MCT mixed micelles as carriers of

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mPEG5k-b-PLGA2k/PCL3.4k/MCT mixed micelles as carriers of disulfiram for improving plasma stability and antitumor effect in vivo Linlin Miao, Jia Su, Xuezhi Zhuo, Lifeng Luo, Yihan Kong, Jingxin Gou, Tian Yin, Yu Zhang, Haibing He, and Xing Tang Mol. Pharmaceutics, Just Accepted Manuscript • DOI: 10.1021/acs.molpharmaceut.7b01094 • Publication Date (Web): 05 Mar 2018 Downloaded from http://pubs.acs.org on March 7, 2018

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Molecular Pharmaceutics

1

mPEG5k-b-PLGA2k/PCL3.4k/MCT mixed micelles as carriers

2

of disulfiram for improving plasma stability and antitumor

3

effect in vivo

4

5

Linlin Miaoa, Jia Sua, Xuezhi Zhuoa, Lifeng Luoa, Yihan Konga, Jingxin Goua, Tian

6

Yinb, Yu Zhanga, Haibing Hea, Xing Tanga,*

7

8

a

9

Road, Shenyang 110016, China

Department of pharmaceutics, School of Pharmacy, Shenyang Pharmaceutical University, NO. 103 Wenhua

10

b

11

110016, China

School of Functional Food and Wine, Shenyang Pharmaceutical University, NO. 103 Wenhua Road, Shenyang

12

13

*

14

E-mail address: [email protected]

15

Telephone number: 86-24-23986343

16

Fax: 86-24-23911736

Corresponding author: Xing Tang

1

ACS Paragon Plus Environment

Molecular Pharmaceutics 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

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Abstract:

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The clinical application of disulfiram (DSF) in cancer treatments is hindered by its

19

rapid

20

poly(ethyleneglycol)-b-poly(lactide-co-glycolide)/poly(ε-caprolactone)

21

(mPEG5k-b-PLGA2k/PCL3.4k) micelles were developed for encapsulation of DSF -by

22

using the emulsification-solvent diffusion method. Medium chain triglyceride (MCT)

23

was incorporated into the mixed polymeric micelles to improve drug loading by

24

reducing the core crystallinity. Differential scanning calorimetry (DSC) results

25

implied that DSF is likely present in an amorphous form within the micelles, and is

26

well dispersed. DSF is encapsulated within the core and the reservoir is stabilized by

27

the hydrophilic shell to prevent rapid diffusion of DSF from the core. The DSF mixed

28

micelles (DSF-MMs) showed good drug loading (5.90%) and a well-controlled

29

particle size (86.4±13.2 nm). The mixed micelles efficiently protected DSF from

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degradation in plasma, with 58% remaining after 48 h, while almost 90% of DSF was

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degraded after the same period for the DSF solution (DSF-sol) which was used as a

32

control. The pharmacokinetics study showed that the maximum plasma concentration

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and bioavailability of DSF were improved by using the DSF-MMs (2 and 2.5 times of

34

the DSF-sol). The TIRs (tumor inhibition rates) of 5-FU, DSF-sol and DSF-MMs

35

were 63.46%, 19.57% and 69.98% respectively, implying that DSF-MMs slowed the

36

growth of a H22 xenograft tumor model effectively.

degradation

in

the

blood

circulation.

In

37 2


this

study,

methoxy

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Keywords: disulfiram; mPEG5k-b-PLGA2k/PCL3.4k; mixed micelles; core crystallinity; plasma

40

stability

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1 Introduction

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Disulfiram (DSF) has been used to treat alcoholism since the 1940's.1 It was reported

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that DSF’s metabolite S-Methyl-N, N-diethylthiocarbamate (Me-DTC) sulfoxide is a

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potent, irreversible inhibitor of low Km mitochondrial aldehyde dehydrogenase

45

(ALDH).2 In recent years, DSF and its metabolites have been demonstrated to possess

46

strong antitumor activity towards various tumor types.3 The anticancer activity of

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DSF is copper (Cu)-dependent,4, 5 and the combination of DSF and Cu triggers the

48

generation of reactive oxygen species (ROS) and inhibits the activity of

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proteasome-NF-kB.6, 7 It has been demonstrated that DSF/Cu induces cell death via

50

the combination of two actions: (1) instant death caused by ROS produced by DSF/Cu

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and (2) delayed and stronger cytotoxicity by diethyldithiocarbamic acid (DDC)-Cu.8

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Furthermore, DSF can also increase the sensitivity of 5-FU and other anticancer drugs

53

towards

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diethyldithiocarbamate-copper complex (CuET),11 metabolite of DSF contributes to

55

the major anti-tumor effects. It could bind NPL4 and induce NPL4 aggregation,

56

consequently inhibiting p97 pathway. However, there has been little progress with

57

regards to in vivo antitumor studies of DSF, which is of particular importance for oral

58

administration. This is because DSF is quickly degraded in the gastrointestinal system,

tumor

cells.9,

10

Recently,

it

has

3


been

reported

that


Page 4 of 31

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with complete degradation occurring within 30 min in simulated gastrointestinal fluid

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(SGF).12 It has been reported that DSF is more stable at pH 6-7, compared to its

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stability in extremely acidic and alkaline conditions.13 The half-life of naked DSF in

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the bloodstream is only 4 min, which severely impedes its applications in the clinical

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field.14 The important metabolites of DSF are shown in Fig. 1. In the blood, DSF

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rapidly degrades to diethyldithiocarbamic acid (DDC), which is also unstable, and is

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further degraded to diethylamine and carbon disulphide. DDC is also a substrate of

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phase Ⅱ metabolism, which involves formation of diethyldithiomethylcarbamate

67

(Me-DDC).

68

diethylthiomethylcarbamate (Me-DTC), which is further oxidized to its corresponding

69

sulphoxide and sulphone metabolites.15 It is for this reason that there is such a

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discrepancy between the anticancer activity of DSF in vitro and in vivo. Hence, it is

71

imperative to develop an efficient drug delivery system to transport DSF to tumor

72

tissues and thereby increase its therapeutic efficacy.

Me-DDC

also

undergoes

oxidative

biotransformation

to

73

Amphiphilic block copolymers, which can self-assemble into structures with

74

hydrophobic cores and hydrophilic shells, have been extensively studied as potential

75

delivery vectors for poorly water-soluble drugs.16,

76

encapsulate the hydrophobic drugs in the core and thus increase their circulation time

77

in the body, facilitating favorable bioavailability and biodistribution. Generally,

78

physical encapsulation or chemical binding to the block copolymer can be used to

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encapsulate active substances into the polymeric micelles.18,

80

encapsulation process requires suitable functional groups between the drug and

17

Such polymeric micelles can

4


19

The conjugated

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polymer, so it is a more complicated strategy. Whereas，the physical encapsulation

82

method is more widely used, and it will be discussed here. However, it is not easy to

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achieve high drug loading and the micelle should be stable enough to deliver an

84

efficient drug dose to the desired site. The amount of drug loading and final stability

85

of polymeric micelles are mainly dependent on the solubility of the drug, type of

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polymers and the interaction between the drug and the hydrophobic core of the

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polymeric micelles.16 To further increase drug stability and drug loading efficiency of

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polymeric micelles, additional hydrophobic materials or physical blending of different

89

polymers can be utilized to alter the properties of the hydrophobic core.20,

21

90

Krishnadas

of

91

egg-phosphatidylcholine into a DSPE-PEG2K system, the mixed polymeric micelles

92

exhibited a greater solubilization of drug in comparison to the DSPE-PEG2K single

93

polymeric micelle.22 Jin X et al. concluded that juglone loaded Poloxamer

94

188/phospholipid mixed micelles exhibited a low toxicity and improved cellular

95

uptake.21,

96

encapsulate and deliver hydrophobic drugs to a specific site.

23

et

al.

reported

that

by

inserting

a

small

percentage

Therefore, mixed micelles offer a promising approach to effectively

97

Song et al. formulated a mixed system of mPEG-PLGA and PCL using the

98

nanoprecipitation method,24 which increased the encapsulation efficiency and the area

99

under the cure (AUC) of DSF 13.5 fold, while the DSF concentration of the

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nanoparticles decreased rapidly in phosphate buffered saline (PBS) containing 10%

101

fetal bovine serum (FBS) after 24 h. Duan et al. successfully entrapped DSF with a

102

drug loading content over 5% by utilizing a shell cross-linked micelle system.25 5



103

However, the materials used in this case were not degradable and the structure was

104

complex, and thus the system was difficult to apply in the clinic. Despite these

105

attempts, it still remains challenging to encapsulate DSF into amphiphilic

106

nanoparticles,26, 27 and these limitations compromise the clinical applications of DSF.

107

Medium chain triglyceride (MCT) is frequently used to prepare lipid-based

108

formulations, such as emulsions, nanocapsules, and lipid nanoparticles. It is

109

considered as a safe and biocompatible excipient, and shows good solubility for some

110

extremely lipophilic drugs. Gou et al. demonstrated the feasibility of improving drug

111

loading through physical blending of MCT into micelle cores.28 Biodegradable

112

polymers such as (lactic-co-glycolic acid) (PLGA) or polycaprolactone (PCL) have

113

been extensively used to prepare nanocarriers for drug entrapment.29 Based on this,

114

the aim of this work was to develop novel DSF-loaded mixed micelles (DSF-MMs) to

115

increase drug stability and bioavailability. The micelles were prepared with

116

PEG5k-b-PLGA2k/PCL3.4k and MCT by the emulsification-solvent diffusion method.

117

To evaluate the potential of the drug delivery system, the optimal formulation was

118

further characterized in terms of its in vitro physicochemical properties, including

119

morphology, particle size, zeta potential, drug-loading capacity, and drug release. In

120

vivo pharmacokinetics, an anti-tumor efficacy test and toxicity tests were also

121

undertaken.

122

2 materials and methods

123

2.1Material 6


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Methoxy poly(ethylene glycol)-b-poly(lactide-co-glycolide) (mPEG-PLGA; Mw:

125

mPEG 5000 Da, PLGA 2000 Da, LA/GA75/25) and poly(ε-caprolactone) (PCL, Mw:

126

3400 Da) were provided by Changchun Institute of Applied Chemistry (Changchun,

127

China). DSF was obtained from the Department of Medicinal Chemistry, Shenyang

128

Pharmaceutical University (Shenyang, China). Medium chain triglycerides (MCT)

129

were purchased from Tieling Pharmaceutical Co. Ltd (Tieling, China). F68 was

130

purchased from BASF AG (Ludwigshafen, Germany). Methanol and acetonitrile were

131

obtained from Tianjin Concord Technology Co. Ltd (Tianjin, China). All other

132

reagents were of analytical grade.

133

2.2 Preparation of DSF-MMs

134

DSF-loaded mixed micelles were prepared by the emulsification-solvent diffusion

135

method using PCL3.4k and mPEG5k-b-PLGA2k as carrier materials. The optimized

136

formulation and preparation process were as follows: PCL3.4k and mPEG5k-b-PLGA2k

137

(5:5, w/w), 15% (of the material) MCT and 10% DSF were dissolved in ethyl acetate:

138

benzyl alcohol (8:2) at 35 °C, and the obtained solution was used as the organic phase.

139

0.3% poloxamer 188, 2% benzyl alcohol and 5% ethyl acetate were dissolved in water

140

and heated to 35 °C. The organic phase was then slowly added to the water phase,

141

followed by high speed shear mixing to obtain the O/W coarse emulsion. The coarse

142

emulsion was subjected to high pressure homogenization at 600 bar for 4 cycles to

143

yield the final emulsion. The same volume of cold water was added into the final

144

emulsion with stirring under 10 °C, and the resulting colloid was purified and 7



145

concentrated with tangential flow filtration by membrane module. Finally, the

146

obtained DSF-MMs were passed through a 0.45 µm membrane syringe filter.

147

2.3 Characterization of DSF-MMs

148

The particle size and zeta potential of the obtained mixed micelles were determined

149

by dynamic light scattering (DLS) using a NicompTM 380 Particle Sizing system (Zeta

150

Potential/Particle Sizer NICOMPTM 380ZLS, Santa Barbara, California, USA).

151

Morphological examination of the micelles was performed using a JEM-2100

152

transmission electron microscope (JEOL, Japan) with an accelerating voltage of 200

153

kV. The drug loading and encapsulation efficiency of DSF in micelles was determined

154

by ultracentrifugation. The drug loading (DL%), drug encapsulation efficiencies

155

(EE%) and the recovery ratio (R%) were calculated by the following equations:

156

DL (%) =

DSF weight ×100% total MMs weight

157

EE (%) =

DSF content in micelles ×100% theoretical DSF content in micelles

158

R (%) =

resulted DSF - MMs weight ×100% total material fed

159

The solution was centrifuged at 3000 rpm for 15 min, and the filtrate was analyzed by

160

HPLC on a Dima C18 column (4.6×200 mm, 5 µm) at 25 °C, using methanol and

161

water at a ratio of 8:2 (v/v) as the mobile phase. The elution was monitored at 254 nm

162

at a ﬂow rate of 1.0 mL/min. The thermal behavior of pure DSF, PCL3.4k,

163

mPEG5k-b-PLGA2k, physical mixtures (PM) and DSF-MMs were evaluated by 8


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164

differential scanning calorimetry. Samples were accurately weighed (3-5 mg) and

165

heated in closed aluminum pans at a heating rate of 10 °C/min from 20 to 100 °C

166

under a nitrogen gas flow of 40 mL/min.

167

2.4 In vitro release and plasma stability

168

In vitro drug release was measured by a dialysis method. In brief, 1.0 mL DSF-MMs

169

or DSF-sol (20 mg DSF was dissolved in Cremophor EL/ethanol of 1mL, and the

170

obtained solution was diluted to 1 mg/mL with distilled water) was sealed in dialysis

171

membranes (MWCO: 12 kDa), and then dispersed in 10 mL phosphate buffered

172

solution (PBS) of pH 7.4 containing 2% w/v sodium dodecyl sulfate (SDS) with

173

continuous shaking at 100 rpm. SDS was used to increase the solubility of DSF in the

174

buffer solution. At pre-determined time intervals (0.5, 1, 2, 4, 6, 8, 12, 24, 48 and 72

175

h), all the release media was removed and replaced with an equal volume of fresh

176

medium to continue the release process. The amount of released DSF was measured

177

using the HPLC method after filtering through a 0.22 µm membrane filter.

178

To determine the plasma stability of encapsulated DSF-MMs and DSF-sol, 0.1

179

mL of each formulation was added into an eppendorf tube containing 0.9 mL blank

180

plasma, with shaking at 37 °C. At set time intervals (0, 0.167, 0.5, 1, 2, 4, 6, 8, 12, 24

181

and 48 h) 50 µL was removed and added to activated cartridges (with 1 mL methanol,

182

1 mL ethylenediaminetetraacetic acid (EDTA, 0.05M). DSF was then eluted from the

183

cartridges using 1 mL water (5% acetonitrile) and 1 mL acetonitrile and measured by

184

HPLC. 9



185

2.5 In vivo pharmacokinetic study and detection of DSF

186

Rats were randomly divided into two groups (n = 3): (1) DSF-sol group and (2)

187

DSF-MMs group. The formulations were injected at a single dose of 20 mg/kg into

188

the tail vein. Blood samples (0.5 mL) were removed from the orbital venous plexus

189

and collected into 1.5 mL EP tubes containing the same volume of stabilizer (0.9%

190

sodium chloride, 0.64% sodium acetate and 0.8% Diethylenetriaminepentaacetic acid,

191

pH 4.5) at 0, 15, 30, 45, 60, 120, 240 and 360 min after dosing. The samples were

192

immediately centrifuged at 6,000 rpm for 5 min and the supernates were collected.

193

DSF was extracted by the solid phase extraction (200 µL sample with 20 µL internal

194

standard diphenhydromine added) and the concentration was determined by the

195

validated UPLC–MS/MS method.17

196

Analytic separation was achieved using a Syncronis C18 column (50×2.1 mm,

197

1.7 mm particle size, Thermo). The mobile phases A and B were acetonitrile and

198

H2O/formic acid 99.9:0.1 (V/V) +1 mM Ammonium acetate. The gradient (0.2

199

mL/min) was as follows: initially at 20% A and reaching 80% A linearly in 0.8 min,

200

followed by 1.5 min at 80% A, then reaching 20% A linearly in 0.2 min. Finally,

201

equilibrating at 20% A for 0.5 min. The column temperature was held at 40 °C.

202

Quantitative determination of DSF was performed using a Waters ACQUITYTM

203

XEVOTQ mass spectrometer. Data was acquired in the electrospray ionization (ESI)

204

mode with positive ion detection and multiple reaction monitoring (MRM). A cone

205

voltage of 14 V and capillary voltage of 3.50 kV were used. The desolvation 10


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206

temperature was maintained at 400 °C and nitrogen was used as the desolvation gas

207

with a flow rate of 600 L/hr. The MS/MS transition of DSF was 296.98 →115.93 and

208

that of diphenhydromine was 256.14 → 167.02.

209

2.6 In vivo anti-tumor efficacy test and toxicity tests

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Kunming mice (18-22 g) were acquired from the animal center, Shenyang

211

Pharmaceutical University. All animal experiments were conducted in accordance

212

with the guidelines of the Laboratory Protocol of Animal Care and Use Committee,

213

Shenyang Pharmaceutical University. H22 tumor cells were intraperitoneally injected

214

into healthy Kunming mice (0.2 mL/mouse), and inoculation of ascites in the mice

215

was observed. After the abdominal cavity could be seen to be filled with ascites, mice

216

were sacrificed and the ascites were diluted 50 times with saline. The H22 xenograft

217

tumor model was prepared by inoculating the right armpit with 0.2 mL diluted ascites

218

after iodophor disinfection and 75% alcohol disinfection. After tumor reached to

219

100-200 mm3, the mice were randomly divided into four groups according to the size

220

of tumor volume and the principle of group consistency. Grouping and dosing

221

regimens are shown in Table. 1. The oral dosing of copper gluconate was at 19.2

222

mg/kg 5-6 hours before tail vein injection of the DSF preparation. The samples were

223

injected into the tail vein of each mouse for 3 weeks (twice a week). Tumor volume

224

and body weight were used to evaluate the treatment efficacy and safety, respectively.

225

At various time intervals, the tumor volume was measured by a vernier caliper and

226

calculated using the following equation: V=1/2ab2, where a and b are the major and 11



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227

minor axes of the tumors. During the course of the experiment, the mice weights were

228

weighed daily. Animals were sacrificed 2 weeks after drug administration using

229

diethyl ether anesthesia, and the final size and weight of tumors were analyzed.

230

2.7 Statistical analysis

231

SPSS17.0 statistical software was used for the Student's t test. Results are given as

232

mean values ± S.D. P values of < 0.05 were considered significant.

233

3 Results and discussion

234

3.1 Preparation of DSF-MMs

235

It has been reported that different structures of PEG-PLGA copolymers with different

236

properties can be used for different drug delivery systems, such as nanoparticles,

237

micelles and hydrogels. When the PEG (fEO) fraction is above 45%, micelles were

238

found

239

polyethyleneoxide-polybutadiene

240

polyethyleneoxide–polyethylethylene

241

mPEG5k-b-PLGA2k was used, and the emulsification-solvent diffusion method shown

242

in Fig. 2 was used to prepare the DSF-MMs. The organic solvents were removed

243

using tangential flow filtration, simultaneously concentrating the formulation.

to

more

easily

form,

shown

(PEOm-PBDn)

in

a

and

(PEOm-PEEn).30

study

of

polymer

hydrogenated

homolog

In

this

study,

244

As summarized in Table 2, all the micelles were in the nanoscale size range, and

245

their size varied from 60 to 200 nm depending on the formulation and preparation

246

process. The ratio of organic phase to water phase had no significant effect on the size 12


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247

and drug loading of the nanoparticles, while increasing the PCL loading significantly

248

increased the size of the nanoparticles and the drug loading. It has previously been

249

reported that the use of PCL is necessary for DSF encapsulation.31

250

It has also been reported that the addition of MCT to the nanoparticle system

251

could improve the drug loading and stability without affecting the drug

252

pharmacokinetics in vivo.28 As MCT can penetrate into the polymer core and disrupt

253

the arrangement of the inner core, the amorphous area is expanded, which increases

254

the drug loading of the nanoparticles. According to our previous work, the maximum

255

drug loading of DSF could be increased from 2.61±0.10% to 8.34±0.20% by

256

incorporating MCT into mPEG-PCL polymeric micelles.28 Furthermore, it is known

257

that the degradation of DSF is less than 1 % in MCT at 60 °C for 10 days,26 and so we

258

prepared micelles with different MCT contents to investigate the effect of MCT on

259

drug loading. There is a gradual increase in drug loading upon MCT incorporation.

260

However, no obvious increase occurred when the MCT amount was over 30%, and

261

the particle size increased sharply. Interestingly, the particle size first decreased as the

262

amount of MCT incorporated was increased, and then increased as addition continued.

263

Glavas et al. reported that the reduction could be caused by the decrease in core

264

crystallinity.32 Polycaprolactone (PCL) is semicrystalline, and its crystallinity may be

265

altered by MCT, after which the core can pack tightly in the non-crystalline regions,

266

which further influences the core size and the nanoparticle size. The increased particle

267

size is thus possibly due to MCT expanding the core volume. Based on the above

13



Page 14 of 31

268

results, 15% MCT of the polymer weight was added to the formulation to acquire the

269

optimum drug loading of 5.90% and the recovery ratio was 82.0%.

270

Based on this, in subsequent experiments the feed ratio of DSF, PCL3.4k,

271

mPEG5k-b-PLGA2k and MCT was set as 2:10:10:3 (w / w) to prepare the DSF-MMs.

272

3.2 Characterization of DSF-MMs

273

The schematic illustration of the DSF-loaded mPEG5k-b-PLGA2k/PCL3.4k/MCT mixed

274

micelles is shown in Fig. 2. The surface morphology of the DSF-MMs was observed

275

and a representative TEM image is presented in Fig. 3. The TEM image confirmed

276

that the micelle formulations were spherical in shape. The particle diameter of the

277

micelles was measured to be 86.4±13.2 nm using DLS analysis, which was consistent

278

with the TEM image. Particles showed good uniformity, and no oil droplets could be

279

seen embedded in the micelle cores when observed in the TEM images, which is

280

expected as the content of MCT is only 15% of the used materials, and is thus not

281

high enough to form an oil core. Zeta potential is correlated to the charge on the

282

surface of a nanoparticle and can reflect particle’s stability to some extent. When the

283

surface charge is high, there is a strong electrostatic repulsion between the

284

nanoparticles, which results in reduced aggregation.33,

285

negatively charged with a zeta potential range of -15 mV to -30 mV. Furthermore,

286

zeta potential is an important feature in relation to the interaction of nanoparticles

287

with plasma proteins, which is known to be influenced by the surface charge of the

288

nanoparticles. The DSC thermograms of DSF powder, PM, PCL3.4k, DSF-MMs and 14


34

The nanoparticles were

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289

mPEG5k-b-PLGA2k are shown in Fig. 4. Pure DSF, PCL3.4k, and mPEG5k-b-PLGA2k

290

have distinct and sharp endothermic peaks at 70.55 °C, 59.46 °C, and 56.45 °C,

291

respectively. The physical mixture showed the 59.45 °C and 70.83 °C peaks. Even

292

though the peaks of PCL3.4k and mPEG5k-b-PLGA2k overlapped, this has no effect on

293

the DSF peak. The disappearance of the DSF peak of the DSF-MMs suggests that

294

DSF may be present in an amorphous or molecular form. As well, there is only a

295

small peak at 52 °C and the peak of PCL3.4k disappears, which may reflect that the

296

state of PCL3.4k is amorphous in the formulation. As the free drug has been removed

297

during the preparation process, the encapsulation efficiency is high, and is thus no

298

longer considered as an assessment indicator during the prescription and process

299

screening.

300

3.3 In vitro release and plasma stability

301

In vitro release and plasma stability tests were carried out to predict the stability of the

302

micelles in vitro. SDS was added to the release medium to increase the solubility of

303

DSF and protect DSF from degradation. The release profile of DSF from DSF-MMs

304

and DSF-sol is presented in Fig. 5, where it can be seen that there was no obvious

305

sustained-release behavior for the DSF-MMs compared to the DSF-sol. As well,

306

approximately 80% of DSF was released within 24 h from DSF-sol. Zhang et al.

307

reported that DSF could be released to 80% at 12 h.35 The difference may be caused

308

by the addition of polyoxyethylene castor oil, which could form micelles,

309

encapsulating DSF and slowing the release. The release curve of DSF-MMs was 15



310

modeled and its release mechanism was subsequently analyzed. It can be seen from

311

Table. 3 that it conforms to the Ritger-Peppas model,36-38 indicating that the release

312

mechanism of DSF-MMs is diffusion and erosion. Diffusion and core degradation are

313

the main release mechanisms of micelles, and physically entrapped drugs are typically

314

released mainly through the diffusion effect, and the characteristics of drug release are

315

also relevant to the degradation of the copolymers. The hydrophobic core was

316

degraded through hydrolysis of the ester linkage of the polymer, leading to sustained

317

release of DSF.

318

DSF is very unstable in the blood, forming DDC after rapid degradation, which

319

soon breaks down into diethylamine and carbon disulfide. Thus, ensuring the stability

320

of DSF in the blood is an important prerequisite for its anti-tumor effect. The stability

321

of DSF-MMs and DSF-sol in the blood was evaluated by monitoring the

322

concentration of DSF in plasma. As shown in Fig. 5b, the DSF-sol group degraded by

323

16.03% after 1 h, while no degradation was seen for DSF-MMs. DSF-MMs and

324

DSF-sol degraded to about 42.57% and 85.57% respectively after 48 h, and the DSF

325

concentration remained above 60% of the total concentration for more than 50 h,

326

demonstrating that DSF-MMs are more stable in plasma. This is in part attributed to

327

the core-shell structure of the polymeric micelles. The outer hydrophilic PEG brush

328

can introduce a steric effect and avoid interaction of the polymeric micelles with

329

biological components, such as proteins and cells, which may lead to early micellar

330

elimination by the RES. Micelles disassembling into single chains under the diluting

331

conditions of blood circulation is another obstacle against plasma stability and 16


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332

development of effective micellar drug carriers. It is possible to prevent dissociation

333

of the micelles by increasing the hydrophobicity of the block copolymer.

334

Song et al.

24

evaluated the stability of mPEG-PLGA/PCL Mixed Nanoparticles

335

in PBS solution containing 10% FBS, and the results showed that the size of the

336

nanoparticles remained mostly constant, with only a slightly decreased intensity after

337

48 h. However, obvious DSF degradation occurred after 24 h, so it could be

338

concluded that the initial release of DSF was mainly due to diffusion, which had no

339

effect on the size of the particles. These results appear to be consistent with the results

340

shown here. The schematic diagram of the dissociation of the micelles and the release

341

of DSF is shown in Fig. 6. According to the release profile, there are two distinct

342

release stages, characterized by the change in the slope of the curves. Burst release

343

may be attributed to the diffusion of DSF absorbed to the hydrophobic core. After this,

344

the system showed a transition to a more gradual release trend, which is related to the

345

diffusion of the drug from the inner side and the dissociation of the polymeric

346

micelles. Regarding the plasma stability, interactions between the drug and protein

347

molecules in solution may be responsible for the rapid release of the drug from the

348

micelles observed. After this, the slow release is mainly controlled by diffusion, with

349

little influence of decomposition.

350

3.4 In vivo pharmacokinetic study

351

A study was undertaken to compare the pharmacokinetics of the DSF-MMs to the

352

DSF-sol. The plasma concentration-time profiles of DSF after injection of DSF-MMs 17



353

or DSF-sol at a dose of 20 mg/kg in rats are shown in Fig. 7, and relevant

354

pharmacokinetic parameters are listed in Table. 4. The plasma concentration of DSF

355

in the mixed micelles group was much higher than that of the control group. The AUC

356

of DSF in the mixed micelles group was nearly 2.5 times of that in the solution group

357

(12084.0 versus 4706.4 µg/L*h). In comparison with the solution group, the

358

DSF-MMs could prolong the half-life of the drug in the plasma (t1/2) from 1.07 h to

359

2.50 h. These results implied that the clearance of DSF in plasma could be reduced by

360

the polymeric micelles, owing to the long-circulating effect of PEG, and the

361

DSF-MMs were able to enhance the stability and availability of DSF in vivo. These

362

results agree with the plasma stability profiles, and address a key barrier against the

363

successful in vivo delivery of DSF. The DSF plasma concentration of DSF-MMs is

364

almost twice that of DSF-sol at all the time points, which could be attributed to the

365

protection of the core-shell structure of the micelles. However, it could be seen to

366

drop rapidly 15 min post-injection, either owing to fast distribution, rapid elimination

367

or micelle dissociation. The outcome is partly consistent with the plasma stability

368

experiment, which showed a sharp decline of DSF concentration after 30 min. As well,

369

the decrease trends of the two curves are nearly the same, possibly because Cremphor

370

EL can affect the disposition and pharmacokinetics of various drugs by changing the

371

unbound drug concentration through micellar encapsulation.8 There are few reports on

372

the in vivo mechanisms of block copolymer metabolism in the literature, however it

373

can be speculated that PEG-PLGA has the same protection effect on DSF. The

374

difference between DSF-MMs and DSF-sol was in the stability and the protection of 18


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375

micelles.

376

3.5 In vivo anti-tumor efficacy test and toxicity tests

377

A tumor growth inhibition study in Kunming mice (using H22 hepatocarcinoma cells)

378

was performed, using a 40 mg/kg DSF equivalent injection of DSF-MMs or DSF-sol

379

into the tail vein of rats. 20 mg/kg of 5-FU or saline were chosen as the positive or

380

negative controls. The results suggested that the TIRs of 5-FU, DSF-sol and

381

DSF-MMs were 63.46%, 19.57% and 69.98% respectively, confirming that

382

DSF-MMs could produce a more pronounced tumor suppressing effect than DSF-sol,

383

with an effect comparable to that of 5-FU. After 17 days of injections, the tumors of

384

each group were removed and photographed, and the tumor growth inhibition rate

385

was calculated, and these results are shown in Fig. 8A-C. The final tumor volume of

386

the DSF-MMs was remarkably lower (P

MCT mixed micelles as carriers of

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