Research Article www.acsami.org
Near-Infrared-Light-Responsive Magnetic DNA Microgels for Photonand Magneto-Manipulated Cancer Therapy Yitong Wang,† Ling Wang,† Miaomiao Yan,‡ Shuli Dong,† and Jingcheng Hao*,† †
Key Laboratory of Colloid and Interface Chemistry & Key Laboratory of Special Aggregated Materials, Shandong University, Ministry of Education, Jinan 250100, People’s Republic of China ‡ Department of Pharmacy, Binzhou Medical College, Yantai 264003, People’s Republic of China S Supporting Information *
ABSTRACT: Functional DNA molecules have been introduced into polymerbased nanocarrier systems to incorporate chemotherapy drugs for cancer therapy. Here is the first report of dual-responsive microgels composed of a core of Au nanorods and a shell of magnetic ionic liquid and DNA moieties in the crosslinking network simultaneously, as effective drug delivery vectors. TEM images indicated a magnetic polymer shell has an analogous “doughnut” shape which loosely surround the AuNRs core. When irradiated with a near-infrared-light (near-IR) laser, Au nanorods are the motors which convert the light to heat, leading to the release of the encapsulated payloads with high controllability. DNA acts not only as a cross-linker agent, but also as a gatekeeper to regulate the release of drugs. The internalization study and MTT assay confirm that these core−shell DNA microgels are excellent candidates which can enhance the cytotoxicity of cancer cells controlled by near-IR laser and shield the high toxicity of chemotherapeutic agents to improve the killing efficacy of chemotherapeutic agents efficiently in due course. KEYWORDS: magnetic DNA microgel, drug release, near-infrared-light responsibility, cancer therapy, Au nanorod
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INTRODUCTION Current cancer treatments heavily rely on chemotherapeutic measures. However, when put into clinical practices, traditional chemotherapeutic drugs are circumscribed by their poor solubility in aqueous media and instability during blood circulation as well as potential cytotoxicity, which could kill “normal” cells and cause toxicity to the patient.1 In order to meet the complicated requirements originating from clinical practices, a massive effort has been made to develop an intelligent drug carrier for efficiently delivering anticancer drugs. An ideal drug carrier should primarily be able to facilely encapsulate a mass of anticarcinogen with high efficiency and deliver it to the tumor sites in security and without premature drug release during the entire voyage. In addition, the controlled release of chemotherapeutic cargoes in a controlled way can be achieved by flexible design of drug carriers, upon suffering the specific stimuli.2,3 Traditionally, because of the microenvironments associated with a pathological situation, the stimuli which include pH,4 temperature,5 redox reactions,6 and the existence of some specific enzyme7 may particularly differ from their healthy surroundings.8 However, it remains a major challenge in the functionality and efficiency of drug carriers when the microenvironment changes little, especially for the sub-nanomolar concentrations of biochemical signals or biomarkers. Fortunately, the remote controlling stimuli, e.g., ultrasound, magnetic field, or light, which possess the facile accessibility and externally manipulative controllability, even without side effects could address these drawbacks easily, © 2017 American Chemical Society
making them appealing in artificial cartilage tissue and clinical diagnostics, as well as in biomedical fields.9−11 Microgels which possess the excellent characters of hydrogels and nanomaterials12 are polymeric colloidal particles with three-dimensional cross-linked network structures. Microgels are similar to hydrogels, possessing high loading capacity, flexible physicochemical properties, biostability, and excellent biocompatibility.13 Compared to hydrogels, microgels are more sensitive to response to external stimuli and have higher load capacity and more drug release.14−16 Furthermore, the nanoscale microgels establish the functionalized applications in biological field such as a high specific surface area for bioconjugation, extended circulation times in blood, and tunable size for realizing active or passive targeting to the tumor sites.17 The sum of these properties, coupled with the presence of an interior network for the incorporation of payloads, make microgels ideal candidates for drug carriers. Integrating different functional materials into one carrier in which every component can act in a synergetic way offers even greater possibilities for clinical practice.18−21 Therefore, recently, multifarious hybrid microgels with multifunctionality and versatile characteristics have attracted the attention of many research fields ranging from materials science to biomedicine.22,23 Received: April 19, 2017 Accepted: August 2, 2017 Published: August 2, 2017 28185
DOI: 10.1021/acsami.7b05502 ACS Appl. Mater. Interfaces 2017, 9, 28185−28194
Research Article
ACS Applied Materials & Interfaces Magnetic microgels can not only have a fast response to external magnetic fields but also be easy to separate and recycle from the mixtures. Moreover, these magnetic nanocomposites make it possible for magnetic resonance imaging guided endovascular intervention or therapy. Wang et al.24 constructed core−shell magnetic microgels by free-radical mild polymerization initiated by the cascade reaction of glucose oxidase and horseradish peroxidase, which can be used for the colorimetric detection of glucose. Hennink et al.25 prepared temperaturesensitive liposomes-loaded alginate microgels. These microgels can release their payload through mild hyperthermia. Furthermore magnetic resonance imaging can be achieved by depositing holmium ion into these microgels. Most magnetic microgels were prepared by introducing the magnetic nanoparticles. However, the external environment has a great influence on the stability of nanoparticles because they are welldispersed in the solvent by adding surfactants. Furthermore, in order to prepare magnetic nanoparticles with uniform sizes and regular morphologies, pyroreaction is usually required. In this regard, novel designs are needed to explore in depth for further development of magnetic microgels. Polymeric ionic liquids (PILs) can be synthesized by ionic liquids used as monomers, which are described as novel materials possessing concurrent character of ionic liquid and polymers.26−28 The most notable feature of PILs is that the nature of the counteranion has a strong influence on their chemical and physical properties. It is particularly significant under the circumstance of cationic PILs with different counteranions. The investigation on ionic liquids with magnetic counterions was pioneered by Eastoe et al.29,30 This kind of ionic liquid can simply be synthesized by coordinating conventional ionic liquids having halide ions with FeCl3 in methanol. Intriguingly, these magnetic ionic liquids are molecular liquids rather than typical magnetic fluids which are composed of magnetic colloidal particles dispersed in liquid. The nanoparticle-free magnetic ionic liquids are themselves paramagnetic nonvolatility, have good stability and tunable physicochemical properties regulated by external magnetic fields. Several literature works have reported that ionic liquids can form microgels as monomers.31,32 Magnetic microgels can be prepared by introducing the magnetic ionic liquids, which realizes the construction of magnetic microgels without nanoparticles. Among the responsive microgels reported in the literature, seldom has anyone succeeded in directly introducing magnetic ionic liquid moieties into the crosslinking network of microgels. Herein, for the first time, we engineer and synthesize a novel dual-responsive core (Aunanorods)−shell (magnetic ionic liquid and DNA moieties in the cross-linking network simultaneously) microgels as effective drug delivery vectors through an interesting approach. As illustrated in Figure 1, a near-infrared-light-responsive drug delivery platform based on gold nanorods (AuNRs) coated with DNA cross-linked magnetic microgel shells is constructed. We design and prepare smart DNA−polymer microgels for regulating the magnetic targeting drug delivery and controllable release of anticancer drugs. This nanocomposite is built from a photothermal AuNRs core, oligonucleotides cross-links, and polymer shells. DNA serves as a valve in smart response to the external stimulus.33−37 Integrating AuNRs with photothermal effects and duplex DNA with the unique temperature-dependent assembly property, near-IR-light-activated supersensitive drug release system is expected. We previously developed cuboid DNA microgels
Figure 1. (a) Schematic illustration of magnetic core−shell DNA microgels. (b) Component and DNA sequences in the microgel shells. (c) Synthesizing process of [VBIM]FeCl3Br.
drug delivery systems with DNA as cross-linking agents which showed excellent biocompatibility and highly selective therapy efficacy for tumor sites.38 In the present work, DNA not only serves as cross-linking agents for improving the biocompatibility but also takes the role as a valve to control the drugs release. The double-stranded oligonucleotide cross-linked magnetic microgel layers are used to realize magnetic targeting drug delivery by introducing the magnetic ionic liquids as the monomers. AuNRs have attracted considerable scientific interest in the field of biomedical applications and are especially suited for thermo-chemotherapeutic agents due to their tunable absorption maxima in the near-IR region and photothermal property.39 When irradiated by a near-IR laser beam, AuNRs in the magnetic microgels are the motors causing a rapid rise in the temperature. Once the heat dissipates into the surroundings, the rising temperature causes the DNA cross-links to separate. As a result, the microgel shells break down and the payloads release. We also demonstrate that such lightresponsive magnetic microgels can be candidates for a remotely controlled drug delivery platform with excellent biocompatibility and highly therapeutic effect for cancer cells.
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EXPERIMENTAL SECTION
Chemicals. The acrydite−DNA and linker DNA were purchased from Shanghai Sangon Co. (Shanghai, China) and HPLC purified. Hydrogen tetrachloroaurate(III) hydrate, 1-vinylimidazole, and 1bromobutane were purchased from J&K Chemical Co. (Beijing, China). Hexadecyltrimethylammonium bromide (CTAB), sodium oleate (NaOL), ascorbic acid (AA), AgNO3, NaOH, NaCl, and 28186
DOI: 10.1021/acsami.7b05502 ACS Appl. Mater. Interfaces 2017, 9, 28185−28194
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ACS Applied Materials & Interfaces
dissolved in 100 μL of methanol and then added into a 10 mL reprepared Au@SiO2 sample, and the mixture was stirred for 5 h. In order to remove the unreacted MPS, the obtained Au@SiO2@MPS samples were centrifuged and re-dispersed by 1 mL of water. Solution DNA A was prepared as follows: 50 μL of MPS-modified Au@SiO2 sample, 21 μL of 142 μM strand A, 14 μL of 5 mM magnetic ionic liquid solution, 7 μL of 1 wt % I2959 solution, and 70 μL of buffer (100 mM Tris-HCl, 500 mM NaCl, 100 mM MgCl2, pH = 8.0) were added to 245 μL of water. The mixture was tirred at 300 rpm and irradiated by UV light (365 nm, portable UV lamp, ZF-7A) for 5 min. Solution DNA B was prepared as follows: 13 μL of 128 μM strand B, 1 μL of 1 mM magnetic ionic liquid solution, 1 μL of 1 wt % I2959 solution, and 5 μL of buffer were added into 30 μL of water. The mixture was stirred at 300 rpm and irradiated by UV light (365 nm, portable UV lamp, ZF-7A) for 5 min. Finally, 500 μL of DNA A solution and 50 μL of DNA B solution were mixed with 40 μL of 3.3 μM linker DNA and incubated in a water bath at 65 °C for 5 min, followed by controlled cooling to room temperature. The prepared microgels were collected by centrifugation and washed with water. Characterizations. UV spectra of magnetic ionic liquids, AuNRs, Au@SiO2, and DNA microgels solutions were examined by a U-4100 UV−visible spectrometer, using a quartz cell at a wavelength range of 400−1000 nm. When the samples were measured for UV−visible spectra, proper background subtraction was conducted for the samples by water as reference. The morphologies of the samples were studied on a JEOL JEM1400 TEM (acceleration voltage, 120 kV) with a Gatanmultiscan CCD for collecting images. About 10 μL of microgel sample was dropped onto carbon membrane support copper grids and dried naturally. ζ potential was measured with a Zeta PALS potential analyzer instrument (Brookhaven, USA) with 5 mm apart spacing between parallel-plate platinum black electrodes and a rectangular organic glass cell with 10 mm path length. The sinusoidal voltage is 80 V, and the frequency is 3 Hz. For SQUID magnetometry, dried samples of the DNA microgels were placed in sealed polypropylene tubes and mounted inside a plastic straw for measuring in a magnetometer with a superconducting quantum interference device (MPMSXL, Quantum Design, USA) and a reciprocating sample option (RSO). The data were collected at 300 K. Drug Loading and in Vitro Release. Dox solution (1 mL, 40 μg/ mL) in PBS buffer was mixed with dry DNA microgels for 12 h for loading the Dox into the magnetic microgels. After being centrifuged (8000 rpm, 15 min) and washed with water to remove unencapsulated Dox. The drug loading content was calculated from the concentration of Dox in the supernatant which was detected by a UV−vis absorption peak at 480 nm. The release profile of Dox was constructed in different buffers (pH = 5.0 and 7.4) whether irradiated by near-IR laser (808 nm, 2 W/cm2) or not. The supernatant was collected by centrifuging at different time intervals, and the same volume of fresh buffers was put back into the mixture. The concentration of Dox released from the nanocomposite was measured by UV−vis spectrophotometry based on the standard curve of Dox at 480 nm. Cell Culture. HepG2 cells were cultured in Dulbecco’s modified Eagle’s medium (DMEM, Fanbo) supplemented with 10% fetal bovine serum (FBS, Fanbo). Cells were maintained in a 5% CO2 atmosphere at 37 °C. Cell Uptake. Dox has unique optical properties with a wide absorption spectroscopy around 480 nm as excitation energy and a strong fluorescence emission around 580 nm. Therefore, cellular uptake of Dox can be monitored by using a laser confocal microscope (Leica TCS SPE, Germany) through the fluorescence of Dox. Before measurements, cells were incubated with 10 μM Dox-loaded DNA microgels for 4 h; then, the supernate was abandoned, and 1 mL of fresh medium was added. The cells were irradiated with or without near-IR laser, respectively. The fluorescence signal of Dox was detected by using a laser confocal microscope (Leica TCS SPE, Germany).
MgCl2 were purchased from Sinopharm Chemical Reagent Co. Ltd. (Shanghai, China). Tetraethyl orthosilicate and (3-(methacryloxy)propyl)triethoxysilane (MPS) were purchased from Aladdin Chemical Reagent Co. Ltd. (Shanghai, China). Doxorubicin hydrochloride (Dox), 2-hydroxy-4′-(2-hydroxyethoxy)-2-methylpropiophe (I2959), trihydroxy(methyl amino)methane (Tris) and sodium borohydride (NaBH4) were purchased from Sigma-Aldrich. All buffers were prepared with ultrapure water with a resistivity of 18.25 MΩ·cm obtained by a UPH-IV ultrapure water purifier (China). All glassware was cleaned using freshly prepared aqua regia followed by washing with plenty of water. Human hepatoma cell lines (HepG2 cells) were gifted from the Experimental Center of Binzhou Medical College. Preparation of Mesoporous Silica Coated Gold Nanorods (Au@SiO2). Gold nanorods were synthesized according to a literature report with minor modifications.40 For preparing the seed solution, 5 mL of 0.5 mM HAuCl4 was mixed with 5 mL of 0.2 M CTAB solution in a 50 mL conical flask. Then 0.6 mL of freshly prepared 0.01 M NaBH4 ice solution was injected into the HAuCl4−CTAB solution with vigorous stirring at 1200 rpm. Then the solution color changed from saffron yellow to brownish yellow, and the stirring was sustained for 2 min. The seed solution was aged at 25 °C for 2 h. For preparing the growth solution, 7.0 g of CTAB and 1.234 g of NaOL were dissolved in 250 mL of warm water (∼50 °C), then the solution was cooled to 30 °C naturally, and 18 mL of 4 mM AgNO3 solution was added. After the mixture was kept undisturbed at 30 °C for 15 min, 250 mL of 1 mM HAuCl4 solution was added. With stirring (700 rpm) for 90 min, the solution became colorless and 1.5 mL of HCl (37 wt % in water, 12.1 M) was then introduced to adjust the pH. After stirring another 15 min at 400 rpm, 1.25 mL of 0.064 M AA was injected and the solution was vigorously stirred (1200 rpm) for 30 s. Finally, 400 μL of the seed solution was added into the growth solution. The resultant mixture was vigorously stirred (1200 rpm) for 30 s and kept undisturbed at 30 °C for 12 h for nanorod growth. The final solutions were centrifuged at 7,000 rpm for 20 min and re-dispersed into the same amount of water. The AuNR solutions were stored at 4 °C for further use. Silica coating AuNR was implemented using a Stöber method with modifications.41 After initial synthesis and purification, 10 mL of AuNR samples were centrifuged for a second time at 10,000 rpm for 20 min and re-dispersed by 10 mL of 0.4 mM CTAB. The mixture solutions were stirred overnight at 300 rpm to allow the CTAB to equilibrate on the surface of the AuNRs. Then, 50 μL of 0.1 M NaOH was added to adjust pH to be 10.6. After the solutions were stirred at 400 rpm for 30 min, 22 μL of 20% TEOS in methanol was added under gentle stirring at 30 min in intervals; this methanol solution was injected three times. The mixture was reacted for 24 h at 26−28 °C and then purified by centrifugation and washed with methanol and water. Synthesis of the Magnetic Ionic Liquids. The magnetic ionic liquids, 3-n-butyl-1-vinylimidazolium trichloromonobromoferrate ([VBIM]FeCl3Br) were synthesized according to reported procedures.29,42 A 16.44 g amount of bromobutane and 9.4 g of 1vinylimidazole were added into a 150 mL of acetonitrile solution. The reaction mixture was stirred at 70 °C for 48 h under the protection of nitrogen and then cooled to room temperature. Then 3-n-butyl-1vinylimidazolium bromide ([VBIM]Br) was obtained by vacuum rotary evaporation, washed three times with ethyl acetate, and dried at 40 °C for 12 h under vacuum. 1H NMR (300 MHz, DMSO-d6): δ 9.70 (s, 1H), 8.28 (s, 1H), 8.01 (s, 1H), 7.35 (m, 1H), 6.02 (m, 1H), 5.43 (m, 1H), 4.24 (t, 2H), 1.83 (m, 2H), 1.28 (m, 2H), 0.92 (t, 3H). The equivalent of iron trichloroferrate was added to the [VBIM]Br and stirred overnight at room temperature, then dehydrating at reduced pressure at 80 °C overnight to yield a viscous brown liquid. Due to the magneto-active metal complex anions [FeCl3Br]−,29 a series of characteristic absorption peaks can be obtained, as shown in Supporting Information (SI) Figure S2c, demonstrating the successful synthesis of magnetic ionic liquid monomers. Preparation of the Magnetic Core−Shell DNA Microgels. Magnetic core−shell DNA microgels were prepared according to our previous work with some modifications.38 First, 1.3 μL of MPS was 28187
DOI: 10.1021/acsami.7b05502 ACS Appl. Mater. Interfaces 2017, 9, 28185−28194
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ACS Applied Materials & Interfaces Cytotoxicity. MTT assay was used to analyze the cell viability of DNA microgels. First, 2000 HepG2 cells were seeded in 96 well microplates and grew in DMEM medium containing 10% fetal bovine serum (FBS) at 37 °C for 24 h. After incubating the cells with various concentrations of magnetic DNA microgels, Dox-loaded magnetic DNA microgels or Dox alone at 37 °C for 4 h, irradiating with or without near-IR laser (808 nm, 2 W/ cm2), the cells were maintained for another 20 h. To measure cytotoxicity, a MTT solution was incubated with cell suspensions for 4 h, and then the medium was carefully removed. By measuring the absorbance of each well at 570 nm using a multidetection microplate reader (Synergy TMHT, BioTek Instruments Inc., USA), cell viability was calculated as the ratio of absorbance of the experimental well and the cell in the control well. The significance level was set as P < 0.05.
higher than the critical micelle concentration (cmc) of CTAB, no silica is formed on the AuNRs surface. AuNRs are centrifuged for a second time after synthesis, and a suitable amount of CTAB is added into the AuNR solution with sufficient mixing to ensure the CTAB is adequately equilibrated on the surface. The amount of AuNRs, the pH, the amount of TEOS, the reaction temperature, and the time are all strictly controlled. The thickness of the silica shell determined by TEM is ∼40 nm (Figure 2b). The mesoporous structure of the silica shell permits exposure of the AuNR cores to the surroundings, enabling a sensitive response of the longitudinal SPR to refractive index changes caused by modified molecules:44,45 a small red shift of the surface plasmon band (Figure 2f). To integrate photothermal effects and magnetic properties in one drug nanoplatform, a magnetic microgel layer consisting mainly of polymeric ionic liquid was coated onto the surface of silica-coated gold nanorods (Au@SiO2). Magnetic ionic liquids, 3-n-butyl-1-vinylimidazolium trichloromonobromoferrate ([VBIM]FeCl3Br) contain magneto-active metal complex anions which can respond to the external magnetic field. [VBIM]FeCl3Br can simply be produced by coordinating [VBIM]Br with FeCl3 in anhydrous methanol at room temperature. Magnetic [VBIM]FeCl3Br molecules are the main components of the microgel backbone, endowing magnetic units to resulting core−shell DNA microgels. The synthesis of [VBIM]FeCl3Br is demonstrated by a variety of techniques (Figure S2 in the SI). The silicane-coupling agent MPS is used to endow the surface of the silica shell with abundant CC bonds, which is the key to coating the polymer shell. The growth of linear polymeric chains was initiated by the MPS groups on the particle surface through photoinitiated freeradical polymerization with the [VBIM]FeCl3Br monomer and acrydite-modified DNA. As a result, single-stranded A was grafted onto the linear polymers (PS-A), which were connected with the surface of the AuNRs. PS-A serves not only as a building block of the microgel shells but also as a near-infraredresponsive unit for the photothermal conversion. Another acrydite-modified single-strand DNA B has also been grafted onto a linear polymer chain (PS-B). As shown in Figure 1b, the sequences of DNA strands A and B are complementary to the cross-linking oligonucleotide (linker DNA). The linker DNA can hybridize to both DNA strands A and B when mixing the PS-A and the PS-B solutions with the linker DNA in a stoichiometric ratio; consequently, the near-infrared-lightresponsive magnetic DNA microgels can be constructed. Representative TEM images of the Au@SiO2@microgels show that the dark gold nanorod core is respectively coated with a uniform grayish polymer shell. Moreover, the relatively uniform size of the core−shell magnetic DNA microgels is about 300 nm in the dry state (Figure 2c−e), which meets the requirement of entering into the cancer cells. The magnetic polymer shell has an analogous “doughnut” shape which is loosely surrounding the AuNRs core. The component change of the as-made DNA microgels was determined by Fourier transform infrared spectra (FTIR; Figure S3 in the SI). In the spectrum of Au@SiO2, the wide absorption peak at 3433 cm−1 is attributed to silanol and adsorbed water. When modified with the coupling agent MPS, this band becomes weak. In the meanwhile, the absorption peak at 1713 cm−1 is associated with the stretching vibration of carbonyl group CO of the ester group, demonstrating the successful binding on the surfaces of SiO2 by coupling agent MPS.42,46 After growth of polymer shells, the new absorption peaks at 3100−3600, 1647, and 1157
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RESULTS AND DISCUSSION Fabrication and Characterization of Magnetic Core− Shell DNA Microgels. AuNRs with the tunable longitudinal surface plasmon resonance (LSPR) in near-IR range are of great interest for potential use for various light-triggered bioimaging and cancer therapy. However, their inherent structures and properties, e.g., nonporous structure, low load capacity, and restricted elasticity, prevent them from acting as drug carriers. In order to overcome these drawbacks, the surfaces of AuNRs-based nanosystems are often modified with cell-specific groups, silica shells, or polyelectrolyte for effective drug delivery.39 In this work, a seed-growth binary surfactantdirected approach was used for preparing monodisperse AuNRs with an aspect ratio of approximately 3.7.40 Transmission electron microscopy images and UV−vis spectra (Figure 2a and
Figure 2. TEM images of (a) AuNRs, (b) Au@SiO2, and (c−e) magnetic core−shell DNA microgels. (f) Extinction spectra of AuNRs, Au@SiO2, and magnetic core−shell DNA microgels.
Figure S1 in the Supporting Information) indicate AuNRs have diameters of 25 nm and LSPR wavelength is located at 790 nm. To prevent the aggregation of AuNRs during the polymerization process, AuNRs were first coated with a silica shell by hydrolysis of TEOS in methanol/water mixtures. During the process of silica coating, a CTAB bilayer formed around the AuNRs serves as a template for the silica layer formation. Murphy et al.43 demonstrated that the most critical factor to formation of a silica layer onto AuNRs is the amount of free CTAB in solution. The silica shell thickness decreases with the increasing of CTAB concentration. When the concentration is 28188
DOI: 10.1021/acsami.7b05502 ACS Appl. Mater. Interfaces 2017, 9, 28185−28194
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ACS Applied Materials & Interfaces
Figure 3. (a) Near-IR-induced heat generation of core−shell DNA microgels aqueous solution with different concentrations at power density of 2 W/cm2. (b) SQUID magnetometry result of core−shell DNA microgels at 300 K.
cm−1, assigned to the stretching vibration of N−H (PNIPAM), the bending vibration of amide carbonyl (PNIPAM),42 and imidazole C−N stretching,47 appeared, respectively. These results demonstrated that the polymer shells were successfully grafted onto the silica surfaces. Dynamic laser scattering measurement showed that the hydrodynamic radius of the microgels is larger than the average diameter calculated from TEM images (Figure S4 in the SI). Compared with blood vessels in normal tissues, blood vessels in tumors are particularly disordered and twisted. It is worth mentioning that the vascular endothelium in tumors has more open channels than normal vessels, ranging from 200 nm to 1.2 μm.48,49 As shown in Figure S5 in the SI, after illumination by a near-IR laser for 5 min, significantly more DNA microgels passed through the filter with pore diameter of 450 nm than those at 25 and 37 °C, indicating the DNA microgels have potential to pass through the vascular endothelium for cancer therapy. Different tumor types have different sizes of tumor leaky vessels ranging from 200 nm to 1.2 μm. The size decrease of drug carriers at higher temperature is more beneficial to extravasate to tumor tissue. Moreover, Figure 2f shows that the LSPR wavelength remains in the near-IR region after polymer coating, which guaranteed the near-IR laser can penetrate into biological tissues deeply. The core−shell DNA microgels can be stably dispersed in a buffer solution of pH 7.4 (Figures S6 and S7 in the SI), while the microgel shells disappear when dispersed in the buffer solution of pH 5.0. As shown in Figure S8 in the SI, no“doughnut” part coated on the surface of Au@ SiO2. It indicates the unstability of DNA in the acidic environment results in the cross-linked structure of microgel shells breaking down, which could meaningfully promote the payloads release from the microgels. Photothermal Conversion and Magnetic Property of Magnetic Core−Shell DNA Microgels. When the AuNRs are illuminated by a near-IR laser with a wavelength that matches with a longitudinal surface plasmon resonance peak of the AuNRs, they can efficiently convert photoenergy to photothermal heat. When the heat dissipated into the surroundings, the temperature would rise and result in the unwinding of the duplex DNA. The microgel shells were spontaneously dissolved, resulting in rapid release of the cargoes from the microgels into the aqueous solution. Therefore, the efficiency of photothermal conversion is a necessary characteristic to evaluate. Temperature−time trends of DNA microgels at different concentrations are shown in
Figure 3a. Once the near-IR laser at 808 nm begins to irradiate, a rapid increase in temperature can be observed in the beginning, after 10 min, the curves reaching a plateau. The rising rate of temperature and the plateau temperature are relevant to the DNA microgel amount. The higher photothermal conversion efficiency of core−shell DNA microgels can make the photothermal heating of the solutions rapid, which can be applied to control the release of the chemotherapy agents to induce cell apoptosis upon exposure to near-IR laser beam. To date, magnetic nanomaterials using dispersions of magnetic nanoparticles have been used with some success in application of controlled drug delivery.50 However, to the best of our knowledge, drug delivery nanoplatforms based on magnetic ionic liquids have been rarely reported. SQUID magnetometry measurement well-reveals that the core−shell DNA microgels are paramagnetic (Figure 3b), which can be intuitively observed in Figure S9 in the SI, where the magnetic core−shell DNA microgels are transported toward and tightly adhered on the surface of the NdFeB magnet rapidly within a few seconds by an applied magnetic field (1 T). The whole migration process was recorded in a video in the Supporting Information (Video S1). The migration process induced by the magnet was evaluated by UV−vis spectra due to intense adsorption bands of AuNRs in the near-IR range. After applying an external magnetic field, the absorbance intensity of DNA microgels longitudinal SPR peak decreased evidently over 4 days, which demonstrated the number of DNA microgels in the upper solution were decreased (the transported DNA microgels were calculated to be approximately 47.4%), indicating the DNA microgels were transported to the magnet (Figure S10 in the SI). Furthermore, as shown in Figure S11 in the SI, the core−shell DNA microgels exhibit ferromagnetic property to some extent. We previously demonstrated that the combination of DNA and cationic magneto-surfactant micelles resulted in the formation of ferromagnetic aggregates with a typical “hysteresis loop”.51 According to classical models describing colloid particles,52,53 for the magnetic ionic liquids alone, counterions exist in both Stern and diffuse layers. In the Stern layer, counterions strongly bind and move as a whole dynamic entity, while, in the diffuse layers, counterions move freely and keep dynamic equilibrium with those in the Stern layer. When the magnetic ionic liquid monomer is incorporated into the polymer network, the counterions are enriched around the microgel. Counterions [FeCl3Br]− are hydrophobic and present 28189
DOI: 10.1021/acsami.7b05502 ACS Appl. Mater. Interfaces 2017, 9, 28185−28194
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Figure 4. (a) Dox release profiles from Mgel-Dox with or without near-IR laser irradiation at different pHs. (b) Controlled release of the Dox from Mgel-Dox by visible and near-IR-light irradiations (808 nm, 2 W/cm2; pH = 7.4, 25 °C).
pH 7.4, and the final drug release percentages are up to 84.2%. A characteristic advantage of our system is the controllable property, which enables one to achieve the more complicated on-demand cargo delivery. As a proof in Figure 4b, the release of Dox molecules is adjusted by applying a near-IR laser. At first, less than 20% Dox releases from Mgel-Dox; when the suspension of Mgel-Dox is illuminated with a near-IR laser beam (808 nm, 2 W/cm2), a more distinct release of Dox molecules occurs, which should contribute to the rapid disassembly of the DNA cross-linked polymer shell coating on the AuNR core. To the best of our knowledge, payloads triggered by a near-IR-light release from AuNRs-based DNA cross-linked microgels is scarcely explored. When the Mgel-Dox is irradiated by a near-IR light, the local temperature rising induced by AuNRs through photothermal effects makes the linker DNA thermal dissociate from their complementary sequences grafted onto the polymer strand chains (PS-A and PS-B). This leads to a rapid dissociation of the microgel layer. In order to avoid cargoes leaking in advance at body temperature (37 °C), we evaluate the Dox release at 37 °C, as shown in Figure S12 in the SI. Less than 30% Dox released from Mgel-Dox for 10 h, demonstrating the DNA microgels can avoid premature drug release at body temperature. This remote controllability of DNA microgels through near-IR light could be smartly applied to on-demand dose in clinical therapy, thereby avoiding adverse effects caused by chemotherapeutic reagents toxicity and could provide enhanced therapeutic efficacy. Cellular Internalization by Magnetic Core−Shell DNA Microgels. In vitro the intracellular release behavior of Dox against human hepatoma cell lines, HepG2 cells, is investigated by confocal laser scanning microscopy (CLSM), and the results are shown in Figure 5. After incubating the HepG2 cells with Mgel-Dox for 4 h without laser treatment, the weak Dox fluorescence signal is observed in the cytoplasm and almost no fluorescent signals are observed in the nucleus (Figure 5b). On the contrary, most Dox are observed in cellular cytoplasm and nuclei after incubating for 4 h with a near-IR laser (Figure 5a), indicating that more Dox are released from Mgel-Dox after irradiating to the laser and then diffused into the nuclei. The amount of released drugs is significantly increased by the photothermal effect. The change in the fluorescence images indicates that the remote heating generated by AuNRs improved the Dox release from the DNA microgels efficiently. Clearly, the results prove that the anticancer drugs can be effectively delivered into living cells by a present nanocarrier controlled by external near-IR light.
strong hydrophobic interaction with the alkyl chains of the magnetic ionic liquid existing in the polymer network, intensifying the combination between counterions [FeCl3Br]− and polymers. The core−shell DNA microgels take on ferromagnetism to some extent. This ferromagnetism demonstrates the potential for controlled orientation of core−shell DNA microgels. It can be useful in transporting drugs to fixed cells or tissues and assist in improving the delivery efficiency in clinical therapy. Drug Loading and Laser-Controlled Release. Doxorubicin hydrochloride (Dox), a well-known chemotherapy drug, is employed as a model drug to evaluate the core−shell DNA microgels as a drug delivery system. Dox is loaded into the core−shell DNA microgels through electrostatic interaction to form a complex (abbreviated as Mgel-Dox). Dox can be easily incorporated into DNA microgels through stirring the DNA microgels solution with Dox solution overnight and removing the unencapsulated Dox by centrifuging and washing for several times. It should be note that the Dox loading content reached up to 26.4%, and Dox entrapment efficiency is 93.8%. Because of the lower cytotoxicity even at high concentrations and easier entry into cells, DNA-based drug nanoplatforms have been extensively studied.54 However, the microgels which used the thermosensitive duplex DNA as drug gatekeepers are scarcely reported. In this work, the duplex DNA undergoes thermal transformation to unwinding into single-stranded DNA at elevated temperatures caused by AuNRs, so that the loaded drugs are released rapidly and completely, which can improve the therapeutic effect of chemtherapy drug efficiently. The release profiles of the captured Dox from Mgel-Dox whether irradiated by near-IR laser or not at different pHs are determined by monitoring the increasing absorbance signals at 480 nm of the supernatants at known time intervals. As shown in Figure 4a, only a small amount of leaking of Dox molecules from the Mgel-Dox occurred during the whole process of testing at pH 7.4 without laser, indicating the core− shell DNA microgels can avoid premature drug release during the entire voyage. The significantly enhanced drug release rate and percentage can be realized by irradiating a laser at 808 nm. It is because the hyperthermal Mgel-Dox suspension dehybridized the duplex DNA and dissociated electrostatic interactions between Dox and the polymer shell. Due to the weak acidic microenvironment of cancer cells, release profiles of the loaded Dox from Mgel-Dox at pH 5.0 are carried. The unstable DNA sequences in the acidic microenvironment promote Dox releasing from the DNA microgels. Therefore, the drug release rate and percentage are higher than those at 28190
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(Figure 6b). It confirms that resulting magnetic core−shell DNA microgels can act as a drug carrier for cancer therapy. Overwhelming efforts to exact point-to-point drug release of chemotherapy, one of the most effective treatments for cancer therapy, have not yet reached clinical trials, making patients suffer heavy adverse effects with long-term treatment. Multidrug resistance and overdose remain a challenge in the clinical situations even though the ideal therapeutic effects can be achieved by employing a small amount of medications. Hence, shielding the high toxicity of chemotherapeutic agents and improving the killing efficacy of chemotherapeutic agents efficiently in due course are expected. As shown in Figure 6c, when treated with Mgel-Dox without a laser, HepG2 cells show dose-dependent cell inactivity and distinct higher cell inactivity than free Dox, indicating that the core−shell DNA microgels shield the high toxicity of chemotherapeutic agents efficiently. With exposure to a near-IR laser, a distinct decrease of cell viability on HepG2 cells is induced by Mgel-Dox with a more than 2-fold lower IC50 of 3.063 μM compared to that without near-IR irradiation (7.461 μM), indicating the near-IR irradiation can be an efficient measure for regulating the therapeutic efficiency of Dox. During the time of transportation, the Dox can be preserved in the network of the microgel shells, avoiding healthy organs and tissues having to suffer severe adverse effects. When the Mgel-Dox is exposed to the near-IR laser, Dox releases from the Mgel-Dox rapidly, resulting in the enhanced therapeutic efficiency of Dox. This
Figure 5. CLSM images of HepG2 cells after incubation with MgelDox for 4 h with (a) or without (b) a near-IR laser (808 nm, 2W/cm2, 10 min). The concentration of Dox in Mgel-Dox complexes is 10 μM.
Cell Cytotoxicity of Dox-Loaded DNA Microgels. As anticancer drug carriers, the first question is to determine the biocompatibility and killing efficacy of our proposed magnetic core−shell DNA microgels for cancer cells. The therapeutic effect of the core−shell DNA microgels on HepG2 cells is evaluated by the MTT assay. First, HepG2 cells are irradiated by 808 nm laser at the same power density of 2 W/cm2 for different times. As shown in Figure 6a, the cell viabilities are all above 90%, demonstrating that near-IR laser irradiation of the HepG2 cells (10 min) have little influence on cells, which is attributed to the low near-IR-light absorption by these cells.55 Furthermore, the cell viabilities of the core−shell DNA microgels are all above 90% for different incubating times, indicating excellent biocompatibility of these nanomaterials
Figure 6. Cytotoxicity assays of (a) cells are exposed to a near-infrared laser at 808 nm (2 W/cm2) for 0, 3, 5, and 10 min, respectively, (b) cells incubated with core−shell DNA microgels at different concentrations, (c) cells incubated with free Dox and Mgel-Dox with or without 808 nm (2 W/cm2), (d) cells incubated with core−shell DNA microgels (6.287 μg/mL), free Dox (10 μM), and Mgel-Dox (the Dox in Mgel is 10 μM) by exposing to 808 nm (2 W/cm2) for 0, 5, and 10 min, respectively. 28191
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high controllability of constructed Mgel-Dox is very important in clinical therapy. Multimodality theranostic platforms, which integrate discrete therapeutics into one platform and achieve higher therapeutic efficacy than that of each single mode, have attracted much attention.56 To certify whether the enhanced therapeutic effects induced by the synergy effects of Dox and photothermal property, HepG2 cells are incubated with free Dox and nonDox-loaded DNA microgels and then exposed to a 808 nm near-IR laser, as shown in Figure 6d. The therapeutic effect of these two cases is dramatically higher than that of the near-IRactivated Mgel-Dox, demonstrating that the synergistic effect is not a key factor in this system. The enhanced killing efficiency of near-IR-activated Mgel-Dox is contributed to the continuous release of Dox via photothermal despiralization of DNA sequences in microgel shells. It has been confirmed that the high controllability of created Mgel-Dox possesses excellent biocompatibility and highly therapeutic effect, which have the potential to act as cancer therapeutic candidates. It is expected that our magnetic ionic liquid-based DNA microgels have the potential to become a clinically viable and versatile carrier for targeting and killing cancers.
Shuli Dong: 0000-0003-3961-7238 Jingcheng Hao: 0000-0002-9760-9677 Notes
The authors declare no competing financial interest.
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ACKNOWLEDGMENTS This work was funded by the National Natural Science Foundation of China (Grant No. 21420102006) and the Specialized Research Fund for the Doctoral Program of Higher Education (Grant No. 20130131130005).
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CONCLUSIONS In summary, we design and construct unprecedented DNA− polymer microgels based on the specificity of ionic liquids with magnetic [FeCl3Br]− ions, near-IR-light-responsive AuNRs for magnetically targeted delivery and controllable release of chemotherapy drugs. It should be realized that the construction of magnetic microgels occurred without magnetic nanoparticles. Because of their facile fabrication, rapid responsiveness, and external controllability, the light- and magnetoresponsive DNA microgels can be used as promising drug vehicles. The DNA not only acts as crossing-linker agent but also can serve as a gatekeeper to regulate the release of drug. Internalization study and MTT assay demonstrate that the Dox-loaded core−shell DNA microgels can efficiently uptake into cancer cells, enhance the cytotoxicity of cancer cells controlled by near-IR laser and shield the high toxicity of chemotherapeutic agents to improve the killing efficacy of chemotherapeutic agents efficiently in due course. It can be controlled well by the near-IR laser. Our results may provide opportunities for broad new functional nucleic acid-based nanostructures and holds great promise as a candidate for applications in the field of nanomaterials and biomedicine.
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ASSOCIATED CONTENT
S Supporting Information *
The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acsami.7b05502. UV−vis, 1H NMR, mass, and FTIR spectra; images of the prefiltered suspension and filtrates; TEM images; photographs of the core−shell DNA microgels; SQUID magnetometry results; and Dox release profiles (PDF) Video of the core−shell DNA microgels by applying the migration toward an applied external NdFeB magnet (AVI)
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