Delivery of Hydrophobic Anticancer Drugs by Hydrophobically

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Delivery of Hydrophobic Anticancer Drugs by Hydrophobically modified Alginate Based Magnetic Nanocarrier Ali Pourjavadi, Shiva Sadat Amin, and Seyed Hassan Hosseini Ind. Eng. Chem. Res., Just Accepted Manuscript • DOI: 10.1021/acs.iecr.7b04050 • Publication Date (Web): 02 Jan 2018 Downloaded from http://pubs.acs.org on January 4, 2018

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Industrial & Engineering Chemistry Research

Delivery of Hydrophobic Anticancer Drugs by Hydrophobically modified Alginate Based Magnetic Nanocarrier Ali Pourjavadi†*, Shiva Sadat Amin†, Seyed Hassan Hosseini‡ †

Polymer Research Laboratory, Department of Chemistry, Sharif University of Technology, Tehran, 11365-9516, Iran

‡

Department of Chemical Engineering, University of Science and Technology of Mazandaran, Behshahr, 01134, Iran Corresponding Author Address: [email protected]: Phone/fax: (982)166165311

ABSTRACT: Since most of the anticancer drugs have low solubility in water, the clinical use of them is limited unless by some modification the solubility increases or the drug is carried with a soluble compartment. To solve this problem, we prepared a magnetic nanocarrier with hydrophobic surface based on oleic acid chains in which the hydrophobic drug molecules such as DOX and PTX are easily adsorbed onto the surface. Since the drug molecules are physically adsorbed on the surface, a large amount of DOX (282 mg.g-1) and PTX (316 mg.g-1) were immobilized onto the nanocarrier. Then, the surface of drug loaded magnetic core was covered by a smart pH-sensitive shell based on sodium alginate. The alginate shell around the magnetic core increased the stability and biocompatibility of drug loaded nanocarrier. The results of drug release profile showed that the drug molecules were released faster in acidic medium than the neutral medium. The MTT assay of nanocarrier showed low toxicity toward MCF-7 and HeLa cells while the drug loaded nanocarrier had high toxicity even higher than the free drugs. Owing to these advantages, the resulting nanocarrier exhibits a promising potential for delivery of hydrophobic drug molecules in clinical applications.

Keywords: Drug delivery; Hydrophobic drugs; Core-shell; Cytotoxicity 1 ACS Paragon Plus Environment

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1. INTRODUCTION Today, cancer has become one of the most fatal diseases which kills many people every year 1

. Among the various treatments for cancer therapy, chemotherapy is one the common ways 2.

Although, the present chemotherapy drugs (such as doxorubicin, paclitaxel, curcumin, camptothecin) are all effective in killing cancer cells, they cannot differ between the cancerous and normal cells 3. So, nonspecific delivery of anticancer agents to the patient's body causes many side effects including organ damages resulted from impaired treatment with lower dose of drug 4. The main challenge in cancer therapy is delivery of anticancer drugs selectively to the cancer cells without interacting with the normal body cells. In last decades, nanocarriers have attracted significant attention in anticancer drug delivery

3b, 5

.

Among the various nanocarriers applied in tumor therapy, magnetic nanocarriers have been more desirable due to their magnetic properties

6

. The anticancer loaded magnetic

nanocarriers can be easily transported to tumor site using an external magnet which reduces the side effects. Moreover, they enhances the contrast of MRI images and can be followed in the body 7. The widespread clinical application of magnetic nanocarriers depends on several parameters such as colloidal stability, cytotoxicity, durability in blood circulation, magnetic intensity and drug release behavior 8. These parameters are strongly connected to the structure of drug loaded magnetic nanocarriers. In this view, the determining factor is the polymer shell around the magnetic nanoparticle core. To date, two types of magnetic nanocarriers have been developed. In the first one, drug molecules are conjugated (chemically bonded or physically adsorbed) to the surface of polymer coated magnetic nanoparticles 9. In the second way, drug molecules are entrapped between the magnetic nanoparticle and polymer shell 10. The simple attachment of drugs to the surface of polymer coated nanocarriers may result to the release of 2 ACS Paragon Plus Environment

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drug molecules in undesirable areas because the dissociation of drug/carrier bond could easily happen

9c

. Moreover, in such systems the release rate of drug is usually high and large

amounts of drug dissociation occur at a short time 11. So, fast release of drug molecule results in high leakage of drug during the transportation and decreases the drug dosage at the target site. On the other hand, coverage of drug loaded magnetic nanoparticles with a smart polymer shell not only protects the drug molecules but also decreases the rate of drug release from the carrier

12

. In these systems, polymer shells should firstly be removed and then the drug

molecules are detached from the surface of magnetic nanoparticles and finally they are released into the target site. Since most of the anticancer drug molecules are insoluble in water medium, they should be delivered to the cancerous cells by attaching to a hydrophilic nanocarrier

13

. The

hydrophilicity of nanocarrier depends on the structure of polymeric shell. Using a negatively charged biocompatible polymer as nanocarrier shell may prevent recognition by the reticuloendothelial system, resulted to long time presence of drug loaded nanocarrier in the blood circulation

14

. The main remained problem here is the attachment of hydrophobic

anticancer drugs to the hydrophilic carrier. Herein, we report a novel magnetic nanocarrier which can be easily loaded with hydrophobic drug molecules. In this system, magnetic nanocarriers are firstly modified by oleic acid chains. This hydrophobic magnetic nanoparticle can easily adsorb the hydrophobic anticancer drugs. Then, the drug loaded magnetic core is covered by the modified alginate polymer. The smart shell is pH-sensitive; therefore in the lower pH medium the alginate shell is removed and drug molecules are released.

2. MATERIALS AND METHODS

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2.1. Materials and instruments. Ferric chloride hexahydrate (FeCl3.6H2O), ferrous chloridetetrahydrate

(FeCl2.4H2O),

ammonium

hydroxide

(30%)

and

3-

aminopropyltriethoxysilane (APTS) obtained from Merck. Oleic acid (OA) and hydrazine monohydrate (100%) were purchased from Aldrich. Sodium alginate (Alg) with medium viscosity was purchased from Chemical Reagent Corp. (Tianjin, China). Doxorubicin hydrochloride (DOX) was obtained from Pfizer, China. Paclitaxel (PTX) powder was purchased from Beijing Norzer Pharmaceutical Co., Ltd. All other materials and solvents were obtained from Merck and double distilled water was used for experiments. MCF-7 human breast carcinoma cells and HeLa cell lines were received from Iran Pasteur Institute, Tehran, Iran. DMEM medium and fetal bovine serum (FBS) were obtained from Biochrom AG, Germany. Dimethyl sulfoxide (DMSO) and [3-(4,5-dimethylthiazol-2-yl)-2,5diphenyltetrazolium bromide] (MTT) were purchased from Sigma. Deionized water was used in all the experiments. UV–vis spectra were collected using a Perkin-Elmer spectrophotometer. FT-IR spectra were recorded on ABB Bommem MB- 100 spectrometer (Canada), KBr was used for making pallet of samples. Thermogravimetric analysis (TGA) was acquired under a nitrogen atmosphere with a heating rate 10 ˚C min-1 using a TGA Q50 thermo-gravimetric analyzer. The XRD pattern was recorded on a RigalcuD/Max-3c X-ray diffractometer. The sizes of samples were observed using transmission electron microscopy (TEM) taken with a Philips CM30 electron microscope. Dynamic light scattering (DLS) measurements were performed by using Zetaplus/ 90plus instrument (Brookhaven Instrument Co., USA). Magnetizations of samples were measured by vibrating sample magnetometer (Meghnatis Daghigh Kavir Co., Kashan, Iran). The PTX concentration was determined using HPLC with C-18 reverse-phase column (Shimadzu, Japan) and acetonitrile/water mixture (ratio 95/5) as mobile phase.


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2.2. Synthesis of magnetic nanocarrier. First, methyl oleate was prepared using esterification of oleic acid in acidic medium. In a round bottom flask, 10 g oleic acid was dissolved in 100 mL methanol and 1 mL H2SO4 was added. The mixture was refluxed for 24 h and then the flask was cooled down. Excess acid was neutralized by addition of NaOH and methanol was evaporated at reduced pressure. Water was added and pure methyl oleate was extracted (MeOA). 1H NMR 15 (500 MHz, CDCl3, δ ppm): 5.32-5.37 ppm (m, 2H); 3.68 ppm (s, 3H), 2,31 (t, 2H, J=7.34); 2.016 (m, 4H); 1.57-1.63 (m, 2H); 1.23-1.37 ppm (m, 23H); 0.89 (t, 3H, J=6.04). Hydrazine oleate was prepared by reaction of MeOA and hydrazine. Typically, 4.0 g MeOA was dissolved in methanol and 5 mL hydrazine was added to the mixture. The mixture was refluxed at 70 ˚C for 24 h. The pure hydrazine oleate (HyOA) was obtained after evaporation of methanol and extraction with water. Alginate dialdehyde (AlgDA) was obtained through the partial oxidation of sodium alginate with HIO4 based on previously reported method

16

. 0.5 g AlgDA was dissolved in 25 mL

water and 10 mL solution of HyOA (0.01 g/mL) in isopropanol was added to the mixture. The mixture was stirred for 6 h at 50 ˚C and a milky solution was obtained. The hydrazine oleate grafted alginate dialdehyde (AlgOA) was precipitated in acetone and washed rapidly with acetone and dried under vacuum at room temperature. Amine functionalized Fe3O4 was prepared based on our previous reported method (MNP@NH2)

9b, 17

. Then MNPs surface was covalently bonded to MeOA. Generally, 1 g

MeOA was dissolved 10 mL methanol and then 0.2 g MNP@NH2 was added to the solution and the mixture was ultrasonically dispersed for 15 min. Then the mixture was refluxed for 24 h until maximum attachment achieve. The OA coated MNPs (MNP@OA) was



magnetically separated and washed several times with methanol (3×15 mL) and dried under vacuum at 50 ˚C. The final nanocarrier was obtained through the coating of MNP@OA by AlgOA. Typically, MNP@OA (0.1 g) was ultrasonically dispersed in 15 mL water/ethanol mixture (1:1). Then, 15 mL solution containing 0.5 g AlgOA in water was dropwise added to this mixture while the flask was vigorously stirred. The final nanocarrier was magnetically collected and upper milky solution mixture was decanted. The magnetic product (MNP@OA@AlgOA) was then washed several times with water (5×15 mL) and ethanol (3×15 mL) and dried at room temperature. 2.3. Drug loading on MNP@OA@AlgOA. DOX was loaded onto the carrier during the preparation of MNP@OA@AlgOA. MNP@OA (3 mg) was ultrasonically dispersed in water/ethanol mixture (2 mL, 1:1). Then, 1 mL doxorubicin hydrochloride (1.2 mg/mL) was added to the solution and the flask was covered with an aluminum foil. After a while, triethylamine was dropwise added to the solution and the mixture stirred for 3 h at room temperature. Then, AlgOA solution was dropwise added to this mixture while the flask was stirred. The final DOX loaded nanocarrier (MNP@OA@DOX@AlgOA) was magnetically collected and upper solution was decanted. The product was carefully washed with water and stored at refrigerator. The amount of free DOX in the solution was determined by measuring the absorbance at 490 nm using a calibration curve. Before DOX measurement in upper solution, the excess AlgOA was removed using a dialysis bag (cutoff 12 kDa). PTX was loaded onto the MNP@OA@AlgOA based on followed method. 1.0 mL PTX (2mg/mL) was dissolved in 10 mL ethanol and then 3 mg MNP@OA was added to the solution. The reaction was stirred for 3 h and then the solvent was evaporated under vacuum. Then, PTX loaded carrier was dispersed in 5 mL water/ethanol mixture (5 mL, 1:1) and AlgOA solution was added to the mixture. The mixture was stirred for 1 h and the final PTX


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loaded nanocarrier was magnetically separated and washed with ethanol and water (MNP@OA@PTX@AlgOA). For determination of loaded PTX, MNP@OA@PTX@AlgOA was dispersed in ethanol/water mixture and the mixture was sonicated for 30 min to complete release of PTX from the nanocarrier. Then, the amount of released PTX was measured by HPLC.

2.4. Drug release experiments. The release of DOX from MNP@OA@DOX and MNP@OA@DOX@AlgOA was investigated at 37 ˚C at pH=5.0 and 7.4 using phosphate buffer. Generally, 3 mg of DOX loaded carrier was dispersed in 5 mL buffer solution and the solution was immersed in a dialysis bag. The dialysis bag contained nanocarrier was then placed in 15 mL of the same buffer solution. The solution was stirred at a constant temperature while the tube was covered by aluminum foil. For each release measurement, 2 mL of solution was separated and analyzed by UV-Vis while 2 mL fresh buffer was added into the tube. Each experiment was triplicated and the DOX release profile was obtained. The release of PTX from MNP@OA@PTX and MNP@OA@PTX@AlgOA was investigated at 37 ˚C at pH=4.5 and 7.4 using PBS buffer. 1 mg of PTX loaded nanocarrier was dispersed in 10 mL buffer solution in a centrifuge tube. The solution was stirred at constant temperature while tube was kept in dark place. At specific intervals, tube was centrifuged and 0.5 mL of solution was separated and tube was refreshed with 0.5 mL fresh buffer. The free PTX in separated solution was extracted with acetonitrile and analyzed by HPLC.

2.5. In vitro cytotoxicity assay. The in vitro cytotoxicity assay was performed on MCF-7 human breast carcinoma and HeLa cells and a MTT assay was used to determine the cell viability of samples. Cells were grown in DMEM containing 5% fetal bovine serum and 100



mg/mL penicillin G and 100 mg/mL streptomycin at 37 ˚C in a humidified 5% CO2 atmosphere. Then, cells were seeded in 96-well plates with a concentration of 8000-9000 cells per wall. The UV sterilized drug nanocarrier was then added to culture wells with different concentrations of DOX or PTX (either as free drugs or drug loaded carriers). The cytotoxicity of MNP@OA@AlgOA was also measured as control cytotoxicity in the same batch. The solution was incubated for 24 and 48 h at 37 ˚C. Cell viability was determined by MTT assay. The cells were washed twice with PBS solution and then 100 µL of MTT solution (0.5 g/L) was added to each well and the plate was incubated for 4 h. In this stage, viable cells reduce the MTT to formazan which is soluble in DMSO. Afterward, the medium containing unreacted dye was removed carefully. The obtained purple formazan crystals were dissolved in 200 mL per well DMSO and the absorbance was measured at a wavelength of 490 nm. All the experiments were performed in triplicate. The following formula was used to calculate the inhibition of cell growth:

Cell viability % =

mean of abs. value of treatment sample × 100 mean of abs. value of control sample

3. RESULTS AND DISCUSSION Magnetic nanocarrier was prepared in a few steps represented in Scheme 1. Magnetic Fe3O4 nanoparticle was used as core of nanocarrier for two reasons: (1) magnetic nanocarriers can be guided to the cancerous sites and (2) potential application of magnetic nanocarriers in MRI. Afterward, the surface of MNP was hydrophobically modified by attachment of OA. This hydrophobic surface easily adsorbs hydrophobic molecules from the water medium. However, non-polar interaction of OA chains with non-polar molecules is weak and adsorbed molecules can easily release in the non-polar medium. The unwanted release of adsorbed molecules can be decreased by coating by a smart polymeric shell. This smart shell was 8 ACS Paragon Plus Environment

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prepared by attachment of OA to sodium alginate. Sodium alginate has a long history of use in various biomedical applications 18. Since, sodium alginate has negative carboxylate groups, coating of nanoparticles with sodium alginate increases the solubility of carrier 19. However, coating of MNP@OA (hydrophobic surface) with sodium alginate (hydrophilic) is not possible and for this reason sodium alginate has to be modified. This modification was introduced to sodium alginate backbone by partial oxidation of sodium alginate and conversion to sodium alginate dialdehyde (AlgDA)

16b

. Then, AlgDA was reacted with

hydrazine oleate and a graft copolymer was obtained through the reaction of hydrazine group of hydrazine oleate with aldehyde part of alginate (AlgOA). The resulted hydrazone bond is pH-sensitive and can easily be dissociated in acidic medium, while it is stable in neutral condition. Addition of AlgOA to drug loaded MNP@OA resulted to a nanocarrier in which hydrophobic OA chains of AlgOA and MNP@OA assembled together as shown in Scheme 1. The outer layer of nanocarrier contains hydrophilic sodium alginate and middle layer contain hydrophobic OA chains which hydrophobic drug molecules are embedded here.



Scheme 1. Preparation of MNP@OA@Drug@AlgOA.

1

H NMR spectra of HyOA (a), NaAlg (b), AlgDA (c) and AlgOA (d) are shown in Figure 1.

The 1H NMR spectrum of HyOA confirms the structure of hydrazine oleate. The 1H NMR spectrum of AlgDA, which is completely matched with the reported spectrum of alginate dialdehyde

16b

, confirms that the partial oxidation of alginate has been successfully done.

Finally, the 1H NMR spectrum of AlgOA shows both characteristic peaks of AlgDA and HyOA at chemical shift ranges of 3.3-6 and 0.8-2.4 ppm, respectively. It is worth noting that the molecular weight of AlgOA could not be measured by GPC because the grafted polymer contain both hydrophilic and hydrophobic polymer chains and there was no suitable solvent or mixed solvent for complete dissolving of AlgOA.

Figure 1. 1H NMR of hydrazine oleate (a), sodium alginate (b), alginate dialdehyde (c) and AlgOA (d). 10 ACS Paragon Plus Environment

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The structure of nanocarrier was confirmed by FT-IR instrument (Figure 2a). The FT-IR spectrum of MNP shows characteristic peak of Fe-O at 580 cm-1. After immobilization of OA onto the surface of MNPs the strong stretching vibration band of C-H was appeared at 2919 and 2851 cm-1. The FT-IR spectrum of AlgOA shows characteristic bands of both OA and sodium alginate, confirming the successful grafting of OA to Alg. The broad peak at 15901660 cm-1 is attributed to stretching vibration of various carbonyl groups in the structure of AlgOA. Because of the several interaction of polymer functional groups the stretching vibration of carbonyl groups cannot be recognized separately. In the FT-IR spectrum of final nanocarrier, the characteristic bands of Fe-O, C-O, C=C, C=O and C-H are clearly observed, confirming the successful coating of MNP@OA with AlgOA. In this spectrum a weak peak at 1657 cm-1 can be attributed to free aldehyde groups of Alg backbone. Moreover, in the FTIR spectrum of MNP@OA@AlgOA, the observed peaks of carbonyl groups at 1590-1660 cm-1 are sharper than bulk AlgOA which can be attributed to this fact that lower polymer interaction occurs at the surface of nanocarrier due to the low concentration of polymer at the surface. Figure 2b shows the FT-IR spectra of MNP@OA@AlgOA (I), free DOX (II), MNP@OA@DOX@AlgOA (III), free PTX (IV) and MNP@OA@PTX@AlgOA (V). The FT-IR spectra of both MNP@OA@DOX@AlgOA and MNP@OA@PTX@AlgOA clearly show the characteristic peaks of free drugs which prove the successful loading of DOX and PTX onto the MNP@OA@AlgOA.



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Figure 2. FT-IR spectra of (a) MNP, MNP@OA, AlgOA and MNP@OA@AlgOA; (b) MNP@OA@AlgOA (I), Free DOX (II), MNP@OA@DOX@AlgOA (III), Free PTX (IV) and MNP@OA@PTX@AlgOA (V).

Thermal gravimetric analyses (TGA) of samples show the amount of organic components in each sample (Figure 3a). The TGA curve of MNP (I) showed a weight loss up to 150 ˚C which is attributed to evaporation of physically adsorbed water and other volatile solvents. APTS coated MNPs (II) shows a main weight loss around 300 ˚C which is corresponding to degradation of propylamine groups of MNP@NH2. The thermal degradation curve of MNP@OA (III) shows 6.1 wt% weight losses at 200-400 ˚C. For MNP@OA@AlgOA (IV), more degradation was observed and 10.4 wt% of sample was degraded at 200-450 ˚C. Moreover, the thermal degradation pattern of MNP@OA@AlgOA is different than MNP@OA which proves the presence of sodium alginate shell. These results confirm the successful preparation of nanocarrier. The DOX and PTX loaded nanocarriers show higher weight

loss at lower temperature

than MNP@OA@AlgOA.

The

TGA

curves

MNP@OA@DOX@AlgOA (V) and MNP@OA@PTX@AlgOA (VI) showed 39 and 42.6 12 ACS Paragon Plus Environment

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wt% weight losses, respectively. These weight losses are corresponding to drug loading of 286 and 322 mg/g for DOX and PTX loading, respectively. The XRD pattern of final nanocarrier (Figure 3b) shows characteristic peaks and relative intensities of Fe3O4 which completely match with standard magnetite (blue lines). Figure 3c shows

the

vibrating

sample

magnetometer

(VSM)

of

MNP,

MNP@OA

and

MNP@OA@AlgOA. As seen, there is no observed coercivity in all samples which shows the superparamagnetic behavior of samples. The pure MNP shows a saturation magnetization around 64.1 emu.g-1 which after coating of MNP with OA its saturation magnetization decreased to 52.3 emu.g-1. More decrease in saturation magnetization occurs when AlgOA was coated onto the MNP@OA. These results demonstrate that coating of MNPs was successfully done. Also, the saturation magnetization of the final nanocarrier is enough large that can be guided in body with an external magnet. Another important aspect of an applicable nanocarrier is surface charge of carrier which determines the recognition by the reticuloendothelial system, colloidal stability and cellular uptake of nanocarrier. Zeta potential analyses of all samples were measured (Figure 3d) at pH=7.4 and the results were as followed; bare MNP (-18.9), MNP@NH2 (+25.1), MNP@OA (+14.9), MNP@OA@AlgOA (-21.4), MNP@OA@DOX (+26.2), MNP@OA@PTX (-10.0) MNP@OA@DOX@AlgOA (-11.7) and MNP@OA@PTX@AlgOA (-31.6). The results showed that after each modification the surface charge of nanocarrier changed which confirms the successful modification in each step.



Figure 3. (a) TGA curves of (I) MNPs, (II) MNP@NH2, (III) MNP@OA, (IV) MNP@OA@AlgOA, (V) MNP@OA@DOX@AlgOA, (VI) MNP@OA@PTX@AlgOA; (b) XRD pattern of MNP@OA@AlgOA; (c) VSM analysis of samples; (d) and zeta potential of samples.

The TEM image of prepared Fe3O4 shows dark nanoparticles with average size of 9 nm (Figure 4a). On the other hand, MNP@OA@AlgOA shows a gray shell around the MNPs which is attributed to polymer shell. The average size of these nanoparticles was estimated 25 nm (Figure 4b). The particle size distribution of MNPs and MNP@OA@AlgOA demonstrated that the mean size of particles were 9 and 25 nm, respectively (Figure 4c,d). However, the size of nanocarrier obtained by DLS analysis is much larger than the results of TEM analysis (Figure 4e). It is worth noted that the observed DLS particle sizes are usually larger than real particle sizes which is due to the fact that DLS analyzer measures the solvated 14 ACS Paragon Plus Environment

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nanoparticles in aqueous solution. These results are actually in agreement with previously published paper

20

. The peak at around 1000 nm in DLS analysis may resulted from some

aggregation of particles or remained AlgOA (in the form of micelles) at the solution. However, the concentration of particles with 1000 nm diameter is low.

Figure 4. TEM images of MNPs (a) and MNP@OA@AlgOA (b); particle size distribution diagram of MNPs (c) and MNP@OA@AlgOA (d); DLS analysis of MNP@OA@AlgOA and drug loaded carrier (e).

Figure 5 shows magnetic dispersion of MNP@OA,MNP@OA@AlgOA and drug loaded samples in various conditions. As it can be seen, MNP@OA was well dispersed in water/ethanol (50/50) mixture. But, MNP@OA is highly hydrophobic and aggregated at the top of the aqueous solution. The high hydrophobicity of magnetic core resulted from the OA chains around the Fe3O4 nanoparticles which lead to high adsorption of hydrophobic molecules from the solution. After coating of AlgOA around the magnetic hydrophobic core, 15 ACS Paragon Plus Environment


nanocarrier MNP@OA@AlgOA was easily dispersed in solution using an ultrasonic. The dispersion was even stable for more than 5 h. This dispersibility came from hydrophilic nature of alginate shell in outer layer of MNP@OA@AlgOA. On the other hand, MNP@OA@AlgOA was precipitated at acidic medium after a while which demonstrated that in an acidic medium Alg shell was removed from the surface of MNP@OA. In acidic solution, the hydrazone bond of hydrazine oleate and alginate dialdehyde was hydrolyzed and alginate shell is free and can be easily dissolved. The result of dispersion showed that the drug loaded nanocarriers were both well dispered in aqueous solution. It was also observed that the final nanocarrier is highly responsive to external magnetic field; so it can be potentially applicable as MRI enhancement agent.

Figure 5. Dispersion of samples in various conditions.

The UV-Vis spectrometer was applied to confirm the presence of drug molecules in the structure of magnetic nanocarrier (Figure 6I and II). The maximum loading of DOX in the 16 ACS Paragon Plus Environment

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nanocarrier was obtained by mixing of various amounts of DOX and nanocarrier. For this reason, various DOX loaded nanocarriers were synthesized by changing the DOX concentration. Each experiment was triplicated and error bars were calculated. The drug loading capacity (DLC) and drug loading efficiency (DLE) was calculated by followed equations:

DLC =

DLE % =

! − ∑ $! %& ! − $! × 100 !

Where MID is the initial DOX mass, MRD is the residual DOX mass in the supernatant solutions and MNC is the nanocarrier mass. The results of DLC and DLE are presented at Figure 6III (The drug concentrations are calculated based on one mg of carrier). The results show that the highest DOX capacity in nanocarrier is 282.3±11 mg.g-1 and in this concentration more than 70% of DOX was adsorbed by nanocarrier. The DLC and DLE were also calculated for loading of PTX on nanocarrier (Figure 6IV). It was found that the maximum loading of PTX on nanocarrier was around 316.4±30 mg.g-1 with loading efficiency of 79% which both DLC and DLE were higher than those for DOX loading. It may be resulted from higher hydrophobicity of PTX than DOX. These results were in good agreement with the results of drug loading amounts obtained from TGA curves. The results prove that the hydrophobic surface of nanocarrier can be efficiently loaded by hydrophobic molecules and a high loaded drug carrier can be achieved by this protocol.



Figure 6. (I) UV-Vis spectra of DOX and (II) PTX loaded nanocarrier; (III) DLC and (IV) DLE curves for DOX and PTX loading on nanocarrier.

DOX release profiles of nanocarriers (with and without AlgOA shell) were then examined at two different pH 5.0 and 7.4 (Figure 7). In the absence of AlgOA shell, MNP@OA@DOX shows a slow release of DOX in both pH=5.0 and 7.4 and more than 70% of DOX released into the solution at 50 h. The slow release of DOX is attributed to this fact that the DOX is poorly soluble in aqueous solution. At pH=5.0 protonation of DOX occurs and therefore a faster release is expected. However, it is clearly observed that the release at pH=7.4 was not very slower than pH=5.0 for MNP@OA@DOX in the absence of any shell. As expected, after coating of AlgOA shell around the MNP@OA@DOX the drug releases became slower at both pHs. The results of DOX release from MNP@OA@DOX@AlgOA showed that after 24 h about 50% of DOX released from the carrier at pH=5.0 while at the same time the DOX release at pH=7.4 was lower than 8%. Even after 50 h, only 27% of 18 ACS Paragon Plus Environment

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DOX was released from the MNP@OA@DOX@AlgOA at pH=7.4, while this value at pH=5.0 was 84%. The results show the strong barrier effect of alginate shell. Since, the release of DOX from MNP@OA@DOX@AlgOA is depends on protonation of carboxylate groups of alginate and dissociation of hydrazone bonds of AlgOA, shell is faster removed at acidic

medium.

These

results

confirm

that

the

release

of

DOX

from

MNP@OA@DOX@AlgOA is highly depends on pH of the solution. This option can be very useful for in vivo examination, where the drug release should be minimal at normal tissues (pH=7.4) to reduces the negative side effects of drugs. The

PTX

release

was

also

investigated

from

MNP@OA@PTX

and

MNP@OA@PTX@AlgOA at two pHs (7.4 and 4.5). The results showed that PTX release from the carrier is slower than DOX release even in more acidic medium. At pH=7.4 the PTX loaded carrier with and without AlgOA shell showed a slow drug release after 72 h. But in the more acidic medium, PTX released faster from the nanocarrier without shell and maximum 65% of PTX released from the MNP@OA@PTX. The release of PTX from MNP@OA@PTX@AlgOA was slowly occurred and after 24 h about 15% of drug was released, but after that the drug release rate became faster and 47 % of PTX released after 72 h. Same release profile was observed for MNP@OA@DOX@AlgOA in acidic medium which can be attributed to removal of AlgOA from the magnetic core. All the results prove that MNP@OA@DOX@AlgOA and MNP@OA@PTX@AlgOA were pH-sensitive and higher concentration of the drug molecules released from the magnetic hydrophobic core in acidic medium.



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Figure 7. DOX and PTX release profile from drug loaded MNP@OA and MNP@OA@AlgOA.

The cytotoxicity of MNP@OA@AlgOA and MNP@OA@DOX@AlgOA were compared with free DOX (at the same concentrations) by MTT assay against MCF-7 cancer cells (Figure 8a,b). The samples were incubated for 24 (a) and 48 h (b) and each experiment was triplicated. As expected, the results showed that free DOX is highly toxic toward the cancerous cells and the cell viability dramatically decreases with increasing in the DOX concentration and incubation time. Hopefully, MNP@OA@AlgOA without any drug showed low toxicity and more than 70% of cells survived after 48 h incubation. On the other hand, MNP@OA@DOX@AlgOA, showed very high toxicity in higher concentrations, interestingly it had higher toxicity than free DOX at same concentrations. This can be attributed to this fact that Fe3+ is slightly toxic for cells due to the tendency of oxygen capture from the cells. These results are in agreement with the previously reported results 20 ACS Paragon Plus Environment

21

. The

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toxicity of MNP@OA@AlgOA and MNP@OA@PTX@AlgOA were also examined toward HeLa cells at 24 and 48 h (Figure 8c,d). The results showed that MNP@OA@AlgOA had no significant toxicity on HeLa cells; even lower toxicity was observed compared with MCF-7 cells. On the other hand, high toxicity was observed when HeLa cells were treated with MNP@OA@PTX@AlgOA. All the results demonstrated that MNP@OA@DOX@AlgOA and MNP@OA@PTX@AlgOA can be potentially applicable for cancer therapy.

Figure 8. Cytotoxicity assay of free DOX, MNP@OA@AlgOA and MNP@OA@DOX@AlgOA toward MCF-7 cancer cells incubated for 24 h (a) and 48 h (b). Cytotoxicity assay of free PTX, MNP@OA@AlgOA and MNP@OA@PTX@AlgOA toward HeLa cells incubated for 24 h (c) and 48 h (d).



4. CONCLUSION In summary, we have successfully designed a novel type of magnetic nanocarrier which contains a hydrophobic surface for adsorption of hydrophobic drug molecules. The designed magnetic carrier could be loaded with large amounts of hydrophobic anticancer drugs DOX and PTX. The coverage of alginate shell around the hydrophobic core improved the stability and biocompatibility of final carrier. Moreover, in more acidic medium the smart alginate shell was fall off from the surface of nanocarrier and the drug molecules released faster in this situation, while the alginate shell was stable in the neutral condition. The results of MTT assay shown that the magnetic nanocarrier was almost safe while the DOX or PTX loaded carriers were highly toxic for cancerous cells. We believe that this new magnetic nanocarrier has the potential for in vivo cancer therapy and various drug molecules can be delivered by this method. Moreover, the magnetic property of nanocarrier makes it possible to use it as MRI enhancement agent.

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Delivery of Hydrophobic Anticancer Drugs by Hydrophobically

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