Magnetic Field-Enhanced Cellular Uptake of Doxorubicin Loaded Magnetic Nanoparticles for Tumor Treatment

Materials Research Express

PAPER Magnetic field-enhanced cellular uptake of doxorubicin loaded magnetic nanoparticles for tumor treatment

To cite this article: Indu Venugopal et al 2016 Mater. Res. Express 3 095010

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1 2 3 4 Magnetic field-enhanced cellular uptake of doxorubicin-loaded magnetic 5 nanoparticles for tumor treatment 6

7 Indu Venugopal1, Sebastian Pernal1, Alexandra Duproz1, Jeromy Bentley2, Herbert Engelhard3, Andreas 8 1 9 Linninger * 10 11 1. University of Illinois at Chicago, Department of Bioengineering, Laboratory for Product and 12 Process Design, 851 S. Morgan St. - 218 SEO, Chicago, IL 60607-7000 13 2. NSF-RET fellow at University of Illinois at Chicago from Naperville Central High School 440 14 Aurora Ave, Naperville, IL 60540 15 3. University of Illinois College of Medicine, Department of Neurosurgery, 912 South Woods St, 16 Chicago, IL 60612 17 18 * Corresponding author (email: [email protected]) 19 20 21 22 23 1.0 Abstract 24 Cancer remains the second most common cause of death in the US, accounting for nearly 1 out of every 25 26 4 deaths. In recent years, several varieties of nanoparticles (NPs) have been synthesized with the intent 27 28 of being utilized as tumor drug delivery vehicles. We have produced superparamagnetic, gold-coated

29 magnetite (Fe3O4@Au) NPs and loaded them with the chemotherapeutic drug doxorubicin (DOX) for 30 31 magnetic drug targeting of tumors. The synthetic strategy uses the food thickening agent gellan gum 32 (Phytagel) as a negatively charged shell around the Fe O @Au NP onto which the positively charged 33 3 4 34 DOX molecules are loaded via electrostatic attraction. The resulting DOX-loaded magnetic nanoparticles 35 36 (DOX-MNPs) were characterized using transmission electron microscopy (TEM), energy dispersive x- 37 ray spectroscopy (EDS), superconducting quantum interference device (SQUID) magnetometry, surface 38 39 area electron diffraction (SAED), zeta potential (ZP) measurements, fourier transform infrared 40 41 spectroscopy (FTIR) as well as UV/Vis and fluorescence spectroscopy. Cytotoxicity of the DOX-MNPs 42 was demonstrated using the MTT [3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide] assay 43 44 on C6 glioma cells. Cellular uptake of DOX-MNPs was enhanced with magnetic fields, which was 45 quantitatively determined using flow cytometry. This improved uptake also led to greater tumor cell 46 47 death, which was measured using MTT assay. These magnetic drug targeting results are promising for a 48 49 new therapy for cancer. 50 51 52 53 54 55 56 57 58 59 60 CONFIDENTIAL - AUTHOR SUBMITTED MANUSCRIPT MRX-101958.R2 Page 2 of 23

1 2 3 4 2.0 Introduction 5 6 Cancer is the second most common cause of death in the US, exceeded only by heart disease. Cancer 7 8 research has been successful as in decreasing in overall death rates between 1991 and 2010, by about 20% 9 1 10 in both men and women . The cancer treatment protocol for most cancers requires chemotherapy. 11 Unfortunately most chemotherapeutic agents exhibit poor specificity in reaching tumor tissue in addition 12 13 to dose-limiting toxicity. As a result, many patients experience adverse side effects due to the 14 administered chemotherapy treatments. 15 16 17 New drug delivery strategies are being developed to improve efficacy of cancer patient treatments, while 18 minimizing systemic toxicity. One promising treatment involves using NPs as a means of targeting 19 20 cytotoxic drugs to specific tumor cells and tissue, thereby limiting systemic exposure. Typically in the use 21 22 of NPs for the purpose of drug therapy, the drug is dissolved, entrapped, encapsulated, or attached to the 23 NPs. The major goals in the design of NPs for drug delivery include: (i) control of particle size, (ii) 24 25 variation of surface properties, and (iii) ability to release of pharmacologically active agents to achieve 26 the site-specification of the drug at the therapeutically optimal rate and dose regimen. Novel NP-based 27 28 applications for cancer treatments are being explored because of their ability to (i) decrease toxicity and 29 30 the occurrence of adverse reactions, (ii) utilize drug more effectively, (iii) control the rate and site of drug 31 release, (iv) provide a more predictable drug delivery system2–4 32 33 34 Currently, there are two different types of targeting utilized by scientists studying NP drug delivery 35 systems for cancer. Passive targeting exploits the enhanced permeability and retention (EPR) effect that 36 37 is defined by leaky vasculature around tumors caused by tumor-induced angiogenesis with incomplete 38 endothelial cell junctions, resulting in the accumulation of the NPs at the tumor site5. This passive 39 40 targeting strategy has been shown to be an efficient method to reduce the toxic side effects and increase 41 6,7 42 the therapeutic index of drug molecules . Active targeting usually involves attaching ligands on the 43 surface of NPs, which bind to over-expressed receptors on tumor cells. Targeting specific surface receptor 44 45 causes selective uptake of NPs by the cancer cells8. Active targeting may also involve utilizing physical 46 47 stimuli for actively localizing NPs to the target site, including temperature, pH, electric charge, light, 48 sound and magnetism9,10. Our research aims at utilizing magnetism as a physical stimulus, for efficient 49 50 targeting of target specific cells. 51 52 Magnetic Drug Targeting (MDT) involves binding tumor drugs to biocompatible magnetic nanoparticles 53 54 (MNPs), which can be guided to specific tumor sites via magnetic fields. The magnetic field can be 55 controlled by magnets placed outside the body and/or by using magnetizable implants inside the body11. 56 57 58 59 60 2

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1 2 3 NPs composed of magnetite (Fe3O4) are biocompatible and biodegradable. The iron in these particles are 4 5 absorbed into the body’s iron stores during metabolism12. 6 7 Synthesis of magnetic microparticles and NPs for the delivery of chemotherapeutics has evolved since the 8 9 1970s. Zimmerman and Pilwat in 1976 studied the delivery of cytotoxic drugs using magnetic 10 erythrocytes13. The targeting of magnetic albumin microspheres encapsulating the anticancer drug (DOX) 11 12 in animals was studied by Widder et al14. Engelhard and Petruska studied the movement of magnetic 13 15 14 microparticles in rat CSF and brain in 1992 . Hafeli et al synthesized biodegradable polylactide 15 microspheres that incorporated magnetite and the beta radiation-emitter Y-90 for radiotherapy targeting in 16 17 199416, and successfully applied it to subcutaneous tumors17. However, these studies employed 18 19 micrometer-sized magnetic particles. Use of MNPs was described for the first time by Lubbe et al in 20 199918. MNPs loaded with epirubicin were administered intra-arterially in a phase-1 clinical trial for 21 22 patients with advanced cancers in 199618. Several groups and start-up companies have synthesized 23 magnetic vectors and shown potential applications of MNPs. These include MNPs used in MRI19, 24 25 magnetic fluid hyperthermia20,21, cell sorting and targeting22,23, bioseparation24, enzyme immobilization25, 26 26,27 28,29 27 immunoassays , and gene transfection and detection systems . 28 29 Different varieties of MNPs have been coated with a “shell” of either organic or inorganic nature30 to 30 31 exploit specific properties. NP coatings often regulate the solubility, hydrophilic or hydrophobic 32 properties, stability, and the targeting ability of the particles. For example, Yang et al synthesized a poly 33 34 [N-(1-one-butyric acid]) aniline shells around Fe3O4 cores that have been further conjugated with 1,3- 35 bis(2-chloroethyl)-1-nitrosourea (BCNU)31. Ito et al have synthesized antibody-conjugated 36 37 magnetoliposomes which can used as carriers to introduce magnetite NPs into target cells, since the 38 32 39 negatively-charged cell surface interacts with the positively-charged surface of the NPs . Inorganic shells 40 have also been incorporated in MNP synthesis, most notably silver, gold or titanium33–35. 41 42 43 DOX is a chemotherapeutic drug, and its mechanism of action involves intercalating into the DNA of 44 cancer cells and preventing cell replication. DOX is widely used to treat various cancers including 45 46 leukemias, Hodgkin’s lymphoma, and cancers of the bladder, breast, stomach, lung, ovaries, thyroid, soft 47 tissue sarcoma, multiple myeloma, and others. It is an FDA approved drug, currently being marketed 48 49 under the names Adriamycin, RUBEX and DOXIL. Various research groups have loaded DOX on 50 36–41 51 organic or inorganic NP shells for the purpose of better drug delivery . 52 53 In this paper, we demonstrate (i) a new technique of loading DOX on superparamagnetic nanoparticles 54 55 and (ii) enhancement of their uptake into tumor cells in the presence of magnetic fields. The resulting 56 DOX-MNP particles is a promising vehicle for tumor treatment applications via MDT. The magnetite 57 58 59 60 3

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1 2 3 core (Fe3O4) of the DOX-MNP, which was synthesized via a co-precipitation method, was coated with a 4 5 gold shell to form the Fe3O4@Au NPs. The polymer gellan gum was coated on the surface of the 6 7 Fe3O4@Au NP which resulted in a negatively charged NP (Fe3O4@Au-GG) due to the carboxylate groups 8 of the gellan gum42,43. The positively-charged amine groups of DOX molecules are electrostatically 9 10 attracted to the negatively-charged gellan gum surface, which resulted in the final DOX-MNPs. 11 12 We also describe the efficacy of our DOX-MNPs to kill glioma cells, and the enhancement in their 13 14 cellular uptake in the presence of a static magnetic field. Increased endocytosis of plain MNPs via 15 magnetic fields has been previously observed, especially to overcome the hurdle of gene delivery across 16 17 cell membranes44–48. In this paper, we show an elevated level of cellular uptake of the tumor drug loaded 18 19 MNPs and consequently, increased drug efficacy and cell death caused by the presence of magnetic 20 fields. This enhancement makes our drug vehicle an ideal candidate for MDT. The targeting and 21 22 augmented cellular uptake using magnetic fields via our DOX-MNPs can be used to target DOX to 23 specific tumor sites, and thus improve efficacy in vivo. 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60 4

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1 2 3 4 3.0 Materials and Methods 5 6 Equipment utilized for synthesis and characterization include a TR-104 Scale (Denver instrument 7 8 Company), 40 Watt DC Motor with variable electric stirrer, IEC Centra MP4R Centrifuge (International 9 10 Equipment Company), Fisher Vortex Genie 2 (Fisher Scientific Company), TH20 Sonicator (Fisher 11 Scientific), Axioskop Fluorescent Microscope with Axiocam (Zeiss), Zetasizer (Malvern), Fourier 12 13 Transform Infrared spectrophotometer (Bruker Tensor 27 FT-IR & OPUS Data Collection Program), 14 Transmission electron microscope (JEOL JEM-1220), Confocal microscope (Zeiss LSM 510 META), 15 16 NanoDrop 1000 UV-Vis spectrophotometer and NanoDrop 3300 Fluorimeter. 17 18 19 All chemicals were purchased from Sigma-Aldrich® with the exception of DOX (LC Laboratories®) and 20 hydrogen tetrachloroaurate (III) hydrate (Alfa Aesar ®). 21 22 23 3.1 Synthesis of Gold-Coated Magnetite Nanoparticles (Fe3O4@Au NP) 24 25 The Fe3O4@Au NPs were synthesized by a coprecipitation technique, described in Mandal et al. (2005) 26 34,49 27 and various other articles , using ferrous and ferric salts to form Fe3O4. These Fe3O4 cores were then 28 coated with a layer of gold by a surface adsorption technique. The detailed procedure is given below. 29 30 31 Iron Stock Solution Preparation 32 33 6.17 g of iron (III) ammonium sulfate (NH4Fe(SO4)2 •12H2O) and 2.51 g of iron (II) ammonium sulfate 34

35 ((NH4)2Fe(SO4)2 • 6 H2O) were dissolved in 100 ml of 0.40 M sulfuric acid aqueous solution (H2SO4 (aq)) 36 37 to make the iron stock solution. The molar concentrations of the two iron salts in the solution were 0.128 38 M iron (III) and 0.064 M iron (II) respectively. 39 40 41 Fe3O4 Nanoparticle Synthesis 42

43 3 ml of 1 M TX-100 solution was added to 250 ml of 1 M aqueous sodium hydroxide (NaOHaq) solution 44 45 and then heated to 75°C in a water bath. 25 ml of the iron stock solution was added dropwise using a 46 syringe into this solution under vigorous non-magnetic stirring. After complete addition of the stock 47 48 solution, the stirring was continued for 20 minutes. During the entire procedure, the temperature of the 49 50 water bath was maintained at 75 °C. The Fe3O4 magnetic cores obtained and were allowed to settle before 51 magnetic separation to remove the supernatant. The particles obtained were then washed multiple times 52 53 with water to remove the excess surfactant. These cores were then resuspended in distilled water. 54 55 Fe O Nanoparticle Coating with Gold (Au) 56 3 4 57 58 59 60 5

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1 2 3 A solution containing the prepared Fe3O4 cores and hydrogen tetrachloroaurate (III) hydrate was taken in 4 5 a 1:2 molar ratio. In addition, 0.5 g of D(+)-glucose was added to the solution followed by sonication for

6 o 7 15 minutes. This solution was then heated in the water bath at 50-60 C for 1 hour while stirring slowly. 8 The NP solution changed to a brown-red color during the coating process. The solution was vortexed, 9 10 followed by subsequent magnetic separation and washed several times before being resuspended in water. 11 12 3.2 Conjugation of DOX to Fe3O4@Au NP via Gellan Gum 13 14 15 An aqueous solution of gellan gum (PhytagelTM) was prepared by slow addition of 5 mg of PhytagelTM to 16 90 ml of water while heating using a magnetic stirring hotplate. Water was added till the solution volume 17 18 reached 100 ml. The pH of the solution was adjusted to 10 by dropwise addition of 0.16 M NaOH 19 o 20 solution. The solution was heated to 70 C while being stirred. After removing from heat, 1.6 mL of 40 21 mg/ml Fe3O4@Au NP solution was added dropwise to the gellan gum solution while it was being non- 22 23 magnetically stirred. The solution was cooled to room temperature and stirred for 1 hour. The 24 25 Fe3O4@Au NPs were recovered using magnetic decantation and washed 3-4 times with distilled water. In 26 a dark room, 3.0 mg of DOX was dissolved in 1 mL of nanopure water, which was then added dropwise 27 28 to the Au@Fe3O4-GG NP solution, while it was being stirred slowly. The solution was stirred non- 29 magnetically for 24 hours. The resulting DOX-MNPs were magnetically decanted and washed several 30 31 times with water. The DOX-MNPs were then resuspended in water to a concentration of 40 mg/ml. 32 33 34 3.3 Estimation of loading of DOX on the Fe3O4@Au-GG nanoparticles 35 36 A known weight of DOX was added to an aqueous dispersion of Fe3O4@Au-GG NPs of known volume 37 38 and concentration. This was then stirred for 24 hours in the dark. The DOX-MNPs formed were 39 magnetically decanted and the supernatant was collected. Weight of DOX in the supernatant was 40 41 determined by both UV absorbance and well as drying/weighing it. 42 43 The percentage loading of DOX was calculated by using the formula shown in equation 1: 44 45 (Total weight of Dox – Weight of Dox in supernatant) 46 % loading = ∗ 100 − (1) 47 Total weight of Dox 48 49 3.4 Confocal Microscopy and Flow Cytometry 50 51 52 DOX-MNPs of various concentrations were incubated with C6 glioma cells grown on (i) confocal dishes 53 with coverglass bottoms for the confocal microscopy experiment and (ii) cell culture multiwall plates for 54 55 the flow cytometry experiment. After treatment with DOX-MNPs, the cells were thoroughly washed with 56 PBS and fixed by incubation in a 4% paraformaldehyde solution for 20 minutes at 37oC. 0.25mL of a 57 58 59 60 6

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1 2 3 1:2000 concentration solution of phalloidin (Santa Cruz Biotechnology Phalloidin CruzFluor™ 488 4 5 Conjugate) was then added to each well and incubated for 2 hours at 4oC. After the phalloidin was 6 7 removed, 0.25 mL of 2.5 µg/mL concentration of DAPI solution was added to each well for 5 minutes at 8 4oC. After washing with PBS to remove the DAPI, the cells were analyzed using a Zeiss LSM 510 META 9 10 confocal microscope. For the flow cytometry experiment, the cells were removed from the multiwall 11 plates using trypsin and suspended in PBS, before analysis using Beckman Coulter CyAn ADP Analyzer. 12 13 14 3.5 Cell TEM 15 16 DOX-MNP treated cells were scraped from the culture ﬂask and centrifuged at 2000 rpm for 20 min, then 17 18 ﬁxed with 2.5% glutaraldehyde. The cell samples were dehydrated by an ethanol series, and embedded in 19 20 epon. Ultra-thin sections of 83 nm were double contrasted with Reynolds lead citrate and 2% uranyl 21 acetate, and imaged in a JEOL JEM-1220 (fitted with a Gatan digital camera) at an accelerating voltage of 22 23 80 kV. 24 25 3.6 MTT assay 26 27 28 DOX-MNPs of various concentrations were incubated with C6 glioma cells grown on a 96 well plates. 29 30 For the cellular uptake experiment involving magnetic fields, a small circular permanent magnet of 0.33 T 31 was placed directly below specific well during the incubation period. The cells were thoroughly washed 32 33 with PBS to remove DOX-MNPs that were not taken up by the cells. 50µL of MTT stock solution (10% 34 v/v solution of MTT and cell growth media) was added to each well. After incubation for one hour, the 35 36 MTT solution was removed and 200µL of DMSO was added to each well. The plate was then incubated 37 38 for 20 minutes before measuring the absorbance from each well using an optical plate reader (BioTek), at 39 a wavelength of 570 nm. 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60 7

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1 2 3 4 4.0 Results 5 4.1 Synthesis and Characterization 6 7 8 TEM, EDS and SQUID magnetometry results of Fe3O4@Au NPs 9 10 The Fe O @Au NPs were characterized using TEM, EDS and SQUID magnetometry. A histogram 11 3 4 12 showing the size distribution of the Fe3O4@Au NPs is given in figure 1A, which indicates that most of the 13 14 Fe3O4@Au NPs have a diameter of 20-25nm. A particle with this diameter range was found to have a 15 magnetite core of 8-12 nm, with a uniform gold shell thickness of 8-15 nm (figure 1B,C). 16 17 18 The gold coating on the magnetite core serves two purposes: (i) It prevents the oxidation of the Fe3O4 19 core into maghemite (Fe O ) by forming an inert biocompatible protective coating layer. Maghemite is 20 2 3 21 found to be less magnetically susceptible than magnetite50,51 (ii) It also forms an excellent platform for 22 23 conjugating drugs to the NP surface. 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54

55 56 Figure 1. Characterization of NPs (A) TEM image showing multiple 57 Fe3O4@Au MNPs along with a histogram of their size distribution; (B) 58 59 60 8

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1 2 3 Schematic of Fe O @Au NP; (C) TEM image of Fe O @Au NP showing the 4 3 4 3 4 total diameter of the particle to be between 20-25 nm; (D) Schematic of a 5

6 quarter of the Fe3O4@Au NP; (E) Detailed TEM image (similar to the 7 schematic shown in 1D) that clearly indicates the Fe3O4 core and gold coating. 8 This image confirms the length of the core radius (4-6 nm); (F) EDS spectrum 9 of the Fe3O4@Au NPs confirming the presence of elements Au, Fe and O; (G) 10 Graph of magnetization versus applied magnetic field (M-H loop) obtained by 11 SQUID magnetometry at 265 K. No remnant magnetization in the Fe3O4@Au 12 NPs was observed. 13 14 15 The EDS results (figure 1F) confirmed the chemical composition of our Fe3O4@Au NPs, indicating the 16 17 presence of the elements gold (Au) and iron (Fe). The magnetic properties of our NPs was determined 18 using SQUID magnetometry. From the plot of magnetization versus applied magnetic field (figure 1G), 19 20 the saturation magnetization (Ms), remnant magnetization (Mr), and coercivity (Hc) were determined for 21 22 our Fe3O4@Au NPs to be as follows: Ms = 29.5 emu/g; Mr = 0.52 emu/g; Hc = 21.1 Oe at 265 K. The 23 hysteresis loop indicates no remnant magnetization, further proving that our NPs are superparamagnetic. 24 25 Magnetite cores of this size exhibit a superparamagnetic property, retaining no net magnetization when 26 removed from an external magnetic field. They can be dispersed without agglomeration until acted upon 27 28 by an external magnetic field, which attracts the NPs and allows for magnetic guidance and localization at 29 30 a desired target site. 31 32 Figure 2A shows a TEM image of DOX-MNPs, in which the gellan gum polymer coating on the 33 34 Fe3O4@Au NPs can be seen clearly. A histogram showing the size distribution of the DOX-MNPs is 35 given in figure 2B, which indicates that most NPs have a diameter of 95-105 nm (figure 1B). The sizes of 36 37 the DOX-MNPs varied between 75 nm and 150 nm. 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 Figure 2: (A) TEM image showing multiple DOX-MNPs. Inset shows HRTEM images of DOX-MNPs; 54 (B) Histogram of the size distribution of DOX-MNPs. 55 56 57 58 59 60 9

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1 2 3 Due to the variability in the size of the DOX-MNPs, there is a possibility that they are multi-core particles 4 5 rather than single core particles, or a mixture of both types. Encapsulation of magnetic cores using a

6 68 7 polymer matrix can result in multi-core particles , which is similar to our synthesis procedure. 8 9 Zeta potential measurements 10 11 12 Measurement of zeta A B 13 potential was one of 14 15 the techniques used to 16 confirm loading of 17 18 DOX molecules on 19 20 the surface of the 21 Fe3O4@Au NPs. The 22

23 Fe3O4@Au NPs were 24 25 coated with the Figure 3: Characterization of DOX-MNPs (A) Schematic of the synthesis of DOX-MNPs; (B) Fluorescence microscope images of Fe O @Au NPs and the 26 biocompatible 3 4 27 fluorescent DOX-MNPs. 28 polymer gellan gum, a 29 food thickening agent. The carboxylate groups of the gellan gum lend a negative charge to the Fe O @Au 30 3 4 31 NP surface, which electrostatically attracts DOX’s positively charged amine groups. 32 33 34 The positive zeta potential of Fe3O4@Au NPs, +11.6, decreases to -27.3 after coating with the negatively 35 charged gellan gum. After loading the positively charged DOX molecules, the zeta potential increases to - 36 37 15.4, confirming DOX loading. The conjugation of DOX to the vehicle was also confirmed by using 38 fluorescence microscopy. As DOX is a fluorescent molecule, DOX-MNPs fluoresce, emitting a red color 39 40 (at 590 nm) when compared to Fe3O4@Au NPs (figure 3B). 41 42 43 Fourier Transform Infrared (FTIR) spectroscopy results 44 45 Loading of the DOX molecules to the NP was also confirmed using FTIR spectroscopy. Figure 4A shows 46 47 the FTIR spectra of the Fe3O4@Au-GG NPs. The presence of the gellan gum polymer on the Fe3O4@Au 48 NP surface is indicated by the presence of a broad O-H peak at 3400 cm-1, glycosidic bonds are indicated 49 50 by the peak at 1640 cm-1, C-H bending peak at 1409 cm-1 and C=C bond at 1500 cm-1. Figure 4B shows 51 52 the FTIR spectra of the DOX-MNPs (blue) superimposed on the free DOX solution spectra (red). Both 53 spectra were found to have regions with similar peaks (circled in pink in figure 4B) with the same x-axis 54 55 (cm-1) values. This confirms the presence of DOX on the DOX-MNP surface. 56 57 58 59 60 10

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1 2 3 4 5 6 A B 7 8 9 10 11 12 13 14 15 16 17 18 Figure 4: FTIR spectra of (A) Fe3O4@Au-GG NPs; (B) Free DOX (red) and DOX-MNPs (blue). The 19 pink circles indicate the common peaks present in both FTIR spectra, which confirms the presence of 20 the DOX molecules on the DOX-MNP surface. 21 Absorption and Fluorescence Emission Spectra 22 23 24 The stability of DOX molecules after loading on Fe3O4@Au-GG NPs was studied by using fluorescence 25 and UV-Vis spectroscopy. The absorption/excitation spectra of free DOX and DOX-MNPs (green and 26 27 brown lines respectively in Figure 5A) indicate that both have an absorption maximum at 480 nm. The 28 emission spectra of DOX in solution and DOX-MNPs were recorded from 500 to 700 nm at a fixed 29 30 excitation wavelength of 480 nm; the spectra recorded are shown in Figure 5B. There was no change in 31 32 the spectral profile of DOX-MNPs when compared to DOX molecules, and the emission peak at 597 nm 33 was retained. 34 35 36 37 A B 38 39 40 41 42 43 44 45 46 47 48 49 50 51 Figure 5: Spectral characterization of DOX-MNPs (A) Excitation spectra of free DOX and 52 DOX-MNPs showing a peak at 480 nm; (B) Fluorescence emission spectra of free DOX and 53 DOX-MNPs showing a peak at 590 nm. Similar emission spectra indicate that the DOX structure has 54 been retained after being loaded on the Fe3O4@Au-GG NP surface. 55 56 57 58 59 60 11

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1 2 3 4 5 6 Confocal microscopy results 7 8 Confocal microscopy was used to observe the 9 10 cellular uptake of DOX-MNPs in C6 glioma 11 12 cells. The actin cytoskeleton and nuclei of the 13 cells were stained using phalloidin (green) and 14 15 DAPI (blue) respectively; the DOX-MNPs are 16 naturally fluorescent. Peak emission 17 18 wavelengths for phalloidin, DAPI, and DOX- 19 MNPs are 460 nm, 518 nm, and 597 nm 20 21 respectively. Figure 6 shows the confocal 22 Figure 6: Confocal microscopy image of C6 glioma 23 microscopy images that confirm DOX-MNPs cells that have endocytosed DOX-MNPs (red). Actin 24 were taken up into the cells. stained by phalloidin is green, and the nucleus 25 stained by DAPI is blue. (A) Z-stack images are 26 acquired to visualize the depth of the cell and 27 DOX loading efficiency 28 confirm MNP uptake. The orthogonal plane borders 29 the main image; (B) DOX-MNPs localize into the 30 The DOX loading efficiency on the Fe3O4@Au- nucleus; (C) Phalloidin highlights the cell boundary 31 GG NPs was determined to be 70% using the showing the drug is present only inside cells; (D) 32 From top to bottom all three fluorescent markers are 33 procedure mentioned in § 3.3. Our shown (DOX-MNP, DAPI, phalloidin and 34 overlapping DOX-MNP and DAPI) in the Z-axis to measurements showed that 1mg of Fe3O4@Au- 35 confirm colocalization with the cell. 36 GG NPs contain 0.31 mg of DOX loaded on its 37 38 surface, resulting in the final DOX-MNPs. 39 40 4.2 DOX-MNP efficacy studies 41 42 43 MTT assay results 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60 12

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1 2 3 Cytotoxicity of DOX-MNPs for killing tumor cells was tested using MTT assay. The assay is dependent 4 5 on the ability of viable cells to metabolize a water-soluble tetrazolium salt into a water-insoluble purple 6 7 colored formazan product. The intensity of the purple color obtained from the resulting absorbance 8 spectrum, is a direct indicator of the cell viability52. Various concentrations of the DOX-MNPs (4 - 9 10 0.125mg/mL) were incubated with C6 glioma cells cultured in 96 well plates. MTT assay was performed 11 to determine the percentage of non-viable cells, which is an indicator of the efficacy of the DOX-MNPs. 12 13 The experiment was also be repeated 14 A B 15 by placing small permanent magnets 16 of 1.47 lbs pull force strength (0.33 T) 17 18 beneath the wells of the cell plate 19 20 during the incubation period with the 21 DOX-MNPs (figure 8A). Confocal 22 23 microscopy images (figure 7) showed 24 that the presence of magnetic field 25 Figure 7: Uptake of 1mg/ml DOX-MNPs by C6 glioma cells 26 significantly increased the DOX-MNP in 30 minutes (A) without magnetic field; (B) in the presence 27 of 0.33 T magnetic field. Cellular uptake is increased in the 28 uptake, as indicated by the higher presence of the magnetic field. 29 emission of fluorescence from within 30 31 the cells. 32 33

34 35 36 37 38 39

40 A B 41 6.9% 42 7.8 % 43 44 9.3 % 45 10 % 46 13.4 % 47 48 49 50 51 52 53 54 55 56 57 58 59 60 13

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1 2 3 4 C D 5 7.6 % 6 10.1 % 9.3 % 7 10.2 % 8 12.2 % 9 10.8 % 13.6 % 10 16.5 % 11.7 % 11 18.6 % 12 13 14 15 16 17 18 Figure 8: MTT assay results (A) Permanent magnet placed below the cell well enhances DOX-MNP uptake. 19 The arrow indicates the direction of movement of DOX-MNPs towards the magnet; MTT assay results 20 showing the decrease in cell viability in the presence of magnetic field (0.33 T) after incubation periods of (B) 21 1.5 hours (C) 3 hours and (D) 4.5 hours using various concentrations of DOX-MNPs. The difference in the % 22 reductions of cell viability between the control and the magnet experiment have been given in red, above the 23 24 bars of the corresponding experiment. 25 The MTT assay results show a reduction in C6 glioma cell viability with increasing concentration and 26 27 treatment duration using DOX-MNPs, in both the magnetic and no magnetic field exposure (control) 28 groups. Also, the glioma cell viability was reduced with exposure to the magnetic field when compared to 29 30 the no magnetic field exposure (control) group, for the same duration of DOX-MNP treatment. 31 32 33 TEM results 34 35 36 A B C D 37 A 38 39 40 41 42 43 44 45

46 E G H I 47 48 49 50 51 52 53 54 55 56 57 58 59 60 14

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1 2 3 Figure 9: TEM images of cells post incubation with DOX-MNPs (A)(B)(C)(D) Cells were incubated with 4 2mg/ml of DOX-MNPs for half hour. (E)(F)(G)(H) Cells were incubated with 2mg/ml of DOX-MNPs for 5 6 half hour under the influence of a 0.33T permanent magnet field. 7 TEM images of the C6 glioma cells were taken after incubation with 2mg/ml DOX-MNP solution for 30 8 9 minutes. The first and second row of images in figure 9 show TEM images of cells that were incubated 10 11 with and without the influence of a 0.33 T magnetic field respectively. The images in the first row show 12 relatively intact tumor cells with less DOX-MNPs uptake. The images in the second row exhibit higher 13 14 DOX-MNP uptake. These cells also have a reduced cytoplasm volume (figure 9 H,I) and show more 15 16 autophagic vacuoles indicative of waste expulsion (figure 9G). Reduced cytoplasm and presence of 17 autophagic vacuoles have been generally taken as indicators of dying cells53–55. 18 19 20 Flow cytometry results 21 22 23 Flow cytometry was used to determine the cellular 24 uptake ratio of DOX-MNPs in the absence (M-) and 25 26 presence (M+) of a magnetic field. Figure 10 shows 27 28 the relative fluorescence in the C6 glioma cells after Avg M+ 29 incubation with DOX-MNP solutions (concentrations Avg M- 30 31 ranging from 0.25 mg/ml to 2 mg/ml). The results 32 suggested an enhancement in the cellular uptake of 33 34 DOX-MNPs in the presence of the magnetic field. 35 36 fluorescence Relative 37 Presence of the magnetic field also increased the total 0.25 0.5 1 2 38 number of cells that showed fluorescence at the end of DOX-MNP concentration (mg/ml) 39 40 the treatment, when compared to the control group (M- 41 ). For example, when the cells were incubated with a 1 Figure 10: Flow cytometry evaluation of the magnetic 42 enhancement of DOX-MNP uptake. The relative 43 mg/ml DOX-MNP solution, fluorescence was detected 44 fluorescence in the C6 glioma cells at different DOX- 45 in 93% of all M+ treated cells, compared to only 78% MNP concentrations after incubating for 45 min. For all concentrations, cellular uptake is higher in the presence 46 in the case of M- treated cells. Similar increase of 10- 47 of magnetic fields (M+), compared to the control group 48 15% was observed in cell uptake ratio between M+ (M-). 49 50 and M- treated cells for all incubation concentrations of DOX-MNPs. 51 52 53 54 55 5.0 Discussion 56 57 58 59 60 15

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1 2 3 In this paper, we synthesized a tumor drug delivery platform for magnetic targeting. A magnetite core, 4 5 was synthesized to enable active targeting using an external magnetic field. The core was coated with a 6 7 gold shell to form Fe3O4@Au NPs. The gold shell facilitated loading of DOX molecules to its surface by 8 utilizing electrostatic interactions with the polymer gellan gum. The resulting DOX-MNPs were 9 10 characterized to confirm the DOX conjugation onto its surface. 11 12 The individual components of the delivery vehicle are not toxic: gold, Fe3O4 and the gellan gum polymer 13 14 are biocompatible. Fe3O4 degrades into its components which are integrated in normal iron homeostasis in

15 56,57 16 the body . Gold is a stable unreactive metal, which has been used in several NP based drug 17 formulations58,59. Gellan gum is a bacterial anionic deacetylated sugar polysaccharide secreted by 18 19 Pseudomonas elodea that has been found to be biodegradable and biocompatible51,60. 20 21 Due to the variability in the size of the DOX-MNPs, there is a possibility that they are multi-core particles 22 23 rather than single core particles, or a mixture of both types. Both hydrogen bonding and Vanderwall’s 24 25 forces are involved in the formation of the gellan gum matrix structure that encapsulates the magnetic 69 26 core(s) . Due to the superparamagnetic nature of the individual Fe3O4@Au cores, the multi-core particles 27 28 are also superparamagnetic, and this does not affect the particle’s ability to deliver therapeutic molecules. 29 Similar experiments conducted by other researchers have also concluded no loss of in the fast relaxation 30 31 time of SPIONs after being entrapped as multi-core particles in a polymer shell 68. Therefore, these 32 33 particles can be successfully used for magnetic drug targeting in tumor treatment. 34 35 Synthesis parameters such as temperature, reagent concentrations, surfactant concentrations and stirring 36 37 speed can affect the sizes of multi-core particles. Production of multi-core particles with good control of 38 the number of magnetic cores per particle, and of the degree of polydispersity of the core sizes, remains a 39 40 difficult task 70. One of the main factors that we have attempted to optimize is the amount of gellan gum 41 42 added to the system. Other papers, such as Dhar et al used smaller quantity of gellan gum (0.02% w/v) to 43 reduce and cap their gold NPs50,51. In our gellan gum coating protocol, we used a 5% w/v solution to 44 45 accommodate for the larger size of our Fe3O4@Au NPs. The amount of gellan gum added to the system 46 47 was adjusted to obtain maximum DOX conjugation on the surface of the NP, while maintaining an 48 adequate particle size for drug delivery purposes. As per the data from our histogram (figure 2), 81% of 49 50 our DOX-MNPs have a mean diameter of around 100 nm. This size is suitable for both systemic and IT 51 delivery. 52 53 54 Confocal microscopy was used to confirm cellular uptake and nuclear localization of the DOX-MNPs. 55 56 According to Figure 6, the DOX-MNPs were successfully endocytosed into the C6 glioma cells, and the 57 58 59 60 16

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1 2 3 z-stack image of the nucleus (stained blue by DAPI) clearly shows co-localization with the DOX 4 5 molecules (red), which is also the site of therapeutic action for DOX. 6 7 FTIR spectroscopy was also used to confirm loading at every step of the synthesis. From the emission 8 9 spectra, we inferred that the fluorescence signature of DOX is preserved after conjugation on the 10 11 Fe3O4@Au-GG NP surface. Therefore, the DOX molecular structure is retained following loading on the 12 MNPs, which is important for its biological activity. 13 14 15 The efficacy of DOX-MNPs to kill tumor cells was determined by using MTT assay. Cell viability graphs 16 showed that DOX-MNPs killed tumor cells, even over short incubation periods (1.5 hours), when a 17 18 magnetic field is present. The efficacy to kill cells under the influence of a magnetic field, was more 19 20 pronounced in greater DOX-MNP concentration experiments, which indicates that this effect is due to 21 increased endocytosis of DOX-MNPs into the C6 glioma cells. Any aggregation of the DOX-MNPs, 22 23 which might have been caused by application of the external magnetic field, did not adversely affect its 24 25 endocytosis and more importantly, the tumor drug efficacy. 26 27 The enhanced DOX-MNP uptake in the presence of magnetic field was also qualitatively and 28 29 quantitatively determined using TEM and flow cytometry respectively. The TEM images indicated 30 increased destruction of C6 glioma cells in the presence of a magnetic field. Flow cytometry clearly 31 32 quantified the enhancement in both the cellular uptake ratio and the DOX-MNP uptake due to the 33 34 magnetic field. 35 36 We believe that the increased transfection in the presence of magnetic fields can be attributed to the 37 38 magnetization of the NPs as well as an additional attractive force (direction of arrow in figure 8A) 39 generated by the magnetic field, which enhances their translocation through the cell membrane61. This 40 41 force augments the transport velocity of MNPs (in addition to the gravitational force). The velocity of the 42 43 particle in the magnetic field (Umag) depends on the volume of the magnetic material (Vmag), the 44 hydrodynamic radius (r), and the magnetic field gradient (dB/dz) as shown in equation 1. 45 46 푉푚푎푔푀 푑퐵 47 푈 = ∗ (1) 48 푚푎푔 12푟 푑푧 49 50 Here, M is the magnetic field strength and  is the viscosity of the cell growth medium. 51 52 53 It has been shown that larger MNPs are more easily taken up by cells62–65. However, for efficient steering 54 55 of NPs in MDT, the NPs need to be superparamagnetic without possessing remnant magnetization on 56 removal of the steering field66. This property requires very small MNPs with a magnetic core diameter of 57 58 59 60 17

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1 2 3 around 20 nm67. When using smaller MNPs that need to deliver the payload before being cleared from the 4 5 body, an additional force can be deployed to enhance transfection into the target cells. This work shows 6 7 that magnetic fields that are used for guidance in the MDT process have the additional beneficial effect of 8 enhancing cellular uptake of drug conjugated superparamagnetic NPs. Magnetically enhanced 9 10 transfection efficiency of DOX-MNPs, in the presence of magnetic fields is a promising effect for future 11 in vivo MDT studies for treatment of cancer. 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60 18

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1 2 3 4 Conclusion 5 6 In conclusion, our results demonstrate adequate physicochemical properties of our biocompatible custom- 7 8 synthesized DOX-MNPs, which show increased cellular uptake into malignant cells in the presence of an 9 10 external magnetic field. The DOX-MNPs were synthesized and extensively characterized to confirm the 11 conjugation of the DOX on its surface. The efficacy of DOX-MNPs to kill tumor cells was tested and 12 13 proven using rat C6 glioma cells. Permanent magnetic fields were found to increase the endocytosis of 14 DOX-MNPs, and consequently decreased cell viability. Understanding the influence of magnetic fields on 15 16 interaction between the cells and MNPs is very important in its use as a drug delivery vehicle, especially 17 18 for MDT. Although the increase in transfection of our DOX-MNPs has been proven in cell culture in this 19 paper, it has yet to be proven in a 3D tissue model or in an in vivo model. Our future work involves 20 21 coupling the MDT technique with augmented cellular uptake of DOX-MNPs for further in vivo tumor 22 23 drug delivery studies. The attractive forces of magnetic gradient fields on drug conjugated 24 superparamagnetic NPs has the ability to potentiate efficacies, improve kinetics and dose response 25 26 profiles of drugs along with site specific delivery by application of magnetic fields. 27 28 29 Acknowledgements 30 31 We would like to thank the National Science Foundation- Research Experience for Teachers Program for 32 33 providing us with support for this project (Grant- CBET 1010621). We would also like to thank Kevin 34 Tangen and the UIC- Research Resources Center for helping us with the confocal microscopy 35 36 experiments. 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60 19

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