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Core–shell–corona doxorubicin-loaded superparamagnetic Fe3O4 nanoparticles for cancer theranostics
Accepted Manuscript Title: Core-Shell-Corona Doxorubicin-Loaded Superparamagnetic Fe3 O4 Nanoparticles for Cancer Theranostics Author: A. Semkina M. Abakumov N. Grinenko A. Abakumov A. Skorikov E. Mironova G. Davydova A.G. Majouga N. Nukolova A. Kabanov V. Chekhonin PII: DOI: Reference:
S0927-7765(15)30292-7 http://dx.doi.org/doi:10.1016/j.colsurfb.2015.11.009 COLSUB 7464
To appear in:
Colloids and Surfaces B: Biointerfaces
Received date: Revised date: Accepted date:
31-7-2015 3-11-2015 5-11-2015
Please cite this article as: A.Semkina, M.Abakumov, N.Grinenko, A.Abakumov, A.Skorikov, E.Mironova, G.Davydova, A.G.Majouga, N.Nukolova, A.Kabanov, V.Chekhonin, Core-Shell-Corona Doxorubicin-Loaded Superparamagnetic Fe3O4 Nanoparticles for Cancer Theranostics, Colloids and Surfaces B: Biointerfaces http://dx.doi.org/10.1016/j.colsurfb.2015.11.009 This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting proof before it is published in its final form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.
Core-Shell-Corona Doxorubicin-Loaded Superparamagnetic Fe3O4 Nanoparticles for Cancer Theranostics A. Semkinaa, M. Abakumova,b, N. Grinenkoc, A. Abakumovd, A. Skorikovb, E. Mironovae, G. Davydovae , A.G. Majougab, N. Nukolovab,c, A. Kabanovb,f, V. Chekhonina,c a
Pirogov Russian National Research Medical University, Ostrovitianov 1, 117997 Moscow, Russia Lomonosov Moscow State University, Lenin Hills 1, 119991 Moscow, Russia c Serbsky Federal Medical Research Centre of Psychiatry and Narcology, Kropotkinskiy 23, 119991 Moscow, Russia d Electron Microscopy for Materials Science (EMAT), University of Antwerp, Groenenborgerlaan 171, 2020 Antwerp, Belgium e Federal State Institution of Science Institute of Theoretical and Experimental Biophysics, Russian Academy of Science, Institutskaya 3, 142290 Pushchino. Russia f University of North Carolina at Chapel Hill, Center for Nanotechnology in Drug Delivery, 125 Mason Farm Road, Chapel Hill, NC 27599-7362 NC, USA b
Corresponding author:
M. A. Abakumov, tel.: +79035864777 e-mail: [email protected]
Pirogov Russian National Research Medical University, Ostrovitianov 1, 117997 Moscow, Russia Lomonosov Moscow State University, Lenin Hills 1, 119991 Moscow, Russia
The paper consists of 5082 words, 1 table and 7 figures.
Graphical abstract:
Highlights
Doxorubicin complex with magnetic nanoparticles was synthesized. Electrostatic attraction were shown to play key role in complex formation. pH sensitive release of Doxorubicin was shown. Formation of complex does not affect Doxorubicin cytotoxic activity.
Abstract Superparamagnetic iron oxide magnetic nanoparticles (MNPs) are successfully used as contrast agents in magnetic-resonance imaging. They can be easily functionalized for drug delivery functions, demonstrating great potential for both imaging and therapeutic applications. Here we developed new pH-responsive theranostic core-shell-corona nanoparticles consisting of superparamagentic Fe3O4 core that displays high T2 relaxivity, bovine serum albumin (BSA) shell that binds anticancer drug, doxorubicin (Dox) and poly(ethylene glycol) (PEG) corona that increases stability and biocompatibility. The nanoparticles were produced by adsorption of the BSA shell onto the Fe3O4 core followed by crosslinking of the protein layer and subsequent grafting of the PEG corona using monoamino-terminated PEG via carbodiimide chemistry. The hydrodynamic diameter, zeta-potential, composition and T2 relaxivity of the resulting nanoparticles were characterized using transmission electron microscopy, dynamic light scattering, thermogravimetric analysis and T2-relaxometry. Nanoparticles were shown to absorb Dox molecules, possibly through a combination of electrostatic and hydrophobic interactions. The loading capacity (LC) of the nanoparticles was 8 wt. %. The Dox loaded nanoparticles release the drug at a higher rate at pH 5.5 compared to pH 7.4 and display similar cytotoxicity against C6 and HEK293 cells as the free Dox. Keywords: magnetic nanoparticles, doxorubicin, drug delivery, anticancer therapy, MRIcontrast agents
Introduction Cancer is one of the most widespread group of diseases causing about 14.6% of all human deaths [1]. Cancer treatments include surgery, radiotherapy, and chemotherapy; the latter being used in nearly 50% of cancer cases [2]. A general drawback of chemotherapy is relatively low efficacy of drug delivery to the target tumor cells accompanied with unintended penetration of the drugs to non-target cells and tissues [3]. This leads to i) a need of using higher doses of anticancer drugs to achieve the desired drug concentration in tumor cells and ii) severe side effects due to off-target toxicity of the chemotherapeutic agents [4]. Another significant disadvantage of the chemotherapy is a development of resistance to the anticancer drugs in tumor cells, which further decreases therapeutic outcomes [2]. In an attempt to overcome the above limitations, various nanobiomaterials, such as liposomes [5], magnetic nanoparticles [6-8], polymers [9], nanogels [10] etc. are used clinically or investigated as nanocarriers for delivery of chemotherapeutic agents. Such carriers increase circulation time of chemotherapeutic agents in blood stream, improve accumulation and retention of the chemotherapeutic agents in the tumor and in some cases increase delivery of chemotherapeutic agents across physiological barriers (for example, the blood-brain, blood-cerebrospinal fluid, blood-tumor) to the disease site [11], [12]. Being of 10 to 100 nm in size, these nanocarriers accumulate in tumors due to the enhanced permeability and retention (EPR) effect which is a passive targeting phenomenon as a result of tumor hypervascularisation, increased vessel permeability and weakened lymphatic drainage [13]. Altogether, this leads to increased therapy efficacy and reduced side effects [14]. Superparamagnetic iron oxide magnetic nanoparticles (MNPs) have recently gained increasing attention as templates for design of nanocarriers for chemotherapeutic agents [3, 15-18]. First, such MNPs have nanoscale size (10-100 nm) [19], which reduces probability of particle’s capture by reticuloendothelial system (RES) [20] and provides localization in tumors via the EPR effect. Second, such MNPs have high specific surface area [20], and this surface can be modified to impart specific properties, such as hydrophilicity/lipophilicity, charge, chemical reactivity, grafting of targeting moieties [17, 21]. Third, the unique mesoscopic superparamagnetic properties of such MNPs allow their detection by magnetic resonance imaging (MRI). By generating its own local magnetic field MNPs can act as contrast agents for MRI by decreasing the signal from the tissue under investigation [22], [23]. The main advantage of MRI is the lack of distance limitations from the tumor mass to the surface of body, unlike other imaging systems [17], which allows detection of any kind of the tumor independently of its localization. Combination of the above imaging and drug delivery modalities provides opportunity for simultaneous imaging and therapy of affected tissues, i.e. theranostics [24], [25] [17, 26-28]. This concept was successfully demonstrated using a variety of anticancer drugs (e.g. paclitaxel [29], doxorubicin (Dox) [30], temozolomide etc. [31]) as well as several visualization techniques such as MRI [29], optical imaging [32], etc. Despite unique combination of theranostic features, use of MNPs in biomedical applications is impeded by their low biocompatibility and tendency to aggregate [14]. Besides that MNPs typically have limited loading capacities [12, 15]. To overcome these limitations, MNPs can be coated by biocompatible materials, thereby providing a shell that i) reduces cytotoxicity and increases biocompatibility [33]; ii) increases dispersion stability of the nanoparticles under physiological conditions; iii) decreases capture and elimination of the nanoparticles by the RES and iv) facilitates further functionalization of MNPs to introduce targeting or other functional moieties [34]. Furthermore, such shells can also be used for drug loading, protecting drug from interaction with non-targeting tissue and enabling controlled release of the drug at the tumor site. [17]. Examples of such coatings include dextran [35], chitosan [36], polyethyleneimine [37] as
well as serum albumins [38],[39]. In some selected cases the MNPs shell can be further grafted with poly(ethylene glycol) (PEG) resulting in core-shell-corona MNP architectures of the nanocarriers with optimized biocompatibility, stealth properties and dispersion stability [18]. The PEG corona increases biocompatibility on the nanocarriers, protects them from nonspecific interactions with the plasma proteins and reduces capture of the nanocarriers by RES [15], [34]. Further improvement of drug delivery to tumors can be achieved by developing stimulussensitive nanoparticles on the basis of MNPs. Such nanoparticles can provide drug release in response to changing environmental parameters (pH, ionic strength, chemical substances). [40]. For example, low oxygen concentration, low permeability of blood vessels and enhanced glucose metabolism in tumor tissues commonly result in a decrease of pH in the tumor [41]. Thus, acidic pH of the tumors can be employed as one environmental cue for tumor-specific drug release from nanostructures. This work describes a new type of pH-responsive core-shell-corona nanocarriers on the base of MNPs. Such nanocarriers have i) a superparamagnetic magnetite (Fe3O4) core, ii) a bovine serum albumin (BSA) shell loaded with Dox as an anticancer agent, and iii) a lyophilizing PEG corona. Physicochemical studies reveal high Dox loading into the MNPs (up to 8 wt.%) and enhanced drug release at acidic pH. In vitro studies using glioma C6 cancer cell cultures show the same cytotoxicity of Dox loaded MNP compared to free Dox.
Materials and methods General Iron acetylacetonate (III) (99.9%), benzyl alcohol (99.8%), bovine serum albumin, glutaraldehyde 25% (w/v) aqueous solution, N-hydroxysuccinimide (NHS), 1-Ethyl-3-(3dimethylaminopropyl) carbodiimide (EDC) were purchased from Sigma-Aldrich, USA. For MNPs synthesis acetone (99.5%) from Samsung Total, Korea, sodium borohydride from Aldrich, USA, monoamino terminated poly(ethylene glycol) hydrochloride (PEG-NH2) from Creative PEGworks, USA) were used. Doxorubicin hydrochloride («Doxorubicin-Teva») was purchased from PHARMACHEMIE B.V., TEVA Pharmaceutical Industries Ltd., Israel. Growth medium DMEM/F12 with 10% fetal bovine serum (Invitrogen, USA) was used as received. Centrifugal filters with pore diameter of 100 and 30 kDa (Amicon, USA) and syringe filters with pore diameter of 0.45 и 0.22 µm (Millipore, USA) were used to sterilize samples. Desalting of polymers was conducted with mini-column for gel filtration NAP-10 (GE HealthCare, USA). MNPs were re-suspended using ultrasonic disintegrator (Laboratory Supplies Co, USA) and acetone was removed with a rotary evaporator equipped with a vacuum pump (Laborota 4000, Heidolph, Germany). Studies of MNP relaxivity were performed with magnetic resonance tomograph ClinScan 7T (Bruker Biospin, USA). Dox concentrations were determined with a UV spectrophotometer UV-VIS U-2800 HITACHI (Japan), λ=490 nm. Synthesis of MNPs, coated by bovine serum albumin (MNP-BSA) Synthesis of MNPs by thermal decomposition of iron acetylacetonate (III) in benzyl alcohol was carried out according to the procedure described in [42]. MNP-BSA were prepared as follows: 5 ml of distilled water were added to 20 mg of MNP and pH was adjusted to 11 using 1M NaOH. This dispersion was sonicated (10 min), and 40 mg of BSA, dissolved in 5 ml of water were added. This mixture was incubated for 4 hours at room temperature and continuous stirring, and
then was dialyzed (25 kDa) against distilled water. 250 µl of 1M NaOH were added to 10 ml of the resulting solution followed by dropwise addition of 1.1 ml of 25% aqueous glutaraldehyde under stirring. The resulting mixture was incubated under stirring for 15 min, then 250 µl of 3M glycine (pH 9.2) were added followed by addition of 0.5 ml of (10 mg/ml) sodium borohydride solution in phosphate buffered saline (PBS) (137 mM NaCl, 2.7 mM KCl, 10 mM Na2HPO4, 1.76 mM KH2PO4, pH=7.4). Then mixture was incubated for 60 minutes. MNP-BSA were separated from excess of glutaraldehyde using cellulose centrifuge filters (Amicon, 100 kDa, 2000 rpm). Purification of MNP-BSA from the excess of unbound BSA was performed by gel filtration (Millipore, Sepharose CL-4B, column 1.8 cm2*40 cm, eluent - 1*PBS, flow rate - 1 ml/min. Two fractions were collected: MNP-BSA (I) - 22-30 min; MNP-BSA (II) 30-42 min. Coating MNP-BSA by the PEG corona (preparation of MNP-BSA-PEG) 5 mg of MNP-BSA were dissolved in PBS buffer to the concentration of Fe3+ of 0.95 mg/ml. Then 266 µl of NHS solution in PBS (10 mg/ml) and 457 µl of EDC solution in PBS (10 mg/ml) were added. The reaction mixture was incubated with stirring for 10 min, after which 470 µl of hydrochloride salt of polyethylene glycol NH2-PEG-OH solution in dH2O (10 mg/ml) was added. Incubation time was 1 h. The resulting MNP-BSA-PEG were separated from excess of PEG by gel filtration using column NAP-10 (Sepadex G25, eluent – PBS). Transmission Electron Microscopy (TEM) The solutions were diluted by nearly equal amount of distilled water and deposited on a holey carbon-film replica. High resolution transmission electron microscopy (HR-TEM) images, high angle annular dark field scanning TEM (HAADF-STEM) images and energy-dispersive X-ray (EDX) compositional maps were recorded using a Tecnai Osiris microscope equipped with a Super-X detector and operated at 200 kV. Dox loading into the MNP-BSA-PEG (preparation of MNP-BSA@Dox-PEG) Solution of MNP-BSA-PEG (100 µg/ml, 1 ml) was loaded into a tube and 50 µl of Dox solution (100 µg/ml) were added. The reaction mixture was incubated for 5 minutes and then the size of formed MNP-BSA-PEG was measured. The procedure was repeated until the total volume of the solution reached 2 ml. Dox loading was calculated as the ratio of drug mass to the total weight of MNP-BSA-PEG and added Dox. LC = [m(Dox)/(m(Dox)+m(MNP-BSA-PEG))]*100% Concentration of released Dox was evaluated by measurement of light absorption at 490 nm wavelength of buffer surrounding dialysis tube. Total Dox amount was calculated by the following formula: M(Dox)=C(Dox)*Mr(Dox)*V where M(Dox) is total mass of Dox in the buffer surrounding dialysis tube, C(Dox) – molar concentration of Dox and V – volume of the buffer outside of dialysis tube.
Study of Dox-loading at different ion strength values
The solutions of Dox loaded MNP-BSA-PEG with different ionic strength values (0.1 – 0.3M) were prepared from PBS solutions by adding appropriate amount of 3M NaCl. The total solutions volume was 1.5 ml. In the next step they were filtered by Amicon filters with 30 kDa cut-off for 5 min (2000 rpm) and Dox concentrations in the filtrates were determined. In vitro Dox release The 266 µl of Dox solution (1.2 mg/ml) were added to 1 ml of MNP-BSA-PEG solution (1.2 mg/ml) to achieve 8% loading of Dox into MNP-BSA-PEG. The reaction mixture was incubated for 15 min at room temperature with stirring, and then transferred to dialysis bags (10 kDa). Dialysis bags were placed in plastic tubes filled with 25 ml of PBS with pH 7.4, 6.5 and 5.5 (room temperature, stirring). The release was carried out at for 25 hours. 500 µl of the external buffer were taken out from tubes at various time intervals and analyzed for Dox concentrations in the solution by spectrophotometry at 490 nm. Amount of released Dox was expressed as a percentage of the initial amount added Dox. Analysis of human fibroblast (HF) cells viability The cells of culture HF were added in 96-well culture plates with concentration of 25000 cells/sm2 in growth medium DMEM/F12 with 5% fetal bovine serum. The cells were cultured at 37°C in a humidified atmosphere with 5% CO2. The growth media was removed after 18 hours from wells and 100 µl per well of sample with appropriate concentration in growth medium DMEM/F12 were added; for each concentration 4 repeats was used. Cells were incubated with the samples for 48 h at 37°C in a humidified atmosphere with 5% CO2. Estimation of cells viability was provided by Axiovert 200 microscope. Method of cells fluorescent staining was used to carry out the analysis by using L-7007 LIVE/DEAD BacLight Bacterial Viability Kit (Invitrogen). MTT cytotoxicity studies on glioma C6 / HEK293 cell lines The 5*103 cells/well of culture C6 or HEK293 were added in 96-well culture plates in 200 ul growth medium DMEM/F12 with 10% fetal bovine serum. The cells were cultured for 48 hours at 37°C in a humidified atmosphere with 5% CO2. The test drugs (free Dox and Dox loaded MNP with 8% LC) were sterilized (0.45 µm nylon filters), 2-fold dilutions of the samples were prepared in the growth medium. The growth media was removed from wells and 100 µl per well of drug with appropriate concentration were added; for each concentration 4 repeats was used. Cells were incubated with the drugs for 24 h at 37°C in a humidified atmosphere with 5% CO2, after which the inoculum was removed, cells were washed with DPBS, and were cultured for 24 h in fresh growth medium (100 µl/well). The growth media was removed and 20 µl of MTS reagent were added into each well After incubation for 4 h at 37°C in the dark in a humidified atmosphere with 5% CO2 the absorbance was recorded with plate analyzer VictorX3 (PerkinElmer, USA) at λ = 490 nm.
Results and Discussion 1. Coating of iron oxide MNPs with BSA to obtain MNP-BSA We used thermal decomposition of iron acetylacetonate (III) in benzyl alcohol as described previously to synthesize magnetite MNPs [42]. As shown earlier this method allows obtaining MNPs with well-defined crystal structure and magnetic properties. For biological applications these MNPs must be coated with biocompatible materials to improve stability and prevent aggregation under physiological conditions. In this study MNP were surface coated with the
BSA molecules that were further cross-linked in order to improve stability. Upon purification of the MNPs from the excess of unbound protein by gel-filtration chromatography the modified MNPs were eluted by a single peak. However this peak was relatively broad suggesting that a mixture of the MNP clusters embedded in a protein shell of variable diameter was obtained after crosslinking. After deconvolution of the peak into two contributions we obtained two nanoparticle fractions - light (MNP-BSA(II) and heavy MNP-BSA(I). These fractions contained nanoparticles of different size and shell composition. The hydrodynamic diameter of the particles in MNP-BSA (I) was 85±10 nm, which was 2.4 times larger than the size of MNP-BSA (II) particles (Table 1). The polydispersity index (PDI) of both fractions was 0.18 ± 0.02. The composition of MNP-BSA (I) and MNP-BSA(II) was evaluated by thermogravimetric analysis (TGA) in air. The protein weight fraction in MNP-BSA(I) nanoparticles was 22% , which is less than that for MNP-BSA(II) (Table 1). Since MNP-BSA (II) display smaller particle size as measured by DLS, they also have greater specific surface area compared to MNP-BSA (I) and, therefore, larger amount of protein was needed for surface coating of the same mass of the iron oxide nanocrystals. Both nanoparticle fractions exhibited negative zeta potential since the BSA molecules have negative charge under physiological conditions [43]. However, the zeta potential of MNPBSA(II) particles was 15 mV greater in the absolute value compared to that of MNP-BSA(I) particles (Table 1). The morphology of core-shell MNP-BSA was investigated by HR-TEM and HAADF-STEM (Fig.1). In the STEM mode the incident electron beam is focused into a very fine probe of ~ 1 A˚ wide. The probe scans over the specimen in a raster pattern. Transmitted electrons are collected with an annular detector located in the diffraction plane and the registered intensity is plotted against the position of the probe. The most commonly used STEM mode is the high-angle annular dark-field STEM (HAADF-STEM). In this mode, the detector collects electrons scattered to very high angles. This way, the contribution of the Bragg scattering (i.e. coherent elastic scattering on the electrostatic potential of the crystal as whole) is minimized. The HAADF-STEM signal is dominated by the Rutherford scattering (i.e. incoherent elastic electron scattering on the nuclei) and the inelastic thermal diffuse scattering. As a result, the HAADF-STEM signal strongly depends on the chemical composition. In this way it was shown by HAADF-STEM method that individual nanoparticles are embedded in protein shell and form agglomerates. The difference in particle size between MNP-BSA(I) and MNP-BSA(II) fractions was also observed by these methods. The T2-relaxivity values of the nanoparticles were equal to 271±6 mM-1s-1 for MNP-BSA (I) and about ~100 mM-1s-1 lower for MNP-BSA (II) (Supp. 1). In the case of superparamagnetic MNPs, used as MRI contrast agents (for example, Feridex IV, Resovist for liver imaging, Lumirem for gastrointestinal tract, etc.), T2-relaxivity values are in the range of 50 mM-1s-1 or higher [44]. Therefore, successful visualization of MNP-BSA in various body tissues using MRI can be expected. The T2-relaxivity of MNP-BSA (I) particles is higher than that of MNP-BSA (II) probably due to the increase of particle size and following increase of local magnetic field inhomogeneity. This leads to the higher surface spin anisotropy of protons around MNP and proton spin dephasing, which decreases T2 relaxation time more effectively in comparison with MNP with smaller size [45]. However, it was found that these particles are less stable in aqueous solution (buffer PBS, 4°C) in comparison with MNP-BSA (II) (Supp. 2, 3). Therefore, despite of higher T2 relaxivity values of MNP-BSA (I) in further experiments we used MNP-BSA (II) nanoparticles.
2. Synthesis of core-shell-corona nanoparticles MNP-BSA-PEG
At the next step core-shell MNP-BSA (II) were conjugated with PEG molecules. This coating is known to improve stability of the nanoparticles in biological fluids and is commonly used in various biomedical applications [18], [46], [47]. In addition PEG provides biocompatibility of the nanoparticles and decreases their elimination by RES [48]. The modification of the MNPBSA nanoparticles resulted in a small increase of the effective diameter, which does not exceed 5 nm in comparison with the diameter of the initial MNP-BSA (II) nanoparticles. After 10 days the hydrodynamic diameter of MNP-BSA-PEG remained constant (41±4 nm) with the PDI value being less than 0.25. The composition of core-shell-corona MNP-BSA-PEG was determined by TGA. As shown in Table 1 the shell-corona content was increased by 8% in MNP-BSA-PEG compared to that in the initial MNP-BSA (II) nanoparticles. The zeta potential of the MNP-BSA-PEG was smaller than that of MNP-BSA (II) before PEG conjugation. The results of T2-relaxivity value measurements of the MNP-BSA-PEG particles shows almost no change in comparison with the initial MNPBSA (II). Thus, we can conclude that coating MNP-BSA (II) with PEG-shell has no significant effect on their magnetic properties (Table 1). Fig.2 shows the viabilities of human fibroblast (HF) cells incubated with medium containing different types of synthesized nanoparticles: MNP-BSA (I), MNP-BSA (II), MNP-BSA-PEG. The viability of HF cells was affected by nanoparticles in dose-dependent manner, but for same Fe3+ concentrations MNP-BSA-PEG was less toxic than both MNP-BSA (I) and MNP-BSA (II). This confirms that PEG-coating decreases the MNP-BSA (II) toxicity with respect to living cells. As a result, PEGylated MNP-BSA with small hydrodynamic diameter are the most promising tool for the anticancer drug loading and following therapy. 4. Magnetization values for MNP, MNP-BSA (I), MNP-BSA (II), and MNP-BSA-PEG To characterize magnetic properties of MNPs with different coatings, magnetization curve of these nanoparticles was obtained. For all MNP samples we did not observe any residual magnetization on the hysteresis curve that indicates superparamagnetic properties of different MNPs (Fig. 3). This indirectly suggests that there is little if any aggregation induced by magnetic attraction in solution after MR imaging. Saturation magnetization of the initial uncoated MNPs was 59±3 emu/g. For comparison, the previously reported values for saturation magnetization of the magnetite nanoparticles varied from as low as 2.5 emu/g [13] to about 50 emu/g [38], or 47 and 57 emu/g for the particles with different coatings [39]. When compared with our initial MNPs the saturation magnetization of MNP-BSA (I), MNP-BSA (II), and MNP-BSA-PEG was decreased by 15%, 66% and 68%, respectively. These values are in good agreement with the content of protein and polymers in our MNP-based nanomaterials. Since the shell and corona materials do not show any significant magnetic properties they are not expected to affect the saturation magnetization value. Therefore the reduction in the saturation magnetization value is explained by decreasing the Fe3O4 mass fraction in the total mass of the modified nanoparticles. Based on the results of the magnetization measurements one can conclude that the obtained coreshell-corona MNPs display magnetic properties that can provide a detectable response to external magnetic field. 5. Loading of MNP-BSA-PEG with Dox At the next stage MNP-BSA-PEG were loaded with Dox. The BSA molecules forming the shell of MNP-BSA-PEG are negatively charged under the physiological conditions (BSA pI=5.5) [43]. Therefore, the shell can absorb Dox molecules that are positively charged at pH 7.4 (pKa=8.3) [49], through a combination of electrostatic interactions with the shell of MNP-BSA-PEG and
possibly, interactions of the anthracycline moieties of the drug with each other, for example, πstacking. This mechanism of drug loading has a distinct advantage over the covalent immobilization as it does not change the drug chemical structure, which fully preserves the functional activity of the drug and is easier for subsequent regulatory approval of successful nanoformulations. Also it is known that the electrostatic coupling strongly depends on the ionic strength in case of weak bases such as the aminoglycan moiety in Dox and acids such as carboxylic groups in the BSA on the pH. As shown further, the Dox loaded MNP-BSA-PEG nanoformulations remain stable even after incubation in PBS. However, the binding of the drug is decreased and the drug release is enhanced upon acidification. Our experiments demonstrate that the Dox loaded MNP-BSA-PEG particles remain stable and practically do not change size up to the Dox loading capacity of 8% wt. Above this point we observed increase in both the size of the complex and the PDI value from 0.20 to 0.45 (Fig. 4). The destabilization effect may be due to neutralization of MNP-BSA-PEG surface charge by increasing amounts of the particle-bound drug. Thus, we defined LC 8% as the optimal Dox loading in MNP-BSA-PEG and used this loading in further experiments. In comparison, covalent crosslinking of Dox and MNP can lead to 13.2 % loading capacity [50]. Formation of a complex of Dox and MNP through hydrophobic interactions, such as in the studies [18] and [15], results in the drug loading capacity ranged from 5.2 to 6.4% and from 3.2 to 3.9%, respectively. The higher Dox loading in the case of covalent binding with MNP can be explained by lower strength of hydrophobic or electrostatic bonding compared to covalent bond, but, as described earlier, these weak interactions do not modify the chemical structure of the drug and do not affect its anti-tumor activity. The T2-relaxivity of the MNP-BSA@Dox-PEG is 160±7 mM-1s-1 shows that Dox loading in the core-shell-corona nanoparticles has no effect on their magnetic properties and does not reduce the diagnostic MRI contrast activity of the developed carrier. Immobilization of the Dox molecules by electrostatic interactions between Dox and MNP-BSA-PEG should result in a decrease of the MNP zeta-potential. The results (Fig. 5) reveal that zeta-potential on the MNP surface increases with the percentage of Dox loading. The unloaded particles were characterized by a zeta potential of -30±3 mV, whereas the MNPBSA@Dox-PEG with 8% Dox content show slightly higher zeta potential of -25±1 mV. In addition, the sample of MNP-BSA@Dox-PEG with 20% Dox content has a much higher zeta potential (-11±1 mV) in comparison with the values of the unloaded nanoparticles and MNPBSA@Dox-PEG with 8% loading. The zeta potential increases, probably due to the shielding the MNP-BSA-PEG negative surface charge by the additional Dox positive ions. As a result, we can see a reduced repulsion between the macromolecular components that can result in agglomeration and sedimentation. In the article [51] increasing the HSA-nanoparticles zeta potential during Dox-loading was also demonstrated. The changes from -43.3±1.1mV to 33.1±2.6 mV were observed. Moreover, great increase in the Dox-loaded HSA-nanoparticles diameter was observed: from 166,5±17,6 nm in the case of unloaded nanoparticles to 379,5±21,5 nm for drug-loaded nanoparticles. This fact confirms that the Dox-loaded nanoparticles tend to enlarge and agglomerate. In addition, the effect of different ionic strength on a Dox loading into MNP-BSA-PEG was investigated. The doxorubicin hydrochloride dissociates in water solutions into the protonized anthracycline and negatively charged chloride ions. The ionic compounds were previously shown to influence the Dox loading into nanoparticles [52]. As it can be seen in Supp. 4, concentration of free unbound DOX released from MNP-BSA@Dox-PEG increases from 0.025 to 0.048 mg/ml at the ion strength increasing from 0.1 to 0.3 M. The shielding of surface charges on MNP-BSA-PEG and Dox weakens the electrostatic interactions between the macromolecules and following drug release from the MNP-BSA@Dox-PEG increases.
6. Dox release from the MNP-BSA@Dox-PEG We analyzed the Dox release under different environmental acidity to confirm the effectiveness of the obtained MNP-BSA@Dox-PEG for the Dox delivery. It is known that the environment acidity of the affected tissue in certain solid tumors increases as a result of hypoxia and subsequent acidosis [41]. According to the data received, at pH=7.4 the amount of released Dox does not exceed 25% by weight of the initial drug content in MNP-BSA@Dox-PEG with LC 8% (Fig. 6). Under acidic conditions, the process is more rapid and at pH=6.5 and pH=5.5 after 25 h the released Dox values amount to 55 and 75%, respectively. We believe that protonation of the negatively charged functional groups in BSA reduces the number of electrostatic interaction centers with the positively charged Dox molecules. As a result, the Dox molecules could not further be bound to the surface of MNP-BSA-PEG and this fact explains more rapid drug release. Thus, MNPBSA@Dox-PEG delivered to the tumor tissue can demonstrate faster Dox release in this area, providing directed and stimulus-sensitive drug delivery. Prolonged Dox release from MNPBSA@Dox-PEG at normal pH values also reduces antitumor drugs side effects, associated with high dose of drugs during conventional chemotherapy. In addition, the fact of intensity dependence of the Dox release on pH-values provides the further evidence of the presence of electrostatic interactions in MNP-BSA@Dox-PEG between BSA core-shell and Dox. The lower the pH, the faster the electrostatic complex is destroyed and consequently drug is released. Release process at different pH values is often analyzed while creating nano-containers with Dox. In the paper [3], the increase of Dox release degree from MNP modified by various polymers with decreasing pH from 7.4 to 5.8 has been shown. The minimum release value was approximately 50% for 250 hours, while the maximum reached value was almost 90%. The same dependence on acidity was found in the study [11], where the Dox release was analyzed at the pH values of 4.5, 5.5, 6.5 and 7.5 and temperature of 37°C. However, even at the lowest pH (4.5), the release value did not exceed 30% after 25 hours. 7. Cytotoxic activity of the MNP-BSA@Dox-PEG To characterize the therapeutic properties of MNP-BSA@Dox-PEG, we investigated its effect on living cells in an MTT cell viability assay, which is widely used to evaluate the cytotoxicity of various biomaterials. The results of the performed MTT-assay are shown in Fig. 7. MNP-BSA@Dox-PEG were still effective and led to inhibition of C6 and HEK293 cells viability. The results demonstrate that the cytotoxic activity of the Dox-loaded nanoparticles is quite similar to that of the respective Dox control in case of both Dox-resistant tumor (C6) and Dox-sensitive (HEK293) cells. Therefore, we can conclude that no loss of drug efficacy after loading to MNP-BSA-PEG was observed and MNP-BSA@Dox-PEG can be used as chemotherapy agent with high efficiency. Other researchers have investigated cytotoxicity of Dox-loaded MNP on other cell lines such as A549 lung cancer cell line [3]. It was found that these particles more actively suppress cell growth, in comparison with free Dox. MNP with covalently immobilized Dox [11] also has greater toxicity as compared with free Dox. In other study [52], the effect of Dox-loaded HSA-nanoparticles on two neuroblastoma cell lines (UKF-NB3, IMR 32) was studied. Different carriers, based on HSA, showed cell growth inhibition, which was comparable or better than that of free Dox. One can suggest that the drug loaded into carrier, based on BSA or HSA, has cytotoxity, which is quite similar with the Dox cytotoxicity due to their high biocompatibility.
Conclusion In this study we describe the preparation of core-shell-corona MNP, loaded with cytostatic drug doxorubicin. The resulting MNP-BSA@Dox-PEG is a promising bifunctional
agent: on the one hand, due to the presence of superparamagnetic properties and high T2relaxivity values, this complex can be used as a contrast agent for MRI diagnostics of tumor tissues. On the other hand, the MNP-BSA@Dox-PEG can be used for therapy of oncological diseases, providing prolonged stimulus-sensitive drug release and reducing side effects of chemotherapy associated with nonspecific drug delivery to the tumors. Thus, the MNPBSA@Dox-PEG can be successfully used in cancer theranostics.
Acknowledgements This work was supported by grant of Russian Federation Ministry of Science and Education grant 11.G34.31.0004 and №14.604.21.0007, RFMEFI60414X0007, grants of RSF 14-15-00698 (cytotoxicity), 14-34-00017 (magnetization measurements) and 14-13-00731 (MRI measurements). We also thank Konstantin Kurilenko (M.V. Lomonosov Moscow State University, Department of Chemistry) for carrying out the TGA analysis.
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Fig.1. HAADF-STEM overview images and HR-TEM images of MNP-BSA(I) (a, b) and MNPBSA(II) (c, d). The HAADF-STEM images (a, c) clearly demonstrate the difference in the size of the NMP-BSA agglomerates in the fractions I and II. The HR-TEM images (b, d) show crystalline Fe3O4 nanoparticles surrounded by the BSA shell. The panel (e) shows the HAADFSTEM image and EDX compositional maps for NMP-BSA
Fig.2. Cytotoxicity study of MNP-BSA (I), MNP-BSA (II), MNP-BSA-PEG in HF cells.
Fig. 3. Magnetization curves of MNP, MNP-BSA (I), MNP-BSA (II), MNP-BSA-PEG.
Fig. 4. Dox-loading on core-shell-corona: the dependence of the MNP-BSA@Dox-PEG hydrodynamic diameter on the amount of loaded Dox.
Fig. 5. Zeta potential values for MNP-BSA-PEG, MNP-BSA@Dox-PEG with 8% LC and MNPBSA@Dox-PEG with 20% LC
Fig. 6. Dox release profile for different pH-values (7.4, 6.5, 5.5).
Fig. 7. Cytotoxicity study of MNP-BSA@Dox-PEG HEK293 cell cultures.
compared to Dox control on C6 and
Table 1. Physical and chemical properties of MNP-BSA(I), MNP-BSA(II) and MNP-BSA-PEG. Hydrodynamic diameter, nm
Organic shell content , mass. %
Т2-relaxivity, mM-1s-1
ζ-potential, mV
Saturation magnetization, emu/g
MNP-BSA (I)
85±10
28±3
271±6
-24±2
50±3
MNP-BSA (II)
36±4
50±3
161±4
-35±3
20±1
MNP-BSAPEG
41±4
58±5
151±3
-30±3
19±1
S0927-7765(15)30292-7 http://dx.doi.org/doi:10.1016/j.colsurfb.2015.11.009 COLSUB 7464
To appear in:
Colloids and Surfaces B: Biointerfaces
Received date: Revised date: Accepted date:
31-7-2015 3-11-2015 5-11-2015
Please cite this article as: A.Semkina, M.Abakumov, N.Grinenko, A.Abakumov, A.Skorikov, E.Mironova, G.Davydova, A.G.Majouga, N.Nukolova, A.Kabanov, V.Chekhonin, Core-Shell-Corona Doxorubicin-Loaded Superparamagnetic Fe3O4 Nanoparticles for Cancer Theranostics, Colloids and Surfaces B: Biointerfaces http://dx.doi.org/10.1016/j.colsurfb.2015.11.009 This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting proof before it is published in its final form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.
Core-Shell-Corona Doxorubicin-Loaded Superparamagnetic Fe3O4 Nanoparticles for Cancer Theranostics A. Semkinaa, M. Abakumova,b, N. Grinenkoc, A. Abakumovd, A. Skorikovb, E. Mironovae, G. Davydovae , A.G. Majougab, N. Nukolovab,c, A. Kabanovb,f, V. Chekhonina,c a
Pirogov Russian National Research Medical University, Ostrovitianov 1, 117997 Moscow, Russia Lomonosov Moscow State University, Lenin Hills 1, 119991 Moscow, Russia c Serbsky Federal Medical Research Centre of Psychiatry and Narcology, Kropotkinskiy 23, 119991 Moscow, Russia d Electron Microscopy for Materials Science (EMAT), University of Antwerp, Groenenborgerlaan 171, 2020 Antwerp, Belgium e Federal State Institution of Science Institute of Theoretical and Experimental Biophysics, Russian Academy of Science, Institutskaya 3, 142290 Pushchino. Russia f University of North Carolina at Chapel Hill, Center for Nanotechnology in Drug Delivery, 125 Mason Farm Road, Chapel Hill, NC 27599-7362 NC, USA b
Corresponding author:
M. A. Abakumov, tel.: +79035864777 e-mail: [email protected]
Pirogov Russian National Research Medical University, Ostrovitianov 1, 117997 Moscow, Russia Lomonosov Moscow State University, Lenin Hills 1, 119991 Moscow, Russia
The paper consists of 5082 words, 1 table and 7 figures.
Graphical abstract:
Highlights
Doxorubicin complex with magnetic nanoparticles was synthesized. Electrostatic attraction were shown to play key role in complex formation. pH sensitive release of Doxorubicin was shown. Formation of complex does not affect Doxorubicin cytotoxic activity.
Abstract Superparamagnetic iron oxide magnetic nanoparticles (MNPs) are successfully used as contrast agents in magnetic-resonance imaging. They can be easily functionalized for drug delivery functions, demonstrating great potential for both imaging and therapeutic applications. Here we developed new pH-responsive theranostic core-shell-corona nanoparticles consisting of superparamagentic Fe3O4 core that displays high T2 relaxivity, bovine serum albumin (BSA) shell that binds anticancer drug, doxorubicin (Dox) and poly(ethylene glycol) (PEG) corona that increases stability and biocompatibility. The nanoparticles were produced by adsorption of the BSA shell onto the Fe3O4 core followed by crosslinking of the protein layer and subsequent grafting of the PEG corona using monoamino-terminated PEG via carbodiimide chemistry. The hydrodynamic diameter, zeta-potential, composition and T2 relaxivity of the resulting nanoparticles were characterized using transmission electron microscopy, dynamic light scattering, thermogravimetric analysis and T2-relaxometry. Nanoparticles were shown to absorb Dox molecules, possibly through a combination of electrostatic and hydrophobic interactions. The loading capacity (LC) of the nanoparticles was 8 wt. %. The Dox loaded nanoparticles release the drug at a higher rate at pH 5.5 compared to pH 7.4 and display similar cytotoxicity against C6 and HEK293 cells as the free Dox. Keywords: magnetic nanoparticles, doxorubicin, drug delivery, anticancer therapy, MRIcontrast agents
Introduction Cancer is one of the most widespread group of diseases causing about 14.6% of all human deaths [1]. Cancer treatments include surgery, radiotherapy, and chemotherapy; the latter being used in nearly 50% of cancer cases [2]. A general drawback of chemotherapy is relatively low efficacy of drug delivery to the target tumor cells accompanied with unintended penetration of the drugs to non-target cells and tissues [3]. This leads to i) a need of using higher doses of anticancer drugs to achieve the desired drug concentration in tumor cells and ii) severe side effects due to off-target toxicity of the chemotherapeutic agents [4]. Another significant disadvantage of the chemotherapy is a development of resistance to the anticancer drugs in tumor cells, which further decreases therapeutic outcomes [2]. In an attempt to overcome the above limitations, various nanobiomaterials, such as liposomes [5], magnetic nanoparticles [6-8], polymers [9], nanogels [10] etc. are used clinically or investigated as nanocarriers for delivery of chemotherapeutic agents. Such carriers increase circulation time of chemotherapeutic agents in blood stream, improve accumulation and retention of the chemotherapeutic agents in the tumor and in some cases increase delivery of chemotherapeutic agents across physiological barriers (for example, the blood-brain, blood-cerebrospinal fluid, blood-tumor) to the disease site [11], [12]. Being of 10 to 100 nm in size, these nanocarriers accumulate in tumors due to the enhanced permeability and retention (EPR) effect which is a passive targeting phenomenon as a result of tumor hypervascularisation, increased vessel permeability and weakened lymphatic drainage [13]. Altogether, this leads to increased therapy efficacy and reduced side effects [14]. Superparamagnetic iron oxide magnetic nanoparticles (MNPs) have recently gained increasing attention as templates for design of nanocarriers for chemotherapeutic agents [3, 15-18]. First, such MNPs have nanoscale size (10-100 nm) [19], which reduces probability of particle’s capture by reticuloendothelial system (RES) [20] and provides localization in tumors via the EPR effect. Second, such MNPs have high specific surface area [20], and this surface can be modified to impart specific properties, such as hydrophilicity/lipophilicity, charge, chemical reactivity, grafting of targeting moieties [17, 21]. Third, the unique mesoscopic superparamagnetic properties of such MNPs allow their detection by magnetic resonance imaging (MRI). By generating its own local magnetic field MNPs can act as contrast agents for MRI by decreasing the signal from the tissue under investigation [22], [23]. The main advantage of MRI is the lack of distance limitations from the tumor mass to the surface of body, unlike other imaging systems [17], which allows detection of any kind of the tumor independently of its localization. Combination of the above imaging and drug delivery modalities provides opportunity for simultaneous imaging and therapy of affected tissues, i.e. theranostics [24], [25] [17, 26-28]. This concept was successfully demonstrated using a variety of anticancer drugs (e.g. paclitaxel [29], doxorubicin (Dox) [30], temozolomide etc. [31]) as well as several visualization techniques such as MRI [29], optical imaging [32], etc. Despite unique combination of theranostic features, use of MNPs in biomedical applications is impeded by their low biocompatibility and tendency to aggregate [14]. Besides that MNPs typically have limited loading capacities [12, 15]. To overcome these limitations, MNPs can be coated by biocompatible materials, thereby providing a shell that i) reduces cytotoxicity and increases biocompatibility [33]; ii) increases dispersion stability of the nanoparticles under physiological conditions; iii) decreases capture and elimination of the nanoparticles by the RES and iv) facilitates further functionalization of MNPs to introduce targeting or other functional moieties [34]. Furthermore, such shells can also be used for drug loading, protecting drug from interaction with non-targeting tissue and enabling controlled release of the drug at the tumor site. [17]. Examples of such coatings include dextran [35], chitosan [36], polyethyleneimine [37] as
well as serum albumins [38],[39]. In some selected cases the MNPs shell can be further grafted with poly(ethylene glycol) (PEG) resulting in core-shell-corona MNP architectures of the nanocarriers with optimized biocompatibility, stealth properties and dispersion stability [18]. The PEG corona increases biocompatibility on the nanocarriers, protects them from nonspecific interactions with the plasma proteins and reduces capture of the nanocarriers by RES [15], [34]. Further improvement of drug delivery to tumors can be achieved by developing stimulussensitive nanoparticles on the basis of MNPs. Such nanoparticles can provide drug release in response to changing environmental parameters (pH, ionic strength, chemical substances). [40]. For example, low oxygen concentration, low permeability of blood vessels and enhanced glucose metabolism in tumor tissues commonly result in a decrease of pH in the tumor [41]. Thus, acidic pH of the tumors can be employed as one environmental cue for tumor-specific drug release from nanostructures. This work describes a new type of pH-responsive core-shell-corona nanocarriers on the base of MNPs. Such nanocarriers have i) a superparamagnetic magnetite (Fe3O4) core, ii) a bovine serum albumin (BSA) shell loaded with Dox as an anticancer agent, and iii) a lyophilizing PEG corona. Physicochemical studies reveal high Dox loading into the MNPs (up to 8 wt.%) and enhanced drug release at acidic pH. In vitro studies using glioma C6 cancer cell cultures show the same cytotoxicity of Dox loaded MNP compared to free Dox.
Materials and methods General Iron acetylacetonate (III) (99.9%), benzyl alcohol (99.8%), bovine serum albumin, glutaraldehyde 25% (w/v) aqueous solution, N-hydroxysuccinimide (NHS), 1-Ethyl-3-(3dimethylaminopropyl) carbodiimide (EDC) were purchased from Sigma-Aldrich, USA. For MNPs synthesis acetone (99.5%) from Samsung Total, Korea, sodium borohydride from Aldrich, USA, monoamino terminated poly(ethylene glycol) hydrochloride (PEG-NH2) from Creative PEGworks, USA) were used. Doxorubicin hydrochloride («Doxorubicin-Teva») was purchased from PHARMACHEMIE B.V., TEVA Pharmaceutical Industries Ltd., Israel. Growth medium DMEM/F12 with 10% fetal bovine serum (Invitrogen, USA) was used as received. Centrifugal filters with pore diameter of 100 and 30 kDa (Amicon, USA) and syringe filters with pore diameter of 0.45 и 0.22 µm (Millipore, USA) were used to sterilize samples. Desalting of polymers was conducted with mini-column for gel filtration NAP-10 (GE HealthCare, USA). MNPs were re-suspended using ultrasonic disintegrator (Laboratory Supplies Co, USA) and acetone was removed with a rotary evaporator equipped with a vacuum pump (Laborota 4000, Heidolph, Germany). Studies of MNP relaxivity were performed with magnetic resonance tomograph ClinScan 7T (Bruker Biospin, USA). Dox concentrations were determined with a UV spectrophotometer UV-VIS U-2800 HITACHI (Japan), λ=490 nm. Synthesis of MNPs, coated by bovine serum albumin (MNP-BSA) Synthesis of MNPs by thermal decomposition of iron acetylacetonate (III) in benzyl alcohol was carried out according to the procedure described in [42]. MNP-BSA were prepared as follows: 5 ml of distilled water were added to 20 mg of MNP and pH was adjusted to 11 using 1M NaOH. This dispersion was sonicated (10 min), and 40 mg of BSA, dissolved in 5 ml of water were added. This mixture was incubated for 4 hours at room temperature and continuous stirring, and
then was dialyzed (25 kDa) against distilled water. 250 µl of 1M NaOH were added to 10 ml of the resulting solution followed by dropwise addition of 1.1 ml of 25% aqueous glutaraldehyde under stirring. The resulting mixture was incubated under stirring for 15 min, then 250 µl of 3M glycine (pH 9.2) were added followed by addition of 0.5 ml of (10 mg/ml) sodium borohydride solution in phosphate buffered saline (PBS) (137 mM NaCl, 2.7 mM KCl, 10 mM Na2HPO4, 1.76 mM KH2PO4, pH=7.4). Then mixture was incubated for 60 minutes. MNP-BSA were separated from excess of glutaraldehyde using cellulose centrifuge filters (Amicon, 100 kDa, 2000 rpm). Purification of MNP-BSA from the excess of unbound BSA was performed by gel filtration (Millipore, Sepharose CL-4B, column 1.8 cm2*40 cm, eluent - 1*PBS, flow rate - 1 ml/min. Two fractions were collected: MNP-BSA (I) - 22-30 min; MNP-BSA (II) 30-42 min. Coating MNP-BSA by the PEG corona (preparation of MNP-BSA-PEG) 5 mg of MNP-BSA were dissolved in PBS buffer to the concentration of Fe3+ of 0.95 mg/ml. Then 266 µl of NHS solution in PBS (10 mg/ml) and 457 µl of EDC solution in PBS (10 mg/ml) were added. The reaction mixture was incubated with stirring for 10 min, after which 470 µl of hydrochloride salt of polyethylene glycol NH2-PEG-OH solution in dH2O (10 mg/ml) was added. Incubation time was 1 h. The resulting MNP-BSA-PEG were separated from excess of PEG by gel filtration using column NAP-10 (Sepadex G25, eluent – PBS). Transmission Electron Microscopy (TEM) The solutions were diluted by nearly equal amount of distilled water and deposited on a holey carbon-film replica. High resolution transmission electron microscopy (HR-TEM) images, high angle annular dark field scanning TEM (HAADF-STEM) images and energy-dispersive X-ray (EDX) compositional maps were recorded using a Tecnai Osiris microscope equipped with a Super-X detector and operated at 200 kV. Dox loading into the MNP-BSA-PEG (preparation of MNP-BSA@Dox-PEG) Solution of MNP-BSA-PEG (100 µg/ml, 1 ml) was loaded into a tube and 50 µl of Dox solution (100 µg/ml) were added. The reaction mixture was incubated for 5 minutes and then the size of formed MNP-BSA-PEG was measured. The procedure was repeated until the total volume of the solution reached 2 ml. Dox loading was calculated as the ratio of drug mass to the total weight of MNP-BSA-PEG and added Dox. LC = [m(Dox)/(m(Dox)+m(MNP-BSA-PEG))]*100% Concentration of released Dox was evaluated by measurement of light absorption at 490 nm wavelength of buffer surrounding dialysis tube. Total Dox amount was calculated by the following formula: M(Dox)=C(Dox)*Mr(Dox)*V where M(Dox) is total mass of Dox in the buffer surrounding dialysis tube, C(Dox) – molar concentration of Dox and V – volume of the buffer outside of dialysis tube.
Study of Dox-loading at different ion strength values
The solutions of Dox loaded MNP-BSA-PEG with different ionic strength values (0.1 – 0.3M) were prepared from PBS solutions by adding appropriate amount of 3M NaCl. The total solutions volume was 1.5 ml. In the next step they were filtered by Amicon filters with 30 kDa cut-off for 5 min (2000 rpm) and Dox concentrations in the filtrates were determined. In vitro Dox release The 266 µl of Dox solution (1.2 mg/ml) were added to 1 ml of MNP-BSA-PEG solution (1.2 mg/ml) to achieve 8% loading of Dox into MNP-BSA-PEG. The reaction mixture was incubated for 15 min at room temperature with stirring, and then transferred to dialysis bags (10 kDa). Dialysis bags were placed in plastic tubes filled with 25 ml of PBS with pH 7.4, 6.5 and 5.5 (room temperature, stirring). The release was carried out at for 25 hours. 500 µl of the external buffer were taken out from tubes at various time intervals and analyzed for Dox concentrations in the solution by spectrophotometry at 490 nm. Amount of released Dox was expressed as a percentage of the initial amount added Dox. Analysis of human fibroblast (HF) cells viability The cells of culture HF were added in 96-well culture plates with concentration of 25000 cells/sm2 in growth medium DMEM/F12 with 5% fetal bovine serum. The cells were cultured at 37°C in a humidified atmosphere with 5% CO2. The growth media was removed after 18 hours from wells and 100 µl per well of sample with appropriate concentration in growth medium DMEM/F12 were added; for each concentration 4 repeats was used. Cells were incubated with the samples for 48 h at 37°C in a humidified atmosphere with 5% CO2. Estimation of cells viability was provided by Axiovert 200 microscope. Method of cells fluorescent staining was used to carry out the analysis by using L-7007 LIVE/DEAD BacLight Bacterial Viability Kit (Invitrogen). MTT cytotoxicity studies on glioma C6 / HEK293 cell lines The 5*103 cells/well of culture C6 or HEK293 were added in 96-well culture plates in 200 ul growth medium DMEM/F12 with 10% fetal bovine serum. The cells were cultured for 48 hours at 37°C in a humidified atmosphere with 5% CO2. The test drugs (free Dox and Dox loaded MNP with 8% LC) were sterilized (0.45 µm nylon filters), 2-fold dilutions of the samples were prepared in the growth medium. The growth media was removed from wells and 100 µl per well of drug with appropriate concentration were added; for each concentration 4 repeats was used. Cells were incubated with the drugs for 24 h at 37°C in a humidified atmosphere with 5% CO2, after which the inoculum was removed, cells were washed with DPBS, and were cultured for 24 h in fresh growth medium (100 µl/well). The growth media was removed and 20 µl of MTS reagent were added into each well After incubation for 4 h at 37°C in the dark in a humidified atmosphere with 5% CO2 the absorbance was recorded with plate analyzer VictorX3 (PerkinElmer, USA) at λ = 490 nm.
Results and Discussion 1. Coating of iron oxide MNPs with BSA to obtain MNP-BSA We used thermal decomposition of iron acetylacetonate (III) in benzyl alcohol as described previously to synthesize magnetite MNPs [42]. As shown earlier this method allows obtaining MNPs with well-defined crystal structure and magnetic properties. For biological applications these MNPs must be coated with biocompatible materials to improve stability and prevent aggregation under physiological conditions. In this study MNP were surface coated with the
BSA molecules that were further cross-linked in order to improve stability. Upon purification of the MNPs from the excess of unbound protein by gel-filtration chromatography the modified MNPs were eluted by a single peak. However this peak was relatively broad suggesting that a mixture of the MNP clusters embedded in a protein shell of variable diameter was obtained after crosslinking. After deconvolution of the peak into two contributions we obtained two nanoparticle fractions - light (MNP-BSA(II) and heavy MNP-BSA(I). These fractions contained nanoparticles of different size and shell composition. The hydrodynamic diameter of the particles in MNP-BSA (I) was 85±10 nm, which was 2.4 times larger than the size of MNP-BSA (II) particles (Table 1). The polydispersity index (PDI) of both fractions was 0.18 ± 0.02. The composition of MNP-BSA (I) and MNP-BSA(II) was evaluated by thermogravimetric analysis (TGA) in air. The protein weight fraction in MNP-BSA(I) nanoparticles was 22% , which is less than that for MNP-BSA(II) (Table 1). Since MNP-BSA (II) display smaller particle size as measured by DLS, they also have greater specific surface area compared to MNP-BSA (I) and, therefore, larger amount of protein was needed for surface coating of the same mass of the iron oxide nanocrystals. Both nanoparticle fractions exhibited negative zeta potential since the BSA molecules have negative charge under physiological conditions [43]. However, the zeta potential of MNPBSA(II) particles was 15 mV greater in the absolute value compared to that of MNP-BSA(I) particles (Table 1). The morphology of core-shell MNP-BSA was investigated by HR-TEM and HAADF-STEM (Fig.1). In the STEM mode the incident electron beam is focused into a very fine probe of ~ 1 A˚ wide. The probe scans over the specimen in a raster pattern. Transmitted electrons are collected with an annular detector located in the diffraction plane and the registered intensity is plotted against the position of the probe. The most commonly used STEM mode is the high-angle annular dark-field STEM (HAADF-STEM). In this mode, the detector collects electrons scattered to very high angles. This way, the contribution of the Bragg scattering (i.e. coherent elastic scattering on the electrostatic potential of the crystal as whole) is minimized. The HAADF-STEM signal is dominated by the Rutherford scattering (i.e. incoherent elastic electron scattering on the nuclei) and the inelastic thermal diffuse scattering. As a result, the HAADF-STEM signal strongly depends on the chemical composition. In this way it was shown by HAADF-STEM method that individual nanoparticles are embedded in protein shell and form agglomerates. The difference in particle size between MNP-BSA(I) and MNP-BSA(II) fractions was also observed by these methods. The T2-relaxivity values of the nanoparticles were equal to 271±6 mM-1s-1 for MNP-BSA (I) and about ~100 mM-1s-1 lower for MNP-BSA (II) (Supp. 1). In the case of superparamagnetic MNPs, used as MRI contrast agents (for example, Feridex IV, Resovist for liver imaging, Lumirem for gastrointestinal tract, etc.), T2-relaxivity values are in the range of 50 mM-1s-1 or higher [44]. Therefore, successful visualization of MNP-BSA in various body tissues using MRI can be expected. The T2-relaxivity of MNP-BSA (I) particles is higher than that of MNP-BSA (II) probably due to the increase of particle size and following increase of local magnetic field inhomogeneity. This leads to the higher surface spin anisotropy of protons around MNP and proton spin dephasing, which decreases T2 relaxation time more effectively in comparison with MNP with smaller size [45]. However, it was found that these particles are less stable in aqueous solution (buffer PBS, 4°C) in comparison with MNP-BSA (II) (Supp. 2, 3). Therefore, despite of higher T2 relaxivity values of MNP-BSA (I) in further experiments we used MNP-BSA (II) nanoparticles.
2. Synthesis of core-shell-corona nanoparticles MNP-BSA-PEG
At the next step core-shell MNP-BSA (II) were conjugated with PEG molecules. This coating is known to improve stability of the nanoparticles in biological fluids and is commonly used in various biomedical applications [18], [46], [47]. In addition PEG provides biocompatibility of the nanoparticles and decreases their elimination by RES [48]. The modification of the MNPBSA nanoparticles resulted in a small increase of the effective diameter, which does not exceed 5 nm in comparison with the diameter of the initial MNP-BSA (II) nanoparticles. After 10 days the hydrodynamic diameter of MNP-BSA-PEG remained constant (41±4 nm) with the PDI value being less than 0.25. The composition of core-shell-corona MNP-BSA-PEG was determined by TGA. As shown in Table 1 the shell-corona content was increased by 8% in MNP-BSA-PEG compared to that in the initial MNP-BSA (II) nanoparticles. The zeta potential of the MNP-BSA-PEG was smaller than that of MNP-BSA (II) before PEG conjugation. The results of T2-relaxivity value measurements of the MNP-BSA-PEG particles shows almost no change in comparison with the initial MNPBSA (II). Thus, we can conclude that coating MNP-BSA (II) with PEG-shell has no significant effect on their magnetic properties (Table 1). Fig.2 shows the viabilities of human fibroblast (HF) cells incubated with medium containing different types of synthesized nanoparticles: MNP-BSA (I), MNP-BSA (II), MNP-BSA-PEG. The viability of HF cells was affected by nanoparticles in dose-dependent manner, but for same Fe3+ concentrations MNP-BSA-PEG was less toxic than both MNP-BSA (I) and MNP-BSA (II). This confirms that PEG-coating decreases the MNP-BSA (II) toxicity with respect to living cells. As a result, PEGylated MNP-BSA with small hydrodynamic diameter are the most promising tool for the anticancer drug loading and following therapy. 4. Magnetization values for MNP, MNP-BSA (I), MNP-BSA (II), and MNP-BSA-PEG To characterize magnetic properties of MNPs with different coatings, magnetization curve of these nanoparticles was obtained. For all MNP samples we did not observe any residual magnetization on the hysteresis curve that indicates superparamagnetic properties of different MNPs (Fig. 3). This indirectly suggests that there is little if any aggregation induced by magnetic attraction in solution after MR imaging. Saturation magnetization of the initial uncoated MNPs was 59±3 emu/g. For comparison, the previously reported values for saturation magnetization of the magnetite nanoparticles varied from as low as 2.5 emu/g [13] to about 50 emu/g [38], or 47 and 57 emu/g for the particles with different coatings [39]. When compared with our initial MNPs the saturation magnetization of MNP-BSA (I), MNP-BSA (II), and MNP-BSA-PEG was decreased by 15%, 66% and 68%, respectively. These values are in good agreement with the content of protein and polymers in our MNP-based nanomaterials. Since the shell and corona materials do not show any significant magnetic properties they are not expected to affect the saturation magnetization value. Therefore the reduction in the saturation magnetization value is explained by decreasing the Fe3O4 mass fraction in the total mass of the modified nanoparticles. Based on the results of the magnetization measurements one can conclude that the obtained coreshell-corona MNPs display magnetic properties that can provide a detectable response to external magnetic field. 5. Loading of MNP-BSA-PEG with Dox At the next stage MNP-BSA-PEG were loaded with Dox. The BSA molecules forming the shell of MNP-BSA-PEG are negatively charged under the physiological conditions (BSA pI=5.5) [43]. Therefore, the shell can absorb Dox molecules that are positively charged at pH 7.4 (pKa=8.3) [49], through a combination of electrostatic interactions with the shell of MNP-BSA-PEG and
possibly, interactions of the anthracycline moieties of the drug with each other, for example, πstacking. This mechanism of drug loading has a distinct advantage over the covalent immobilization as it does not change the drug chemical structure, which fully preserves the functional activity of the drug and is easier for subsequent regulatory approval of successful nanoformulations. Also it is known that the electrostatic coupling strongly depends on the ionic strength in case of weak bases such as the aminoglycan moiety in Dox and acids such as carboxylic groups in the BSA on the pH. As shown further, the Dox loaded MNP-BSA-PEG nanoformulations remain stable even after incubation in PBS. However, the binding of the drug is decreased and the drug release is enhanced upon acidification. Our experiments demonstrate that the Dox loaded MNP-BSA-PEG particles remain stable and practically do not change size up to the Dox loading capacity of 8% wt. Above this point we observed increase in both the size of the complex and the PDI value from 0.20 to 0.45 (Fig. 4). The destabilization effect may be due to neutralization of MNP-BSA-PEG surface charge by increasing amounts of the particle-bound drug. Thus, we defined LC 8% as the optimal Dox loading in MNP-BSA-PEG and used this loading in further experiments. In comparison, covalent crosslinking of Dox and MNP can lead to 13.2 % loading capacity [50]. Formation of a complex of Dox and MNP through hydrophobic interactions, such as in the studies [18] and [15], results in the drug loading capacity ranged from 5.2 to 6.4% and from 3.2 to 3.9%, respectively. The higher Dox loading in the case of covalent binding with MNP can be explained by lower strength of hydrophobic or electrostatic bonding compared to covalent bond, but, as described earlier, these weak interactions do not modify the chemical structure of the drug and do not affect its anti-tumor activity. The T2-relaxivity of the MNP-BSA@Dox-PEG is 160±7 mM-1s-1 shows that Dox loading in the core-shell-corona nanoparticles has no effect on their magnetic properties and does not reduce the diagnostic MRI contrast activity of the developed carrier. Immobilization of the Dox molecules by electrostatic interactions between Dox and MNP-BSA-PEG should result in a decrease of the MNP zeta-potential. The results (Fig. 5) reveal that zeta-potential on the MNP surface increases with the percentage of Dox loading. The unloaded particles were characterized by a zeta potential of -30±3 mV, whereas the MNPBSA@Dox-PEG with 8% Dox content show slightly higher zeta potential of -25±1 mV. In addition, the sample of MNP-BSA@Dox-PEG with 20% Dox content has a much higher zeta potential (-11±1 mV) in comparison with the values of the unloaded nanoparticles and MNPBSA@Dox-PEG with 8% loading. The zeta potential increases, probably due to the shielding the MNP-BSA-PEG negative surface charge by the additional Dox positive ions. As a result, we can see a reduced repulsion between the macromolecular components that can result in agglomeration and sedimentation. In the article [51] increasing the HSA-nanoparticles zeta potential during Dox-loading was also demonstrated. The changes from -43.3±1.1mV to 33.1±2.6 mV were observed. Moreover, great increase in the Dox-loaded HSA-nanoparticles diameter was observed: from 166,5±17,6 nm in the case of unloaded nanoparticles to 379,5±21,5 nm for drug-loaded nanoparticles. This fact confirms that the Dox-loaded nanoparticles tend to enlarge and agglomerate. In addition, the effect of different ionic strength on a Dox loading into MNP-BSA-PEG was investigated. The doxorubicin hydrochloride dissociates in water solutions into the protonized anthracycline and negatively charged chloride ions. The ionic compounds were previously shown to influence the Dox loading into nanoparticles [52]. As it can be seen in Supp. 4, concentration of free unbound DOX released from MNP-BSA@Dox-PEG increases from 0.025 to 0.048 mg/ml at the ion strength increasing from 0.1 to 0.3 M. The shielding of surface charges on MNP-BSA-PEG and Dox weakens the electrostatic interactions between the macromolecules and following drug release from the MNP-BSA@Dox-PEG increases.
6. Dox release from the MNP-BSA@Dox-PEG We analyzed the Dox release under different environmental acidity to confirm the effectiveness of the obtained MNP-BSA@Dox-PEG for the Dox delivery. It is known that the environment acidity of the affected tissue in certain solid tumors increases as a result of hypoxia and subsequent acidosis [41]. According to the data received, at pH=7.4 the amount of released Dox does not exceed 25% by weight of the initial drug content in MNP-BSA@Dox-PEG with LC 8% (Fig. 6). Under acidic conditions, the process is more rapid and at pH=6.5 and pH=5.5 after 25 h the released Dox values amount to 55 and 75%, respectively. We believe that protonation of the negatively charged functional groups in BSA reduces the number of electrostatic interaction centers with the positively charged Dox molecules. As a result, the Dox molecules could not further be bound to the surface of MNP-BSA-PEG and this fact explains more rapid drug release. Thus, MNPBSA@Dox-PEG delivered to the tumor tissue can demonstrate faster Dox release in this area, providing directed and stimulus-sensitive drug delivery. Prolonged Dox release from MNPBSA@Dox-PEG at normal pH values also reduces antitumor drugs side effects, associated with high dose of drugs during conventional chemotherapy. In addition, the fact of intensity dependence of the Dox release on pH-values provides the further evidence of the presence of electrostatic interactions in MNP-BSA@Dox-PEG between BSA core-shell and Dox. The lower the pH, the faster the electrostatic complex is destroyed and consequently drug is released. Release process at different pH values is often analyzed while creating nano-containers with Dox. In the paper [3], the increase of Dox release degree from MNP modified by various polymers with decreasing pH from 7.4 to 5.8 has been shown. The minimum release value was approximately 50% for 250 hours, while the maximum reached value was almost 90%. The same dependence on acidity was found in the study [11], where the Dox release was analyzed at the pH values of 4.5, 5.5, 6.5 and 7.5 and temperature of 37°C. However, even at the lowest pH (4.5), the release value did not exceed 30% after 25 hours. 7. Cytotoxic activity of the MNP-BSA@Dox-PEG To characterize the therapeutic properties of MNP-BSA@Dox-PEG, we investigated its effect on living cells in an MTT cell viability assay, which is widely used to evaluate the cytotoxicity of various biomaterials. The results of the performed MTT-assay are shown in Fig. 7. MNP-BSA@Dox-PEG were still effective and led to inhibition of C6 and HEK293 cells viability. The results demonstrate that the cytotoxic activity of the Dox-loaded nanoparticles is quite similar to that of the respective Dox control in case of both Dox-resistant tumor (C6) and Dox-sensitive (HEK293) cells. Therefore, we can conclude that no loss of drug efficacy after loading to MNP-BSA-PEG was observed and MNP-BSA@Dox-PEG can be used as chemotherapy agent with high efficiency. Other researchers have investigated cytotoxicity of Dox-loaded MNP on other cell lines such as A549 lung cancer cell line [3]. It was found that these particles more actively suppress cell growth, in comparison with free Dox. MNP with covalently immobilized Dox [11] also has greater toxicity as compared with free Dox. In other study [52], the effect of Dox-loaded HSA-nanoparticles on two neuroblastoma cell lines (UKF-NB3, IMR 32) was studied. Different carriers, based on HSA, showed cell growth inhibition, which was comparable or better than that of free Dox. One can suggest that the drug loaded into carrier, based on BSA or HSA, has cytotoxity, which is quite similar with the Dox cytotoxicity due to their high biocompatibility.
Conclusion In this study we describe the preparation of core-shell-corona MNP, loaded with cytostatic drug doxorubicin. The resulting MNP-BSA@Dox-PEG is a promising bifunctional
agent: on the one hand, due to the presence of superparamagnetic properties and high T2relaxivity values, this complex can be used as a contrast agent for MRI diagnostics of tumor tissues. On the other hand, the MNP-BSA@Dox-PEG can be used for therapy of oncological diseases, providing prolonged stimulus-sensitive drug release and reducing side effects of chemotherapy associated with nonspecific drug delivery to the tumors. Thus, the MNPBSA@Dox-PEG can be successfully used in cancer theranostics.
Acknowledgements This work was supported by grant of Russian Federation Ministry of Science and Education grant 11.G34.31.0004 and №14.604.21.0007, RFMEFI60414X0007, grants of RSF 14-15-00698 (cytotoxicity), 14-34-00017 (magnetization measurements) and 14-13-00731 (MRI measurements). We also thank Konstantin Kurilenko (M.V. Lomonosov Moscow State University, Department of Chemistry) for carrying out the TGA analysis.
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Fig.1. HAADF-STEM overview images and HR-TEM images of MNP-BSA(I) (a, b) and MNPBSA(II) (c, d). The HAADF-STEM images (a, c) clearly demonstrate the difference in the size of the NMP-BSA agglomerates in the fractions I and II. The HR-TEM images (b, d) show crystalline Fe3O4 nanoparticles surrounded by the BSA shell. The panel (e) shows the HAADFSTEM image and EDX compositional maps for NMP-BSA
Fig.2. Cytotoxicity study of MNP-BSA (I), MNP-BSA (II), MNP-BSA-PEG in HF cells.
Fig. 3. Magnetization curves of MNP, MNP-BSA (I), MNP-BSA (II), MNP-BSA-PEG.
Fig. 4. Dox-loading on core-shell-corona: the dependence of the MNP-BSA@Dox-PEG hydrodynamic diameter on the amount of loaded Dox.
Fig. 5. Zeta potential values for MNP-BSA-PEG, MNP-BSA@Dox-PEG with 8% LC and MNPBSA@Dox-PEG with 20% LC
Fig. 6. Dox release profile for different pH-values (7.4, 6.5, 5.5).
Fig. 7. Cytotoxicity study of MNP-BSA@Dox-PEG HEK293 cell cultures.
compared to Dox control on C6 and
Table 1. Physical and chemical properties of MNP-BSA(I), MNP-BSA(II) and MNP-BSA-PEG. Hydrodynamic diameter, nm
Organic shell content , mass. %
Т2-relaxivity, mM-1s-1
ζ-potential, mV
Saturation magnetization, emu/g
MNP-BSA (I)
85±10
28±3
271±6
-24±2
50±3
MNP-BSA (II)
36±4
50±3
161±4
-35±3
20±1
MNP-BSAPEG
41±4
58±5
151±3
-30±3
19±1