Microporous and Mesoporous Materials 200 (2014) 46–51

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Microporous and Mesoporous Materials

journal homepage: www.elsevier .com/locate /micromeso

Dual-responsive drug delivery system with real time tunable releasebehavior

http://dx.doi.org/10.1016/j.micromeso.2014.07.0601387-1811/� 2014 Elsevier Inc. All rights reserved.

⇑ Corresponding author at: Shanghai Institute of Ceramics, Chinese Academy ofSciences, Shanghai 200050, PR China. Fax: +86 21 5241 2632.

E-mail address: lifang02013@163.com (F. Li).

Fang Li a,b,⇑, Yingchun Zhu a, Yunli Wang a

a Shanghai Institute of Ceramics, Chinese Academy of Sciences, Shanghai 200050, PR Chinab Graduate School at Shenzhen, Tsinghua University, Shenzhen 518055, PR China

a r t i c l e i n f o a b s t r a c t

Article history:Received 31 December 2013Received in revised form 16 July 2014Accepted 17 July 2014Available online 15 August 2014

Keywords:Drug delivery systemMacromolecular coatingRemote controlElectro- and thermo-responsive

A dual-responsive drug delivery system (DDS) was successfully established by coating functional macro-molecules onto the outlets of mesoporous silica nanospheres. Electro-sensitive and thermally responsiveunits were copolymerized into the macromolecular networks to control the drug release under differentcircumstances. The simulated drug delivery experiments conducted in vitro revealed that the drug releaserate could be regulated continuously and remotely by one or both stimuli sources. The drug release rateunder 0.5 Hz alternating electric field was almost twice as much as that without an alternating electricfield. This combination of electro-responsive function and temperature sensitive characteristics couldlead to a real time tunable drug delivery system with wide application area.

� 2014 Elsevier Inc. All rights reserved.

1. Introduction

The continued development of modern medicine, combinedwith the pursuit of pharmaceuticals and treatments useful tohuman health and well-being, reflect the civilization and advance-ment of a society. However, developing new drugs is a tedious pro-cess, and very few drug candidates can pass through the necessarypre-clinical and clinical evaluations [1]. The primary reason forlack of drug candidates is the discordance between drug deliverykinetics and chronopharmacology. Therefore, various drug deliverysystems (DDS) have been developed in order to tailor drug releasekinetics and improve the delivery efficiencies according to theactual physiological needs [2]. Many delivery systems use smartresponsive units that respond to environmental stimuli (such aspH, temperature, light or magnetic fields) to enhance the deliveryefficiency and accuracy [3–8]. However, relatively inflexible singleresponsiveness is inadequate for the requirements of complexsituations in human bodies.

Nowadays, functionalized nanocarriers that are responsive tomultiple stimuli are being explored as the potential solutions tothis problem [9,10]. In such systems, the cooperation of eachfunctional unit with the others is required to improve the overall

efficacy. Among these functional co-stimuli, thermo-responsive-ness is of particular interests as it can easily harness the tempera-ture difference between the normal and pathological tissues. Theencapsulated drugs can be released conveniently by either hyper-thermic or hypothermic therapies, which have been widely studiedfor the treatment of cancer and other diseases [11–13]. Significantimprovements in the therapeutic efficiencies have been reportedby combining thermo-responsive materials with other sensitiveunits [14,15].

It is well known that human body temperature deviates fromthe typical physiological temperature in the presence of pathogensor pyrogens [16,17]. Therefore, thermo-sensitive units are veryefficient for the treatment of such diseases. The aim of this studywas to further improve the responding efficiencies of the DDS bycombining the electro-sensitive units with thermo-sensitive units,and dual-functionalized mesoporous silica (DFMS) system wasthus established. The thermo-responsive moiety uses N-isopropyl-acrylamide (NIPAAm), which is copolymerized with the electro-responsive moiety 4-nitrophenyl methacrylate (NPMA) to form amacromolecular coating on the mesoporous silica nanoparticles(MSNs) surface. The electro-sensitive unit NPMA, which has a per-manent electric dipole moment of 2.32 D, gives the system tunablerelease behavior because of its rotation and reorientation move-ments when exposed to an external electric field [18]. For in vivoapplications, the thermo-sensitive units would give an immediateresponse at the lesion site, while the external electric field wouldbe applied afterwards to tune the drug release rate according tothe specific requirements.

F. Li et al. / Microporous and Mesoporous Materials 200 (2014) 46–51 47

2. Experimental

2.1. Preparation and surface modification of MSN–MPS microspheres

Mesoporous silica nanospheres (MSNs) were synthesized simi-lar to the method reported in Ref. [19]. In detail, N-cetyltrimethyl-ammonium bromide (CTAB, 1.0 g, 2.7 mmol) was dissolved in asolution of 480 mL distilled water. NaOH (2 M, 3.5 mL) was thenadded, and the result solution temperate was raised to 80 �C. Afterstirring vigorously for 2 h, tetraethyl orthosilicate (TEOS, 5 mL,21.9 mmol) was then introduced dropwise to the above solution.The reaction was taken at 80 �C for 8 h to obtain the template con-taining MSN (t-MSN). The surfactant free MSN was achieved byremoving the templates completely through acidic extraction.

The template containing MSNs (t-MSNs) were isolated andwashed thoroughly for preparation of silane agent modified MSNs.The 3-methacryloxypropyltrimethoxysilane (c-MPS, 0.7 mL) wasadded to a toluene suspension of CTAB-containing MSN (0.8 g).After mechanically stirred for 48 h at 40 �C, the resulted materials(MSN–MPS) were filtered off, washed thoroughly with methanoland dried under room temperature for the next step.

2.2. Synthesis of DFMS drug delivery system

The macromolecules coated MSNs were synthesized via aseeded radical polymerization method. Typically, MSN–MPS wasredispersed in 10 mL toluene in a round bottom flask (25 mL)under nitrogen protect. After bubbling with dried nitrogen gasfor 30 min, the initiator AIBN was added to the suspension. Then,the mixture was stirred for another 30 min named solution I. Atthe same time, a toluene solution of NPMA (0.001 mol) and NIP-AAm monomers (0.001 mol) was bubbled with high purity nitro-gen gas for 30 min named solution II. The functional monomerwas fabricated according to the method described in followingsupporting references [26]. 1HNMR (for NPMA, CDCl3): d = 2.12(s, 3H, CH3), 5.85 and 6.34 (m, 2H, CH2), 7.27 and 8.28 (m, 4H, CH2).

The solution II was added afterwards to solution I with a syr-inge, followed by reacting at 70 �C for 24 h. The polymerizationwas terminated by opening tube to the atmosphere and then dilut-ing with acetone. The solid product was isolated by filtration andwashed with methanol and acetone to remove the unreactedmonomers and physisorbed oligomer. The acidic extractionmethod was used to remove the CTAB surfactants. In detail, theas-synthesized solid products were added to 100 mL ethanol and5 mL concentrated hydrochloric acid. After stirring for 24 h at roomtemperature, DFMS samples were obtained by filtration exten-sively with absolute ethanol and methanol. The dynamic behaviorof the permanent electric dipoles on the surface of mesoporous sil-ica was investigated on the electrochemistry workstation (Solar-tron impedance/gain-phase analyzer 1287/1260). The C is definedas C ¼ 1

ðZ0þiZ00Þxi¼ Z00þZ0 i

xðZ02þZ002Þ; the dielectric loss is described as

tand = tan(90 � h), where h is the phase angle.

2.3. Drug loading and release experiments

For drug release, ibuprofen was used as a model drug. The driedDFMS systems (0.3 g) were dispersed in 20 mL ethanol solution ofibuprofen (1.6 g) for 7 days at room temperature. The drug loadedmicrospheres were filtrated with acidic ethanol solution. The drugstorage capacity was analyzed by TG-DSC analysis (See SupportingInformation Fig. S1), which is about 14.75% (173.01 mg of ibupro-fen per gram of DMDR system). Then, DFMS-IBU was obtained bydried in oven at 40 �C for 48 h. The 0.05 g of ibuprofen loadedDFMS-IBU was conformed into plates (U10 � 0.5 mm) under anappropriate stress (3 MPa) for drug release.

Release experiments were carried out using SBF (simulatedbody fluid, pH = 7.4) at 25 and 42 �C. Aliquot volume was with-drawn from the release mediums at a given time and equal amountof fresh solution was added simultaneously. The drug releaseamount was measured by UV–Vis spectroscopy at about 264 nm.The cumulative fraction of the IBU was calculated by the followingequation:

Cumulative drug release ¼ Mt=M1 ð1Þ

where Mt and M1 corresponds to the cumulative mass of IBUreleased at time t and the total amount of IBU loaded, respectively.

For electric field drug release, the plates were enclosed in a cellwhich is composed of two sheet platinum electrodes which arecoated with insulator to prevent electrochemical reactions in thesolution and placed away from 8 mm with ±8V voltage. Squarewave voltage of variable frequency was applied to the system,and the release behavior was detected by tracking the amount ofibuprofen release into the solution by UV/Vis spectrometry.

2.4. Characterization

Transmission electron microscopy (TEM) was obtained on aJEM-2010 microscope, operating with an accelerating voltage of200 kv. Negative staining was performed using a droplet of 2 wt%phosphotungstic acid. X-ray diffraction (XRD) was recorded on D/max 2550 V diffractometer with CuKa radiation (k = 1.5418Å). 1HNuclear Magnetic Resonance (1HNMR) Spectroscopy was carriedout on varian Mercury vx at 300 MHz. Sample was dissolved indeuterochloroform (CDCl3) for removing the influence of solution.Fourier transform infrared spectroscopy (FTIR) analysis was carriedout using KBr discs in the region 4000–500 cm�1 by a SHIMADZU(IR Prestige-21). pH value was measured on Model PHS-25 pHMeter Instruction (with the precision of 0.1). UV–Vis spectroscopywas performed on U-2300 spectroscopy using a pair of quartz cuv-ettes (12.4 * 12.4 * 45 mm). The TG-DSC analysis was conduct onSTA-449C thermal analyzer at 10 �C/min in air. N2 adsorption–desorption isotherms were obtained on a Micrometrics Tristar3000 pore analyzer at 77 K under continuous adsorption condi-tions. DLS measurements were performed using Malven Instru-ments Zetasizer Nano ZS.

3. Results and discussions

The synthesis of macromolecule coated MSNs employed aseeded radical polymerization method using 3-methacryloxypro-pyltrimethoxysilane (c-MPS) as the linker. The thermo- and elec-tro-sensitive comonomers were copolymerized together to forma rate-modulating macromolecular coating (poly(nitrophenylmethacrylate-co-isopropylacrylamide-co-methacryloxypropyltri-methoxysilane) (PNNMS)) on the MSNs surface. In this manner, adual-functionalized mesoporous silica (DFMS) system wasobtained to control the delivery of payloads through the functionalcoating polymer (Fig. 1a). TEM was used to investigate the surfacemorphology of the synthesized MSN and PNNMS coated MSN. Bycomparison with the images of MSN, most of functionalized MSNswere ca. 130 nm with a thin PNNMS coating of ca. 3 nm (Fig. 1band c), which was in agreement with the negative staining imagefor DMFS drug delivery system (Fig. 1c insert). The thickness ofthe polymer coating was calculated to be 4.89 nm by the TG-DSCresults (See Supporting Information), which was quite similar tothe results obtained by TEM analysis. The thickness of the func-tionalized MSN layer was controlled very thin to improve the load-ing capacity of the carriers [20]. The DFMS nanoparticles exhibitedwell-defined pore channel structures even after applying the poly-mer coating, which benefits the loading of therapeutic agents.

Fig. 1. The synthesized procedure of the polymer film on the outlets of the mesoporous silica nanoparticles (MSNs) (a) and the TEM images of MSNs (b) and DFMS drugdelivery system (c); insert is the negative staining TEM image for DMFS drug delivery system.

Fig. 2. (a) Low angle XRD patterns of MSN, MSN–MPS, DFMS and DFMS-IBU; (b) FITR spectra of MSN, MSN–MPS and DFMS.

48 F. Li et al. / Microporous and Mesoporous Materials 200 (2014) 46–51

XRD patterns of various samples with or without a polymercoating are displayed in Fig. 2a. Typical MCM-41-type Bragg dif-fraction peaks of (100), (110) and (200) faces were observed priorto the organic modification. After the MPS and PNNMS coatings,

the intensity of the characteristic diffraction peaks reducedbecause of the adsorption of the organic layers. In particular, thediffraction capability of the drug loaded DFMS system reducedgreatly, showing the full storage capability of the payload. Fig. 2b

Fig. 3. UV–Vis absorption spectra of the DFMS system under differenttemperatures.

F. Li et al. / Microporous and Mesoporous Materials 200 (2014) 46–51 49

illustrates the corresponding FTIR spectra; the absorption peaks at1090 cm�1 for all samples exhibited the stretching vibration of Si–O–Si bonds. The peaks at 1630 and 3430 cm�1 were assigned to theabsorbed water and hydroxyl groups. The presence of peaks at 797and 1467 cm�1 after the MPS coating were assigned to the vibra-tion of (CH2)3 and Si–C bonds. The absorption peaks at around1700 and 2920 cm�1 were attributed to the vibration of C@O andC–H bonds. Lastly, peaks at 1530 and 1461 cm�1 after the PNNMSmodification revealed the presence of N–H and C–N bonds. Theseobservations indicated that the polymerization took place exclu-sively on the MSN surface. Thus, we confirmed the successful mod-ification of MSNs with functional macromolecules, which was alsoconsistent with the results obtained by N2 adsorption–desorptionisotherms (See Supporting Information Fig. S2).

It is reported that the lower critical solution temperature (Tt) ofthe thermo-responsive polymers can be easily adjusted by inter-penetrating or copolymerizing with hydrophilic or hydrophobiccomponents [21,22]. The difference between the Tt and body tem-perature (Tb) greatly influences the responding sensitivity of thesystem. After incorporation of hydrophobic NPMA, the Tt of thesmart delivery system DFMS were determined by examining theUV–Vis absorption spectra at different temperatures: 25, 30, 34and 42 �C (Fig. 3). We observed that at these temperatures theabsorbance of the samples at 274 nm were 0.75, 1.09, 1.05, and1.22, respectively. An abrupt increase in the absorbance at around30 �C was observed, indicating an acute phase transition corre-sponding to the Tt of the system. Therefore, a successfully tuning

Fig. 4. Drug release profile of IBU loaded DFMS drug delivery system under different el42 �C.

of the DFMS system provides a characteristic Tt much lower thanthe Tb: (37–38 �C). The modification of Tt to values lower than Tb

is required for two reasons: one, for improving the response sensi-tivity; second, for hyperthermic or hypothermic therapies. Often-times for in vivo applications, a sustained drug release is needed,and a slower rate of drug release under hyperthermia is more use-ful. In this case, an external electric field can be applied to releasethe remaining drug that could not be released by the concentrationgradient. On the other hand, a fast drug release under hypothermiais more useful when a large amount of drugs is needed. In this sit-uation, an external electric field can be applied to regulate therelease rate; drug release can be made more rapid simply bychanging the frequency.

The drug release behavior of the DFMS system was investigatedunder alternating electric fields in release mediums at differenttemperatures (Fig. 4). We chose two typical temperatures: 25 �C(below the Tt) and 42 �C (above the Tt) to study the thermal sensi-tivity of the samples. In the 25 �C SBF solution without an alternat-ing electric field, 45.8% drugs were released in 8 h; whereas ca.67.7% drugs were released in the same time interval under analternating electric field (0.5 Hz). A similar drug release behaviorwas observed for the 42 �C SBF solution (Fig. 4b). The drug releaserate from the prepared DFMS system could be further improved byapplying an alternating electric field. For drug release at 42 �C withand without alternating electric field, ca. 57.3% (under 0.5 Hz elec-tric field) and 25.8% (without electric field) of the total loaded IBUreleased from the system in 8 h, respectively. Comparison of drugrelease behavior at 25 �C under alternating electric field, the drugrelease rate at 42 �C under alternating electric field is a littledepressed in the first 8 h. The reason behind this might be the col-lapse of PNIPAAm component after elevating temperature. How-ever, under alternating electric field the drug release rate at 42 �Cis higher than release rate at 25 �C after 8 h, which leading to theslope changes of the drug delivery curves for these four samples(42 �C & 0 Hz < 25 �C & 0 Hz < 25 �C & 0.5 Hz < 42 �C & 0.5 Hz). Thisphenomenon might be owing to the invalidation of the PNIPAm asthe incubation time going, and also the motivated activity at hightemperature.

The release kinetics under different conditions are analyzed byfitting the Peppas–Sahlin equation [23,24]:

Mt=M1 ¼ k1tn þ k2t2n ð2Þ

where Mt/M1 represents the fraction of release; t is the releasetime; k1 and k2 are kinetic constants related to Fickian and non-Fic-kian diffusion cases, respectively; and n is the diffusion exponent. Inthe 25 �C SBF solution without an alternating electric field, the

ectric frequencies in the SBF solution (pH = 7.4) at temperature of (a) 25 �C and (b)

Table 1The release kinetic parameters of DFMS drug delivery system.

Release circumstances k1 k2 n

25 �C & 0 Hz 0.1797 0.0052 0.4225 �C & 0.5 Hz 0.1032 0.0434 0.6342 �C & 0 Hz 0.1037 0.0062 0.3942 �C & 0.5 Hz 0.1648 0.0278 0.42

50 F. Li et al. / Microporous and Mesoporous Materials 200 (2014) 46–51

release kinetic parameter k1 (0.1797) was much higher than k2

(0.0052), which indicates a drug release rate that is controlled bythe concentration gradient (Table 1). The diffusion exponent nwas 0.42 under these situations, indicating a Fickian diffusionbehavior. After applying an alternating electric field, the kineticexponent n was apparently increased to 0.63, resulting in a

Fig. 5. Illustration of the responsive nature of the system: (a and d) change the morpholomedium temperature; (b and c) change the morphology of electro sensitive units from sfrequency (blue layer coating: the shrinkage state of the electro-responsive units in the punits in the polymer network; Blue coil: the electro-responsive units in the polymerinterpretation of the references to colour in this figure legend, the reader is referred to

Fig. 6. (a) The capacitance and (b) tangent of the dielectri

non-Fickian diffusion controlled drug release process. The releasekinetic parameters k1 and k2 were changed to 0.1032 and 0.0434;the increase in k2/k1 ratio revealed an increased mechanically con-trolled drug release behavior caused by the movements of the dipo-lar moieties. In the 42 �C SBF solution without an alternatingelectric field, the release kinetic parameters k1, k2 and n were deter-mined as 0.1037, 0.0062 and 0.39, respectively, indicating that thediffusion caused by the concentration gradient became depressedbecause of the deswelling of PNNMS at higher temperatures. Thekinetic parameter k2 increased to 0.0278 under the stimulus of analternating electric field because of the dipolar reorientation.

It is well understood that the thermo-responsive units pre-serve globule-to-coil transitions when exposed to temperatureslower than the Tt of the system. In the absence of an alternatingelectric field, the electro-sensitive units exhibit shrinkage state as

gy of thermo sensitive units from swelling state to shrinkage state by elevating thehrinkage state to swelling state by applying alternating electric field with differentolymer network; yellow layer coating: the shrinkage state of the thermo-responsivenetwork; yellow coil: the thermo-responsive units in the polymer network). (Forthe web version of this article.)

c loss as the functions of frequency for DFMS system.

F. Li et al. / Microporous and Mesoporous Materials 200 (2014) 46–51 51

well because of the p–p stacking interactions between the phe-nyl rings [25]. At room temperature (25 �C, T < Tt), the tempera-ture sensitive NIPAAm segments in the polymer networkexhibit a swelling state. An elevation in the temperature resultsin a drop in the drug release rate (Fig. 5a); a higher release ratecould be achieved by applying an alternating electric field withdifferent frequencies (Fig. 5c). As a result, the release kineticscould be shifted between Fickian and non-Fickian diffusion, thusexhibiting an improved drug delivery behavior (Table 1 andFig. 4). In addition, the shrinkage of the thermo-responsive unitsat high temperature collapses the polymer coating and leads to aself-driven reduction of dosage, which is helpful for easing hyper-thermia caused by tissue infection or certain drug injection[16,17]. In a release medium with T > Tt (42 �C), the drug releaserate could also be tuned to provide a nearly two-fold increase inthe rate by applying an alternating electric field (Fig. 5b). FromFig. 4b, it was observed that the drug release rate under 0.5 Hzalternating electric field was almost twice that without the field.Therefore, the delivery of the drugs can be easily tuned through acooperative interaction between the two stimuli according to dif-ferent needs, which potentially widens the application area of thesystem.

The dynamic responsive behavior of the dipolar nitrophenylmoieties was further examined on a Solartron impedance/gain-phase analyzer. The electrical relaxation processes displayed inFig. 6 were caused mainly by the electro-sensitive units of theDFMS system under the electric field. The relaxation at a frequencyof approximately 103 was attributed to the reorientation of the sidegroup of nitrophenyl moieties, which was caused by the polariza-tion of the molecular permanent dipole moments [26,27]. Therelaxation observed below a frequency of 102 was attributed tothe supercapacitor effect caused by the mesoporous silica [18].These results indicated that dipolar nitrophenyl moieties mightwork together with thermo-sensitive units to regulate the drugrelease.

4. Conclusions

In conclusion, a remotely control DFMS system responsive toboth temperature and electric field stimuli was successfully pre-pared and implemented in drug release studies. The thin macro-molecular coating was found to act as a rate modulator forregulating the diffusion kinetics of the drugs loaded in the channelsof inorganic nanocarriers. The release quantities could be tunedcontinuously by changing the frequency of the applied externalelectric field under different temperatures. This smart and flexiblemodulation of the drug delivery process provides another optionfor regulating the release kinetics after drug administration andholds great potential in the context of biomedical applications.

Acknowledgments

We thank the support of the National Natural Science of China,Grant Number: 50772125, 50732202 and National High Technol-ogy Research and Development Program of China, Grant Number:2008AA03Z303, and the Chinese Academy of Sciences – China(KGCX2-YW-210-03) and the Chinese Postdoctoral Science Foun-dation, Grant Number: 04180040.

Appendix A. Supplementary data

Supplementary data associated with this article can be found, inthe online version, at http://dx.doi.org/10.1016/j.micromeso.2014.07.060.

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