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\begin{document}

\title{Design and preparation of a new multi-targeted drug delivery system using multifunctional nanoparticles for co-delivery of siRNA and paclitaxel}



\author[1]{Sara Hosayni Nasab}%
\author[1]{Amin Amani}%
\author[2]{Hossein Ali Ebrahimi}%
\author[3]{Ali Asghar Hamidi}%
\affil[1]{University of Mohaghegh Ardabili}%
\affil[2]{Ardabil University of Medical Sciences}%
\affil[3]{Tabriz University of Medical Sciences}%


\vspace{-1em}



  
  \date{May 08, 2024}


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\selectlanguage{english}
\begin{abstract}
Drug resistance is a great challenge in cancer therapy using
chemotherapeutic agents. Administration of these drugs with siRNA is an
efficacious strategy in this battle. Here, it was tried to incorporate
siRNA and paclitaxel simultaneously into a novel nanocarrier. The
selectivity of carrier to target cancer tissues was optimized through
conjugation of folic acid (FA) and glucose onto its surface. The
structure of nanocarrier formed from ternary magnetic copolymers based
on FeCo-polyethylenimine (FeCo-PEI) nanoparticles and polylactic
acid-polyethylene glycol (PLA-PEG) gene delivery system.
Biocompatibility of FeCo-PEI-PLA-PEG-FA(NPsA), FeCo-PEI-PLA-PEG-Glu
(NPsB) and FeCo-PEI-PLA-PEG-FA/Glu (NPsAB) nanoparticles and also
influence of PTX-loaded nanoparticles on in vitro cytotoxicity were
examined using MTT assay. Besides, siRNA-FAM internalization was
investigated by fluorescence microscopy. The results showed the blank
nanoparticles were significantly less cytotoxic at various
concentrations. Meanwhile, siRNA-FAM/PTX encapsulated nanoparticles
exhibited greater cytotoxicity on MCF-7 cells. NPsAB/siRNA/PTX
nanoparticles showed greater effect on MCF-7 cell viability than
NPsA/siRNA/PTX and NPsB/siRNA/PTX. Also, they induced significantly
higher cytotoxic effects on cancer cells compared with NPsA/siRNA/PTX
and NPsB/siRNA/PTX due to their multi-targeted properties using folic
acid and glucose. We concluded that NPsAB nanoparticles have great
potential for co-delivery of both drugs and genes for use in gene
therapy and chemotherapy.%
\end{abstract}%



\sloppy


\textbf{Introduction}

Cancer has been identified as the second main cause of mortality
worldwide , after heart diseases (Sant, Iyer, Gaharwar, Patel, \&
Khademhosseini, 2013). Chemotherapy is one of the most common treatment
procedures for cancer therapy which is used after surgery and in a
combination with radiotherapy (Cao et al., 2011; Lv et al., 2014; Zou et
al., 2012). Although the chemotherapy has been proved to be useful in
many cases, nonetheless, the severity of its side effects and also the
drug resistance to these drugs lead to reducing the efficacy of
chemotherapeutic agents (Eggenberger et al., 2010; Zintchenko, Philipp,
Dehshahri, \& Wagner, 2008). During the previous few decades,
nanoparticles were employed to optimize the effect of drugs and genes,
reduce their side effects and deliver the medicines in a targeted manner
for efficient therapy (Hami, Amini, Ghazi-Khansari, Rezayat, \& Gilani,
2014a). Folic acid (FA) is an essential nutrient for the human body
which can enter cells through folate receptors (FRs) (Zwicke, Mansoori,
\& Jeffery, 2012). In the human body, folic acid is used in the
synthesis of the DNA. Therefore, the expression level of folate receptor
in cancer cells is higher compared to normal cells, due to theirs need
to high level of folic acid (Alibolandi et al., 2016; Mojica Pisciotti
et al., 2014). Glucose is another important nutrient in the human body
which provides the energy that is needed in many metabolic functions and
synthetic processes of different compounds in cells, serum and tissues
(Chatterjee, 2013; Talpur, Echard, Ingram, Bagchi, \& Preuss, 2005). So,
glucose transporters (Gluts) are over-expressed on the cell membrane of
many malignant cells (Smith, Volkert, \& Hoffman, 2005; Qiang Wang et
al., 2010). In 1927, Warburg et al. reported that glucose uptake was
significantly enhanced in cancer tissues compared to those in normal
tissues because of their high metabolic activity (Warburg, Wind, \&
Negelein, 1927). After that, glucose has been used to target cancer
cells for different diagnostic and therapeutic applications (Li, Ma,
Dang, Liang, \& Chen, 2014; Mamaeva et al., 2016).

It is well known that the targeted endocytosis pathway of nanoparticles
is based on the interaction between cell surface receptors and ligands
(S. Wang et al., 2013; Xia et al., 2013). Therefore, the efficiency of
drug delivery was limited up to saturation of cell surface receptors by
specific ligands of nanoparticles. Therefore, it seems that specific
targeting of tumor tissues could be improved through simultaneous
incorporation of several ligands such as glucose and folic acid inside
the structure of nanoparticles. So, in this study it was aimed to
optimize the surface of nanoparticles with co-addition of folic acid and
glucose.

In recent years, drug resistance was emerged as one of the biggest
therapeutic challenges in cancer therapy and so it attracted many
research attentions (Bentley-Goode, Newton, \& Thompson, 2017; Marques
et al., 2014). Simultaneous co-targeting of cancer cells using two or
more drugs seems to be a novel and effective approach to overcome the
problem of drug resistance (Camirand, Lu, \& Pollak, 2002; Q. He, Liu,
Sun, \& Zhang, 2004; Mancini et al., 2014).

The use of siRNA to suppress expression of oncogenes has been considered
as a promising solution for cancer treatment in recent years (Kumar,
Yigit, Dai, Moore, \& Medarova, 2010; Patil \& Panyam, 2009). Although,
this method can be helpful lonely, however, due to the complexity of
human cancer, other complementary therapies may be needed to improve the
efficacy of cancer therapy by siRNA (Peng, Hsu, Lin, Cheng, \& Hsu,
2017; X. Z. Yang et al., 2011). Combination therapy of siRNA and
chemotherapeutics such as paclitaxel has been considered as a useful
strategy for increasing the effectiveness of cancer therapy
(Kapse-Mistry, Govender, Srivastava, \& Yergeri, 2014).

Various carriers were used to deliver genetic agents. Cationic polymers
such as polyethyleneimine (PEI) are of the most common ones and have the
ability to interact with RNA and DNA and form complexes with high
transfection efficiency. However, studies showed that application of PEI
is associated with critical cytotoxicity effects (Abebe et al., 2015b;
Moret et al., 2001). In addition, hydrophilic polymers such as PEI have
low capacity to use as a nano-vector for delivering hydrophobic and
neutral drugs such as PTX into cancer cells.

Several methods have been developed to transfer different compounds
using nanoparticles to cancer cells (Amani, Zare, Asadi, \&
Asghari-Zakaria, 2018; Kircheis et al., 1999). Synthesis of copolymers
by binding of hydrophobic polymers such as PLA to cationic hydrophilic
polymers allows simultaneous co-delivery of siRNA and PTX to cancer
cells.

Biodegradable polymers such as polylactide (PLA),
poly(D,L-lactide-co-glycolide) (PLGA), and polyglycolide (PGA) have
great application potential as drug carriers due to their good
biocompatibility and biodegradability (Lu et al., 2014; Perez et al.,
2001a; J. Wang, Xu, Liu, Sun, \& Yang, 2016).

Researchers have developed tri-block copolymers using PLA, PEI, and PEG
such as tri-block PEI-PLA-PEG copolymers, as potential gene delivery
nano-vectors (Abebe et al., 2015a, 2015b; Sim et al., 2017). The
hydrophobic PLA segment affords a degree of biodegradability to
nanoparticles. It also increases the stability of nano-carrier through
charge shielding effects. In recent decades, much study has been done on
multifunctional nanoparticle. Multifunctional nanoparticles are
attractive for cancer treatment due to high potential to overcome the
barriers in various extracellular and intracellular conditions (Chen et
al., 2015; Liu et al., 2016; Yan, Li, Li, Zhu, Shen, Yi, Wu, Yeung, Xu,
Xu, \& Chu, 2014). This study presents multifunctional biodegradable
magnetic nanoparticles including PLA, PEG, PEI, FeCo nanoparticles,
folic acid and glucose for co-delivery and targeting of siRNA and PTX.

\textbf{Materials and methods}

\textbf{2.1. Materials}

2-Azidoethyl \selectlanguage{greek}β-\selectlanguage{english}D-galactopyranoside was purchased from Synthose Inc.
(Canada), polyethyleneimine (Mw 1800 Da) was purchased from
Polysciences, Inc. (Germany), polyethyleneimine (PEI), 1.8 kDa, branched
was obtained from polyscience (USA), stannous octoate, tetraethyl
orthosilicate (TEOS), 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide
(EDC), polyvinyl alcohol (PVA), dichloromethane (DCM), MTT, RPMI 1640,
DMSO, penicillin/streptomycin solution, N-hydroxylsulfosuccinimide
(NHS), fetal bovine serum (FBS) were purchased from Sigma-Aldrich (USA).
HO-PEG-COOH was obtained from Rapp Polymere GmbH (Germany). N-
hydroxysuccinimide, Folic acid and dicyclohexylcarbodiimide were
purchased from Acros (Germany). Agarose gel, PVP, MgCL2, Tris-HCl, and
EDTA were purchased from Merck (Germany).

\textbf{Synthesis of Azido-Functionalized Folate}

Preparation of PLA-PEG-FA copolymer was started with the synthesis of
OH-PLA-PEG-COOH according to the method described by Hami et al (Hami,
Amini, Ghazi-Khansari, Rezayat, \& Gilani, 2014b). Briefly, definite
amounts of lactide monomers was added to HO-PEG-COOH under dry argon,
then the reaction was continued at 140 \selectlanguage{ngerman}°C for 24 h in the presence of
stannous octoate (Sn(Oct)2) as catalyst. The mixture was dissolved in
chloroform and precipitated by a mixture of ether and methanol (v/v=
1:1). Then, the HO-PLA-PEG-COOH obtained as precipitate was dried. This
obtained polymer was dissolved in DCM with DCC (1.2 mmol) and NHS (0.6
mmol). The product was stirred at 0 °C and 25°C for 1 h and 24 h
respectively. Triethylamine (TEA) and propargylamine (0.6 mmol) were
added into the above product. Next, the mixture was stirred at 25°C for
another 24 h. The product (HO-PLA-PEG-alkyne copolymer) was precipitated
in cold ether, filtered and dried at 30 °C under vacuum. Then, the
HO-PLA-PEG-alkyne copolymer (0.1 mmol) was conjugated with acryloyl
chloride (C3H3ClO) (0.2 mmol) in dry toluene containing triethylamine
(C6H15N) (0.2 mmol). The sample was stirred for 10 h at 80 °C. After
that, the product was cooled to 25°C, filtered and precipitated in
n-hexane then dried in vacuum oven to produce acrylate-PLA-PEG-alkyne
copolymer. Acrylate-PLA-PEG-Folic acid and Acrylate-PLA-PEG-Glucose were
synthesized stepwise by the addition of Acrylate-PLA-PEG-alkyne
copolymer (0.1 mmol) and Azido-functionalized folate and
Azido-functionalized glucose in 30 mL aq. NH4HCO3 (10 mM) separately.
Then, sodium ascorbate and CuSO4 solutions were added to
acrylate-PLA-PEG-alkyne copolymer and stirred with a rotator at 25 °C
for 24 h. Next, the product was filtered with a 0.22 \selectlanguage{greek}μ\selectlanguage{english}m membrane filter.
The reaction solutions were diluted with sufficient NaCl aqueous
solution, and then extracted several times by dichloromethane. The
resulting products were concentrated under vacuum and precipitated in
cold ether. Finally, acrylate-PLA-PEG-FA copolymer and
acrylate-PLA-PEG-Glu copolymers were dried under vacuum.

\textbf{Synthesis of FeCo nanoparticles}

First, FeO(OH) and Co3O4 in appropriate amount were dissolved in
Trioctylamine (TOA), then the mixture was loaded into a flask and
dehydrated at 160 \selectlanguage{ngerman}°C for 2 h under a flow of argon. Afterward, the
mixture was quickly heated to 370 °C while vigorous stirring is
maintained using a mechanical stirrer. Next, the product was cooled to
room temperature and collected by centrifugation at 10,000 rpm. Finally,
the product washed with hexane and ethanol, and dried under vacuum to
obtain FeCo nanoparticles (FeCoNPs).

\textbf{Synthesis of FeCo@SiO2-SS-PEI nanoparticles}

To synthesize the FeCo@SiO2-SH, 200 mg of FeCo NPs dispersed in ethanol
by ultrasound, followed by sequential addition of 1 mL NH3.H2O, 20 mL
TEOS and 1 mL MPTMS in the ultrasound for 3 hours (Zhang, 2014).

Then, the nanoparticles were collected using magnet, and washed with
ethanol and water for 5 times and then the magnetic nanoparticles were
dried under vacuum at 25 ºC. To synthesize FeCo@SiO2-SS-COOH
nanoparticles, 200 mg of Sulfhydrylated FeCo nanoparticles was dispersed
in 12 mL methanol, and then, the 2- carboxyethyl 2-pyridyl disulfide
(200 mg) was added and reaction performed for 36 hours, the obtained
nanoparticles were washed and then dried as mentioned above. Finally, to
synthesize FeCo@SiO2-SS-PEI nanoparticles, 200 mg of FeCo@SiO2-SS-COOH
nanoparticles was dispersed in 16 mL PBS buffer (pH = 7.4) and then, the
activation of the carboxyl groups was performed with addition of 200 mg
of EDC and 100 mg NHS at 25 Cº for half an hour. After that, 4 mL PBS
buffer containing 200 mg PEI (1800 Da) was added and stirred at 25 Cº
for 2 days. The resulting nanoparticles were washed with distilled water
and ethanol, then the nanoparticles were lyophilized (Lifilizator Alpha
model 1-2 LD plus, Christ, Germany).

\textbf{Synthesis of FeCo@SiO2-SS-PEI- PLA -PEG-FA and
FeCo@SiO2-SS-PEI-PLA-PEG-Glu and FeCo@SiO2-SS-PEI-PLA-PEG-Glu-FA}

FeCo@SiO2-SS-PEI, acrylate-PLA-PEG-FA and acrylate-PLA- PEG-Glu were
dissolved in 6ml of chloroform, separately. The chloroform solution of
acrylate-PLA-PEG-FA and acrylate-PLA-PEG-Glu was added dropwise into
FeCo@SiO2-SS-PEI solution separately. Then the mixture was stirred for
24 h at 45 °C and stirred for 24 h. The resulting nanoparticles were
washed with distilled water and ethanol, then the dispersion was
centrifuged at 11068 ×g for 30 minutes and the supernatant was removed.
The nanoparticles were lyophilized. To synthesize
FeCo@SiO2-SS-PEI-PLA-PEG-Glu-FA nanoparticles, the same amounts of
acrylate-PLA-PEG-FA and acrylate-PLA-PEG-Glu were added into
FeCo@SiO2-SS-PEI solution. Then the FeCo@SiO2-SS-PEI-PLA-PEG-Glu-FA
nanoparticles were synthesized as protocol mentioned above.

\textbf{Encapsulation of siRNA-FAM and PTX into NPsA, NPsB and NPsAB}

The drug-loaded nanoparticles NPsA/siRNA/PTX, NPsB/siRNA/PTX and
NPsAB/siRNA/PTX NPsAB/siRNA/PTX were prepared based on the previous
report with minor modification (Perez et al., 2001a).

Briefly, to synthesis of NPsA/siRNA/PTX, 0.5 ml water solution
containing siRNA-FAM (250\selectlanguage{greek}μ\selectlanguage{english}g) was added dropwise into 4 ml of aceton/DCM
solution containing NPsA (30 mg) and PTX (5 mg)The reaction mixture was
sonicated on ice for 30 seconds (Sonicator\selectlanguage{ngerman}® XL, Misonix, NY), then
slowly added to 6 mL of PVA solution (1\% w/v) and sonicated again for
two minutesNext, the reaction mixture was poured dropwise into a stirred
PVA solution (30ml, 0.3\% w/v).

After 5 min, the solvent (DCM and acetone) from the nanoparticles
fraction was evaporated using a rotary vacuum evaporator. Then, the
NPsA/siRNA/PTX were collected were collected from the solution by
centrifugation at 15000 rpm (16602 ×g) for 30 min (Sigma 1-14K), and
after washed several times with deionized water to remove the unloaded
drug, the the resultant were filtered using a 0.1 and 0.45 micrometer
membrane filter to remove large scattering particles and free NPsA, and
finally, the product was lyophilized. The NPsB/siRNA/PTX and
NPsAB/siRNA/PTX were also prepared as described above.

The encapsulation efficiency of siRNA-FAM and PTX in nanoparticles was
determined by measuring the amount of siRNA-FAM and PTX that was not
encapsulated in nanoparticles. Therefore, the amount of siRNA-FAM and
PTX in the supernatant of the nanoparticle were measured using
spectrophotometer at 490 (siRNA-FAM) and 230 (PTX) nm respectively.
Then, the amount of siRNA-FAM and PTX in the supernatant was compared
with the total amount of siRNA-FAM and PTX used in the encapsulation
process (Son \& Kim, 2010). encapsulation efficiency (EE) of siRNA-FAM
and PTX was determined as follow Eq. (1):

(\%EE) =\(\ \frac{\ Amount\ of\ used\ drug-Amount\ of\ drug\ in\ supernatant\ }{\text{Amount\ of\ used\ drug}}\)×100 (1)

The PTX and siRNA release kinetics from the nanoparticles was assessed
in PBS buffer (pH 7.4). The nanoparticles were treated separately with 5
ml of PBS buffer (1 mg/ml) then, collected via centrifugation (16602×g
for 30 minutes) at predetermined time intervals (i.e. 30 minutes - 30
days after inoculation). After each collection, the supernatants were
employed to determine the amount of siRNA and PTX released from the
nanoparticles. Afterwards, nanoparticles were re-suspended in 5 ml of
fresh PBS buffer (pH 7.4), and incubated at 37 °C until the next
scheduled time interval. The total DNA release ratio for each sample was
calculated using the following Eq. (2):

Cumulative Release (\%) =\(\frac{\mathrm{\text{Total}}\mathrm{\text{\ drug}}\mathrm{\text{\ content\ in}}\mathrm{\ }\mathrm{\text{supernatant}}}{\mathrm{\text{The\ amount\ of\ encapsulated\ drug}}}\mathrm{\times 100}\)(2)

\textbf{Characterization of nanoparticles}

The morphology of NPsA/siRNA/PTX, NPsB/siRNA/PTX and NPsAB/siRNA/PTX
nanoparticles were examined using transmission electron microscopy (TEM,
JEM-2100).

The average size, zeta potential and magnetic properties of
nanoparticles were investigated using dynamic light scattering (DLS,
Malvern Instruments, Westborough, MA, USA) and vibrating sample
magnetometer (VSM, a Standard 7403 Series, Lakeshore) respectively.

\textbf{In vitro cytotoxicity and transfection assay}

The in vitro cytotoxicity assay for blank nanoparticles (NPsA, NPsB and
NPsAB) and siRNA-FAM/PTX encapsulated nanoparticles (NPsA/siRNA/PTX,
NPsB/siRNA/PTX and NPsAB/siRNA/PTX) was evaluated on MCF-7 cell lines.
The cells were seeded in 96-well plates at a density of approximately
7000 cells per well, in 100 \selectlanguage{greek}µ\selectlanguage{english}L of complete growth medium (RPMI 1640
culture medium supplemented with 1\% Penicillin/Streptomycin and 10\%
FBS) and were incubated at 37\selectlanguage{ngerman}°C in 5\% CO2 for 24 hours. Then the cells
were exposed to various concentrations of the blank nanoparticles (100,
200, 400, 600, 800 and 1000 \selectlanguage{greek}µ\selectlanguage{english}g/ml) for 48 hours at 37\selectlanguage{ngerman}°C. Next, each well
was supplemented with 20 \selectlanguage{greek}µ\selectlanguage{english}L of MTT solution (5 \selectlanguage{greek}μ\selectlanguage{english}g/mL in culture medium)
for 5 hours at 37\selectlanguage{ngerman}°C. The supernatant was removed and 200 \selectlanguage{greek}µ\selectlanguage{english}l of dimethyl
sulfoxide (DMSO) was added per well to dissolve the Formazan crystals.
Following this step, the plates were incubated at 37\selectlanguage{ngerman}°C for 30 minutes.
Next, the absorbance level at 570\selectlanguage{english} nm was recorded by a multiwall plate
reader (BioTek Instruments; Winooski, VT, USA) (Qian Wang et al., 2017).
Finally, the percentage for the cell viability was determined according
to the following Eq. (3):

Cell viability (\%) =\(\ \frac{OD\ of\ each\ sample\ at\ 570\ nm\ }{OD\ of\ negative\ control\ at\ 570\ nm}\)x100 (3)

The in vitro cytotoxicity assay for the nanoparticles encapsulated with
siRNA-FAM/PTX (NPsA/siRNA/PTX, NPsB/siRNA/PTX and NPsAB/siRNA/PTX) was
evaluated with equal amount of PTX (equivalent PTX concentration of 40
nM for each sample) in each sample on the MCF-7 cell lines as described
above. To investigate the transfection ability of nanoparticles, the
cells were incubated with NPsA/siRNA/PTX, NPsB/siRNA/PTX and
NPsAB/siRNA/PTX, then were washed twice using PBS, and the siRNA-FAM
transfection efficiency was determined using an inverted-fluorescent
microscope (Nikon TE200).

\textbf{Results and discussion}

\textbf{Synthesis of PLA-PEG-FA and PLA-PEG-Glu copolymers}

The chemical structure of the PLA-PEG-FA and PLA-PEG-Glu copolymers were
examined with proton nuclear magnetic resonance spectroscopy (1H NMR).
Fig. 1 showed the 1H-NMR spectrum of PEI-PLA-PEG-FA copolymer, where the
presence of CH2 protons of PEG in PEI-PLA-PEG-FA copolymer was observed
around 3.6 ppm (-CH2CH2O-). Moreover, the appearance of signals at 5.1
and 1.6 ppm are attributed to the methine (CH) protons and methyl
protons -- (CH3) of PLA, respectively (Fig. 1 A) (Heald et al., 2002;
Sim et al., 2017).

The appearance of chemical shifts at 6.5-8.6 ppm (aromatic protons
associated with FA) and 7.9 ppm (s, weak, 1H, triazoles) provided
evidence for successfully obtaining PEI-PLA-PEG-FA copolymer (S.-J. Yang
et al., 2010)(Fig. 1 A).

Also in the Fig. 1 B The multiple sharp peaks at 3.6-3.8 ppm and weak
peak at 5.4 ppm which were assigned to methylene protons in PEG and
aromatic proton in glucose respectively, appeared on the spectrum after
the reaction of PLA-PEI with PEG-Glu (Fig. 1 B)(Sadeghi et al., 2015).

The structures of PLA-PEG-FA and PLA-PEG-Glu were ascertained by the
FTIR spectrum (Fig. 2). The sharp peak appearing at 1767 cm-1 was
assigned to the carbonyl (C=O) group in the PLA-PEG copolymer (Xiong et
al., 2011).

The peaks that appeared in the regions of 1193 cm-1 and 1460 cm-1 are
related to stretching C-O and the bending of -CH2- groups in PLA-PEG
copolymer, respectively (Amani, Kabiri, Shafiee, \& Hamidi, 2019;
Danafar, Rostamizadeh, Davaran, \& Hamidi, 2017).

Moreover, the peaks at 1640 and 1095 cm-1 corresponds to the C=O and
C-O-C stretching of COOH in the FTIR spectrum of PEG, PLA- PEG-FA and
PLA-PEG-Glu suggested that PEG was grafted to the nanoparticles (Ahmed,
Arfat, Castro-Aguirre, \& Auras, 2016; Rafat et al., 2010).

The FTIR spectra of PLA-PEG-FA copolymer showed that there is also a
broad peak in the region between 1580 and 1648 cm\selectlanguage{english}-1 belonging to the
amine group of FA. Moreover, the peak observed around 2108 and 2116 cm-1
were derived from azido group of FA and Glu respectively. The peak at
3580 cm-1 corresponds to the O-H stretching of OH in the FTIR spectrum
of glucose suggested that glucose was grafted to the PLA-PEG.

The FT-IR spectra of FeCo, FeCo-PEI, FeCo-PEI-PLA-PEG-FA, and
FeCo-PEI-PLA-PEG-Glu are illustrated in Fig. 2.

After conjugation of PEI with FeCo nanoparticles, the PEI peaks from
1050 cm-1 to 1250 cm-1 were attributed to the amide and amide amino
groups , which indicated that the PEI was attached on the surface of
FeCo nanoparticles (Fig. 2) (Gultekinoglu et al., 2015).

As also evident in Fig. 2, almost all peaks related to PLA-PEG,
FeCo-PEI, glucose and folic acid were observed in the NPsA and NPsB,
therefor it is indicated the successful synthesis of these
nanoparticles.

Thermogravimetric analysis (TGA) of NPsA and NPsB was performed to find
out different chemical compounds in the samples. The TGA curves of PEG,
PEI, PLA, PLA-PEG, PLA-PEG-FA and CBP-FA copolymers, are shown in Fig.
3.

TGA analysis of the pure polymers (PEI, PLA and PEG) shows single step
degradation while the spectrum of PLA-PEG copolymer exhibits three steps
degradation related to the presence of three components in the PLA-PEG
copolymer. The initial mass loss in the range of 175 degC to 305 degC, may
be due to the loss of adsorbed water molecules from the copolymer
(Jadhav et al., 2013).

The second stage of the PLA-PEG degradation occurs between 320 degC and
380 degC, which is related to the thermal degradation reaction of the PLA
polymer (Shih \& Huang, 2011). The third stage of thermal degradation
occurs at temperatures above 380 degC, is associated with degradation of
PEG polymer (Kwon \& Kim, 2006).

The thermogram of the --FA and --Glu shows several mass losses during
the temperature rise from 200 to which proposes that the glucose has
higher temperature resistance, compared to other pure PEI, , and folic
acid and therefore, temperature resistance of PLA-PEG copolymer
increased after binding to glucose.

As it is seen in Fig. 3 the FeCo-PEI nanoparticles has a much better
thermal stability than that of PLA-PEG-FA and PLA-PEG-Glu. Therefore,
improvement in temperature resistance of PLA-PEG-FA and PLA-PEG-Glu was
observed after modification with FeCo nanoparticles, which indicated the
successful conjugation of PLA-PEG-FA and PLA-PEG-Glu with FeCo.

\textbf{Preparation and characterization of nanoparticles}

The results of DLS indicated that there was no significant difference
between the particle size of NPsA, NPsB and NPsAB (the hydrodynamic
diameter of NPsA, NPsB and NPsAB was about 82, 94 and 88 nm
respectively). However, the hydrodynamic diameter of the nanoparticles
increased after being conjugated with siRNA and PTX (Fig.~4). For
example, the particle size of the NPsA increased from 88 nm to 120 nm
after being encapsulated with siRNA and PTX. The surface charge of NPsA,
NPsB and NPsAB was significantly reduced after being encapsulated with
siRNA and PTX. It seems that the negative charge of phosphate backbone
of the siRNA leads to decrease in the charge of the nanoparticles. A
previous study found that the nanoparticles with a slightly negative
surface charge and average particle sizes ranging between 100 and 150 nm
have much higher transfer efficiency than larger nanoparticles (Bala,
Hariharan, \& Kumar, 2004; C. He, Hu, Yin, Tang, \& Yin, 2010; Panyam \&
Labhasetwar, 2003). Although nanoparticles smaller than 100 may have
higher transfection efficiency than nanoparticles larger than 100 nm but
the high nonspecific uptake of these nanoparticles could increase the
side effects on healthy tissue (Puppo et al., 2014).

The transmission electron microscopy (TEM) of the nanoparticles showed
that the NPsA/siRNA/PTX, NPsB/siRNA/PTX and NPsAB/siRNA/PTX
nanoparticles have spherical shape with uniform size distribution and
smooth surface (Fig. 5). There was no significant difference between the
morphology of the nanoparticles. The results also showed that the
average size of the nanoparticles was about 100 nm.
These~results~were~consistent~with~DLS~data (Fig. 4).

\textbf{Magnetic property determination}

Saturation magnetization of nanoparticles ranged from 70.6 to 165.2
emu/g. The FeCo nanoparticles showed high saturation magnetization
(165.2 emu/g), however these quantities decreased after coating the FeCo
with PEI. The values for saturation magnetization of FeCo and FeCo-PEI
were determined as 165.2 and 131.4 emu/g respectively (Figure 6 A).

The saturation magnetization values for NPsA, NPsB, and NPsAB were lower
than the FeCo and FeCo-PEI nanoparticles (99.41, 107.84 and 94.8 emu/g
for NPsA, NPsB, and NPsAB nanoparticles respectively) (Fig. 6 A). Such
attenuation pattern of saturation magnetization is often seen in
nanoparticles after coating the nanoparticles with non-magnetic
compounds.(Pouponneau, Leroux, \& Martel, 2009). In general, the
saturation magnetization of nanoparticles was reduced with encapsulation
of siRNA and PTX into NPsA, NPsB and NPsAB nanoparticles (Fig. 6 A).
However, there was no significant difference between the magnetic
properties of the nanoparticles after PTX and siRNA encapsulation.
Magnetite properties of nanoparticles is an important factor in imaging
and targeting using magnetic nanoparticles. Previous studies showed that
nanoparticles with magnetite properties up to 9 emu/g are suitable for
use in imaging and targeting nanoparticles (Chorny et al., 2010; Yan,
Li, Li, Zhu, Shen, Yi, Wu, Yeung, Xu, Xu, \& others, 2014). Therefore,
it seems that the nanoparticles synthesized in this experiment appear to
be suitable for such purposes.

The results of siRNA-FAM and PTX encapsulation efficiency in NPsA, NPsB
and NPsAB are shown in Fig. 7. Generally, the encapsulation efficiency
of siRNA-FAM was significantly higher than that of PTX. The
encapsulation efficiency of siRNA-FAM in various nanoparticles was in
the range of 78 to 85\% while, in the case of PTX, it was in the range
of 38 to 42\%. There is no significant difference between encapsulation
efficiency of the nanoparticles in each group (Fig. 7). The maximum
difference between the high and low encapsulation~efficiency values of
siRNA-FAM and PTX~loaded samples was 7 and 4\% respectively (Fig. 6 B).
It seems that electrostatic interactions between the negative charge of
siRNA-FAM molecules and the positive charge of polyethyleneimine in the
nanoparticles is responsible for the increased values of siRNA-FAM
encapsulation efficiency compared to PTX.

There was no significant difference between release behavior of PTX and
also siRNA from NPsA/siRNA/PTX, NPsB/siRNA/PTX and NPsAB/siRNA/PTX
nanoparticles. However, the cumulative release percentage of siRNA from
nanoparticles were significantly lower than PTX after 30 days. The
cumulative release percentage of siRNA was found to be significantly
lower than that of PTX after 30 days incubation at 37 degC in PBS buffer
(Fig. 7). It seems that the electrostatic interaction between siRNA and
PEI of the nanoparticles may be responsible to this phenomenon.
Therefore, it is plausible that electrostatic interactions between the
phosphate groups of siRNA and the NH2 groups of the nanoparticles
resulted in a reduction in the release rate of siRNA from the
nanoparticles.

The~PTX and siRNA~release~from~the~nanoparticles~was observed in two
different stages. In the first stage, more than 50\% of the PTX-loaded
into the nanoparticles was released in the first three days. It was
found that the rate of PTX release decreased significantly with
increasing the incubation time more than three days. A similar release
pattern was observed for siRNA. These results are in agreement with
previous studies (Abebe et al., 2015b; Perez et al., 2001b). Those
reports have attributed the sustained release property of the
nanoparticles to the PLA segment in the structure of the nanoparticles.
This property of nanoparticles can be used for intracellular controlled
drug release in tumor cells.

\textbf{In vitro cytotoxicity and transfection assay}

In vitro cytotoxicity of blank nanoparticles and siRNA-FAM/PTX
encapsulated nanoparticles were compared after 48 h. As it was seen in
Fig. 8A, the siRNA/PTX-free nanoparticles (NPsA, NPsB, and NPsAB) showed
good biocompatibility, and no remarkable cytotoxicity was observed after
48\selectlanguage{english} h of incubation with MCF-7 cells. However, its toxicity significantly
increased after encapsulation of siRNA/PTX in the nanoparticles
(NPsA/siRNA/PTX, NPsB/siRNA/PTX, and NPsAB/siRNA/PTX). As it can be seen
in Fig. 8B, siRNA has no toxicity effect on MCF-7 cells. Therefore, it
can be concluded that the decrease in cell viability might be due to the
effect of PTX on MCF-7 cells. The IC50 value of NPsAB/siRNA/PTX was
significantly lower than NPsA/siRNA/PTX and NPsB/siRNA/PTX after 48 h of
incubation. Several studies have demonstrated the effect of folic acid
and glucose conjugated nanoparticles to increase drug transfer
efficiency (Jain, Rathi, Jain, Das, \& Godugu, 2012; Macheda, Rogers, \&
Best, 2005; Zhang et al., 2013).

However, this improvement could be continuing until the saturation of
folic acid receptors on the surface of cancer cells by nanoparticles. In
other word, after saturating the folic acid receptors by nanoparticles,
nonspecific uptake of nanoparticles into healthy tissues increases. The
higher cytotoxicity of free PTX may be due to unrestricted and rapid
diffusion of PTX into MCF-7 cells, while the nanoparticles are
internalized into cancer cells via the receptor-mediated endocytosis
pathway (S. Wang et al., 2013). In addition, sustained release of PTX
from these nanoparticles due to the presence of polylactic acid in their
structure may be another reason for low cytotoxicity of the
nanoparticles compared to free PTX.

Fluorescence microscopy was used to evaluate the ability of the
nanoparticles for co-delivery of nucleic acid and drug into MCF-7 cells.

Fig. 8 C-E shows fluorescence images of MCF-7 cells after 48 h of
incubation with NPsA/siRNA/PTX, NPsB/siRNA/PTX and NPsB/siRNA/PTX
nanoparticles. The fluorescence emission observed within the cells
indicated the ability of the nanoparticles to transfer siRNA-FAM into
MCF-7 cells. Moreover, the circular morphology of MCF-7 cells after
incubation with the nanoparticles may be happened since the induction of
apoptosis in MCF-7 cells byNPsA/siRNA/PTX, NPsB/siRNA/PTX and
NPsB/siRNA/PTX nanoparticles.

\textbf{Conclusions}

In this study, conjugated copolymers of NPAB/siRNA/PTX were successfully
synthesized for targeted gene delivery to breast cancer cells. Their
chemical structure was confirmed through NMR and FTIR analysis and
magnetic properties determination. The results obtained from TEM and DLS
revealed that the nanoparticles have spherical morphology, with the
hydrodynamic diameter in the range between 100-150 nm, which is suitable
for intravenous injection and \emph{in vivo} applications. Invitro drug
release showed a sustained release manner of PTX and siRNA, however the
cumulative release of PTX is more than siRNA. The cytotoxicity assay
demonstrated no significant toxicity of unloaded nanoparticles on MCF-7
cells, however PTX/siRNA loaded nanoparticles caused significant
toxicity. In addition, the ability NPAB to improve the transfection
efficiency was confirmed through increased uptake of siRNA-FAM in MCF-7
cells via fluorescence microscopy imaging.

\textbf{Acknowledgements}

This study has been supported by the Deputy Research and Technology,
Ardabil University of Medical Sciences.

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Fig. 1. 1H NMR spectra of PLA-PEG-FA (A) and PLA-PEG-Glu (B) in D2O
solvent

Fig. 2. FT-IR spectrum of various copolymers and magnetic nanoparticles

Fig. 3. TGA thermograms of magnetic nanoparticles and different
copolymers

Fig. 4. the particle size and zeta potential of the both PTX/siRNA
encapsulated and blank nanoparticles

Fig. 5. TEM image of NPsA/siRNA/PTX, NPsB/siRNA/PTX and NPsAB/siRNA/PTX
nanoparticles

Fig. 6. Magnetic behavior (A) and encapsulation efficiency (B) of the
nanoparticles

Fig. 7. DNA release profiles of the nanoparticles in PBS buffer (pH=7.4)
(Mean ± standard deviation) (n = 3).

Fig. 8. Biocompatibility of nanoparticles (A) and the effect of PTX and
siRNA loaded into the nanoparticles on MCF-7 cells (B) Fluorescence
microscopy images of MCF-7 cells were treated with NPsA/PTX/siRNA,
NPsB/PTX/siRNA and NPsAB/PTX/siRNA after 48 hours incubation.

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