=kHt

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Qt/Q

80

Qt/Q

60

Time (hour)

Time (hours)

60

K-P model n Qt/Q= kRt

40 20

20 MSNs-Ibu@APTMS-FITC MSNs-Ibu@APTMS-FITC@PDA

0 0

2

4

6

8

10

MSNs-Ibu@APTMS-FITC@PDA

0 0

12

time

20

40

60

80

100

120

140

Time (hour)

Fig. 8. (a) Release profiles of ibuprofen absorbance from MSNs-Ibu@APTMS-FITC and MSNs-Ibu@APTMSFITC@PDA in PBS at 37 oC, (b) the correlation between ibuprofen amounts and release times, (c) the model fits of release amounts versus square root time by the Fickian model and Higuchi model, respectively, and (d) the model fits of release amounts versus cumulative times by the K-P model.

To be summarized, our MSNs-drug systems show the various release kinetics depending on three main factors: i) molecular interactions (via surface charges), ii) size of drug molecules (via PDA coating), 20

and iii) anchoring effects of APTMS-FITC conjugates (via pore closure effect). The mechanistic schemes corresponding to three main factors are illustrated in Fig. 9. As shown in Fig. 9A, the release rate of ibuprofen from the MSNs shows the dependence on the hydrogen-bonding interactions between drug molecules and the pore walls of MSNs. The release rate of MSNs-Ibu was retarded by the molecular interactions between COOH groups of ibuprofen and OH groups of pore walls. On the other hand, MSNs-Ibu@APTMS-FITC exhibited the faster release rate in comparison to that of MSNs-Ibu. The reason may be that the APTMS-FITC conjugates significantly reduced the number of OH groups on the pore surface, resulting in the decreased interactions between the pore surface and the drug. Furthermore, the decrease of solution pH (from pH7.4 to pH5) led to a slight decrease of hydrogen-bonding interactions between ibuprofen and the pore surface so that ibuprofen exhibited a slightly faster release rate at lower pH5. As shown in Fig. 9B, the release kinetics of our MSNs-drug systems is strongly influenced by the molecular size of the drugs (Ibuprofen and FITC). For example, FITC molecules tended to inter-associate to a larger size so that the release of the FITC was significantly impeded and/or blocked by the PDA coating layer with a reduced pore size. As a result, the MSNs-FITC@PDA exhibited much slower release rate which was well predicted by the Higuchi model, whereas the MSNs-FITC without the PDA coating layer exhibited the rapid release kinetics which was predicted by the Fickian model. Furthermore, the MSNs-Ibu@PDA exhibited the slower release rate in comparison to that of MSNs-Ibu, indicative of the retardation effect of the PDA coating layer on the release of ibuprofen from the MSNs-Ibu@PDA. As shown in Fig. 9C, the FITC-APTMS conjugates demonstrated the highly sustained release of ibuprofen from the MSNs-Ibu@APTMS-FITC@PDA, despite that the conjugates significantly reduced the number of OH groups on the pore surface. With the assistance of the PDA coating layer, the APTMS-FITC conjugates acted as a brush-like hurdle which impelled the drug to take a convoluted path during the release process. More importantly, the pore mouth was concentrated with many electrostatic and/or hydrogen-bonding interactions of various functional groups (-COOH, -OH, -NH-CS-NH-, -NH2, and -OH) 21

stemmed from both APTMS-FITC conjugates and PDA coating layer. That is, the molecular interactions are intensively localized at the pore mouth, inducing a pinch-effect on the pore channel (so-called a prominent anchoring effect), which was highly effective in trapping (or blocking) drug molecules at the pore mouth of MSNs. Ibuprofen release

A

Ibuprofen (Ibu)

Hydrogenbonding Effect

PDA

Retarded release of Ibu from MSNs

MSNs

OH

OH

OH

OH

MSNs

O

Molecular interaction between Ibu and –OH

B

O

O O

O O

O

O O

O

MSNs

FITC release

PDA

FITC

Retarded release of FITC from MSNs

Molecular Size Effect

MSNs

OH

OH

OH

OH O

MSNs

Controlled Release

Inter-associated FITC

FITCbig molecular size

Ibuprofen release MSNs@APTMS-FITC FITC

O

O

O O

MSNs

MSNs@PDA

C

MSNs

PDA

Pore Closure Effect

APTMS

Sustained release of Ibu from MSNs

O

Ibuprofensmall molecular size MSNs

OH

OH

MSNs

APTMS-FITC Conjugates

MSNs@APTMS-FITC@PDA

Fig. 9. Schematic representation of three influencing factors on controlled release of model drugs (Ibuprofen, FITC) from our MSNs systems: A: Hydrogen bonding effect on the retarded release of ibuprofen from the MSNs; B: Size effect on the retarded release of FITC from the MSNs@PDA; C: Pore closure effect on the sustained release of ibuprofen from the MSNs@APTMS-FITC@PDA.

Conclusions Based on the previous reports of MSNs for drug delivery [28, 29, 55, 65, 80], we successfully conjugated APTMS-FITC anchors on the pore walls of MSNs with a remarkable increase in surface area and pore volume. Our MSNs demonstrated the kinetically controlled release of the model drugs (ibuprofen 22

and FITC) depending on their surface charges and molecular sizes. With regard to the release kinetics of ibuprofen, the decrease of solution pH (from 7.4 to 5) induced a slightly increased hydrogen-bonding interactions between ibuprofen and the pore walls of MSNs [106, 107], resulting in the slightly prolonged release of the drug from the MSNs. The PDA coating layer was quite effective to retard the FITC molecules from MSNs-FITC@PDA, when compared with the slight retardation effect on the release of relatively small ibuprofen from the MSNs-Ibu@PDA. Most of all, the APTMS-FITC conjugates grafted on the pore walls of MSNs@PDA, i.e., MSNs@APTMS-FITC@PDA, exhibited a prominent anchoring effect (via pore closure effect), which was highly effective in trapping (or blocking) diffusing molecules at the pore mouth of MSNs [53-55, 69]. As a result, the MSNs-Ibu@APTMS-FITC@PDA exhibited the highly sustained release of ibuprofen over a long-period of time (~7 days). The rapid release kinetics was predicted by Fick’s law, and the slow release kinetics was predicted by the Higuchi model. Our MSNsdrug system is an efficient drug vehicle platform because it can accommodate various drugs and confer kinetically controlled release of the drugs with different surface charge and molecular size.

Acknowledgments: This work was supported through the National Research Foundation of Korea (NRF2016R1D1A1B03931486) and the GRRC program of Gyeonggi Province (GRRC Gachon 2014-B04, Development of Nanomaterials for Biomedical Sensing Applications).

References [1] Y. Malam, M. Loizidou, A.M. Seifalian, Liposomes and nanoparticles: nanosized vehicles for drug delivery in cancer, Trends in Pharmacological Sciences 30(11) (2009) 592-599. [2] M. Çağdaş, A.D. Sezer, S. Bucak, Liposomes as Potential Drug Carrier Systems for Drug Delivery, Application of Nanotechnology in Drug Delivery 57028 (2014) 1-51. [3] A.D. Tagalakis, R. Maeshima, C. Yu-Wai-Man, J. Meng, F. Syed, L.P. Wu, A.M. Aldossary, D. McCarthy, S.M. Moghimi, S.L. Hart, Peptide and nucleic acid-directed self-assembly of cationic nanovehicles through giant unilamellar vesicle modification: Targetable nanocomplexes for in vivo nucleic acid delivery, Acta Biomaterialia 51 (2017) 351-362. [4] E. Blanco, C.W. Kessinger, B.D. Sumer, J. Gao, Multifunctional micellar nanomedicine for cancer therapy, Experimental Biology and Medicine 234(2) (2009) 123-131. 23

[5] Z.L. Tyrrell, Y. Shen, M. Radosz, Fabrication of micellar nanoparticles for drug delivery through the self-assembly of block copolymers, Progress in Polymer Science 35(9) (2010) 1128-1143. [6] M.S. Alai, W.J. Lin, S.S. Pingale, Application of polymeric nanoparticles and micelles in insulin oral delivery, Journal of Food and Drug Analysis 23(3) (2015) 351-358. [7] R. Trivedi, U.B. Kompella, Nanomicellar formulations for sustained drug delivery: strategies and underlying principles, Nanomedicine Journal 5(3) (2010) 485-505. [8] A.M. Jhaveri, V.P. Torchilin, Multifunctional polymeric micelles for delivery of drugs and siRNA, Frontiers in Pharmacology 5 (2014) 77-103. [9] Y. Fang, Y. Jiang, Y. Zou, F. Meng, J. Zhang, C. Deng, H. Sun, Z. Zhong, Targeted glioma chemotherapy by cyclic RGD peptide-functionalized reversibly core-crosslinked multifunctional poly(ethylene glycol)-b-poly(ε-caprolactone) micelles, Acta Biomaterialia 50 (2017) 396-406. [10] B.K. Nanjwade, H.M. Bechra, G.K. Derkar, F.V. Manvi, V.K. Nanjwade, Dendrimers: emerging polymers for drug-delivery systems, European Journal of Pharmaceutical Sciences 38(3) (2009) 185-196. [11] K. Madaan, S. Kumar, N. Poonia, V. Lather, D. Pandita, Dendrimers in drug delivery and targeting: Drugdendrimer interactions and toxicity issues, Journal of Pharmacy & Bioallied Sciences 6(3) (2014) 139-150. [12] J.R. Baker, Jr., Dendrimer-based nanoparticles for cancer therapy, Hematology ASH Education Program 2009 (2009) 708-719. [13] N.-C. Cheng, W.-J. Lin, T.-Y. Ling, T.-H. Young, Sustained release of adipose-derived stem cells by thermosensitive chitosan/gelatin hydrogel for therapeutic angiogenesis, Acta Biomaterialia 51 (2017) 258-267. [14] Z. Liu, Y. Jiao, Y. Wang, C. Zhou, Z. Zhang, Polysaccharides-based nanoparticles as drug delivery systems, Advanced Drug Delivery Reviews 60(15) (2008) 1650-1662. [15] S.R. Mudshinge, A.B. Deore, S. Patil, C.M. Bhalgat, Nanoparticles: Emerging carriers for drug delivery, Saudi Pharmaceutical Journal 19(3) (2011) 129-141. [16] W.H. De Jong, P.J.A. Borm, Drug delivery and nanoparticles: Applications and hazards, International Journal of Nanomedicine 3(2) (2008) 133-149. [17] S. Parveen, R. Misra, S.K. Sahoo, Nanoparticles: a boon to drug delivery, therapeutics, diagnostics and imaging, Nanomedicine: Nanotechnology, Biology and Medicine 8(2) (2012) 147-166. [18] T. Sun, Y.S. Zhang, B. Pang, D.C. Hyun, M. Yang, Y. Xia, Engineered Nanoparticles for Drug Delivery in Cancer Therapy, Angewandte Chemie International Edition 53(46) (2014) 12320-12364. [19] D.G. Yu, X. Wang, X.Y. Li, W. Chian, Y. Li, Y.Z. Liao, Electrospun biphasic drug release polyvinylpyrrolidone/ethyl cellulose core/sheath nanofibers, Acta Biomaterialia 9(3) (2013) 5665-5672. [20] D.-G. Yu, J.-J. Li, M. Zhang, G.R. Williams, High-quality Janus nanofibers prepared using three-fluid electrospinning, Chemical Communications 53(33) (2017) 4542-4545. [21] C. Yang, D.-G. Yu, D. Pan, X.-K. Liu, X. Wang, S.W.A. Bligh, G.R. Williams, Electrospun pH-sensitive core–shell polymer nanocomposites fabricated using a tri-axial process, Acta Biomaterialia 35 (2016) 77-86. [22] P. Ghosh, G. Han, M. De, C.K. Kim, V.M. Rotello, Gold nanoparticles in delivery applications, Advanced Drug Delivery Reviews 60(11) (2008) 1307-1315. [23] M. Arruebo, Drug delivery from structured porous inorganic materials, Journal of Nanomedicine and Nanobiotechnology 4(1) (2012) 16-30. [24] C. Wu, Y. Zhu, J. Chang, Y. Zhang, Y. Xiao, Bioactive inorganic-materials/alginate composite microspheres with controllable drug-delivery ability, Journal of Biomedical Materials Research 94(1) (2010) 32-43. [25] R.V. Pinto, F. Antunes, J. Pires, V. Graça, P. Brandão, M.L. Pinto, Vitamin B3 metal-organic frameworks as potential delivery vehicles for therapeutic nitric oxide, Acta Biomaterialia 51 (2017) 66-74. [26] S. Kwon, R.K. Singh, R.A. Perez, E.A. Abou Neel, H.-W. Kim, W. Chrzanowski, Silica-based mesoporous nanoparticles for controlled drug delivery, Journal of Tissue Engineering 4 (2013) 1-18. [27] I.J. Majoros, C.R. Williams, A. Becker, J.R. Baker, Jr., Methotrexate delivery via folate targeted dendrimerbased nanotherapeutic platform, Nanomedicine and Nanobiotechnology 1(5) (2009) 502-510. [28] Slowing, II, J.L. Vivero-Escoto, C.W. Wu, V.S. Lin, Mesoporous silica nanoparticles as controlled release drug delivery and gene transfection carriers, Advanced Drug Delivery Reviews 60(11) (2008) 1278-1288. [29] S. Wang, Ordered mesoporous materials for drug delivery, Microporous and Mesoporous Materials 117(1–2) (2009) 1-9. 24

[30] K.-N. Yang, C.-Q. Zhang, W. Wang, P.C. Wang, J.-P. Zhou, X.-J. Liang, pH-responsive mesoporous silica nanoparticles employed in controlled drug delivery systems for cancer treatment, Cancer Biology & Medicine 11(1) (2014) 34-43. [31] M. Talelli, M. Iman, A.K. Varkouhi, C.J. Rijcken, R.M. Schiffelers, T. Etrych, K. Ulbrich, C.F. van Nostrum, T. Lammers, G. Storm, W.E. Hennink, Core-crosslinked polymeric micelles with controlled release of covalently entrapped doxorubicin, Biomaterials 31(30) (2010) 7797-804. [32] S. Mura, J. Nicolas, P. Couvreur, Stimuli-responsive nanocarriers for drug delivery, Nature Materials Journal 12(11) (2013) 991-1003. [33] T.L. Andresen, S.S. Jensen, K. Jorgensen, Advanced strategies in liposomal cancer therapy: problems and prospects of active and tumor specific drug release, Journal of Progress in Lipid Research 44(1) (2005) 68-97. [34] W.H. Chen, G.F. Luo, Q. Lei, F.Y. Cao, J.X. Fan, W.X. Qiu, H.Z. Jia, S. Hong, F. Fang, X. Zeng, R.X. Zhuo, X.Z. Zhang, Rational design of multifunctional magnetic mesoporous silica nanoparticle for tumor-targeted magnetic resonance imaging and precise therapy, Biomaterials 76 (2016) 87-101. [35] Y. Deng, D. Qi, C. Deng, X. Zhang, D. Zhao, Superparamagnetic High-Magnetization Microspheres with an Fe3O4@SiO2 Core and Perpendicularly Aligned Mesoporous SiO2 Shell for Removal of Microcystins, Journal of the American Chemical Society 130(1) (2008) 28-29. [36] K.C. de Souza, G.F. Andrade, I. Vasconcelos, I.M. de Oliveira Viana, C. Fernandes, E.M. de Sousa, Magnetic solid-phase extraction based on mesoporous silica-coated magnetic nanoparticles for analysis of oral antidiabetic drugs in human plasma, Materials Science and Engineering: C 40 (2014) 275-280. [37] Y. Wang, H. Gu, Core–Shell-Type Magnetic Mesoporous Silica Nanocomposites for Bioimaging and Therapeutic Agent Delivery, Advanced Materials 27(3) (2015) 576-585. [38] V. Torchilin, Multifunctional and stimuli-sensitive pharmaceutical nanocarriers, European Journal of Pharmaceutics and Biopharmaceutics 71(3) (2009) 431-44. [39] L. Zhu, V.P. Torchilin, Stimulus-responsive nanopreparations for tumor targeting, Integrative Biology 5(1) (2013) 96-107. [40] F. Danhier, O. Feron, V. Preat, To exploit the tumor microenvironment: Passive and active tumor targeting of nanocarriers for anti-cancer drug delivery, Journal of Controlled Release 148(2) (2010) 135-46. [41] L. Xing, H. Zheng, Y. Cao, S. Che, Coordination polymer coated mesoporous silica nanoparticles for pHresponsive drug release, Advanced Materials 24(48) (2012) 6433-6437. [42] X. Zeng, G. Liu, W. Tao, Y. Ma, X. Zhang, F. He, J. Pan, L. Mei, G. Pan, A Drug-Self-Gated Mesoporous Antitumor Nanoplatform Based on pH-Sensitive Dynamic Covalent Bond, Advanced Functional Materials 27(11) (2017) 1605985-1605994. [43] C. Gimenez, C. de la Torre, M. Gorbe, E. Aznar, F. Sancenon, J.R. Murguia, R. Martinez-Manez, M.D. Marcos, P. Amoros, Gated mesoporous silica nanoparticles for the controlled delivery of drugs in cancer cells, Langmuir 31(12) (2015) 3753-3762. [44] X. Xu, S. Lü, C. Gao, C. Feng, C. Wu, X. Bai, N. Gao, Z. Wang, M. Liu, Self-fluorescent and stimuli-responsive mesoporous silica nanoparticles using a double-role curcumin gatekeeper for drug delivery, Chemical Engineering Journal 300 (2016) 185-192. [45] C. Coll, A. Bernardos, R. Martínez-Máñez, F. Sancenón, Gated Silica Mesoporous Supports for Controlled Release and Signaling Applications, Accounts of Chemical Research 46(2) (2013) 339-349. [46] W. Cheng, J. Nie, L. Xu, C. Liang, Y. Peng, G. Liu, T. Wang, L. Mei, L. Huang, X. Zeng, pH-Sensitive Delivery Vehicle Based on Folic Acid-Conjugated Polydopamine-Modified Mesoporous Silica Nanoparticles for Targeted Cancer Therapy, ACS Applied Materials & Interfaces 9(22) (2017) 18462-18473. [47] S. Giri, B.G. Trewyn, M.P. Stellmaker, V.S.Y. Lin, Stimuli-Responsive Controlled-Release Delivery System Based on Mesoporous Silica Nanorods Capped with Magnetic Nanoparticles, Angewandte Chemie International Edition 44(32) (2005) 5038-5044. [48] C.-Y. Lai, B.G. Trewyn, D.M. Jeftinija, K. Jeftinija, S. Xu, S. Jeftinija, V.S.Y. Lin, A Mesoporous Silica NanosphereBased Carrier System with Chemically Removable CdS Nanoparticle Caps for Stimuli-Responsive Controlled Release of Neurotransmitters and Drug Molecules, Journal of the American Chemical Society 125(15) (2003) 4451-4459. [49] T.D. Nguyen, K.C.F. Leung, M. Liong, C.D. Pentecost, J.F. Stoddart, J.I. Zink, Construction of a pH-Driven Supramolecular Nanovalve, Organic Letters 8(15) (2006) 3363-3366. 25

[50] R. Hernandez, H.-R. Tseng, J.W. Wong, J.F. Stoddart, J.I. Zink, An Operational Supramolecular Nanovalve, Journal of the American Chemical Society 126(11) (2004) 3370-3371. [51] N.K. Mal, M. Fujiwara, Y. Tanaka, Photocontrolled reversible release of guest molecules from coumarinmodified mesoporous silica, Nature 421(6921) (2003) 350-353. [52] H. Guo, H. Qian, S. Sun, D. Sun, H. Yin, X. Cai, Z. Liu, J. Wu, T. Jiang, X. Liu, Hollow mesoporous silica nanoparticles for intracellular delivery of fluorescent dye, Chemistry Central Journal 5(1) (2011) 1-10. [53] M. Bouchoucha, M.-F. Côté, R. C.-Gaudreault, M.-A. Fortin, F. Kleitz, Size-Controlled Functionalized Mesoporous Silica Nanoparticles for Tunable Drug Release and Enhanced Anti-Tumoral Activity, Chemistry of Materials 28(12) (2016) 4243-4258. [54] S.C. Shen, W.K. Ng, L. Chia, J. Hu, R.B.H. Tan, Physical state and dissolution of ibuprofen formulated by cospray drying with mesoporous silica: Effect of pore and particle size, International Journal of Pharmaceutics 410(12) (2011) 188-195. [55] Y. Hu, J. Wang, Z. Zhi, T. Jiang, S. Wang, Facile synthesis of 3D cubic mesoporous silica microspheres with a controllable pore size and their application for improved delivery of a water-insoluble drug, Journal of Colloid and Interface Science 363(1) (2011) 410-417. [56] Q. Gao, Y. Xu, D. Wu, Y. Sun, X. Li, pH-responsive drug release from polymer-coated mesoporous silica spheres, Journal of Physical Chemistry C 113(29) (2009) 12753-12758. [57] S. Kapoor, R. Hegde, A.J. Bhattacharyya, Influence of surface chemistry of mesoporous alumina with wide pore distribution on controlled drug release, Journal of Controlled Release 140(1) (2009) 34-39. [58] F. Qu, G. Zhu, S. Huang, S. Li, S. Qiu, Effective controlled release of captopril by silylation of mesoporous MCM41, ChemPhysChem 7(2) (2006) 400-406. [59] S.W. Song, K. Hidajat, S. Kawi, Functionalized SBA-15 materials as carriers for controlled drug delivery: Influence of surface properties on matrix-drug interactions, Langmuir 21(21) (2005) 9568-9575. [60] Y. Yan, J. Fu, T. Wang, X. Lu, Controlled release of silyl ether camptothecin from thiol-ene click chemistryfunctionalized mesoporous silica nanoparticles, Acta Biomaterialia 51 (2017) 471-478. [61] S. Huh, J.W. Wiench, B.G. Trewyn, S. Song, M. Pruski, V.S. Lin, Tuning of particle morphology and pore properties in mesoporous silicas with multiple organic functional groups, Chemical Communications (18) (2003) 2364-2365. [62] S. Huh, J.W. Wiench, J.-C. Yoo, M. Pruski, V.S.Y. Lin, Organic Functionalization and Morphology Control of Mesoporous Silicas via a Co-Condensation Synthesis Method, Chemistry of Materials 15(22) (2003) 4247-4256. [63] N. Lang, A. Tuel, A Fast and Efficient Ion-Exchange Procedure To Remove Surfactant Molecules from MCM-41 Materials, Chemistry of Materials 16(10) (2004) 1961-1966. [64] E. Aznar, F. Sancenón, M.D. Marcos, R. Martínez-Máñez, P. Stroeve, J. Cano, P. Amorós, Delivery Modulation in Silica Mesoporous Supports via Alkyl Chain Pore Outlet Decoration, Langmuir 28(5) (2012) 2986-2996. [65] Y. Wang, Q. Zhao, N. Han, L. Bai, J. Li, J. Liu, E. Che, L. Hu, Q. Zhang, T. Jiang, S. Wang, Mesoporous silica nanoparticles in drug delivery and biomedical applications, Nanomedicine: Nanotechnology, Biology and Medicine 11(2) (2015) 313-327. [66] T. Li, X. Shen, Y. Geng, Z. Chen, L. Li, S. Li, H. Yang, C. Wu, H. Zeng, Y. Liu, Folate-Functionalized MagneticMesoporous Silica Nanoparticles for Drug/Gene Codelivery To Potentiate the Antitumor Efficacy, ACS Applied Materials & Interfaces 8(22) (2016) 13748-13758. [67] J.B. Wood, M.W. Szyndler, A.R. Halpern, K. Cho, R.M. Corn, Fabrication of DNA Microarrays on PolydopamineModified Gold Thin Films for SPR Imaging Measurements, Langmuir 29(34) (2013) 10868-10873. [68] C.C. Ho, S.J. Ding, Structure, properties and applications of mussel-inspired polydopamine, Journal of Biomedical Nanotechnology 10(10) (2014) 3063-3084. [69] W. Cheng, C. Liang, L. Xu, G. Liu, N. Gao, W. Tao, L. Luo, Y. Zuo, X. Wang, X. Zhang, X. Zeng, L. Mei, TPGSFunctionalized Polydopamine-Modified Mesoporous Silica as Drug Nanocarriers for Enhanced Lung Cancer Chemotherapy against Multidrug Resistance, Small 13(29) (2017) 1-12. [70] J. Park, T.F. Brust, H.J. Lee, S.C. Lee, V.J. Watts, Y. Yeo, Polydopamine-based simple and versatile surface modification of polymeric nano drug carriers, ACS Nano 8(4) (2014) 3347-3356.

26

[71] D. Chang, Y. Gao, L. Wang, G. Liu, Y. Chen, T. Wang, W. Tao, L. Mei, L. Huang, X. Zeng, Polydopamine-based surface modification of mesoporous silica nanoparticles as pH-sensitive drug delivery vehicles for cancer therapy, Journal of Colloid and Interface Science 463 (2016) 279-287. [72] Q. Zheng, T. Lin, H. Wu, L. Guo, P. Ye, Y. Hao, Q. Guo, J. Jiang, F. Fu, G. Chen, Mussel-inspired polydopamine coated mesoporous silica nanoparticles as pH-sensitive nanocarriers for controlled release, International Journal of Pharmaceutics 463(1) (2014) 22-26. [73] P. DeMuth, M. Hurley, C. Wu, S. Galanie, M.R. Zachariah, P. DeShong, Mesoscale porous silica as drug delivery vehicles: Synthesis, characterization, and pH-sensitive release profiles, Microporous and Mesoporous Materials 141(1–3) (2011) 128-134. [74] Q. He, J. Shi, Mesoporous silica nanoparticle based nano drug delivery systems: synthesis, controlled drug release and delivery, pharmacokinetics and biocompatibility, Journal of Materials Chemistry 21(16) (2011) 58455855. [75] D.K. Shravan Kumar Patel, Amol P. Waghmode and Avinash S.Dhabale, Solubility enhancement of ibuprofen using hydrotropic agents, International Journal of Pharmacy & Life Sciences 2(2) (2011) 524-545. [76] S. Wang, Z. Song, J. Wang, Y. Dong, M. Wu, Solubilities of Ibuprofen in Different Pure Solvents, Journal of Chemical & Engineering Data 55(11) (2010) 5283-5285. [77] S.M.V. Noor Ahmed VH, Rafi B, Venkaiah A, Manohar, Jameerullah,Veeranjeneyulu, Naga Sri ND, Soheib Afroz , Zinab Parvez Hussain, Determination of Effect of Polymers on Solubility of Poorly Water Soluble Drug, International Journal of Advanced Pharmaceutics 5(2) (2015) 72-80. [78] F.J. Green, The Sigma-Aldrich handbook of stains, dyes, and indicators, Aldrich Chemical Company 2 (1990) 1776. [79] D. Wu, A. Ensinas, B. Verrier, C. Primard, A. Cuvillier, G. Champier, S. Paul, T. Delair, Zinc-stabilized colloidal polyelectrolyte complexes of chitosan/hyaluronan: a tool for the inhibition of HIV-1 infection, Journal of Materials Chemistry B 4(32) (2016) 5455-5463. [80] A. Liberman, N. Mendez, W.C. Trogler, A.C. Kummel, Synthesis and surface functionalization of silica nanoparticles for nanomedicine, Surface Science Reports 69(2–3) (2014) 132-158. [81] S.-H. Wu, C.-Y. Mou, H.-P. Lin, Synthesis of mesoporous silica nanoparticles, Chemical Society Reviews 42(9) (2013) 3862-3875. [82] V. Cauda, C. Argyo, D.G. Piercey, T. Bein, “Liquid-Phase Calcination” of Colloidal Mesoporous Silica Nanoparticles in High-Boiling Solvents, Journal of the American Chemical Society 133(17) (2011) 6484-6486. [83] Y.-H. Liu, H.-P. Lin, C.-Y. Mou, Direct Method for Surface Silyl Functionalization of Mesoporous Silica, Langmuir 20(8) (2004) 3231-3239. [84] X.W. Zeng, H.B. Chen, Y. Zheng, W. Tao, Y.G. Fan, L.Q. Huang, L. Mei, Enhanced adsorption of puerarin onto a novel hydrophilic and polar modified post-crosslinked resin from aqueous solution, Journal of Colloid and Interface Science 385 (2012) 166-173. [85] J. Huang, R. Deng, K. Huang, Equilibria and kinetics of phenol adsorption on a toluene-modified hyper-crosslinked poly(styrene-co-divinylbenzene) resin, Chemical Engineering Journal 171(3) (2011) 951-957. [86] S.-M. Lai, T.-Y. Tsai, C.-Y. Hsu, J.-L. Tsai, M.-Y. Liao, P.-S. Lai, Bifunctional Silica-Coated Superparamagnetic FePt Nanoparticles for Fluorescence/MR Dual Imaging, Journal of Nanomaterials 2012 (2012) 1-7. [87] B. Chen, Z. Wang, G. Quan, X. Peng, X. Pan, R. Wang, Y. Xu, G. Li, C. Wu, In vitro and in vivo evaluation of ordered mesoporous silica as a novel adsorbent in liquisolid formulation, International Journal of Nanomedicine 7 (2012) 199-209. [88] A. Tadjarodi, F. Zabihi, S. Afshar, Experimental investigation of thermo-physical properties of platelet mesoporous SBA-15 silica particles dispersed in ethylene glycol and water mixture, Ceramics International 39(7) (2013) 7649-7655. [89] I. Ursachi, A. Vasile, A. Ianculescu, E. Vasile, A. Stancu, Ultrasonic-assisted synthesis and magnetic studies of iron oxide/MCM-41 nanocomposite, Materials Chemistry and Physics 130(3) (2011) 1251-1259. [90] P. Horcajada, A. Rámila, J. Pérez-Pariente, M. Vallet-Regí, Influence of pore size of MCM-41 matrices on drug delivery rate, Microporous and Mesoporous Materials 68(1-3) (2004) 105-109. [91] J.B.S. Ng, P. Kamali-Zare, H. Brismar, L. Bergström, Release and Molecular Transport of Cationic and Anionic Fluorescent Molecules in Mesoporous Silica Spheres, Langmuir 24(19) (2008) 11096-11102. 27

[92] I. Izquierdo-Barba, E. Sousa, J.C. Doadrio, A.L. Doadrio, J.P. Pariente, A. Martínez, F. Babonneau, M. Vallet-Regí, Influence of mesoporous structure type on the controlled delivery of drugs: release of ibuprofen from MCM-48, SBA-15 and functionalized SBA-15, Journal of Sol-Gel Science and Technology 50(3) (2009) 421-429. [93] P. Costa, J.M. Sousa Lobo, Modeling and comparison of dissolution profiles, European Journal of Pharmaceutical Sciences 13(2) (2001) 123-133. [94] N.A. Peppas, J.J. Sahlin, A simple equation for the description of solute release. III. Coupling of diffusion and relaxation, International Journal of Pharmaceutics 57(2) (1989) 169-172. [95] J. Siepmann, N.A. Peppas, Modeling of drug release from delivery systems based on hydroxypropyl methylcellulose (HPMC), Advanced Drug Delivery Reviews 48(2-3) (2001) 139-157. [96] J. Siepmann, N.A. Peppas, Higuchi equation: Derivation, applications, use and misuse, International Journal of Pharmaceutics 418(1) (2011) 6-12. [97] N.L. Thomas, A.H. Windle, A deformation model for Case II diffusion, Polymer 21(6) (1980) 613-619. [98] S. Lee, K.S. Knaebel, Effects of mechanical and chemical properties on transport in fluoropolymers. I. Transient sorption, Journal of Applied Polymer Science 64(3) (1997) 455-476. [99] H. Fujita, Diffusion in polymer-diluent systems, Springer Berlin Heidelberg 3 (1961) 1-47. [100] T. Higuchi, Mechanism of sustained-action medication. Theoretical analysis of rate of release of solid drugs dispersed in solid matrices, Journal of Pharmaceutical Sciences 52 (1963) 1145-1149. [101] P. Gribbon, B.C. Heng, T.E. Hardingham, The analysis of intermolecular interactions in concentrated hyaluronan solutions suggest no evidence for chain-chain association, Biochemical Journal 350 (1) (2000) 329-335. [102] W.-L. Liu, S.-H. Lo, B. Singco, C.-C. Yang, H.-Y. Huang, C.-H. Lin, Novel trypsin-FITC@MOF bioreactor efficiently catalyzes protein digestion, Journal of Materials Chemistry B 1(7) (2013) 928-932. [103] A.G. Bajpayee, C.R. Wong, M.G. Bawendi, E.H. Frank, A.J. Grodzinsky, Avidin as a model for charge driven transport into cartilage and drug delivery for treating early stage post-traumatic osteoarthritis, Biomaterials 35(1) (2014) 538-549. [104] C.J. Breen, M. Raverdeau, H.P. Voorheis, Development of a quantitative fluorescence-based ligand-binding assay, International Journal of Scientific Reports 6 (2016) 25769-25777. [105] B.N. Bhadra, I. Ahmed, S. Kim, S.H. Jhung, Adsorptive removal of ibuprofen and diclofenac from water using metal-organic framework-derived porous carbon, Chemical Engineering Journal 314 (2017) 50-58. [106] M. Zhang, J. Liu, Y. Kuang, Q. Li, D.W. Zheng, Q. Song, H. Chen, X. Chen, Y. Xu, C. Li, B. Jiang, Ingenious pHsensitive dextran/mesoporous silica nanoparticles based drug delivery systems for controlled intracellular drug release, International Journal of Biological Macromolecules 98 (2017) 691-700. [107] P. Bilalis, L.-A. Tziveleka, S. Varlas, H. Iatrou, pH-Sensitive nanogates based on poly(l-histidine) for controlled drug release from mesoporous silica nanoparticles, Polymer Chemistry 7(7) (2016) 1475-1485.

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