1. Introduction Breast cancer is the most frequently diagnosed cancer in women and is ranked as the second cause of death after lung cancer, so there are a large number of studies performed to find effective treatments. Chemotherapy is the most useful method for treatment of metastatic cancers. Presently, the combination of chemotherapeutic agents has been developed for breast cancer therapy [1-4].

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Among the most active anticancer agents, doxorubicin is widely used in treatment of solid tumors and leukemia [1] with cytotoxic, cytostatic and anti-neoplastic effects. It works by attacking cells that grow quickly such as cancer cells, but because of its non-selective action in induction of cell death several side effects such as bone marrow depression, reduced immunity, cardiovascular toxicity are arisen [5]. Specific and targeted drug delivery can decrease their toxic effects on healthy cells and also increase their performance. So much research has been done for targeting doxorubicin to decrease its side effects, increase its toxicity in targeted tissues and deliver the drug exclusively [6-10]. Peptide receptors mostly expressed on tumor cells are used for selective drug delivery to malignant tumors. Luteinizing Hormone-Releasing Hormone (LHRH) which is produced in hypothalamus gland is a decapeptide hormone that regulates pituitary-gonadal axis affecting reproduction. LHRH receptors are over expressed in breast, ovarian, endometrial and prostate cancers but these receptors do not increase significantly in healthy organs [11, 12], and can be used as targeting agents to increase the absorption of anticancer drugs to LHRH positive tumor cells. Many attempts have been made to use LHRH and its analogues as targeting agents for nanoparticles [13-20]. In Schally and Nagy’s work [13, 14] LHRH and its analogues were used as vectors of cytotoxic drugs and cisplatin and doxorubicin were conjugated to them. All of these conjugated drugs had better performances and anti-proliferative effects compared to free drugs [15-17]. Drugs that are conjugated with LHRH analogues have shown to improve specific affinity to LHRH receptors and are taken up by the malignant cells [18-21]. This ligand

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may also be conjugated to colloidal systems or nanoparticles with high capacity of drug delivery. Minko et al. [20] produced the poly-amidoamine dendrimers that were targeted by LHRH conjugated paclitaxel. Taheri et al. [22] used albumin conjugated methotrexate that was then targeted with LHRH for specific treatment of breast cancer. Another method for improving the anti-tumor selectivity and toxicity effect of cytotoxic agents is the use of magnetic carriers [23-27]. Super paramagnetic iron oxide nanoparticles (Fe3O4) with core size between 10-100 nm are powerful targeted delivery vehicles in various biomedical applications [24]. These particles have organic or inorganic coating, on or within which a drug may be loaded and they are delivered by an external magnetic field to their target tissue. If the external magnetic field is removed, they do not exhibit any residual magnetic interaction at room temperature. Therefore, they do not agglomerate or being uptake by phagocytes and after injection they remain in circulation and pass through the capillary systems of organs and tissues without any vessel embolism and thrombosis [25, 28]. Recently, various anticancer drugs including paclitaxel, methotrexate, mitoxantrone and doxorubicin have been conjugated with magnetic nanoparticles to enhance tumor targeting [27, 29-31]. In another research magnetic nanoparticles targeted by LHRH have been used for diagnosing and drug delivery to breast cancer. This is due to their capability to be recognized by magnetic field and MRI and also their ability to improve the cytotoxic effects of drug by producing local heat. In this regard gold nanoparticles have been targeted by triptorelin, the LHRH analogue, and then have been loaded by doxorubicin [32]. These nanoparticles have been used with the 100-200 times lower concentration than free drug to show the same toxicity effect. In another

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application, targeted iron oxide nanoparticles with LHRH, have been used for diagnosing of metastatic breast cancer [33]. Kakar et al. [34] studied the cytotoxicity effects of LHRH conjugated Fe3O4 on MCF-7 and MDA-MB231 breast cancer cell lines. The aim of the present study was to produce a dual targeted drug delivery system for more effective treatment of breast cancer by using magnetic nanoparticles loaded with doxorubicin which were able to be localized in target tissues and internalized within the desired cells by application of an external magnetic field. Behdadfar et al. [35] reported that Zn-substituted magnetic nanoparticles have more heat dissipation capacitance compared to commercially available nanoparticles. These nanoparticles were successfully used in production of magnetic chondroitin targeted nanoparticles for targeting of montelukast in prevention of in-stent restenosis [36]. Therefore, in the present study Fe-Zn magnetic nanoparticles were used to be localized in the desired site of tumor by an external magnetic field and produce local heat while simultaneous releasing the loaded drug. Due to the over-expression of LHRH receptors in breast cancer cells and low or none expression in normal cells [11, 12], the goal of this attempt was an improved targeted drug delivery to tumor site by conjugating LHRH with nanoparticle vehicles. The poly (methyl vinyl ether maleic acid)/chitosan copolymer was developed for functionalization by LHRH and used to trap the magnetic nanoparticles and doxorubicin.

2. Materials and methods

2.1. Materials Chitosan, ferric nitrate [Fe(NO3)3.9H2O], ammonium hydroxide 25%, zinc chloride (ZnCl2), dichloromethane, citric acid (C6H8O7.H2O), poly )methyl vinyl ether-co-maleic acid(, N-hydroxysuccinimide

(NHS), glacial acetic

acid,

1-ethyl-3-(3-dimethylaminopropyl

carbodiimide (EDC), polysorbate 80 (Tween® 80), acacia gum, sodium alginate, bovine serum albumin (BSA), anhydrous dimethyl sulfoxide (DMSO), and Triton X100 were purchased from Merck Chemical Company (Germany). Triptorelin acetate as the LHRH analogue was purchased from Tofigh Daru Research and Engineering Co. (Tehran, Iran) and chondroitin sulfate from Farabi Pharmaceutical Company (Iran). Carbomer P934, Pluronic® F-68, Dulbecco™ phosphate buffered saline (PBS), 0.25% trypsin, and 3-,5-dimethylthiazol-2-yl]-2,

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5-diphenyl tetrazolium bromide (MTT) were from (Sigma-Aldrich, US) and Soluplus® from BASF (Germany). Sodium hydroxide was provided from Panreac Quimica Co., Ltd. (Barcelona, Spain). Doxorubicin HCl was purchased from Hangzhou ICH Biopharm Co., Ltd. (Zhejiang, China). FITC labeled LHRH receptor antibody was obtained from Biorbyt (Cambridge, UK). Deionized water freshly purged with nitrogen gas was used in all steps of the synthesis of magnetic nanoparticles and preparing all aqueous solutions. MCF-7 (LHRH receptor positive) and SKOV3 (LHRH negative) human cancer cell lines were obtained from the Pasteur Institute (Iran). RPMI-1640 medium was from PAA (Austria) and penicillin-streptomycin mixtures 50 IU/ml from GIBCO Laboratories (Scotland). All the chemicals and reagents were used without further purification.

2.2. Magnetic Fe-Zn nanoparticles Zn-substituted iron oxide nanoparticles were synthesized by the method reported before by Behdadfar et al. [35] via single step hydrothermal reduction route in the presence of citric acid as a reducing agent and stabilizer. Briefly, ferric nitrate [Fe(NO3)3.9H2O] and zinc chloride (ZnCl2) were dissolved in distilled water. After 10 min stirring, a 25% solution of NH4OH was added slowly to reach a pH medium of 9. Vigorous stirring continued for another 10 min and a reddish brown slurry was formed. The slurry was then centrifuged and washed three times with deionized distilled water to remove excess ions and reach a pH medium of 7. Citric acid (CA) was then added to the mixture and it was stirred vigorously for another 10 min and then transferred into a 500 ml volume teflon-lined autoclave. The autoclave was kept at 180°C for 20 hours and then free-cooled to room temperature. Ferrofluids were prepared by washing the precipitate with acetone, mixing with 10 ml of deionized distilled water and sonicated for 10

6

min in an ultrasonic bath. The black suspensions were centrifuged for 10 min at 6000 rpm. The supernatants were the desired ferrofluids and the precipitation was discarded. For powder characterizations, the precipitate was washed with acetone via magnetic decantation several times and then freeze dried at 70°C and 0.2 mbar pressure. Substitution of Zn affects critically the magnetic properties and crystallinity of the synthesized nanoparticles and hence their heating efficiency. This is a cheap, facile, low energy and environmental friendly method, which leads to the formation of ferrofluids with high intrinsic loss powers, so nanoparticles prepared by this route have potential applications in magnetic hyperthermia [35]. Therefore, the magnetic nanoparticles used in this work were Zn-substituted MNPs coated by CA that were dispersed in water by the concentration of 49 mg/ml. The presence of citrate ions on the surface of the magnetic nanoparticles was confirmed by infrared spectroscopy (FTIR) and quantified by TGA (thermogravimetric analysis) and DTA (differential thermal analysis) as reported before in our previous work [1](B. Behdadfar, A. Kermanpur, H. Sadeghi-Aliabadi, M. d. P. Morales, M. Moza_ari, Synthesis of aqueous ferrouids of zn< sub> x fe< sub> 3- x o< sub> 4 nanoparticles by citric acid assisted hydrothermal-reduction route for magnetic hyperthe). The crystallinity of the Zn-substituted MNPs were obtained by the X-ray diffraction (XRD), also differential scanning calorimetric analysis (DSC). The morphology and uniformity in size and shape of the MNPs were evaluated by TEM (Zeiss, EM10C, Germany). The results are shown in our previous work [35].

2.3. Preparation of poly (methyl vinyl ether maleic acid) (PMVMA)/chitosan copolymer

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The schematic representation of the chemical reactions for synthesis the Chi-PMVEMA copolymer is depicted in Fig. 1. Firstly, PMVMA was activated using NHS and EDC. For this work PMVMA (0.4 mmol, 40 mg) was dissolved in acetate buffer (pH 5.5). Then the solution of NHS (0.6 mmol, 67 mg) and EDC (0.6 mmol, 111 mg) in phosphate buffer (pH 7.4) were added and reacted for 3 h at 4°C. The resultant solution was added to chitosan solution (0.025 or 0.05 mmol, equal to 250 or 500 mg) in 20 ml of distilled water (pH 4-6) and the reaction was further stirred at room temperature for 24 h. The molar ratio of chitosan and PMVMA copolymer were 4:1. The reaction mixture was centrifuged and supernatant was removed and dialyzed for one day against phosphate buffer (pH 7.4), one day against acetate buffer (pH 6) and two days against distilled water. The resulting product was freeze dried and kept in dry state for further studies. The chemical structure of the synthesized copolymer was characterized by FTIR and 1HNMR spectroscopy. (1HNMR) spectra were recorded on Bruker Biospin AC-80 400 MHz 1HNMR spectrophotometer (Germany) and Fourier transform infrared (FTIR) spectra were obtained by FTIR spectrophotometer (Rayleigh, WQF-510/520, China).

2.4. Conjugation of LHRH peptide to (PMVMA)/chitosan copolymer After production of the copolymer different ratios of LHRH were conjugated to the copolymer and their particle size was examined by Zetasizer (Zetasizer-ZEN 3600 Malvern Instrument Ltd., Worcestershire, UK) based on dynamic light scattering principle technique. For this purpose 25 mg of (PMVMA)/chitosan copolymer was dissolved in 1 ml of deionized water and the pH of the solution was adjusted to 4.5 by glacial acetic acid while stirring on a magnetic stirrer (IKA PT Lopower, Germany). Then different amounts (2.5, 5, and 10 mg) of LHRH were added to the copolymer solution. Separately, EDC in equal weights of LHRH was

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dissolved in 1 ml of deionized water and added to the solution of the copolymer. The resulted solution was kept for 15 hours in dark at 4-10°C while stirring on a magnetic stirrer (IKA PT Lopower, Germany). Then it was centrifuged (Microcentrifuge Sigma 30k, UK) at 10000 rpm for 15 min in micro-centrifuging filter tubes (Amicon Ultra, Ireland) with 10 kDa molecular weight cutoff to separate the un-reacted LHRH and EDC. The copolymer-LHRH conjugate was separated and lyophilized (Christ, 24 LDPlus, Germany) for further application. The successful conjugation of the copolymer to LHRH was confirmed by FTIR spectroscopy (JASCO, FT/IR-6300, Japan). Data was acquired in range of 400-4000 cm-1.

2.5. Coupling efficiency of LHRH to the chitosan/PMVMA copolymer Capillary electrophoresis method (Agilent 7100 capillary electrophoresis system, Agilent Technologies, Germany) with UV detector, was used fordetermining the coupling efficiency of LHRH to chitosan/PMVMA by Robin et al. method [37] with some modifications. The analysis was performed at 215 nm using a fused silica capillary with the internal diameter of 50 m, an effective length of 40 cm and a total length of 48.5 cm. The running buffer was formic acid 75 mM with the pH adjustment on 3 by ammonium hydroxide. The samples were injected at a pressure of 50 mbar for 5 seconds (hydrodynamic injection). The temperature was kept at 21±1C and the separation potential was +15 KV [37]. All samples, buffers and solutions were filtered through 0.22 m syringe filters before use. The total analysis time was 8 min. Agilent chem. Station Rev. B.04-02 software was used to perform integration of the peaks. A standard curve using 6 different concentrations (12.5, 25, 50, 62.5,

9

100, 125 g/ml) of LHRH in deionized water was initially plotted using the same method as above, prior to LHRH assay. To find the LHRH conjugation efficiency 33 mg of the freeze-dried conjugate was dissolved in 30 ml of deionized water (the weight ratio of LHRH to chitosan/PMVMA was 1:10 so for producing 33 mg of the conjugate, 3 mg of LHRH was used and the concentration of LHRH in 30 ml was 0.1 mg/ml nominally). Then 2 ml of this sample was centrifuged (Micro-centrifuge Sigma 30k, UK) at 10000 rpm for 10 min in micro-centrifuging filter tubes (Amicon Ultra, Ireland) with a 10 kDa molecular weight cutoff and the supernatant was used to determine the concentration of the free LHRH. After applying the sample the retention time of 4.6 min was obtained as the peak of LHRH, then the concentration of free LHRH was calculated by using the area under the peak in the equation of LHRH standard curve.

2.6. Preparation of layer-by-layer nanoparticles of doxorubicin A layer-by-layer coating technique was used to entrap the water soluble DOX HCl inside the core of the nanoparticles. Some points were considered in architecting the nanoparticles: 1) DOX was predominantly a positively charged amphoteric drug, containing amino group in the danoseamine moiety of its molecule [38], 2) zeta potential of MNPs was negative due to the presence of citrate coating [35] and 3) LHRH-chitosan/PMVMA conjugate had a positive zeta potential surface charge. As the LHRH was used as a targeting agent, the LHRH-chitosan/ PMVMA conjugate was the last layer and the first one was MNPs and DOX. For this purpose 48 mg of Soluplus (equal or less than 0.8% of the final product) was dissolved in 2 ml of

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deionized water and shaken on magnet stirrer (IKA PT Lopower, Germany) for 15 minutes, 3 mg of DOX was dissolved in 2 ml of deionized water and added to the Soluplus solution while stirring. Then 25 l of MNPs dispersion (containing 49 mg MNPs/ml) was diluted in 2 ml of deionized water and sonicated in a bath sonicator (Hawashin 505, Korea) for 10 minutes and added to the solution of DOX and Soluplus on an orbital shaker (DAIKI SCIENCES, Korea) while shaken at 150 rpm rate for 30 minutes. At this stage the MNPs were electrostatically attached to DOX and the final zeta potential of their conjugate was still positive. Therefore, a negative layer was needed between the positive layer of MNPs conjugated to DOX and the positive final layer of LHRH-chitosan/PMVMA conjugate. For this purpose three different polymers with negative charge including; acacia gum, PMVMA or sodium alginate were selected [39]. For each formulation a solution of 1 mg/ml of the negative charged polymer was prepared and added drop by drop to the first layer (containing 3 mg DOX+1.2 mg MNPs+48 mg Soluplus) until reaching the highest zeta potential and the least particle size. To meet this situation 20 ml of acacia gum, 15 ml of PMVMA, and 2 ml of sodium alginate were used. At last the obtained nanoparticles were coated with the LHRH-chitosan/PMVMA conjugate. For this purpose 33 mg of LHRH-chitosan/PMVMA conjugate containing 2.16 mg LHRH (approximately 3.3% of final weight of formulation) was dissolved in 22 ml of deionized water and added to the nanoparticles. The mixture was shaken on the orbital shaker at 150 rpm for 30 minutes. In this way three different formulations were obtained (Table 1): Schematic representation of the magnetic lay-by-layer nanoparticles loaded with doxorubicin HCl is shown in Fig. 2.

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To confirm the presence of magnetic nanoparticles in the core of layer-by-layer assembled nanoparticles the FTIR analysis was done.

2.7. Characterization of the layer-by-layer assembled nanoparticles

2.7.1. Particle Size The particle size and particle size distribution of MNPs were determined using Zetasizer (Zetasizer-ZEN 3600 Malvern Instrument Ltd., Worcestershire, UK) based on dynamic light scattering principle technique. For these measurements 1 mg of the freeze-dried powder of MNPs was added to 10 mL of deionized water and ultra-sonicated for 10 min (Bath sonicator, HW ASH IN505, Republic of Korea), then particle size of MNPs dispersion was measured at room temperature. Furthermore, the mean particle sizes of three different formulations of layer-by-layer LHRH coated nanoparticles were measured by the same device after suitable dilution with deionized water (1:5) to measure the mean particle size and poly dispersion index (PDI) of the nanoparticles. The average diameter and particle size distribution were reported from 3 measurements.

2.7.2. Determination of iron content of the magnetic layer-by-layer nanoparticles The iron content of the magnetic layer-by-layer nanoparticles was determined by inductively coupled plasma optical emission spectrometry (ICP-OES, PERKIN ELMER-7300 DV, USA) using 10 ml of nanoparticles dispersion loaded with MNPs. The analysis of sample was done in comparison with the ICP-MS standard (Sigma Perkin Elmer Iron (Fe) Pure Grade Atomic Spectroscopy Calibration Standard was supplied with a comprehensive Certificate of Analysis that documented the quality and reliability. Concentration; 1000 mg/l; matrix is 2% HNO3, Volume is 500 ml).

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2.7.3. Determination of doxorubicin loaded in the targeted magnetic layer-by-layer nanoparticles Drug loading efficiency (LE%) was determined by measuring the concentration of free drug in aqueous medium. For this purpose 400 l of the LHRH coated magnetic layer-by-layer nanoparticles was centrifuged (Microcentrifuge Sigma 30k, UK) at 10000 rpm for 15 min in micro-centrifuging filter tubes (Amicon Ultra, Ireland) with a 10 kDa molecular weight cutoff, and the concentration of free drug in the aqueous medium diluted 1:5 with deionized water was measured by a UV-visible spectrophotometer (UV-mini 1240, Shimadzu, Kyoto, Japan) at max=274 nm. Unloaded targeted nanoparticles were used as control. The amount of entrapped drug was determined through the difference between the total and the free drug. Loading efficiency (LE%) was calculated by the following equation: Drug loading efficiency % =

Total drug  free drug 100 eq. 1 Total drug

2.7.4. In vitro release of doxorubicin from targeted magnetic layer-by-layer nanoparticles The in vitro release of doxorubicin from targeted nanoparticles was monitored in PBS 0.2 M (pH 7.4) containing 2% of Tween® 20. One ml of aqueous dispersion of each formulation was transferred into the dialysis membrane bags (Mw cutoff 12000, Membra-Cel, Viskase, USA) and the end-sealed dialysis bags were sunk fully in 10 ml of release medium at room temperature. At appropriate time intervals 600 l samples were taken and the concentration of doxorubicin released in the medium was determined by UV spectrophotometry method at max=499.4 nm. The parameter of release efficiency within 48 hours (RE48%) was used to compare the release profiles:

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t

RE48% =

 y . dt 100 0

y100 . t

eq. 2

2.7.5. Thermal analysis of the assembled layer-by-layer nanoparticles To determine the coated chitosan-PMVMA-LHRH conjugate on the assembled layer-by-layer nanoparticles thermo-gravimetric analysis (TGA) was performed. Thermo-analytical technique of nanoparticles was accomplished on a simultaneous Thermal Analysis device (STA) (LINSEIS L81/1750-PLATINUM, Germany) from 12-725°C at a heating ramp of 10°C, under a constant flow of nitrogen gas (100 ml/min). In this regard, TGA and DTA analyzes were performed.

2.7.6. Study the particles morphology by transmission electron microscope (TEM) The morphology of the LHRH targeted layer-by-layer nanoparticles were studied by TEM (Zeiss, EM10C, Germany). The samples for TEM study were prepared by placing a droplet of the nanoparticles dispersion onto a 300 mesh carbon coated copper grid and allowing it to dry in air naturally. Finally, micrographs were taken with different levels of magnification with an accelerating voltage of 80 kV.

2.7.7. Magnetic properties of the LHRH targeted layer-by-layer nanoparticles Magnetic properties of the prepared MNPs and the LHRH targeted layer-by-layer nanoparticles were obtained by a vibrating sample magnetometer (VSM) (AGFM/VSM 3886 Kashan, Iran) at room temperature in a magnetic field strength of 1 Tesla.

2.8. Protein binding measurements The protein binding study illustrates the behavior of nanoparticles in circulation. To determine the protein binding interaction with targeted layer-by-layer nanoparticles containing MNPs,

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an in vitro test was performed for measuring the interaction of bovine serum albumin (BSA) with nanoparticles dispersion. To determine the protein binding of the optimized targeted nanoparticles, at first the UV spectrophotometer absorption of a BSA solution in distilled water was determined in the range of 200-400 nm and the standard curve was obtained by the concentration of BSA between 62.5-500 g/ml in the max=277.5 nm. Then 2 ml of BSA solution with concentration of 500 g/ml was added to 2 ml of the targeted nanoparticles at the concentration of 50 g/ml. They were mixed together on an orbital shaker at the rate of 150 rpm for 1 hours and then centrifuged at 10000 rpm for 5 min and the absorption of the supernatant solution was measured at max=277.5 nm. The amount of free BSA and the protein binding percent was achieved by subtraction of blank absorption from sample absorption by using the standard curve.

2.9. Stability test The particle size of the LHRH targeted nanoparticles was measured every 5 days till one month for evaluation of the stability of the optimized formulation and the particle size of the nanoparticles was determined as mentioned earlier. 2.10. Evaluation of the expression of the LHRH receptor on MCF7 tumor cells The expression of the LHRH receptors was studied on MCF7 breast cancer cells in vitro. Primary cultures of MCF7 breast cancer cells as LHRH positive cells [40, 41], and SKOV3 tumor cells (as LHRH receptor negative cell) [19, 42] were prepared in 6-well plates (Costar, IL, USA) at a density of 300,000 cells per well. FITC labeled LHRH antibody (10 g antibody/106 cells) was directly added to cells and incubated for 45 min at 37°C. After 45 min,

15

the cells were then thoroughly washed three times with PBS (pH 7.4) and evaluated by fluorescence microscopy (NIKON eclipse TI-U, Japan).

2.11. In vitro cytotoxicity studies MCF-7 cells (as LHRH receptor positive breast cancer cells) were seeded in 96-well plates at 5×104 cells/ml and incubated in a CO2 incubator (Napco 6500, French) for 24 h to grow. Then the cells were treated with free DOX, LHRH targeted layer-by-layer nanoparticles loaded with DOX, and non-targeted layer-by-layer nanoparticles loaded with DOX. Two types of blank layer-by-layer nanoparticles were also used: LHRH targeted layer-by-layer nanoparticles but without DOX and non-targeted layer-by-layer nanoparticles without DOX. All these groups were tested in four different concentrations based on DOX concentrations of 0.2, 0.4, 0.8 and 1.6 M in the presence and absence of magnetic field (0.420 Tesla) at 37°C for 48 h. After this period, each well was exposed to 20 l of MTT solution (5 mg/ml of PBS) and plates were incubated for an additional 3 h. Then the culture medium (containing free and unreacted MTT) of the wells was removed and blue-violet formazan crystals were dissolved by adding 150 𝜇l of DMSO. The color intensity was measured at wavelength of 570 nm using an ELISA plate reader (Awareness, USA). Then to show that the cellular uptake of targeted layer-by-layer nanoparticles was achieved by LHRH receptors binding in MCF7 cells, in a separate experiment the LHRH receptors were saturated by free LHRH analog in different concentrations of 0, 10, 100, 250 or 500 nM/well [43], 1 h prior to incubation with the nanoparticles in concentration of 0.4 M based on DOX

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concentrations in the absence of magnetic field at 37°C for 48 h. Then cell survival was calculated using the aforementioned procedure in the absence of magnetic field. Untreated cells and cells treated with free DOX were used as negative and positive controls, respectively. Standard deviations were obtained from 3 replicate for each cell line.

2.11. Cellular uptake Cellular uptake of the targeted and non-targeted drug loaded layer-by-layer nanoparticles in the 2 studied cell lines in the presence and absence of magnetic field (0.420 Tesla) was demonstrated by fluorescent microscope (NIKON eclipse TI-U, Japan, =560 nm). For this purpose 1 ml of MCF-7 cells were seeded in 6-well plates at 105 cells/ml concentration and incubated in a CO2 incubator (Napco 6500, French) for 24 h to grow. Then the cells were treated with free DOX, LHRH targeted nanoparticles loaded with DOX, and non-targeted nanoparticles loaded with DOX in the concentration of 0.4 M based on DOX in the presence and absence of magnetic field (0.420 Tesla) at 37°C for 1 h. After this period, the medium of wells was removed and the cells were washed by PBS, then each well was observed by fluorescent microscope based on DOX fluorescence. 2.12. Statistical Analysis Values were processed using Microsoft Excel 2010 and IBM SPSS Statistics (Ver. 21, US) using analysis of variance (ANOVA) followed by the post hoc test of LSD and the level of significance was set at p<0.05.

3. Results and discussion FTIR spectra of chitosan, PMVEMA, and chitosan-PMVEMA conjugate are shown in Fig. 3. For chitosan (Fig. 3a), the peaks at 3436 cm-1, 1631 cm-1 and 1521 cm-1 were attributed to the

17

O-H and C=O and amino stretching vibration bond. In the spectrum of PMVEMA (Fig. 3b); a broad absorption bond at 3420 cm-1 was corresponded to the hydroxyl group of this polymer, and the absorption bond at 1716 cm-1 was assigned to its carbonyl group. Fig. 3c shows a peak in 3452 cm-1 which relates to NH bound of primary and secondary amine and amide groups, the aliphatic C-H bounds of chitosan or PMVMA show a peak in 2923 cm-1. The C=O of the carboxylic acid of the copolymer is also seen in 1705 cm-1 The 1HNMR spectrum of the chitosan (Fig. 4a) in D2O exhibited 3 separated peaks appearing at δ=2.0 ppm (COCH3), δ= 3 ppm (NH2) and δ=3.4-3.8 (CH group in sugar ring). For PMVEMA (Fig. 4b), the signals at 3.7, 3.36 and 1 ppm were assigned to Ha, Hb and Hc. The presence of the reference peaks of chitosan (2.7 ppm), PMVEMA (3.4 ppm) and as well as shift of H groups of NH2 of chitosan from δ= 3 ppm to lower field convinced us that chemical bonding had occurred between ingredients. 1HNMR spectra of PMVEMA-chitosan copolymer also showed specific peaks for both PMVEMA and chitosan. The degree of substitution of PMVEMA on chitosan was calculated by 1HNMR spectrum from the peak area Hc of PMVEMA observed at σ= 3.4 ppm and H2 of chitosan at σ=2.7 ppm. Accordingly, the substitution degree of PMVEMA to chitosan was about 0.08 to 1.

3.1. FTIR spectra of LHRH conjugated chitosan-PMVMA copolymer To confirm the conjugation of LHRH to chitosan-PMVMA copolymer, FTIR analysis of LHRH, chitosan-PMVMA copolymer, and LHRH-chitosan-PMVMA conjugate were studied. The results are shown in Fig. 5. The spectrum of LHRH (Fig. 5a) shows the stretching OH of COOH group in 3310 cm-1, the stretching C-O of COOH group is seen in 1240 cm-1 and the

18

C-N bound of amine group is in 1102 cm-1. Also there is a peak in 1662 cm-1 which stands for the carboxyl (C=O) bound of amide group. The FTIR of chitosan-PMVMA copolymer (Fig. 5b) shows a peak in 3452 cm-1 which relates to NH bound of primary and secondary amine and amide groups, the aliphatic C-H bounds of chitosan or PMVMA show a peak in 2923 cm-1. There is not any peak in the FTIR spectrum of LHRH (Fig. 5a) at this region while it appears in the conjugate at 2934 cm-1 (Fig. 5c). Also there is a peak in 1705 cm-1 which stands for C=O bound of carboxylic acid of PMVMA. The OH bound of this COOH group is seen in 2493 cm-1. As it is shown in FTIR of the conjugate of LHRH and copolymer (Fig. 5c) the C=O of the carboxylic acid of the copolymer which was seen in 1705 cm-1 (Fig. 5b) is shifted to 1662 cm-1 indicating its conversion to C=O of the amide group. The sharp peak in 1580 cm-1 is also related to the NH bound of the amide group (Fig. 5c) which appears very sharply in the conjugate. These changes convinced us that the COOH of LHRH is conjugated to the NH2 of the chitosan-PMVMA copolymer.

3.2. LHRH conjugation efficiency Capillary electrophoresis method was used for determining the LHRH content of LHRH-chitosan-PMVMA conjugate. The standard curve plotted using 6 different concentrations (12.5, 25, 50, 62.5, 100, 125 g/ml) of LHRH in deionized water, led to the equation of: y=2.5651x+19.45 with regression coefficient of r2=0.99, by which the conjugation efficiency of LHRH to the copolymer was estimated. The concentration of free LHRH in the filtered sample was 27.47 g/ml which means that from the total concentration of the LHRH

19

applied 72.51% was conjugated to chitosan-PMVMA copolymer. Fig. 1 of supplements shows the capillary electrophoresis chromatograms of free and conjugated LHRH.

3.3. Particle size and zeta potential The particle size and particle size distribution measurements of MNPs coated with citric acid were obtained using a dynamic light scattering (DLS) instrument. The results showed the mean particle size of 70 nm with a polydispersity index of 0.2 which indicates low diversity of the particle size. The results of physicochemical properties of different formulations of LHRH targeted nanoparticles are shown in Table 2. Table 2 shows that the polydispersity index (PDI) of the prepared layer-by-layer nanoparticles was between 0.3 and 0.4 which means low diversity of the particle size. As PDI is calculated from the square of the standard deviation/mean diameter, less value of PDI indicates enhanced homogeneity of the magnetic micelles [2](M. Nayebsadrian, J. Varshosaz, F. Hassanzadeh, H. Sadeghi, M. Banitalebi, M. Rostami, Screening the most effective variables on physical properties of folate-targeted dextran/retinoic acid micelles by taguchi design, Journal of Nanomaterials 2012 (2012) 1). In this study three formulations were produced using constant amount of MNPs, DOX, Soluplus and chitosan-PMVMA/LHRH conjugate. The only difference between them was the type and the amount of negative middle polymer. However, Table 2 indicates that all of the formulations have acceptable particle size. The least particle size and zeta potential were seen in acacia containing formulation.

3.4. ICP analysis

20

The ICP analysis confirmed the entrapment of super-paramagnetic Zn-Fe oxide nanoparticles into the assembled ones. As seen in Table 2 the ICP of acacia formulation was quite high and about 84%. In sodium alginate formulation the efficiency percent of MNPs entrapment was fairly acceptable and about 35% but for PMVMA formulation this percentage was very lower and about 3.6%. Therefore, the acacia and sodium alginate formulations had more acceptable iron contents than the other formulation. Also in our previous study about magnetic folate-dextran-retinoic acid micelles for dual targeting of doxorubicin the ICP of different formulations was between 2.5-11% [44].

3.5. Determination of doxorubicin loading in LHRH-coated nanoparticles The results of Table 2 indicates that all formulations had acceptable loading efficiency between 62-87%. The highest drug loading efficiency was 87.6% which related to sodium alginate containing formulation. There are different effective factors on the loading of drugs in different carriers that may be generally divided into two categories, ie., (I) environmental factors and (II) the manner in which the drug interacts with the specific polymeric carriers such as polymeric nanoparticles, liposomes, and micelles [24]. Although our previous study [7] showed that increased polymer content resulted in less drug loading efficiency, in the present study the effect of polymer content on this parameter may be the most important environmental factor, as it is shown in Table 1 the weight of the intermediate layer in sodium alginate formulation is significantly lower than the other two formulations. The nature and amount of the drug are the other factors that affected the loading efficiency. Morales et al. [45] reported that the high hydrophobicity of paclitaxel made it difficult to develop into an effective drug delivery system.

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In the current study a constant concentration of doxorubicin was used in all formulations therefore, the effect of the drug content on the loading efficiency is not shown, but obviously as mentioned before, the positive charge of doxorubicin caused its attachment to the MNPs. However, Yoo and Park [6] reported that doxorubicin loading efficiency in polyethylene glycol-folate conjugate polymer gradually decreased as the initial applied drug was increased. As a result it may be concluded that the environmental factors along with the polymer type and concentration, the way that the drug binds to the polymer, and the nature and amount of the drug play an important role in controlling the drug loading efficiency and lead to success release of the drug molecule at the target site.

3.6. In vitro release of doxorubicin from targeted nanoparticles Fig. 2 of supplements shows the release profile of DOX versus time for the prepared nanoparticles. All formulations released 100% of the loaded drug within 48 hours and in all of them the release profile included three phases, the first phase with a very steep slope in which between 60-80% of the loaded drug was released during the first 7 hours, the second phase had a slower slope during 7-24 hours that almost 10-30% of the drug was released in this step and the final phase in which the remained 10% of the drug was released. This may be interpreted as doxorubicin HCl was a very water soluble drug and was released from the near surface of the nanoparticles and leaked promptly into the release medium while the next slow release phases were due to the drug diffusion through the inner layers of the nanoparticles [29,46]. Also our previous study showed the folate-dextran/retinoic acid magnetic micelles prepared for dual targeted delivery of doxorubicin in breast cancer cells released almost 100% of the loaded drug

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within 120 min and after a rapid burst release the remaining drug was released with a near zero order kinetic model [44]. Previous studies [46, 40] show the release of DOX from magnetic micelles is pH dependent. Yang et al. [46] reported that DOX release from both water soluble polymers like poly (ethylene glycol) and insoluble polymers like poly(3-caprolactone) at pH 5.0 was much faster than that at pH 7.4. As seen in Table 2 all three developed nanoparticles had acceptable release efficiency within 48 hours in PBS (0.2 M, pH 7.4). Considering the results obtained from Table 2 and the release profiles of DOX from developed layer-by-layer nanoparticles the best formulation from particle size point of view was nanoparticles containing acacia, the best zeta potential was seen in sodium alginate formulation but the best loading efficiency (87.6%) related to the sodium alginate formulation which was significantly greater than the other formulations (p<0.05). The RE48% of the three formulations were almost the same and the difference was not significant (p>0.05). The highest MNPs loading was also seen in the acacia formulation. Considering that the sodium alginate formulation had the highest stability due to the relatively large zeta potential (Table 2) and had the highest drug loading efficiency, acceptable particle size and RE48%, it was selected as the optimized formulation and was used for further studies.

3.7. Morphology of targeted layer-by-layer assembled nanoparticles A typical TEM photograph of targeted nanoparticles containing alginate is shown in Fig. 6. Results indicates that nanoparticles were fairly smooth and spherical in shape, and the particle size was ranging between 180 and 220 nm. Considering that the TEM micrographs were taken

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two weeks after the production of the nanoparticles containing alginate the results of the TEM micrographs are a little bit larger than that obtained by DLS method (Table 2). TEM micrograph shows the presence of the magnetic nanoparticles in the core of the layer-by-layer nanoparticles.

3.7. Magnetic properties of the targeted nanoparticles containing alginate Fig. 7 displays the superparamagnetic property of CA-coated MNPs and LHRH targeted MNPs in the powder state at room temperature. The saturation magnetization was determined to be 48 and 12 emu/g for CA-coated MNPs and LHRH targeted nanoparticles, respectively. The reduction of saturated magnetism of LHRH targeted nanoparticles compared to MNPs (Fig. 7http://www.hindawi.com/journals/bmri/2013/680712/fig11/ - a), confirms the presence of an organic coating on their surface [47, 48].

3.8. Protein binding of the targeted nanoparticles containing alginate Nanoparticle composition, hydrophobicity, presence of specific functional groups, pH and temperature have been shown to affect protein adsorption on the surface of nanoparticles [49]. The MNPs are developed for preclinical and human application and should be formulated into a suitable delivery system with appropriate pharmacokinetic parameters [50]. One limitation of the MNPs is their destabilization because of adsorption of plasma proteins which leads to nonspecific uptake by the reticuloendothelial system (RES). To evade clearance by RES and to improve the circulation time of particles, MNPs are coated with polymers [24]. In the present study, LHRH targeted nanoparticles were used to encapsulate MNPs and reduce protein adsorption. To prove this concept, optimal formulation was tested for in vitro protein (BSA) adsorption. The BSA adsorption was measured by the UV-visible spectrophotometer.

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At first BSA solution with the concentration of 500 g/ml was used for determination of maximum wavelength of absorption spectrum in distilled water in the range of 200-400 nm and then the standard curve for albumin was plotted in this wavelength between the concentrations of 50625 g/ml and the line equation was obtained to be: y=0.0004x+0.0275 with regression coefficient of R2=0.9999. After that with the same procedure mentioned in the Methods section 2.8. For measuring the the protein binding of the nanoparticles the mixture of BSA and nanoparticles was prepared (according to the method mentioned in section 2.8) and centrifuged at 10000 rpm for 5 min, then the absorption of supernatant solution was measured in 277.5 nm and the absorbance of the sample was changed to concentration using the equation obtained from the standard curve. The results showed that 69% of the protein was free or just 31% was adsorbed by the MNPs entrapped in the nanoparticles.

3.9. Stability test The stability study was performed by measuring the particle size of the alginate containing LHRH targeted nanoparticles every five days until one month under room temperature. Fig. 3 of supplements shows that the particle size of the targeted nanoparticles increased over time which is probably the result of their aggregation, but fortunately the dispersion was very stable for almost one month. This formulation showed no aggregation or sedimentation at least for one month and a gradual increase in particle size from 127 nm to 394 nm was observed during this period, which indicated the high stability of the dispersion. In most studies on the dispersion of colloidal nanoparticles like in the study reported by Chandrasekharan et al. [48] on the

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d-alphatocopheryl-co-poly (ethylene glycol) 1000 succinate (TPGS) micelles the particle size of the micelles was stable just for 12 days at its best condition. 3.10. Evaluation of the expression of the LHRH receptor on MCF7 tumor cells The binding of FITC-labeled LHRH receptor antibodies on the surface of LHRH receptor positive cells could exhibit membrane fluorescence. As shown in Fig. 8, MCF7 LHRH receptor positive cells exhibited membrane fluorescence after 45 min of incubation with FITC-labeled LHRH receptor antibody at 37°C. But, SKOV3 LHRH receptor negative cells did not show any fluorescence when incubated with FITC-labeled LHRH receptor antibodies (Fig. 8). Thus the membrane fluorescence exhibition of MCF7 tumor cells after incubation with FITC-labeled LHRH receptor antibody could confirm the expression of LHRH receptors on the surface of MCF7 tumor cells. 3.10. In vitro cytotoxicity tests Cell survival of MCF-7 cells are shown in Figs. 9. The cytotoxicity of DOX loaded LHRH targeted nanoparticles containing alginate was compared with free drug and non-targeted DOX loaded nanoparticles in the absence and presence of magnetic field, also the probable cytotoxicity of the targeted and non-targeted blank nanoparticles were checked. Fig. 9 shows that in MCF-7 cells the cell survival percentage was decreased significantly (p<0.05) in targeted nanoparticles group compared to free DOX and non-targeted nanoparticles in most drug concentrations both in absence and presence of magnetic field. In the absence of magnetic field the LHRH targeted nanoparticles showed the lowest IC50 of 0.4 M compared to free DOX and non-targeted nanoparticles that showed the cell survival percentages of 68% and 66%

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in this concentration, respectively. In the presence of magnetic field also LHRH targeted nanoparticles showed the lowest IC50 of 0.4 M compared to free DOX and non-targeted nanoparticles which showed the cell survival percentages of 62% and 60% in this concentration, respectively (Fig. 9). Moreover, both targeted and non-targeted nanoparticles showed the lower cell survival percentages in concentrations of 0.2, 0.4 M of DOX in the presence of magnetic field compared to the same groups in the absence of magnetic field (p<0.05). In the concentration of 0.8 M just targeted nanoparticles showed significant difference between the presence and absence of magnetic field and other groups were not different (p>0.05). At concentration of 1.6M none of the studied groups showed significant difference between the presence or absence of the magnetic field (p>0.05). This may be interpreted because of the high concentration of the drug that cause high cell cytotoxicity even without the presence of the magnetic. This effect indicates that at concentrations of 0.2-0.4 M and somehow at 0.8M the presence of magnetic external fields besides the targeting ligand led to increasing cytotoxic effects of the nanoparticles. To show the effect of LHRH targeting in MCF-7 (LHRH positive) cells over expressing LHRH receptors [40, 41], in absence of magnetic field the receptors of LHRH were saturated with different concentrations of LHRH before contacting the cells with the targeted nanoparticles [19, 42]. Fig. 10 shows that in the absence of LHRH the cell viability was significantly (p<0.05) lower than other concentrations while by increasing the LHRH concentration before exposure of the cells with the targeted nanoparticles, the cell viability increased and at last at the concentration of 250 nM of LHRH the receptors were saturated so that there was not any significant difference between 250 and 500 nM of LHRH and the cell viability reached to almost 94%. In other words, this test confirms the entrance of targeted nanoparticles is probably facilitated by active endocytosis through the LHRH receptors so that after their saturation the nanoparticles were not able to enter actively and the cell viability increased.

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3.11. Cellular uptake of nanoparticles The microscopy pictures of MCF-7 cells treated by three different groups of free DOX, targeted DOX loaded nanoparticles and DOX loaded non-targeted nanoparticles are shown in Fig. 11. As doxorubicin itself has fluorescent property, the cells with fluorescent emissions show the cellular uptake of the drug, also more fluorescent emissions shows more uptake and also better drug entry to the cells. Weber et al. [51] also studied cellular uptake of DOX as a function of cholesterol in MCF-7 cells incubated with doxorubicin (2 μM) for 2 h and showed the same fluorescence color of green or blue color at λ em of 520 nm. Fig. 11 shows that the highest fluorescent emission was seen in MCF-7 cells treated by targeted DOX loaded nanoparticles, while the fluorescent emissions of MCF-7 cells treated by non-targeted DOX loaded nanoparticles and free DOX were weaker than the targeted nanoparticles although the non-targeted ones still emitted considerably stronger fluorescence than those treated with free DOX. This confirms that the presence of LHRH as the targeting moiety has enhanced cellular uptake of DOX in LHRH-positive receptor cells (MCF-7).

4. Conclusion A novel layer-by-layer technique was used for production of targeted nanoparticles for dual targeting of DOX in breast cancer. MNPs containing Zn-Fe were synthesized by hydrothermal method and loaded in LHRH-targeted nanoparticles containing DOX. The developed nanoparticles consisted of three layers: inner positive layer containing of DOX and MNPs, the middle negative polymer layer composed of acacia or sodium alginate or PMVMA, and the most outer layer was the LHRH-chitosan-PMVMA conjugate. Among the studied formulations the best results were obtained from sodium alginate containing formulation. The highest growth inhibitory effect was observed in MCF-7 cells treated with targeted DOX loaded nanoparticles

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in the presence of magnetic field at 0.4 M concentration. The results demonstrated the potential of the LHRH targeted layer-by-layer nanoparticles to achieve dual tumor targeting by magnetic field-guided in breast cancer cells. This may reduce the required dose of DOX and consequently reduction the side effects of this drug. The results should be checked in vivo to confirm the promising results obtained from the cell culture tests. Acknowledgement The authors appreciate the financial support of this work by Research Vice Chancellor of Isfahan University of Medical Sciences. References [1] Z.G. Surmeli, U. Varol, B. Cakar, M. Degirmenci, C. Arslan, G.D. Piskin, B. Zengel, B. Karaca, U.A. Sanli, R. Uslu,;1; Capecitabine aintenance therapy following docetaxel/capecitabine combination treatment in patients with metastatic breast cancer, Oncol. Lett. 10(4) (2015) 2598-2602. [2] M. Li, Y.Fan, Q. Li, P. Zhang, P. Yuan, F. Ma, J. Wang, Y. Luo, R. Cai, S. Chen, Q. Li, B. Xu,;1; Vinorelbine plus platinum in patients with metastatic triple negative breast cancerand prior anthracycline and taxane treatment, Medicine (Baltimore). 94(43) (2015) e1928. [3] N.E. Avis, S. Crawford, J. Manuel,;1;Quality of life among younger women with breast cancer, J. Clin. Oncol. 23(15) (2005) 3322–3330. [4] A.V. Schally, A.M. Comaru-Schally, A. Nagy, M. Kovacs, K. Szepeshazi, A. Plonowski,;1; Hypothalamic hormones and cancer, Front. Neuroendocrinol. 22(4) (2001) 248–291. [5] J.J. Wang, E. Cortes, L.F. Sinks, J.F. Holland,;1;Therapeutic effect and toxicity of adriamycin in patients with neoplastic disease, Cancer. 28(4) (1971) 837–843. [6] H.S. Yoo, T.G. Park,;1; Folate-receptor-targeted delivery of doxorubicin nano-aggregates stabilized by doxorubicin--PEG--folate conjugate, J. Control. Rel. 100(2) (2004) 247–256. [7] M. Nayebsadrian, J. Varshosaz, F. Hassanzadeh, H. Sadeghi, M. Banitalebi, M. Rostami,;1; Screening the most effective variables on physical properties of folate-targeted dextran/retinoic acid micelles by taguchi design, J. Nanomater. 2012 (2012) 133.

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Fig. 1. Schematic representation of the chemical reactions for synthesis the chitosan-PMVEMA copolymer

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Fig. 2. Schematic representation of the layer-by-layer structure of doxorubicin loaded in LHRH-chitosan/PMVAM coated magnetic nanoparticles

Fig. 3. FTIR spectra of a) chitosan, b) PMVEMA, c) chitosan-PMVEMA conjugate

Fig. 4. 1HNMR spectra of a) chitosan, b) PMVEMA, c) chitosan-PMVEMA conjugate

Fig. 5. FTIR spectra of a) LHRH, b) chitosan-PMVMA copolymer, c) LHRH-chitosan PMVMA

Fig. 6. TEM micrographs of a) CA coated MNPs and b) LHRH targeted nanoparticles prepared by alginate middle layer

7.

Magnetization curve measured using vibrating sample magnetometer (VSM):

(a) CA-coated magnetic MNPs, (b) LHRH targeted nanoparticles

Fig. 8. Fluorescence microscopy images of (a) MCF7 and (b) SKOV3 tumor cells after incubation with FITC labeled LHRH receptor antibody for 45 min at 37◦C.

Fig. 9. Cell viability of MCF-7 cells in the absence or presence of magnetic field after treatment with different concentrations of alginate containing LHRH targeted nanoparticles loaded with doxorubicin in comparison with free drug and non-targeted nanoparticles by MTT assay. *: shows significant difference with targeted nanoparticles loaded with DOX in absence of magnetic field in comparable concentrations (P<0.05), #: shows significant difference with targeted nanoparticles loaded with DOX in presence of

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magnetic field in comparable concentrations (P<0.05), ^: shows significant difference with the absence of magnetic field in comparable concentrations (P<0.05).

Fig. 10. The effect of the concentration of unconjugated LHRH analog (0, 10, 100, 250 and 500 nM) for 1 h exposition on the antiproliferative efficacy of targeted doxorubicin loaded nanoparticles (0.4 M according to doxorubicin concentration) in MCF7 cells for 48 h. The viabilities were expressed as a percentage of untreated controls (100%).

Fig. 11. Visible and fluorescent microscopy pictures of MCF-7 cells treated by targeted doxorubicin loaded nanoparticles, non-targeted doxorubicin loaded ananoparticles and free doxorubicin after 1 h incubation.

Table 1. Formulation of different magnetic layer-by-layer nanoparticles loaded with doxorubicin HCl Intermediate negative layer type

DOX HCl (mg)

MNP (mg)

Soluplus (mg)

Acacia gum Sodium alginate PMVMA

3 3 3

1.2 1.2 1.2

48 48 48

Weight of intermediate layer 20 2 15

LHRH-chitosanPMVMA conjugate

33 33 33

Table 2. Particle size, zeta potential, loading efficiency, release efficiency of 48 h (RE48%), and magnetic nanoparticles loading efficiency of targeted layer-by-layer assembeled nanoparticles containing doxorubicin with three different middle layers (mean±SD, n=3)

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Middle layer of nanoparticles Acacia Sodium alginate PMVMA

Particle size PDI (nm)SD 88.12.9 128.02.6 182.68.6

0.34 0.3 0.39

Zeta potential (mV) 27.6 30 10

TDENDOFDOCTD

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Loading efficiency (%)SD 67.01.0 87.61.5 62.32.3

RE48 Magnetic loading efficiency (%)SD (%)SD 84.0 79.80.6 35.0 81.90.1 3.6 83.40.9

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