PEGylated Chitosan Double Layer toward a More

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Natural Salep/PEGylated Chitosan Double Layer towards a More Sustainable pH-Responsive Magnetite Nanocarrier for Targeted Delivery of DOX and Hyperthermia Application Nasrin Zohreh, Sakineh Alipour, Seyed Hassan Hosseini, Morena S. Xaba, Reinout Meijboom, Mahdi Fasihi Ramandi, Nazila Gholipour, and Mehdi Akhlaghi ACS Appl. Nano Mater., Just Accepted Manuscript • DOI: 10.1021/acsanm.8b02076 • Publication Date (Web): 11 Jan 2019 Downloaded from http://pubs.acs.org on January 13, 2019

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ACS Applied Nano Materials

1

Natural Salep/PEGylated Chitosan Double Layer towards a More Sustainable pH-

2

Responsive Magnetite Nanocarrier for Targeted Delivery of DOX and Hyperthermia

3

Application

4 5

Nasrin Zohreh,a* Sakineh Alipour,a Seyed Hassan Hosseini,b Morena S. Xaba,c Reinout

6

Meijboom,c Mahdi Fasihi Ramandi,d Nazila Gholipour,e,f Mehdi Akhlaghig

7 8

aDepartment bDepartment

of Chemical Engineering, University of Science and Technology of Mazandaran, Behshahr, 48518,

9 10

Iran cDepartment

of Chemistry, University of Johannesburg, P.O. Box 524, Auckland Park, 2006 Johannesburg, South

11 12

Africa dMolecular

Biology Research Center, system biology and poisonings institute, Baqiyatallah University of Medical

13 14

Sciences, Tehran,1435116471, Iran eChemical

Injuries Research Center, Systems Biology and Poisonings Institute, Baqiyatallah University of Medical

15

Sciences, Tehran, 1435116471, Iran

16 17

of chemistry, Faculty of Science, University of Qom, P. O. Box: 37185-359, Qom, Iran

fFaculty gResearch

of Pharmacy, Baqiyatallah University of Medical Sciences, Tehran, 1435116471 Iran

Center for Nuclear Medicine, Tehran University of Medical Sciences, Tehran, 1414713135, Iran

18 19

Abstract

20

Two pH responsive polysaccharide-based magnetite nanocarriers have been developed from

21

modified salep and PEGylated chitosan decorated onto the surface of magnetite nanoparticles for

22

active loading and targeted delivery of doxorubicin. The first nanocarrier was formed via imine

23

bond formation and electrostatic hydrogen bonding interactions between dialdehyde salep-

24

modified magnetite nanoparticles and DOX in a one-layered nanocarrier fashion. Subsequently,

25

PEGylated chitosan was used as a second layer on the surface of the first nanocarrier, mainly by

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the same interactions, producing a double-layered polysaccharide-based nanocarrier. In vitro

27

release studies indicated that the use of PEGylated chitosan as the second shell provides more

28

control on the rate and amount of DOX release and makes double-layered nanocarrier more pH-

29

sensitive. Magnetic heating capacity of double-layered nanocarrier was investigated and release

30

profile under AMF (42 C, magnetically induced) showed a good improvement in the time and

31

amount of DOX release. In vitro MTT assays depicted that the DOX-loaded double-layered

32

nanocarrier produces a large cytotoxic response to HeLa cells, which is comparable with free DOX

33

in higher concentration. DOX-free double-layered nanocarrier also exhibited low cytotoxicity

34

against normal cells, an indication of their excellent biocompatibility. Application of AMF also

35

showed a large effect on the cytotoxicity of the nanocarrier rather than without AMF condition.

36

Investigation on cellular uptake of nanocarrier revealed the high targeting ability of nanocarrier

37

toward HeLa cells, especially in the presence of AMF. Blood compatibility investigations

38

indicated that hemolysis of RBCs and coagulation times fall in to the normal range for blood in

39

the presence of DOX-free and DOX-loaded double-layered nanocarrier.

40 41

Keywords

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pH responsive nanocarrier; smart delivery; doxorubicin; salep; PEGylated chitosan; magnetic

43

hyperthermia

44 45

1. Introduction

46

Along with the dramatic appearance of cancer in human societies, the use of chemotherapy drugs

47

has been increased. Accordingly, design and application of different delivery methods for

48

chemotherapeutic drugs are rapidly developed.1-5 In this regard, fabrication of a powerful drug


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49

delivery system (DDS) is vital to reduce the side effects caused by chemotherapy drugs. Smart

50

drug carriers are able to distinguish the pathophysiological differences between normal and tumor

51

cells.6-8 The delivery of anticancer drugs by incorporating it to a smart carrier reduces the side

52

effects and enhances the solubility, stability and circulation time of drugs.7, 9-11

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Among all the features that a suitable nanocarrier should have, biocompatibility and

54

biodegradability are two more important aspects that always should be considered.12-14 Until now,

55

many biodegradable and biocompatible synthetic and natural polymers including polyesters,15

56

polyamides,16

57

polysaccharides,22-23 have been used in the structure of smart nanocarriers. It seems that the

58

application of natural polysaccharides in DDS is more attractive24-25 because they are easy

59

accessible (they are non-synthetic) and exhibit renewability, high stability, non-toxicity, good

60

hydrophilicity, and biodegradability in the human body.24, 26 Moreover, the unique structure of

61

polysaccharides allows many non-covalent interactions with biological tissues, which resulted in

62

higher cellular uptake for polysaccharide based carriers.27-28 On the other hand, most of the natural

63

polysaccharides contain only hydroxyl and carboxyl groups which cannot easily react with drug

64

molecules.29 Therefore, for practical application of polysaccharides in DDS, their backbone should

65

be tailored by some chemical modifications.30-32

66

So far, different polysaccharides such as chitosan,31, 33-34 starch,30 hyaluronic acid,22 dextran,35 and

67

etc. have been used in different types of drug nanocarriers such as core-shell nanoparticles,36

68

hydrogels,25, 37 micelles,15, 38 vesicles,23 and nanocapsules.39 Salep, a commonly additive in the

69

food industries, is a low-cost and newfound polysaccharide in DDSs.40-41 A few researchers have

70

reported on the use of salep in the design and synthesis of smart responsive drug nanocarriers.42-45

polyethylene

imines,17-18

acrylic

polymers,19-20


polyglycerols,21

and


71

In the last decade, there has been an increasing attention in the exploring of core-shell nanocarriers

72

so that they have emerged as unique class of nanocarriers, which is along with the concept of

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“surface engineering”.46-48 However, the fabrication of nanocarriers with a magnetite core and

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polysaccharide shells has been given less attention.49 This is while such a nanocarrier could

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provide high degree of biocompatibility, low-cost accessibility and the possibility of using

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combination of cancer therapies such as hyperthermia along with the MRI imaging and

77

chemotherapy.14, 36, 50-51

78

In this work, we developed a magnetite-based nanocarrier in which MNP nanoparticles were

79

initially coated by dialdehyde salep to prepare a new class of pH-responsive nanocarrier to targeted

80

delivery of doxorubicin. Then the drug loaded nanoparticles were functionalized with

81

poly(ethylene glycol) modified chitosan, PEGylated chitosan, as a second protective shell to

82

reduce the drug release and obtain a better control on the rate and amount of DOX release.

83 84

2. Experimental

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2.1. Materials and methods

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Iron chloride tetrahydrate (FeCl2.4H2O), iron chloride hexahydrate (FeCl3.6H2O), 3-

87

aminopropyltriethoxysilane (APTS), and NH4OH (30%) and were obtained from Merck. Sodium

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metaperiodate (NaIO4), monomethoxy poly(ethylene glycol) (mPEG, Mw=2000), and chitosan

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(medium molecular weight) were obtained from Sigma Aldrich. Salep (Mn = 1.17 × 106 g/mol, Mw

90

= 1.64 × 106 g/mol (high Mw), PDI = 1.39, eluent = water, flow rate = 1 mL/min, acquisition

91

interval = 0.43 s from GPC results) was obtained from a supplier in Kordestan, Iran. DOX.HCl

92

was purchased from Pfizer, China. HeLa cells were obtained from Iran Pasteur Institute, Tehran,

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Iran fetal bovine serum (FBS) and DMEM medium were obtained from Biochrom AG, Germany.


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94

[3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide] (MTT) and DMSO were

95

purchased from Sigma.

96

UV-vis spectra were recorded using a Perkin-Elmer spectrophotometer. An ABB Bommem MB-

97

100 spectrometer (Canada), and KBr pellet of samples were used to perform FT-IR spectra.

98

Thermogravimetric analysis (TGA) was acquired under using a TGA Q50 thermo-gravimetric

99

analyzer (N2 atmosphere, heating rate 10 ˚C min-1). Powder XRD analyses were performed using

100

a Rigaku MiniFlex-600 diffractometer with Cu Kα radiation (λ = 1.5406 Å) at room temperature.

101

High resolution transmission electron microscopy (HR-TEM) was performed on a JEOL JEM-

102

2100F electron microscope with an accelerating voltage of 200 kV. Hydrodynamic size

103

measurements were performed by using Zetaplus/90plus instrument (Brookhaven Instrument Co.,

104

USA). Molecular weight of polymers were determined by Agilent 1100 gel permeation

105

chromatograph (GPC) equipped with a refractive index detector using DI water as the eluent at a

106

flow rate of 1 mLmin-1 at 30 C. Dextran was used as standard sample for molecular weight

107

measurement in GPC. Magnetizations of samples were measured by vibrating sample

108

magnetometer (Meghnatis Daghigh Kavir Co., Kashan, Iran).

109 110

2.2. Synthesis of dialdehyde salep (DAS)

111

DAS obtained from oxidation of natural salep by sodium metaperiodate (NaIO4) according to the

112

previously reported method.52 Briefly, 1.5 g of sodium periodate dissolved in 10 mL DI water and

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then, sulfuric acid solution (0.01 M) was dropwise added to the mixture until the pH adjusted to

114

1.5. This solution applied for oxidation reaction of salep. In this regard, 1.0 g of salep dissolved in

115

50 mL DI water under vigorously stirring at room temperature and 3 mL NaIO4 solution and 15

116

mL DI water were added to the solution. Oxidation reaction proceed for 3.5 h at 40 °C in a dark



117

place. At the end, oxidized salep (oxidation degree of 35%)52 precipitated by adding acetone to the

118

reaction mixture. The precipitate was separated and washed with water and acetone. Pure product

119

was obtained using dialysis bag technique. Oxidation percentage was calculated according to

120

literature methods.52

121 122

2.3. Synthesis of chitosan grafted methoxypolyethylene glycol (CSP)

123

Jones reagent was prepared53 for oxidation of methoxypolyethylene glycol (mPEG-OH) to mPEG-

124

CO2H. Briefly, 330 mg (3.3 mmol) CrO3 added into the 10 mL flask containing 2.3 mL DI water.

125

After complete dissolution, the flask placed into ice-water bath and stirred for 10 minutes.

126

Concentrated H2SO4 (0.24 mL) was added into the mixture and stirring continued for 10 minutes.

127

Afterward, the as prepared Jones reagent was slowly added to mPEG solution (5.0 g, 3.3 mmol, in

128

50 mL acetone) in a 100 mL round bottom flask and the mixture was stirred at room temperature.

129

After 24 h the reaction was quenched by adding 1 mL isopropyl alcohol. Reaction mixture

130

containing solid green chromium salts was filtered and purified by silica gel column

131

chromatography with methanol as eluent (4.25 g of mPEG-CO2H).

132

The obtained mPEG-CO2H dissolved in 50 mL methanol and p-toluene sulfonic acid (5%) was

133

added to the solution. Then, the reaction mixture was refluxed to complete esterification for one

134

week. Afterward, the reaction mixture was neutralized by Na2CO3 and concentrated via rotary

135

evaporator. The residual viscose fluid dissolved in 10 mL DI water and extracted with 50 mL

136

(5×10 mL) dichloromethane. The organic phase was concentrated by rotary evaporator and mPEG-

137

CO2Me precipitated after drying in open-air flow. The precipitate washed three times with cold

138

diethyl ether (4.0 g of pure mPEG-CO2Me).


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In the final step, 4.0 g of chitosan dissolved in 120 mL acetic acid (0.02%) and 4.0 g of mPEG-

140

CO2Me was added slowly for 15 minutes. The amidation process continued at 60-70 °C under

141

magnetic stirrer for one week. Next, acetone was added into the reaction mixture and mPEG-g-CS

142

(CSP) was precipitated. The precipitate was separated by decanting the supernatant solution and

143

washed several times with DI water followed by acetone and dried under air flow (5.7 g).

144 145

2.4. Synthesis of amine functionalized magnetite nanoparticles

146

The Fe3O4 magnetite nanoparticles (MNP) were prepared by co-precipitation method according to

147

our previous methods.54-55 In this regard, FeCl2.4H2O (4.3 g, 21.5 mmol) and FeCl3.6H2O (11.4 g,

148

43 mmol) were dissolved in 300 mL DI water in a 500 mL three necked round bottom flask under

149

nitrogen at room temperature. A solution containing 2.0 g NaOH and 250 mL ammonia was added

150

dropwise until the pH of mixture adjusted to 12. After continuously stirring for 2 h, obtained black

151

precipitate separated by an external magnet and washed with DI water until the pH of the

152

supernatant reached to 8 and then, washed two times with ethanol and dried at 60 C under vacuum.

153

In order to functionalize the MNPs surface with amine groups, 2.0 g of MNP was well dispersed

154

in 100 mL H2O:ethanol mixture (1:4) in a 100 mL round bottom flask. Next, 1.0 mL of ammonia

155

and 5 mL of (3-aminopropyl) triethoxysilane (APTS) was added into the mixture and the mixture

156

was vigorously stirred under reflux at 80 C. After 24 h, MNP@APTS nanoparticles were

157

separated (by magnet), washed with EtOH (3×100 mL) and dried at 60 C under vacuum.

158 159

2.5. Synthesis of MNP@DAS@DOX (nanocarrier 1) and MNP@DAS@DOX@CSP (nanocarrier

160

2)



161

Magnetite nanocarriers were synthesized using the following procedures. MNP@APTS (0.2 g)

162

was well dispersed in 200 mL of DI water and aqueous solution of DAS (200 mg) was slowly

163

added into the mixture under vigorous stirring at room temperature. After 5 h the MNP@DAS was

164

separated with an external magnet, washed with DI water and EtOH, and dried under vacuum at

165

60 C.

166

For DOX loading, 5 mg of MNP@DAS was ultrasonically dispersed in 3 mL PBS. Then 7 mL

167

solution of DOX with different concentration (0.05, 0.1, 0.2, 0.3, 0.5 mg.mL-1 DOX) was added.

168

The mixture was stirred in the dark at room temperature for 24 h. Afterwards, magnetite

169

nanocarrier 1 (MNP@DAS@DOX) was separated by magnetic decantation and washed with PBS

170

solution to remove unloaded drug.

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The doxorubicin concentration in supernatants was tracked by UV-Vis absorption at a wavelength

172

of 480 nm. The DLC (drug loading content) and DLE (drug loading efficiency) were calculated

173

by:

174 175

DLC =

𝐼𝑛𝑖𝑡𝑖𝑎𝑙 𝑎𝑚𝑜𝑢𝑛𝑡 𝑜𝑓 𝐷𝑂𝑋 ― 𝐹𝑟𝑒𝑒 𝑎𝑚𝑢𝑜𝑛𝑡 𝑜𝑓 𝐷𝑂𝑋 𝑖𝑛 𝑠𝑢𝑝𝑒𝑟𝑛𝑎𝑡𝑎𝑛𝑡 𝑠𝑜𝑙𝑢𝑡𝑖𝑜𝑛 𝑁𝑎𝑛𝑜𝑐𝑎𝑟𝑟𝑖𝑒𝑟 𝑚𝑎𝑠𝑠

176

177

DLE (%) =

𝐼𝑛𝑖𝑡𝑖𝑎𝑙 𝑎𝑚𝑜𝑢𝑛𝑡 𝑜𝑓 𝐷𝑂𝑋 ― 𝐹𝑟𝑒𝑒 𝑎𝑚𝑢𝑜𝑛𝑡 𝑜𝑓 𝐷𝑂𝑋 𝑖𝑛 𝑠𝑢𝑝𝑒𝑟𝑛𝑎𝑡𝑎𝑛𝑡 𝑠𝑜𝑙𝑢𝑡𝑖𝑜𝑛 𝐼𝑛𝑖𝑡𝑖𝑎𝑙 𝑎𝑚𝑜𝑢𝑛𝑡 𝑜𝑓 𝐷𝑂𝑋 × 100

178 179

Then, the surface of MNP@DAS@DOX nanoparticles (nanocarrier 1) was coated with CSP as a

180

second shell. Briefly, 8 mg MNP@DAS@DOX was dispersed in a flask containing 3 mL DI water.

181

Then, CSP solution (10.0 mg in 5 mL) was added into the mixture and allowed to stir at room

182

temperature for 2 h. Afterwards, MNP@DAS@DOX@CSP nanoparticles (nanocarrier 2) was 8 ACS Paragon Plus Environment

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magnetically separated and washed four times with DI water and dried at room temperature under

184

vacuum.

185 186

2.6. In vitro pH-triggered drug release at 37 C and 42 C

187

To investigate the in vitro release of DOX, the release profiles of DOX-loaded nanocarriers were

188

investigated at 37 or 42 C in two different media; 5 mg of DOX-loaded nanocarrier 1 or 2 were

189

suspended in 5 mL of PBS (pH=7.4) or acetate buffer (pH=5.5), and the mixture was immediately

190

poured into a dialysis bag (MWCO 3500 Da). The end-sealed dialysis bag was placed in 10 mL of

191

the same buffer medium in a heating bath to initiate the release experiments (37 or 42 C and speed

192

of 100 rpm). At a specified time, 3 mL of solution was removed from outside the dialysis bag and

193

fresh buffered solution of the same volume was replaced with. The amount of released DOX was

194

determined using a UV-Vis spectrometer at a wavelength of 480 nm.

195 196

2.7. Investigation on magnetically induced heat generation of nanocarrier 2

197

To evaluate magnetically induced heating effect, magnetite nanoparticles was dispersed in 1 mL

198

water and the suspension was subjected to alternating magnetic field (AMF) of MagneTherm

199

system (400 kHz, 23.9 kA.m-1 or 300 Oe, 10 min) equipped with a fluoroptic fiber thermometer.

200

The temperature differences in the sample holder was estimated within ±1.0 C. The sample

201

temperature was measured as a function of heat induced by applied AMF.

202

AMF-induced heating ability of magnetite nanoparticles was defined by specific absorption rate

203

(SAR). SAR value was calculated by the initial linear increase in temperature (dT) per time interval

204

(dt) according to:

205



206

𝑆𝐴𝑅

(𝑊 𝑔) = 𝐶 ×

𝑀𝑑 𝑀𝐹𝑒

×

Page 10 of 45

𝑑𝑇 𝑑𝑡

Eq. 1

207 208

where C is the specific heat of the medium (Cwater=4.18 J/g℃), Md the summation of water and

209

nanoparticle mass, and MFe (g/L) is the magnetite mass (Fe3O4), and dT/dt is the slope at initial

210

times after applying the magnetic fields. The intrinsic loss power (ILP) was determined by:

211 212

SAR

Eq. 2

ILP = 𝐻2𝑓

213 214

2.8. In vitro pH/AMF-triggered drug release

215

To investigate AMF-stimulated drug release, 1 mg of Fe3O4@DAS@DOX@CSP was dispersed

216

in 1 mL buffer solution (pH=7.4 or 5.5) transferred into the eppendorf inside the coil of the

217

MagneThermTM. The tube were then subjected to an AMF (23.9 kA/m, 400 kHz) for 10 h. The

218

sample temperature increased up to (42 ± 1) °C through 6 min (hyperthermia temperature). The

219

temperature was then controlled by the field amplitude (H) to be maintained constant. . At the

220

specified time intervals, the AMF was switched off and 0.5 mL supernatant was immediately

221

extracted (magnetic separation) to determine the amount of released DOX by UV-Vis

222

spectroscopy. Before analysis the DOX content of the extracted solution, fresh-buffered solution

223

of the same volume was replaced while AMF was switching on.

224 225

2.9. In vitro evaluation of cell viability

226

In vitro cytotoxicity of nanoparticles was evaluated using HeLa cells (Cervical cancer) and HEK-

227

293 (human embryo kidney, normal cells). Cells were grown in DMEM containing 100 mg/mL

228

penicillin G, 5.0% fetal bovine serum and 100 mg/mL streptomycin at 37 C (humidified 5.0% 10 ACS Paragon Plus Environment

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229

CO2 atmosphere). Then, cells were seeded in 96-well plates with a concentration of 8×103-9×103

230

cells per wall. The samples (different concentrations of nanoparticles) were then sterilized by UV

231

and added to the culture wells. For control experiment, cells were regarded without any further

232

treatment. The solution was incubated for 16 or 24 h at 37 ˚C and detached from the culture wells.

233

For AMF application, the cell suspension was then placed into plastic tubes, and exposed to an

234

AMF (23.9 kA/m, 400 kHz) for 2 h. Cells were re-seeded in 96-well plates containing fresh

235

medium for another 24 h. For static magnetic field treatment, one set of 24 well plates was placed

236

under the effect of magnetic field during the incubation. MTT assay was used to determine cell

237

viability. The cells were washed twice with saline and then 100 μL of MTT solution (0.5 g/L) was

238

added to each well and the plate was incubated for 4 h. In this condition, viable cells reduce the

239

MTT to DMSO-soluble formazan. Then, the medium containing unreacted dye was removed

240

carefully. The obtained purple formazan crystals were dissolved in 200 mL per well DMSO and

241

the absorbance was measured at a wavelength of 490 nm. All the experiments were performed in

242

triplicate. The inhibition of cell growth was calculated by:

243 244

Cell viability (%) =

Mean of abs.value of treatment sample Mean of abs.value of control sample

× 100

245 246

2.8. Hemolysis assays

247

The hemolysis of human erythrocytes was studied in the presence of nanoparticle. In this assay

248

human erythrocytes were collected from healthy individuals in EDTA containing tubes. The

249

erythrocytes were washed three times to clear the supernatant from hemoglobin released from

250

lysed erythrocytes and then 20% (v/v) erythrocytes/PBS suspension was prepared. This suspension

251

was diluted to 10% by adding 100 μL of diluted series of nanoparticles (32-1000 mg/L) in triplicate 11 ACS Paragon Plus Environment


252

in a 96-well microliter plate. Complete hemolysis was achieved by suspending erythrocytes in

253

PBS containing 0.2% Triton X-100 (Ryadnov 2002). The plates were incubated for 15 min at 37

254

°C and after centrifugation; the absorbance of the supernatant fluid was measured at a wavelength

255

of 415 nm. The percentage of hemolysis was calculated by following formula: Hemolysis (%) =

256

[

(𝐴𝑛𝑝 ― 𝐴𝑃𝐵𝑆) (𝐴𝑇𝑟𝑖𝑡𝑜𝑛 ― 𝐴𝑃𝐵𝑆)

] × 100

257 258

2.9. In vitro coagulation assay

259

Nanoparticle-induced blood coagulation disorders were studied by measuring the prothrombin

260

time (PT) and activated partial thromboplastin time (APTT). The fresh blood was decalcified (by

261

collecting it into a tube with citrate ions) to prevent the clotting process from starting before the

262

test. A platelet poor plasma was obtained by centrifugation of blood at 1500 g (16 minutes).

263

Nanoparticle solutions (7.25-200 µg/mL) were prepared in platelet poor plasma. The mixture of

264

500 µL of the platelet poor plasma and 50 µL of nanoparticle solution was heated at 37 C for two

265

minutes and then thromboplastin (for the PT assay) or the kaolin and Cephalin (for the APTT

266

assay) and the calcium chloride (excess quantities of ionized calcium) was added to the mixture

267

and the time to fibrin clot formation recorded. The citrated plasma was used as control.

268 269

2.10. Cell uptake studies

270

The intracellular internalization of MNP@DAS@DOX@CSP was evaluated on HeLa cell lines.

271

HeLa cells were seeded in a well plate (105 cells per well) and incubated for 24 h (37 ℃, 5%

272

humidified CO2 atmosphere) until the 70% confluence is etched. Then, MNP@DAS@DOX@CSP

273

was added to each well ([DOX] = 10 μg·mL-1). The culture medium was then incubated for 1, 3

274

and 6 h at 37 ℃. The cells were rapidly washed with phosphate buffer to eliminate any unbound 12 ACS Paragon Plus Environment

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nanocarrier. The cell nuclei were stained blue by DAPI. Then the cells were again washed with

276

saline and fixed with 4% formaldehyde for 15 min. The fluorescence of samples was detected by

277

confocal laser scanning microscopy (Leica TCS SPE). DAPI and DOX were excited at 405 and

278

488 nm, respectively and the emission fluorescence was collected at wavelengths of 440-480 nm

279

and 550-600 nm, respectively.

280 281

3. Result and discussion

282

3.1. Synthesis and characterization of nanocarriers

283

Magnetite polysaccharide-based nanocarriers were fabricated according to scheme 1. Salep was

284

considered as the main polymeric natural backbone because of its cheapness, availability and

285

biocompatibility.45, 56-57 In addition, salep could provide a high degree of non-covalent interaction,

286

mainly hydrogen bonding interactions, with DOX. Initially, salep was partially oxidized to

287

dialdehyde salep (DAS). This would provide aldehyde groups, which is applicable for covalent

288

binding of salep to the magnetite core. Then, magnetite nanoparticles (MNP) were synthesized

289

according to our previous works via co-precipitating of iron salts in basic solution.58 Subsequently,

290

MNPs were functionalized with APTS to obtain amine groups on the surface of magnetite

291

nanoparticles (MNP@NH2). Afterwards, MNP@NH2 were covalently coated by DAS via

292

formation of imine bonds between NH2 and aldehyde groups which was further improved by

293

hydrogen bonding interactions to produce MNP@DAS. Finally, DOX was loaded on the surface

294

of as-prepared MNP@DAS through the imine bond formation and hydrogen bonding interactions

295

to afford nanocarrier 1. In the next step, we thought that nanocarrier 1 could be coated with a

296

second shell to obtain a nanocarrier with further control on DOX release in neutral pH. Natural

297

polysaccharide chitosan was selected because of its cheapness, availability, and biocompatibility



298

that provides many amine and hydroxyl groups.59-60 Accordingly, chitosan was first PEGylated via

299

the reaction of chitosan and methoxycarbonyl poly(ethylene glycol) (mPEG-CO2Me) in acidic

300

solution, denoted as CSP, and then was used as second shell to obtain nanocarrier 2

301

(MNP@DAS@DOX@CSP). The main interactions between nanocarrier 1 and CSP was assumed

302

imine bond and hydrogen bonding interactions. It has been proved that the presence of PEG chains

303

on the surface of nanocarriers strongly increases durability and stability of the nanocarriers in

304

blood circulation and makes them more resistant against the immune system.61-63


Page 14 of 45

Page 15 of 45 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


305 306

Scheme 1 Preparation of nanocarrier 1 and 2.

307 308

FT-IR was used as a technique to analyze the transformations of functional groups (Fig. S1).

309

Functionalization of MNP with APTS was readily confirmed by comparison of FT-IR spectrum

310

of (a) and (b) in the stretching vibrations of C-N bonds area. FT-IR spectrum of DAS (Fig. S1-c)

311

showed distinct bands at 1734, 1617, and 1000-1150 cm-1 attributed to the aldehyde groups, C-O 15 ACS Paragon Plus Environment


312

linkages, and bending O-H, respectively. Synthesis and characterization of DAS from salep was

313

also followed by FTIR and NMR analyses (Fig. S2 and S3, respectively). The FT-IR spectrum of

314

MNP@DAS (Fig. S1-d) showed absorption bands of C-O linkages besides Fe-O (580-620 cm-1)

315

confirming DAS coating; however, the intensity of absorption bands of unreacted aldehyde groups

316

and newly formed imine C=N groups are low and overlapped with the other bands. The FT-IR

317

spectrum of CSP (Fig. S1-e) showed amidic C=O band at 1682 cm-1 accompanied with other

318

related bands. Synthesis of CSP from chitosan and mPEG was also followed by FT-IR, GPC, and

319

NMR analyses (Fig. S2, S4, and S5, respectively). Loading of CSP shell on MNP@DAS or

320

MNP@DAS@DOX (Fig. S1-f and g) mainly resulted in increasing the intensity of absorption

321

bands through 1000-1200 cm-1. More evidence on CSP loading were further ruled out from other

322

analyses (TGA analysis or release experiments). DOX loading also resulted in some minor changes

323

in the FT-IR spectrum of MNP@DAS@DOX@CSP (Fig. S1-g) compared to MNP@DAS@CSP

324

(Fig. S1-f).

325

TGA analyses of MNP@NH2 and MNP@DAS (Fig. 1-I) strongly confirmed loading of APTS and

326

DAS onto the surface of MNP and MNP@NH2. The related weight losses were calculated to be

327

3.5 and 10.8 wt%, respectively. Different ratio of MNP@NH2:DAS were also examined to obtain

328

optimum loading amount for DAS according to IR (Fig. S6) and TG (Fig. S7) analyses (Table S1).

329

TGA analysis of MNP@DAS@CSP showed 2.7 wt% weight loss compared to MNP@DAS

330

attributed to the successful coating of CSP as the second shell. Loading of DOX was confirmed

331

from the large difference in weight loss of MNP@DAS@DOX@CSP compared to

332

MNP@DAS@CSP (Fig. 1-I). Consequently, approximate content of DOX was calculated to be

333

about 112 mg/g.


Page 16 of 45

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334

The XRD pattern of MNP nanoparticles and MNP@DAS@CSP (Fig. 1-II) showed comparable

335

distinct lines align with the reference lines of MNP. This clearly indicates that crystalline structure

336

of MNP retains unchanged during preparation of the nanocarriers. Magnetically induced heating

337

efficacy, which is important for hyperthermia application, is highly dependent on the magnetic

338

property of magnetic nanoparticles. Therefore, the magnetic properties of MNP and

339

MNP@DAS@CSP were investigated using a vibrating sample magnetometer (VSM) at room

340

temperature (Fig. 1-III). The results depicted that the saturation magnetization (Ms) value of

341

MNP@DAS@CSP decreases by 10.7 emu/g after loading of DAS and CSP organic shells. The

342

diminished Ms is related to the decreased effective weight fraction of the magnetite components.

343

However, the magnetization of nanocarrier is still enough to be attracted to an external magnet

344

(Fig. 1-IV).

345

346 17 ACS Paragon Plus Environment


347

Fig. 1. (I): TG analyses of MNP@APTS (a), MNP@DAS (b), MNP@DAS@CSP (c), and

348

MNP@DAS@DOX@CSP (d); (II): p-XRD diffraction patterns; (III): VSM analyses; (IV):

349

magnetic responsibility of MNP and MNP@DAS@CSP.

350

351

The morphology and size distribution of nanocarriers were further investigated by TEM analyses

352

(Fig. 2, a-d). TEM images of MNP (a,b), and MNP@DAS@CSP (c,d) clearly revealed the

353

formation of almost spherical nanoparticles with size distribution of 6-10 nm and 42-52 nm,

354

respectively (Fig. 2e). An Increase in the average size of nanocarriers is due to the coating of MNP

355

with the first and second shells. Dynamic light scattering (DLS) was also performed to investigate

356

the hydrodynamic size of the nanoparticles (Fig. 2-f). The hydrodynamic radius (Rh) of the MNP,

357

MNP@DAS, and MNP@DAS@CSP exhibited narrow width with average of 43 and 58 and 105

358

nm, respectively. Loading of DAS and CSP was also evidenced by the increase in the size of drug-

359

free nanocarriers from 43 (in MNP) towards 58 and 105 in MNP@DAS and MNP@DAS@CSP,

360

respectively. The pH-dependent analysis of hydrodynamic size of MNP@DAS and

361

MNP@DAS@CSP (Insert Fig. 2f) showed the pH sensitivity of both carrier; however,

362

MNP@DAS@CSP showed to be more sensitive as discussed in section 3.2.

363 364


Page 18 of 45

Page 19 of 45 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


365 366

Fig. 2 (a,b) TEM images of MNP and (c,d) MNP@DAS@CSP; (e) particle size distribution; (f)

367

DLS analysis (insert: pH-dependent hydrodynamic size).

368

To investigate the loading of DOX onto the nanocarriers, UV-vis spectroscopy was used (Fig. 3-

369

I). As expected, MNP@DAS@DOX@CSP exhibited a broad absorption band at a wavelength of

370

490 nm due to the characteristic absorption of  in DOX, while MNP, MNP@DAS, and

371

MNP@DAS@CSP did not show considerable absorption in this area. It could be concluded that

372

DOX is properly loaded onto the nanocarriers. To obtain optimum amount of DOX loading on

373

nanocarrier 1, MNP@DAS (5 mg in 3 mL PBS) was mixed with 7 mL of various concentrations 19 ACS Paragon Plus Environment


374

of DOX solution (0.05, 0.1, 0.2, 0.3, 0.5 mg·mL-1). Consequently, DLE and DLC were calculated

375

based on the results of UV-vis absorption of samples (Fig. 3-II). The results indicated that

376

increasing in the concentration of DOX from 0.3 to 0.5 mg·mL-1 does not considerably change the

377

DLC while the DLE is dramatically decreased from 38.7 to 24.8%. Thus, the optimum DLC and

378

DLE of DOX loading for the nanocarrier 1 (MNP@DAS@DOX) were selected as 162.6 mg/g and

379

38.71%, respectively. The high value of DLC could be attributed to the strong hydrogen bond

380

interactions and imine bond formation between DOX and DAS.

381 382

383 384

Fig. 3 UV analyses to evaluate DOX loading on MNP@DAS (I), and DLC/ELC curves vs. DOX

385

concentration (II).

386 387

3.2. In vitro drug release studies triggered by pH and heating bath hyperthermia

388

Imine bonds, and to a lesser extent hydrogen bonding interactions, are stable under physiological

389

condition (pH = 7.4) while they break in lysosomal condition (pH = 5.5) resulted in the release of

390

DOX.64-67 Firstly, in vitro release of DOX from nanoparticles 1 and 2 was evaluated by the dialysis

391

bag diffusion technique at pH=7.4 (phosphate buffer, PBS) and pH=5.5 (acetate buffer) under

392

heating bath 37 C. A solution of dispersed nanocarriers 1 or 2 (5 mg nanocarrier in 5 mL buffer) 20 ACS Paragon Plus Environment

Page 20 of 45

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393

was placed in the dialysis bag. The dialysis bag was immersed in 10 ml of buffer and was stirred

394

in dark place for 80 h maintained at 37  0.5 °C. Samples were withdrawn at regular time intervals

395

(0, 0.5, 1, 2, 3, 4, 5, 6, 7.5, 9.5, 25, 31.5, 49, 55.5, 73, and 80 h), and the same volume was replaced

396

by fresh buffer medium. The samples were analyzed by UV-Vis spectroscopy to determine the

397

amount of DOX release. The DOX release profiles related to nanocarrier 1 and nanocarrier 2 are

398

presented in Fig. 4-I. Some important results could be obtained from the release curves. First, both

399

nanocarriers are pH-sensitive because the total release of DOX from nanocarriers 1 and 2 at

400

pH=5.5 is more than that of pH=7.4. Second and more important, nanocarrier 2 is more pH-

401

sensitive because the difference of DOX-release in acidic and neutral condition is higher (34.5)

402

than that of nanocarrier 1 (5.3 %). Third, DOX release from nanocarrier 2, especially in the early

403

hours, is relatively slower than that from nanocarrier 1 (at both of pH). This is a vital factor for a

404

DOX-nanocarrier to prevent rapidly increasing the DOX concentration in the cellular environment.

405

Fourth, nanocarrier 2 releases less than 45% of loaded DOX in neutral condition (80 h) while this

406

value is about 84.3% for nanocarrier 1. This is also an important factor to prevent the appearance

407

of side effects in non-cancerous cells, which have normal neutral pH. According to literature

408

reports (Table S2), 45% DOX release at pH = 7.4 is within the acceptable range of DOX release

409

from nanocarriers with inorganic core/organic shell (through 80 h). It is important to note that the

410

second shell also prevents the release of DOX from nanocarrier 2 at pH = 5.5;68 however, this

411

reduction is low and the total release reaches to 91% within 135 h which is comparable with the

412

total release of DOX from nanocarrier 1 within 80 h (Fig. 4-I). The results clearly demonstrate the

413

efficiency of nanocarrier 2 relative to nanocarrier 1 and this was our main goal for designing

414

nanocarrier 2 from 1.



Page 22 of 45

415

Scheme S1 shows the proposed mechanisms for DOX-release from both of nanocarriers. At pH =

416

5.5, imine bonds breaks down to produce free DOX consequently protonated in acidic medium.

417

Strong hydrogen bonding interactions between drug and the carrier also weakened and replaced

418

with newly formed hydrogen bonding between H3O+/H2O and drug or shells. Finally, enhanced

419

solubility of DOX and repulsion of positively charged species could decompose the carrier led to

420

the enhanced release of DOX. The release mechanism explained above occurs slowly for

421

nanocarrier 2, because the PEGylated chitosan surrounded the surface of nanocarrier 2 solves and

422

protonates (NH2 to NH3+) slowly in this condition and thus prevents DOX to release rapidly.

423

Indeed, the release condition should firstly effect on PEGylated chitosan led to the delayed release

424

of DOX. This is very important in neutral medium where the release of DOX from nanocarrier 2

425

is significantly reduced compared to nanocarrier 1. Indeed, at pH = 7.4, PEGylated chitosan does

426

not protonate and solve easily led to the significant reduced release of DOX while such an

427

interpretation could not be accounted for nanocarrier 1. In the case of nanocarrier 1, hydrogen

428

bonding interactions between H2O or any basic groups of medium and hydrophilic groups of DAS

429

(OH) leads to the release of high percent of DOX from nanocarrier 1.69-72 It is important to note

430

that the acid-triggered drug release behavior is one of the most important designed mechanism for

431

anti-cancer

432

endosomes/lysosomes are slightly acidic.

433

Temperature/pH dependent doxorubicin release behaviors was then evaluated at 42 C

434

(hyperthermia temperature) using a heating bath. All experimental conditions were similar to the

435

conditions explained in section 3.2 (except temperature). The comparative releasing patterns of

436

nanocarrier 2 are shown in Fig. 4-II. The release profiles at 42 C clearly follow the

437

temperature/pH-dependent manner. As shown in Fig. 4-II, the nanocarrier 2 showed substantially

drug

delivery

systems

as

the

tumor

microenvironment,


and

the

cell

Page 23 of 45 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


438

prolonged release profiles up to approximately 6 days at 37 C, while burst release occurred within

439

2 days at 42 C (both of pH). It seems that the higher temperature accelerates the interruption of

440

hydrogen bonding interactions between DOX and nanocarrier led to the faster release at pH=7.4.

441

In acid condition, the higher temperature also catalyzed the breaking of imine bonds led to the 50.2

442

and 80.1% release of DOX through the first 10 and 50 hours at 42 C while the corresponding

443

values are 21.4% and 62.6% at 37 C.

444

445 446

Fig. 4 (I): Release profile of nanocarrier 1 and 2 at 37 C (inset: enlarged first 10 h); (II):

447

Comparative release profile of nanocarrier 2 under heating bath 37 and 42 C.

448 449

3.3. Magnetic heating capacity and in vitro drug release studies triggered by magnetic

450

hyperthermia

451

Magnetic nanoparticles, such as magnetic drug nanocarriers, generate heat being exposed to an

452

alternating magnetic field, named as “magnetic hyperthermia”. In magnetic hyperthermia therapy,

453

the temperature of the tumors is increased to 41-47 C.73-74 In this condition, healthy cells



454

undergoes reversible damage whereas cancer cells are destroyed. Accordingly, hyperthermia

455

property or magnetic heating capacities of MNP, MNP@DAS@CSP and nanocarrier 2 (0.3, 0.5,

456

1, and 2 mgnanocarrier/mLPBS) were evaluated under an AMF at a frequency of 0.4 MHz (H=300 Oe,

457

H×f = 9.56×109 Am-1s-1). The evolution of the temperature was monitored in situ using a non-

458

metallic fluoroptic fiber thermometer. Fig. 5-I shows the temperature vs. time plot for magnetic

459

nanoparticles with different concentrations. As depicted in Fig. 5-I, increasing the sample

460

temperature is dependent to the concentration and the temperature curve exhibits almost linear

461

shape with increasing the concentration of nanocarrier. The content of MNP in each sample

462

(calculated based on TGA analysis) is cited in Table S3. SAR is generally used to study the ability

463

of a ferrofluid sample to absorb energy from AMF per unit mass, which reflects the heating

464

efficiency of magnetic nanoparticles. According to the initial slope of the resulted curves (Fig. 5-

465

I), the SAR values for 1 mg of MNP, MNP@DAS@CSP, and nanocarrier 2 were calculated to be

466

303, 297, and 280, respectively. It can be concluded that functionalization of MNP with

467

polysaccharides and DOX increases the size of magnetic nanoparticles and reduces the Brownian

468

relaxation led to the reduction of SAR value. On the other hand, higher concentrations of

469

nanocarrier 2 results in the increasing of SAR values and thus the time required to raise the

470

temperature of the sample from 26 to 42 ℃ reduces. Since, SAR is directly related to the frequency

471

and field amplitude (eq. 1) in each experiment,, it could not be compared to other SAR values

472

obtained in a different condition. Intrinsic loss parameter (ILP), on the other hand, is obtained by

473

normalizing the SAR by eq. 2.75 Consequently, ILP is an intrinsic and comparable property. The

474

ILP of nanocarrier 2 for all concentration (Table S3) fall into the normal range (0.15-3.1) of Fe-

475

based nanomagnetic materials specially prepared for hyperthermia applications,76 and

476

commercially available ferrofluids.77-78 The calculated ILP of MNP (1.32) is also comparable with


Page 24 of 45

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477

the reported value for MNP synthesized by coprecipitation methods.73, 79 As shown in Fig. 5-I, the

478

time required to raise the temperature of the sample from 26 to 42 ℃ is about 174, 289, 347, and

479

550 s for samples of 2, 1, 0.5, and 0.3 mgnanocarrier/mLPBS concentration, respectively. According to

480

the SAR value and the temperature increment ability, we selected the concentration of 1.0

481

mgnanocarrier/mLPBS for magnetic hyperthermia experiments.

482

In vitro DOX release studies were then carried out to evaluate the drug release behavior under the

483

hyperthermia conditions induced by AMF. DOX release from nanocarrier 2 at the concentration

484

of 1 mg·mL-1 (SAR = 281 Wg-1) and application of AMF (10 h, both pH) was tracked by UV-Vis

485

spectroscopy. The temperature was kept constant (42±1 C) by controlling the magnetic field

486

strength (H) during the experiment (Fig. 5-I). Release curves were recorded using the absorption

487

intensities of DOX as a function of time. Fig.5-II shows release curves of DOX release under AMF

488

condition compared to release curves under heating bath 37 and 42 C. As shown in Fig. 5-II, at

489

pH 5.5, total release of DOX (10 h) reaches to 89.8% under AMF hyperthermia which is 39.6 and

490

68.3% more than total release of DOX under heating bath 42 and 37 C through the same time (10

491

h). This shows the unique combination effect of AMF hyperthermia and pH on DOX release from

492

magnetic nanocarrier 2. On the other hand, at pH 7.4, the effect of AMF hyperthermia is low and

493

when it compares to heating bath condition at of 37 and 42 C shows only 6.4 and 14.6% for

494

difference release. Apparently, at pH = 7.4 (physiological pH), the imine bond is quite stable, and

495

its hydrolysis is slow. The burst release of DOX at (in all conditions) is because of physically

496

adsorbed DOX onto the surface of polysaccharide shell (by hydrogen bonding interactions) that

497

releases faster (through 10 h) at higher temperature produced by heating bath or AMF

498

hyperthermia.



499

The results of temperature- and pH-dependence drug delivery curves prove out the advantages of

500

exploiting the magnetic properties of the magnetite cores to increase the temperature by an

501

alternative MF. Thus, the effect of endogenous stimuli (pH) could be enhanced by the application

502

of magnetic hyperthermia to induce temperature increase on the tumor site.

503

504 505

Fig. 5 (I): Magnetically induced thermal response curve of MNP (Fe3O4), MNP@DAS@CSP, and

506

nanocarrier 2 under AMF (H=300 Oe, H×f = 9.56×109 Am-1s-1); (II) Release curves of DOX from

507

nanocarrier 2 under AMF compared to heating bath 37 and 42 C.

508 509

3.3. In vitro cytotoxicity studies

510

In biomedical applications, there are serious concerns on the safety and biocompatibility of

511

nanomaterials, considering the possibility of greater interactions between nanomaterials and

512

biological system. Therefore, the cytotoxicity of blank carrier MNP@DAS@CSP was investigated

513

against HEK-293 (normal cells) and HeLa cells by MTT assay, with and without application of

514

AMF (Fig. 6a). As shown in Fig. 6a, without application of AMF, the blank carrier

515

MNP@DAS@CSP showed negligible cytotoxicity (up to 85% vialibility) on both of HeLa and

516

normal cells even at high concentration of 400 µg mL-1. Cytotoxicity results of MNP@DAS@CSP 26 ACS Paragon Plus Environment

Page 26 of 45

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517

(400 µg mL-1) revealed that the hyperthermia generated by AMF heating decreases cell viability

518

of HeLa Cells to 65 ± 2% while normal cells do not affect non-reversible damage and retain their

519

viability up to 85%. This effect is less when using lower concentrations of MNP@DAS@CSP

520

because the AMF could not effectively raise the temperature of the sample to 42 C (Fig. 6a).

521

Cytotoxicity of nanocarrier 2 (MNP@DAS@DOX@CSP) was also evaluated (with and without

522

static magnetic field (MF) against HeLa Cells (Fig. S8). As shown in Fig. S8, the cell viability

523

obviously decreased as the DOX concentration increased from 0.01-2 μg/mL. However, the

524

cytotoxicity of nanocarrier 2 under static magnetic field (MF) condition (2 μgDOX/mL) has shown

525

only a slight improvement when it compared to the without MF condition because the static MF

526

only increases the absorption of magnetite nanoparticles into the cells and does not effect on the

527

amount of drug release.

528

To evaluate the rule of hyperthermia treatment, a new cytotoxicity assay was designed using higher

529

concentration range (1-200 μg/mL) of nanocarrier 2 (0.11-22.4 μgDOX/mL). Interestingly, the

530

results (Fig. 6b) showed that there is an important reduction of live cells as the alternating magnetic

531

field (AMF) is applied beside the nanocarrier 2 so that MNP@DAS@DOX@CSP with AMF

532

application produced about 20% higher cytotoxicity than that of without AMF (at 100 and 200

533

μg/mL of nanocarrier 2) which is comparable with cytotoxicity of free DOX at the same

534

concentration. It is obvious that AMF has no effect on the cytotoxicity of free DOX. The results

535

based on the study prove that nanocarrier 2 produced a cytotoxic response to HeLa cell lines largely

536

and DOX-free nanocarrier 2 can exhibit little cytotoxicity, an indication of their excellent

537

biocompatibility.

538



539

540 541

Fig. 6 Relative in vitro cell viability of MNP@DAS@CSP, Free DOX and nanocarrier 2 using

542

MTT assay. (a) HeLa or normal cells incubated with MNP@DAS@CSP at different concentration

543

for 24 h. (b) HeLa cells incubated with free DOX or nanocarrier 2 at different concentration for 16

544

h. To evaluate the effect of hyperthermia, further 2 h incubation was applied under AMF with

545

H=300 Oe and f = MHz.

546 547

3.4. In vitro hemocompatibility studies

548

It is necessary to evaluate the blood compatibility of a drug carrier because it will be finally

549

injected into blood vessels. In this work, hemolysis of red blood cells (RBCs) has been investigated

550

as a measure of the ability of a drug carrier to rupture lysosomes. The hemolytic activity of

551

nanocarrier 2 on hRBCs were evaluated under neutral (pH=7.4) condition using different

552

concentrations of nanocarrier 2 and DOX-free nanocarrier 2 as presented in Fig. 7a. Both of DOX-

553

free and DOX-loaded nanocarrier 2 showed no significant hemolytic effect over a broad

554

concentration range of 0-500 µgmL-1. Even at the high concentration of 500 µgmL-1, mean

555

hemolytic activity as low as 0.04-0.35% was detected for both which is in the acceptable range

556

when it compares to the literature reports (Table S4). Therefore, it can be concluded that 28 ACS Paragon Plus Environment

Page 28 of 45

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557

nanocarrier 2 has negligible hemolytic activity, which is a crucial feature for future in vivo

558

biomedical applications.

559

In vitro plasma coagulation induced by nanoparticles is mostly used as a model to estimate the in

560

vivo thrombogenicity. The blood coagulation processes are based on different pathways that are

561

determined frequently by measuring the partial thromboplastin time (PTT) and prothrombin time

562

(PT). Fig.8c-d show that almost all the PTT and PT values of both MNP@DAS@CSP and

563

MNP@DAS@DOX@CSP fell within the normal ranges, and two parameters did not show any

564

difference compared to their respective control and it appear to be independent of concentration in

565

a broad range of 7.25-200 μg/mL. In addition, there is no significant difference of PTT and PT

566

values between MNP@DAS@CSP and MNP@DAS@DOX@CSP at the same concentrations.

567

This suggests that both of MNP@DAS@CSP and MNP@DAS@DOX@CSP do not activate the

568

coagulation pathways.

569



Page 30 of 45

570 571

Fig.

7

The

in

vitro

hemocompatibility

studies

of

MNP@DAS@CSP

and

572

MNP@DAS@DOX@CSP (nanocarrier 2). (a) Hemolysis assay on red blood cells using water

573

and PBS as positive and negative controls. (b) Intuitive images of hemolysis assay. (c) PTT and

574

PT coagulation analyses, the time for the concentration of 0 µg·mL-1 refers to blood control.

575 576

3.5. Cellular uptake study

577

The intracellular DOX releases with and without AMF were also tracked by confocal laser

578

scanning microscopy (Fig. 8). We first examined the targeting ability for nanocarrier 2 against

579

HeLa cells at different incubation time 1, 3, and 6 h at [DOX] = 10 µgmL-1 without application of

580

AMF. As shown in Fig. 8a, the intracellular internalization of nanocarrier 2 is time dependent so

581

that significant red fluorescence intensity is appeared after 6 h incubation confirming that the

582

doxorubicin is effectively internalized into the cancer cell through the endocytosis process. The 30 ACS Paragon Plus Environment

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583

most striking observation concerns the nuclei DOX fluorescence in HeLa cells incubated with

584

nanocarrier 2 when they submitted to AMF for 1 h (Fig. 8b). The CLSM images (Fig. 8b) show

585

that the cellular uptake of nanocarrier 2 treated by AMF clearly increases rather than when they

586

do not being influenced by AMF (Fig. 8a) demonstrating the effect of AMF application to

587

accelerate intracellular DOX release markedly. All these results provided strong evidence that

588

nanocarrier 2 could use to be an effective selective targeted molecule.

589



590 591

Fig. 8 Fluorescence microscopic images of HeLa cells treated with MNP@DAS@DOX@CSP

592

(nanocarrier 2) where nucleus was stained by DAPI (blue). (a) [DOX] = 10 µg/mL without AMF.

593

(b) [DOX] = 10 µg/mL and each incubation time includes 1 h AMF condition.

594 32 ACS Paragon Plus Environment

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595

4. Conclusions

596

We have synthesized pH-responsive magnetite nanocarriers (MNP) based natural polysaccharides

597

for targeted anticancer DOX delivery. Loading of DOX onto the salep-modified MNP obtained a

598

high-loaded DOX-nanocarrier, subsequently covered by the second layer PEGylated chitosan.

599

After characterization of both of nanocarriers, they were evaluated under a series of in vitro studies.

600

Release studies showed that the double layered DOX-nanocarrier is more pH-sensitive and

601

exhibited excellent efficiency for DOX release. Apparently, the use of second layer PEGylated

602

chitosan prevents the nanocarrier to release DOX in neutral condition in one hand and postpones

603

the release of DOX in acidic condition in the result of its slow protonation and solvation led to the

604

higher control on the speed and amount of DOX release. Nanocarrier 2 showed a good magnetic

605

heating capacity and strong responsiveness to AMF so that the DOX release (42 C) into HeLa

606

cells were significantly improved. MTT cytotoxicity also showed high biocompatibility of DOX-

607

free nanocarrier and excellent killing effect of DOX-loaded nanocarrier on HeLa cells. MTT

608

cytotoxicity also showed the double-layered DOX-nanocarrier have comparable activity with free

609

DOX by AMF. Hemolysis and coagulation studies also revealed high blood compatibility of the

610

double layered DOX-nanocarrier. Investigation on cellular internalization revealed the high

611

targeting ability of nanocarrier 2 toward HeLa cells, especially under AMF. The most promising

612

aspect of this study is the strategy of using a second layer on nanocarrier to control drug release.

613

The use of natural polysaccharide also makes the nanocarrier biocompatible. Preparation method

614

is simple, and the shells are inexpensive polysaccharides. Finally, the possibility of using a

615

combination of cancer therapy such as hyperthermia along with the chemotherapy is a highlighted

616

aspect of the designed nanocarrier that could be expanded the biomedical applications.

617 618 33 ACS Paragon Plus Environment


619 620

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796

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with pH and light dual response for anticancer drug delivery. Biomacromolecules 2013, 14 (8),

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objects through RAFT Dispersion Polymerization and Influences of Morphology on Drug

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Li, Y.; Wan, J.; Zhang, Z.; Guo, J.; Wang, C., Targeted Soft Biodegradable

Chen, Y.; Su, M.; Li, Y.; Gao, J.; Zhang, C.; Cao, Z.; Zhou, J.; Liu, J.; Jiang, Z., Enzymatic

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Li, D.; Feng, X.; Chen, L.; Ding, J.; Chen, X., One-Step Synthesis of Targeted Acid-Labile

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