- Email: [email protected]
Ultrasound triggered phase-change nanodroplets for doxorubicin prodrug delivery and ultrasound diagnosis: An in vitro study
Accepted Manuscript Title: Ultrasound triggered phase-change nanodroplets for doxorubicin prodrug delivery and ultrasound diagnosis: an in vitro study Authors: Jinbiao Gao, Baiqing Yu, Chao Li, Ming Xu, Zhong Cao, Xiaoyan Xie, Wei Wang, Jie Liu PII: DOI: Reference:
S0927-7765(18)30829-4 https://doi.org/10.1016/j.colsurfb.2018.11.046 COLSUB 9811
To appear in:
Colloids and Surfaces B: Biointerfaces
Received date: Revised date: Accepted date:
10 June 2018 24 October 2018 19 November 2018
Please cite this article as: Gao J, Yu B, Li C, Xu M, Cao Z, Xie X, Wang W, Liu J, Ultrasound triggered phase-change nanodroplets for doxorubicin prodrug delivery and ultrasound diagnosis: an in vitro study, Colloids and Surfaces B: Biointerfaces (2018), https://doi.org/10.1016/j.colsurfb.2018.11.046 This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting proof before it is published in its final form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.
Ultrasound triggered phase-change nanodroplets for doxorubicin prodrug delivery and ultrasound diagnosis: an in
SC RI PT
vitro study
Jinbiao Gao a,†, Baiqing Yu a,†, Chao Li a,†, Ming Xub, Zhong Caoa, Xiaoyan Xieb, Wei Wangb,*, Jie Liua,* a
School of Biomedical Engineering, Sun Yat-sen University, Guangzhou, Guangdong 510006,
China
Department of Medical Ultrasonics, Institute of Diagnostic and Interventional Ultrasound,
U
b
A
N
The First Affiliated Hospital of Sun Yat-sen University, Guangzhou, Guangdong, 510080, China
These authors contribute equally to this work.
*
E-mail: [email protected]
M
†
A
CC
EP
TE
D
[email protected]
Graphical Abstract
Highlights Phase-change contrast agent and acid-cleavable prodrug were co-loaded in a single nanodroplet system
Ultrasound triggered phase-change nanodroplets for enhanced ultrasound imaging
Enhanced uptake of drug and therapeutic effect were achieved with ultrasound irradiation
SC RI PT
ABSTRACT
U
Ultrasound-triggered delivery system is among the various multifunctional and stimuli-
N
responsive strategies that hold great potential as a robust solution to the challenges of
A
localized drug delivery and gene therapy. In this work, we developed an ultrasoundtriggered delivery system PFP/C9F17-PAsp(DET)/CAD/PGA-g-mPEG nanodroplet, which
M
combined ultrasound responsive phase-change contrast agent, acid-cleavable doxorubicin prodrug and cationic amphiphilic fluorinated polymer carrier, aiming to achieve both high
D
imaging contrast and preferable ultrasound-triggered anti-cancer therapeutic effect. The
TE
optimized nanodroplets were characterized as monodispersed particles with a diameter of about 400 nm, slightly positive surface charge and high drug-loading efficiency. The functional augmenter PGA-g-mPEG provided the nanodroplets good sustainability, low
EP
cytotoxicity and good serum compatibility, as confirmed by stability and biocompatibility tests. In ultrasound imaging study, the nanodroplets produced significant contrast with
CC
ultrasound irradiation (3.5 MHz, MI=0.08) at 37 ℃. Cell uptake and cytotoxicity studies in HepG2 and CT-26 cells showed the enhanced drug uptake and therapeutic effect with the
A
combination of nanodroplets and ultrasound irradiation. These results suggest that the PFP/CAD-loaded phase change nano-emulsion can be utilized as an efficient theranostic agent for both ultrasound contrast imaging and drug delivery.
KEYWORDS: ultrasound; perfluorocarbon; phase-change contrast agent; nanodroplets; doxorubicin;
1. Introduction
SC RI PT
Ultrasound-triggered delivery system is among the various multifunctional and stimuliresponsive strategies that have been extensively investigated and employed for localized
drug delivery and gene therapy[1]. As an external stimulus, ultrasound outperforms many other physical and biochemical stimuli with its low cast, noninvasiveness, temporal and
spatial control, depth of stimulation and good patient acceptance[2]. The robust properties of an ideal ultrasound-triggered drug delivery system include in vivo stability of drug
U
carrier prior to ultrasound irradiation, enhanced localized therapeutic effect under the
N
acoustic field, and contrast effect for imaging[3]. These advantages of the ultrasound-
A
triggered drug delivery vehicles are attributed to the unique mechanics of ultrasound irradiation that mainly include 1) cavitation caused by ultrasound shockwave that disrupts
M
the drug carriers and nearby tissue membrane to promote release and extravasation, 2) acoustic streaming that drives particles into the target site and 3) hyperthermia that can act
D
as an additional stimulus[4, 5].
TE
The most explored ultrasound-trigger drug delivery vehicles include ultrasoundresponsive polymer or lipid nanocarriers and gas-filled microbubbles, which actively
EP
respond to cavitation and enhance signal echoes for imaging[5]. Early studies exploited the ultrasound responsiveness of microbubbles by co-administrating microbubbles with
CC
therapeutic agents or nanocarriers to facilitate the delivery and uptake[6, 7]. The application of microbubbles was taken one step forward with the development of functionalized shelled microbubbles, including lipid-microbubbles and polymer-
A
microbubbles that carry drug or gene payload[8-10]. However, the large-size microbubbles have limited permeability, short circulation time and relatively low payload[11-13]. Nanobubbles emerged recently with refined fabrication strategies, aiming to strengthen passive targeting and increase the half-life of the system, and have achieved positive results[14, 15].
As an alternative to gas bubbles, phase-change agent has gained inflated attention with its combined strength: the high tissue permeability due to the small size and high acoustic responsiveness due to the phase-changing attribute[16, 17]. Phase-change agents can be encapsulated and stabilized by polymer micelles in the form of nanodroplets in normal physiological environment. When exposed to ultrasound, acoustic droplet vaporization
SC RI PT
(ADV) of the phase-change agent is triggered, producing contrast effect and promoting
nanodroplet extravasation[18]. Perfluoropentane (PFP) is one of the pioneering phase-
change agents applied in ultrasound-triggered drug delivery vehicles. In depth reports
indicate that PFP encapsulated in drug-loaded nanodroplets has a phase-changing temperature as high as 37 ℃, providing contrast effect, enhancing the uptake of therapeutic
agents in vivo and thus showing prospective value in drug delivery and gene therapy[19,
U
20]. In our previous works, we developed an amphiphilic cationic polymer C 11F17-
N
PAsp(DET), which has a fluorinated end group affinitive to PFP and cationic side chains for gene loading. The PFP-containing nanocarrier showed encouraging enhancement of
A
gene transfection efficacy with insonation and promising potential as an efficient drug
M
delivery vehicle[21].
Doxorubicin(DOX), an anthracycline topoisomerase II inhibitor, has been long applied
D
as an effective chemotherapy agent in cancer treatments. Free DOX is not tumor-specific,
TE
and causes damage to healthy tissues[22], requiring the assistance of efficient drug delivery systems. The engineering of DOX delivery systems is inspired by one of the most explored
EP
site-specific stimulus: the acidic environment of tumor tissue (pH≈6.5) and intracellular endosomes (pH=5.0-6.0). A promising approach in the design of pH-responsive systems
CC
involves conjugating the drug molecule with acid-sensitive linkages like hydrazone, carbamate or cis-aconityl bond. Drug conjugated remains stable at neutral pH during transportation and is triggered to release after arriving at tumor site or intracell, thus
A
promoting target-specificity, cell uptake and reducing toxicity in irrelevant sites[1, 23]. Prodrugs with conjugated DOX are reported to have high loading efficiency, longer circulation half-life and decent pH-responsiveness[24-26]. It is worth noting that besides acid-sensibility, cis-aconityl doxorubicin (CAD) is highly soluble in water and negatively charged, and is thus a superior choice of drug payload for the cationic amphiphilic carrier system.
In this study, we developed a PFP-containing and CAD-loaded ultrasound-triggered drug delivery system to achieve optimized imaging, uptake and therapeutic effect. As shown schematically in Scheme 1, The phase-change agent PFP is encapsulated in the amphiphilic cationic polymer C9F17-PAsp(DET) in the form of nanodroplets. The pretreated DOX, cis-aconityl doxorubicin (CAD) is absorbed onto the nanodroplet
SC RI PT
electrostatically. The nanodroplet is further coated with PGA-g-mPEG, an augmenting coating used in our previous study[27]. This coating material consists of polyethene glycol
(PEG) chains to stabilize the nanodroplet and gamma-polyglutamic acid (γ-PGA), an anionic polymer known to specifically interact with gamma-glutamyl transpeptidase
(GGT) in mammalian cell lines to enhance cell uptake via a GGT-mediated pathway[28,
29]. The obtained ternary nanodroplets, PFP/C9F17-PAsp(DET)/CAD/PGA-g-mPEG
U
(PFP-TNDs), was measured with dynamic light scattering (DLS) and observed with
N
transmission electron microscopy (TEM) to determine the physical characteristics. Biocompatibility tests and competitive inhibition experiments were conducted to
A
investigate the stability of the TNDs and the effect of γ-PGA, respectively. Furthermore,
M
the ultrasound responsiveness and therapeutic efficacy of the PFP-containing, CAD-loaded
2. Methods
EP
2.1. Materials
TE
cytotoxicity assays.
D
TNDs against cancer cells were evaluated by in vitro ultrasound imaging, cell uptake and
β-benzyl-L-aspartate N-carboxyanhydride (BLA-NCA) was purchased from Beijing
CC
HWRK Chem (Beijing, China) and used as received. 2,2,3,3,4,4,5,5,6,6,7,7,8,8,9,9,9heptadecafluorononylamine (C9F17-NH2) and fluorescence isothiocyanate (FITC) were purchased from Sigma-Aldrich. N-methyl-2-pyrrolidone (NMP) and diethylenetriamine
A
(DET) were purchased from the Aladdin Industrial (Shanghai, China) and used after distillation. Perfluoro-n-pentane (PFP) was purchased from Strem Chemicals. Doxorubicin hydrochloride (DOX·HCl) was purchased from Beijing HVSF United Chemical Materials (Beijing, China). Cis-aconitic anhydride (CA) was purchased from Alfa Aesar China (Tianjin, China). Dimethyl sulfoxide-d6 (DMSO-d6) and deuterium oxide (D2O) were purchased from Adamas Reagent. Poly(γ-glutamic acid) (γ-PGA; 5000 Da) was purchased
from Nanjing Sai Taisi Biotechnology (Nanjing, China). HepG2 cells, CT-26 cells and HUVEC cells were purchased from the Shanghai Cell Bank of the Chinese Academy of Sciences (Shanghai, China). HepG2 and HUVEC cells were cultured in Dulbecco’s modified Eagle’s medium (DMEM, Hyclone) and CT-26 cells were cultured in RPMI Medium 1640 (Hyclone), supplemented with 1% (w/v) penicillin-streptomycin and 10%
SC RI PT
(v/v) fetal bovine serum (FBS) and in a 5% CO2 environment at 37 ℃.
2.2. Synthesis of C9F17-poly{N-[N’-(2-aminoethyl)]-aspartamide} [C9F17-PAsp(DET)]
C9F17-PAsp(DET) was prepared stepwise according to Scheme S1. C9F17-PBLA was
synthesized with the BLA-NCA ring opening polymerization initiated by C9F17- NH2. BLA-NCA (0.5 g) and C9F17- NH2 (37.5 mg) was dissolved in 5 mL of anhydrous dimethyl
U
formamide. Dichloromethane (5 mL) was added dropwise into the above mixture solution.
N
The polymerization was kept in a dry N2 atmosphere for 3 days at 25 ℃. Then, the solution was dialyzed against dimethyl sulfoxide for 1 day and against deionized water for 2 days.
A
C9F17-PBLA (white powder) was obtained by lyophilization. C9F17-PAsp(DET) was
M
synthesized with aminolysis of the side chains. Briefly, C9F17-PBLA (150 mg) and was dissolved in NMP (5 mL) and DET (1 mL, mixed with 4 mL of NMP) was added dropwise
D
into the solution. 1 M HCl (9.7 mL) was injected to the mixture in an ice bath after 6 h of
TE
stirring, followed by dialysis against 0.01 M HCl for 1 day and against water for 2 days. C9F17-PAsp(DET) was obtained as a yellowish solid after lyophilization. C9F17PAsp(DET)-FITC was synthesized by adding FITC (12.3 mg dispersed in 4 mL deionized
EP
water) into C9F17-PAsp(DET) solution (50 mg in 6 mL deionized water) and stirring for 12 h, followed by dialysis against water and lyophilization. All products in the workflow were
CC
stored in -20 ℃.
A
2.3. Synthesis of cis-aconityl-doxorubicin (CAD) The CAD was synthesized according to the previous study [30]. Briefly, Doxorubicin
hydrochloride (DOX·HCl, 50 mg) was dissolved in 6 mL of 0.1M carbonate buffer (pH=9.4) and cooled in an ice bath. CA (70 mg) in 0.6 mL absolute 1,4-dioxane solution was added dropwise with stirring, and the pH was kept at 9 by careful addition of cold 0.5 M NaOH and the following 15 min of stirring. Insoluble contents were removed by
centrifugation of the mixture (13000 rpm, 5 min, 4 ℃). CAD was precipitated by addition of cooled 1 M HCl to the mixture (the pH was reduced to 1) and was collected with centrifugation. The collected CAD was re-dissolved in deionized water with pH of the solution adjusted to neutral by the addition of 0.2 M NaOH. To further purify the product, the solution was washed with 4 mL chloroform/methanol (95:5, v/v) for five times. The
SC RI PT
organic solvent was removed on the rotary evaporator and the solution was lyophilized. CAD was obtained as a red powder and was stored in dark environment at -20 ℃. 2.4. Characterization of the synthetic products
C9F17-PBLA (in DMSO-d6) and C9F17-PAsp(DET) (in D2O) were characterized by proton nuclear magnetic resonance spectroscopy (1H-NMR) using a Bruker AvanceIII
U
spectrometer (400 MHz). Furthermore, the molecular weight and of the polymers was
N
calculated from the spectra.
A
The DOX and CAD were characterized by 1H-NMR spectroscopy (both in DMSO-d6), Fourier transform infrared spectroscopy (FTIR) and electrospray ionization mass
M
spectrometry (ESI-MS). The FTIR spectra were obtained with the potassium bromide method. The ESI-MS was detected with a Thermo Finnigan TSQ Quantum Ultra AM EMR
D
Mass Spectrometer. The mass value of CAD was determined by negative ion mode (ESI-)
was 0.2 mL/min.
TE
and DOX was determined by positive ion mode (ESI+). The flow rate of the mobile phase
EP
2.5. Preparation of nanodroplets The PFP/C9F17-PAsp(DET) nanodroplets (PFP-NDs) were fabricated with oil/water
CC
emulsification. Briefly, 5 mg of C9F17-PAsp(DET) was dissolved in 1 mL of deionized water and cooled at 4 ℃ for 5 min. PFP (60 μL) was slowly added to the solution, followed
A
by sonication with a Misonix Sonicator (S-4000, with probe) for 90 s in a cycle of 1 s sonication at 22% amplitude and 1 s standby in an ice bath. The resulting PFP-NDs were kept at 4 ℃ and used in no more than 1 h. The PFP/C9F17-PAsp(DET)/CAD binary nanodroplets (PFP-BNDs) and modified PFP/C9F17-PAsp(DET)/CAD/PGA-g-mPEG ternary nanodroplets (PFP-TNDs) were prepared stepwise as shown in Scheme 1. The PFP-BNDs were prepared in different C9F17-
PAsp(DET)/CAD mass ratios by gently mixing the PFP-NDs aqueous suspension with the CAD aqueous solution. After incubation at 4 ℃ for 30 min, a certain amount of PGA-gmPEG (prepared according to the method previously reported by Yang et al) aqueous solution was added to obtain the ternary nanodroplets PFP-TNDs in various C/N ratios (the molar ratio of the dissociative carboxyl in PGA-g-mPEG to the primary amine groups in
SC RI PT
the N-substituted poly-aspartamides), followed by incubation at 4 ℃ for 30 min. Finally,
the mixture was dialyzed against 5% glucose solution for 12 h at 4 ℃. For biocompatibility test and uptake experiments, we fabricated PFP/C9F17-PAsp(DET)/PGA-g-mPEG and PFP
C9F17-PAsp(DET)-FITC/PGA-g-mPEG nanodroplets, respectively, in which CAD solution was replaced by the same amount of deionized water. Nanodroplets were used
U
upon dialysis.
N
2.6. Characterization of nanodroplets
The particle size of PFP-NDs, PFP-BNDs, and PFP-TNDs were determined with
A
dynamic light scattering (DLS, Malvern Zetasizer). The zeta potential of the nanodroplets
M
were determined by phase analysis light scattering (Malvern Zetasizer). The morphology of the nanodroplets (stained with phosphotungstic acid) was observed by a transmission
D
electron microscope (TEM, JEOLJEM 1400).
TE
For quantification of CAD in the following tests, the UV-vis absorbance spectra of PFPTNDs, PFP-NDs/PGA-g-mPEG and free CAD was measured by the UV-vis
EP
spectrophometer (Beckman DU730) and the fluorescence spectra with the excitation wavelength of 480 nm was measured by the spectrofluorometer (Horiba FluoroMax 4). For calculation of drug loading content (DLC) and drug loading efficiency (DLE) of the
CC
fabricated PFP-TNDs, 300 μL of acetonitrile was added to 100 μL of the dialyzed PFP TNDs solution to destroy the nanodroplets. The light absorbance of the mixture at 480 nm
A
(the characteristic absorption peak of CAD[31]) was measured by the microplate reader (BioTek Synergy 4) and a standard curve was applied to calculate the concentration of CAD in the samples. DLC and DLE were calculated as: DLC% =
WCAD loaded WPFP−TNDs
× 100% and DLE% =
WCAD loaded WCAD added
× 100%
(WCAD loaded : the weight of the loaded CAD. WCAD added : the weight of CAD added in fabrication of PFP-TNDs. WPFP−TNDs : the weight of PFP-TNDs.). 2.7. Stability and biocompatibility study The stability of the PFP-BNDs and PFP-TNDs was evaluated by monitoring the change
SC RI PT
of particle size in a PBS solution (0.01 M, pH = 7.4, 10% FBS) at 25 ℃ and 37 ℃ on a shaker (80 rpm). Samples were collected at various time points and the size distribution of the nanodroplets was analyzed by DLS. To evaluate the stability related to vaporization,
PFP-TNDs was incubated in PBS at various temperatures and the microscopic pictures of bubble generation were captured. The stability of CAD loading was evaluated by
measuring the accumulative release of CAD from PFP-TNDs over time in stirring PBS
U
(0.01 M, pH = 7.4 and 5.0) at 37 ℃ with samples treated with ultrasound (frequency 3.5
N
MHz, MI=1.5, 60 s of irradiation) in 30 min.
A
To investigate the biocompatibility of PFP-TNDs, HepG2, CT-26 cells (tumor cell lines) and HUVEC (normal cell line) were seeded in 96-well plates at a density of 5×105
M
cells/well and incubated at 37 ℃ in a 5% CO2 atmosphere for 12 h. PFP-NDs/PGA-gmPEG was added in various concentrations of C9F17-PAsp(DET) and the cells were
D
incubated for 24 h. The cell viability was determined by MTT assay. Briefly, the cells were
TE
rinsed twice with PBS (0.01 M, pH = 7.4), followed by addition of 100 μL fresh medium containing MTT (1 mg/mL) and 4 h incubation. The medium was replaced and the
EP
formazan salt crystals were dissolved by 200 μL of DMSO. After incubation for 5 min, the light absorbance of the samples was determined using a microplate reader (BioTek Synergy A
CC
4) at 570 nm. The cell viability was calculated as: Viability% = Asample × 100% (Asample control
and Acontrol represent the absorbance of the sample and the untreated control, respectively).
A
2.8. Erythrocyte agglutination and hemolysis The serum compatibility of the PFP-TNDs was evaluated with erythrocyte agglutination
and hemolysis tests. Erythrocytes were prepared by centrifuging healthy human blood (1500 rpm, 10 min) and washing the depositing erythrocytes with PBS thrice. Erythrocytes were resuspended with PBS at a density of 1×105 cells/mL and various concentrations (5-
80 μg/mL) of the PFP-BNDs and PFP-TNDs were mixed with the erythrocyte suspension for 2 h at 37 ℃. For the agglutination assay, 10 μL of the mixture was observed with a microscope. For hemolysis assay, the mixture was centrifuged (1500 rpm for 5 min) and the light absorbance of supernatant was measured with a microplate reader at 413 nm. Photographs of hemolysis in the samples were taken. Erythrocytes were also treated with
SC RI PT
Triton X-100 (0.1%, w/v) as a positive control or PBS as a negative control. 2.9. In vitro ultrasound imaging
PFP-TNDs were diluted in degassed water (500 μg/mL), loaded in a plastic pipette with
its head submerged in degassed water. A commercial ultrasound platform (Aplio500, Toshiba Medical Systems) with a transducer head (375BT, 3 cm × 1 cm) was used to
U
produce focused ultrasound. Contrast-enhanced ultrasonography imaging was generated
N
with pulsed ultrasound (pulse length 250 ms and mean frequency 4 kHz) at room temperature (25 ℃) and body temperature (37 ℃) with different resonance frequencies and
A
mechanical index (MI) values. For the samples measured at body temperature, the imaging
M
was performed after heating for 20 s. Ultrasound contrast was quantified digitally by
D
measuring the grey-scale intensities on the images captured.
TE
2.10. Cellular uptake with ultrasound Flow cytometry was performed to study the cellular uptake profile of PFP-TNDs triggered by ultrasound irradiation. Briefly, HepG2 cells and CT-26 cells were seeded in
EP
6-well plates at a density of 3×105 cells/well for 12h. The CAD or PFP-TNDs solution was added to each well (net DOX concentration 4 μg/mL) for 15 min, followed by ultrasound
CC
irradiation (frequency 3.5 MHz, MI=1.5, 60 s of irradiation) from an experimental ultrasound platform. Ultrasonic couplant was used to fill the gap between the bottom of
A
plastic plates and the diagnostic probe to facilitate the transmission of sound energy. The uptake of CAD in both cells at 2 h and 4 h following ultrasound irradiation was analyzed with flow cytometry. Briefly, the cells were washed thrice with PBS (0.01 M, pH = 7.4), digested with trypsin, centrifugated (1500 rpm, 5 min) and resuspended in cooled PBS (0.5 mL). The harvested cells were analyzed by flow cytometry (FACS Callibur, Becton Dickinson) at the wavelength of 480 nm. To investigate the cellular uptake of PFP-TNDs
themselves, the PFP-TNDs conjugated with FITC were prepared and used for flow cytometry study using the same method described before. 2.11. Inhibition study The targeting efficacy of γ-PGA was investigated with competitive inhibition cell
SC RI PT
uptake and cytotoxicity study on HepG2 cells (GGT-positive cell line). In the cell uptake
study, HepG2 cells was seeded in a 6-well plate at a density of 2×105 cells/well and
incubated for 12 h. Serum-free culture medium dissolved with various concentrations of free γ-PGA were added to each well for 1 h, followed by replacement with the medium of
10% FBS containing the same amount γ-PGA. PFP-TNDs (4 μg/mL net concentration of
DOX) were then added and the cells were incubated for another 4 h. The uptake of the
U
CAD in HepG2 cells was analyzed using flow cytometry with the method described above.
N
In the cytotoxicity study, HepG2 cells were seeded in 96-well plates at a density of 3×103 cells per well and pre-incubated with serum-free culture medium containing various
A
concentrations of free γ-PGA. After 1 h incubation, the serum-free medium was replaced
M
with the medium containing 10% FBS and the same amount of γ-PGA. Then, the PFPTNDs (4 μg/mL net concentration of DOX) were added to each well, followed by 24 h
D
incubation. The cell viability was measured with MTT assay and calculated with the same
TE
method as mentioned above.
2.12. Cytotoxicity with ultrasound
EP
The cytotoxicity of the PFP-TNDs with the effect of ultrasound irradiation was evaluated by MTT assay. Briefly, HepG2 and CT-26 cells were seeded in 96-well plates at
CC
a density of 3×103 cells per well and then incubated for 24 h. the cells were incubated in medium (DMEM for HepG2 cells and 1640 for CT-26 cells) with free CAD, free
A
DOX·HCl or PFP-TNDs in various net DOX doses for 15 min at 37°C, followed by US irradiation produced by an ultrasonic probe beneath the bottom of wells (frequency 3.5 MHz, MI=1.5, 60 s of irradiation). To ensure accurate transmission of ultrasound energy, all wells adjacent to the sample wells were added with PBS and the gap between the bottom of plastic plates and the diagnostic probe was filled with ultrasonic couplant. The control groups were added with fresh medium instead and were not treated with US. After 6h
following ultrasound exposure, the medium was replaced with 200 μL of fresh medium for the groups that required medium change. And all the cells were incubated for a total of 24 h. The cell viability was measured with MTT assay and calculated according to the methods described above.
SC RI PT
2.13. Statistical Analysis All data are presented as the mean ± standard error of the mean of experiments repeated no less than three times. Comparisons between two groups were statistically analyzed using
the Student’s t test. When the P value was <0.05, the difference was considered statistically
U
significant.
N
3. Results
A
3.1. Polymers and prodrug Synthesis
M
The synthesis strategy of amphiphilic fluorinated copolymer with high affinity to perfluorocarbon and capacity for drug loading was designed and developed based on our previous work[21, 32]. In this work, C9F17-NH2 was used as the ROP initiator to synthesize
D
fluorinated polymer precursor C9F17-PBLA, followed by aminolysis to produce the
TE
cationic C9F17-PAsp(DET). From the 1H-NMR spectrum of C9F17-PBLA, a successful ROP was confirmed by the appearance of aromatic peak (Fig. S1A, 7.34 [COOCH2Ph]).
EP
The degree of polymerization of the benzyl-L-aspartate unit was calculated to be 40.38 and the molecular weight to be 8,776 Da. From the spectrum of C9F17-PAsp(DET), complete
CC
conversion of the benzyl esters to DET moieties in the polymer side chains was confirmed by the disappearance of aromatic peak and the appearance of the DET peaks (Fig.S1B, 2.75
[NHCOCH2CH],
3.20
[NHCH2CH2NH],
3.30
[NH2CH2CH2NH,
A
NHCH2CH2NH], 3.49 [NH2CH2CH2]). DOX was modified by cis-aconityl linkage to form acid-sensitive CAD. The anthracene
protons (Fig. S2A ) was detected in both the spectrums of DOX and CAD, while an additional peak (Fig. S2B ) appeared in the spectrum of CAD, likely corresponding to the proton of -COCH=. The synthesized CAD was further characterized with FTIR (Fig.
S3). Spectrum of DOX marks the light absorbance of the -NH2 group at 3525 and 3329 cm-1, while spectrum of CAD showed an absorbance of -NH- at 3440 cm-1 only. The mass values of CAD and DOX were investigated by ESI- and ESI+, respectively (Fig. S4). Four signals were detected in the spectrum of CAD: (mass to charge ratio, m/z) 676.1, 698.2, 720.1 and 721.1. The theoretical molecular weight of CAD is 699.6 g/mol. Apparently, the
SC RI PT
signal at m/z=676.1 was for [M-CO2-H]-, m/z=698.2 was for [M-H]- and m/z=720.1 and 721.1 were caused by isotope molecules. In the spectrum of DOX, the signal at m/z=544.2
was for [M+H]+ and the rests were caused by isotope molecules. The aforementioned data indicated the successful synthesis of CAD.
Synthesis of PGA-g-mPEG was confirmed by 1H-NMR spectrum (Fig. S5), which showed the characteristic group -OCH3 () in the PGA main chain and -
U
CH2CH2CONH- () in the PEG side chain. The grafting ratio of PEG is calculated to
N
be 14%, which is close to the balanced ratio 10% [21] that offers decent stability and
A
biocompatibility with minimized the hindrance effect of the interaction of nanodroplets to
M
cell membranes.
D
3.2. Fabrication and Optimization of Nanodroplets
TE
The PFP loaded NDs (PFP/ C9F17-PAsp(DET), PFP-NDs) were fabricated using ultrasonic emulsification. The emulsification was clearly observed as the immiscible
EP
mixture of PFP and polymer solution changed into white milk-like dispersion. The dispersion was maintained for at least 12 h, indicating that the PFP was successfully
CC
stabilized by the fluorinated amphiphilic polymer. DLS analysis of the PFP-NDs sample showed a uniformed distribution of particle size that was 340 ± 23 nm and the zeta potential
A
that was 68.5 ± 1.84 mV. CAD was loaded onto the PFP-NDs electrostatically, forming CAD/PFP/ C9F17-
PAsp(DET) binary nanodroplets (PFP-BNDs). For optimization of drug loading efficacy, PFP-BNDs with various mass ratios of C9F17-PAsp(DET) to CAD were fabricated. Measurement results of the PFP-BNDs (Error! Reference source not found.) showed a reduce of zeta potential compare with PFP-NDs at the mass ratio of 20/1 (49.1 ± 2.1 mV
compared with 68.5 ± 1.84 mV), indicating the presence of negatively charged CAD absorbed on the positively charged NDs. Zeta potential of the PFP-BNDs further reduced with the increase of the content of CAD. The size of the nanodroplets became less uniform with the decrease of the mass ratio. Judging from the dispersity index (PDI), the polydispersity increased with the increase of the content of CAD. The sample became non-
SC RI PT
uniform when the mass ratio is below 1/1 and seemingly precipitated. This change of the
size distribution of the nanodroplets is probably attributed to the electrostatic force by the
loaded and oppositely charged CAD that compresses and destabilizes of the particle surface. This is also justified by the reduce of average diameter of the PFP-BNDs when the mass ratio decreased from 20/1 to 5/1. According to the results described above, mass ratio of
C9F17-PAsp(DET) to CAD of 5/1 was selected for all subsequent experiments to maximize
U
the loading of CAD and ensure the integrity of nanodroplets.
N
For further optimization of the nanodroplets, PGA-g-mPEG was applied in this study
A
as a stabilizing coating. PGA-g-mPEG offers superior chemo-physical properties with two functional segments: 1) the γ-PGA segment that mitigates the strong surface positive
M
charge of the PFP-BNDs, which is known to attribute to in vivo aggregation and nonspecific protein binding, and 2) the PEG segments that reduces the clearance in vivo by
D
avoiding opsonization and degradation[33]. The amount of PGA-g-mPEG added was
TE
relative to the amount of the cationic C9F17-PAsp(DET), simplified as the C/N ratio (the total mass of carboxyl groups on PGA chains vs. the total mass of primary amine groups
EP
on PAsp(DET) chains). In order to fabricate the most stable and effective nanodroplet, multiple C/N ratios were tested and the resulting PFP-TNDs were evaluated with DLS
CC
(Error! Reference source not found.). The results showed that at C/N=0.5, the PFPTNDs has a slightly positive surface potential (18.1 ± 3.78 mV), which is both electrostatically stable and preferable for interaction with negatively charged cell
A
membranes. Therefore, the C/N ratio of 0.5 was applied in the fabrication of nanodroplets for the rest of the study. Standardized characterization of the CAD content (UV-vis spectra and fluorescence spectra shown in Fig. S6) showed that the optimized PFP-TNDs had a DLC of 7.56 ± 1.65% and a DLE of 45.13 ± 9.79%. The PFP-BNDs and PFP-TNDs appeared to be spherical in TEM images (0). The size distribution of both nanodroplets appeared were
consistent to the DLS results (Table 1). It is noteworthy that PFP-BNDs tend to aggregate, possibly due to strong surface charge, while PFP-TNDs tend to disperse, indicating the stabilizing effect of PGA-g-mPEG coating.
SC RI PT
3.3. Stability and biocompatibility study The time-dependent stability of the PFP-TNDs was evaluated by in vitro serum test (0A). PFP-TNDs maintained an average size below 550 nm in FBS/PBS solution for at least 48 h at 25℃ or at least 8 h at 37℃. The temporal size increase of PFP-TNDs within 1 h at 37℃ was likely due to partial vaporization as visualized in Fig. S7. By contrast, the size of the PFP-NDs and PFP- BNDs increased dramatically in 5 h and seemingly collapsed
U
afterwards even at lower temperature 25℃. From the unshown DLS data, in addition to the
N
increase of size, PDI of the PFP-NDs and PFP-BNDs rapidly increased from less than 0.2 at 15 min to 1 at 5 h, while the count rate of the samples decreased from above 200 to less
A
than 100. The increase of PDI and decrease of count rate were signs of particle aggregation
M
and vaporization, suggesting the instability of PFP-NDs and PFP-BNDs. In contrast, PFPTNDs are protected from serum protein adsorption by the PGA-g-mPEG coating. Results
D
from CAD release studies (Fig. S8) suggested the stability of CAD loading prior to
PFP-TNDs.
TE
ultrasound irradiation, the ultrasound-triggered CAD release and acid-sensitiveness of
The cytotoxicity of PFP-TNDs was evaluated in cell viability test with HepG2, CT-26
EP
and HUVEC cells incubated with non-loaded nanodroplets (contained no CAD) for 24 h. PFP-NDs/PGA-g-mPEG showed low cytotoxicity (cell viability higher than 80%) in both
CC
cells with concentration as high as 80 µg/ml (0B). In erythrocyte agglutination test, erythrocytes incubated with PFP-TNDs maintained normal morphology comparable to that
A
of the negative control (PBS). In contrast, PFP-BNDs group showed significant erythrocyte agglutination similar to the positive control (Triton X) (0C). The serum biocompatibility of PFP-TNDs was quantified in polymer concentration-dependent hemolysis analysis, where hemolysis activity stayed low in PFP-TNDs group for concentrations as high as 80 µg/ml but skyrocketed in PFP-BNDs group with varying polymer concentration (0D).
3.4. Ultrasonic imaging study The ultrasound contrast capability of PFP-TNDs was tested using a clinical ultrasound imaging device. For an ultrasound theranostic system, a balanced setting is necessary to achieve desirable imaging and therapeutic effects. Diluted PFP-TNDs suspension was first
SC RI PT
examined at 25 ℃ and at 37 ℃ (0). PFP-TNDs noticeably vaporized and produced significant ultrasound contrast at 37 ℃, indicating that the body temperature is preferable for imaging.
We further explored several ultrasound parameters, including resonance frequency and
mechanical index (MI), to acquire the optimal setting that triggers phase transition of PFP,
produces the best backscatter for imaging and is clinically preferable. PFP-TNDs were
U
irradiated with ultrasound at frequencies of 2.5 MHz to 8.0 MHz and a set mechanical
index at 0.08. As visualized by ultrasound imaging (0A), the contrast generated in PFP-
N
TNDs samples enhanced with the increase of frequency, peaked at 3.5 MHz, and depleted
A
at higher frequencies. The results are consistent to the quantification test that showed
M
highest grayscale intensity at the frequency of 3.5 MHz (0B). In the imaging test with different MI, enhanced contrast was clearly observed at an MI of 0.08 (0C, D), signaling
D
the occurrence of microbubble cavitation. The ultrasound contrast depleted with MI above 0.4, which can be explained by bubble coalescence and escape due to high MI [4]. From
TE
the results shown above, we concluded that medium frequency (3.5 MHz) and low MI (0.08) is preferable for imaging upon the administration of PFP-TNDs. This optimized
EP
ultrasound setting, according to previous studies, facilitates ADV and microbubble cavitation[4] and also ensures clinical safety[34]. Time-dependent imaging test was
CC
performed using the optimal setting concluded above (3.5 MHz, MI = 0.08). Strong contrast persisted for at least 40 min (0E), confirming the capability of PFP-TNDs to
A
provide sustained ultrasound imaging.
3.5. Cell uptake study The uptake performance of CAD-loaded PFP-TNDs was examined in cell uptake experiment with HepG2 cells (GGT-positive) and CT-26 cells (GGT-negative), and the
uptake of CAD in the cells was quantified utilizing the characteristic light absorption of CAD (at 480 nm). In addition to the uptake profile, the difference of CAD uptake level was emphasized by comparisons of the mean fluorescence intensity (MFI) as shown in 0. HepG2 cells presented greater sensitivity to ultrasound with PFP-TNDs and overall enhanced uptake compared with CT-26 cells, as the MFI of groups with PFP-TNDs were
SC RI PT
apparently higher than groups with free CAD, and the group treated with ultrasound
showed enhanced uptake both at 2 h and 4 h upon irradiation (0B and 5D). The CT-26 cells even showed slight enhancement effect of ultrasound, however, there was no significant increase of uptake with PFP-TNDs (0A and 5C). The difference in the uptake profile
between HepG2 and CT-26 cells is possibly attributed to the specific nature of HepG2 cell,
which is γ-glutamyl transpeptidase (GGT) positive. As highlighted by previous studies, γ-
U
PGA specifically interacts with GGT receptor, resulting in a significant increase in the
N
uptake of the PFP-TNDs [28]. Furthermore, in order to confirm the intracellular fluorescence signal came from CAD-loaded nanodroplets, the FITC conjugated PFP-TNDs
A
(C9F17-PAsp(DET)(FITC)) was used for additional cell uptake study. Fig. S9 showed
M
similar intracellular fluorescence trend compared with the intracellular CAD fluorescence, which suggested the PFP-nanodroplets could be efficiently uptaked by tumor cells. The
D
above results indicated PFP-TNDs combined with ultrasound irradiation enhances the
TE
uptake of CAD loaded, especially with GGT-positive HepG2 cells.
EP
3.6. Competitive inhibition study To evaluate this enhancement effect of γ-PGA cell targeting delivery, PFP-TNDs were
CC
investigated with receptor competition experiments in GGT-positive HepG2 cells. Flow cytometry results showed a decreased uptake of CAD in samples pretreated with high
A
concentration (30 µM) of the competitive reagent γ-PGA. (Error! Reference source not found.A). Further cytotoxicity study showed that the cytotoxicity of PFP-TNDs steadily decreased with the addition of γ-PGA (Error! Reference source not found.B), confirming the effect of PGA-g-mPEG as an augmenting coating on the uptake of the nanodroplets.
3.7. Cytotoxicity study The therapeutic effect of the ultrasound triggered PFP-TNDs was investigated with in vitro cell viability of HepG2 and CT-26 cells. To highlight the effect of ultrasound, a half of the groups had the medium refreshed at 6 h following ultrasound irradiation so that the
SC RI PT
sustained influence of PFP-TNDs remnants was removed. For comparison, the other half of the groups were incubated for 24 h straight with the medium unchanged.
As shown in 0, at high level of DOX net concentration (above 2 µg/mL) free CAD had lower cytotoxicity than free DOX in both cells, which was explained by the anionic CAD
being less interactive to cell membrane and non-cleaverable in ECM environment. The cytotoxicity of PFP-TNDs without ultrasound treatment was higher than that of free CAD
U
(highlighted in Fig. 6B, D), especially at higher DOX concentrations (above 2 µg/mL).
This result was possibly attributed to the properties of the nanocarrier, including small size,
N
positive surface charge and the PGA-g-mPEG coating, which enhanced drug uptake.
A
Ultrasound irradiation noticeably enhanced the therapeutic effect, as the PFP-TNDs treated
M
with ultrasound presented higher cytotoxicity than untreated PFP-TNDs. This effect appeared on groups with refreshed medium (0A, C) and was more significant on groups
D
with unchanged medium (0B, D). Ultrasound triggered PFP-TNDs presented enhanced cytotoxicity comparable to free DOX on HepG2 cells (0C, D), indicating the augmenting
TE
effect of PGA-g-mPEG coating. The results above show that ultrasound triggered PFPTNDs achieve decent therapeutic effect and it is contributed by factors including
EP
ultrasound cavitation, nanodroplet formulation and the functional prodrug. From the results on cellular uptake and cytotoxicity, we noticed a large room for
CC
improving the ultrasound triggered therapeutic effect of the PFP-TNDs. The intense bubble generation from PFP-TNDs in higher temperatures (Fig. S7) suggested a large window for
A
improvement in the vaporization of PFP and ultrasound cavitation. Strategies of promoting phase-change efficiency and ultrasound cavitation, for example introducing photothermal mechanisms[35], are in consideration in future study to further enhance the ultrasound trigger therapeutic effect of the PFP-TNDs.
4. Conclusions In this study, we formulated an ultrasound triggered, perfluorocarbon and drug loaded polymeric nanodroplet for ultrasound imaging and tumor treatment. The optimized nanodroplet presented good serum stability and biocompatibility. The PFP stabilized in the
SC RI PT
nanodroplets indulged ultrasound sensitivity, enhanced contrast and sustained imaging. Furthermore, the nanodroplets achieved decent cellular uptake of drug and therapeutic effect with the integration of ultrasound cavitation, functional modification of drug and augmenting coating of PGA-g-mPEG. The imaging and therapeutic capability of the ultrasound triggered nanodroplet system can be further improved by applying addition
U
strategies of promoting phase-change efficiency and ultrasound cavitation in future study.
N
5. Acknowledgement
A
This work was supported by the National Natural Science Foundation of China
M
(51773231, 81572726), the National Natural Science Foundation of Guangdong Province (2016A030313315, 2014A030312018), the Tip-top Scientific and Technical Innovative
D
Youth Talents of Guangdong special support program (2014TQ01R651), and the Project of Key Laboratory of Sensing Technology and Biomedical Instruments of Guangdong
A
CC
EP
TE
Province (2011A060901013).
References
A
CC
EP
TE
D
M
A
N
U
SC RI PT
[1] V.P. Torchilin, Nature reviews Drug discovery, 13 (2014) 813. [2] H. Xia, Y. Zhao, R. Tong (Eds.), Therapeutic Ultrasound; Springer, 2016, Chapter 21. [3] S.R. Sirsi, M.A. Borden, Advanced drug delivery reviews, 72 (2014) 3. [4] N. Rapoport, D.A. Christensen, A.M. Kennedy, K.H. Nam, Ultrasound in medicine & biology, 36 (2010) 419. [5] K.W. Ferrara, Advanced drug delivery reviews, 60 (2008) 1097. [6] R.J. Price, D.M. Skyba, S. Kaul, T.C. Skalak, Circulation, 98 (1998) 1264. [7] Q.L. Lu, H.D. Liang, T. Partridge, M.J. Blomley, Gene therapy, 10 (2003) 396. [8] A. Kheirolomoom, P.A. Dayton, A.F. Lum, E. Little, E.E. Paoli, H. Zheng, K.W. Ferrara, Journal of controlled release, 118 (2007) 275. [9] P. Li, Y. Zheng, H. Ran, J. Tan, Y. Lin, Q. Zhang, J. Ren, Z. Wang, Journal of controlled release, 162 (2012) 349. [10] C. Niu, Z. Wang, G. Lu, T.M. Krupka, Y. Sun, Y. You, W. Song, H. Ran, P. Li, Y. Zheng, Biomaterials, 34 (2013) 2307. [11] K. Ferrara, R. Pollard, M. Borden, Annual review of biomedical engineering, 9 (2007) 415. [12] T.M. Krupka, L. Solorio, R.E. Wilson, H. Wu, N. Azar, A.A. Exner, Molecular pharmaceutics, 7 (2010) 49. [13] S. Ibsen, M. Benchimol, D. Simberg, C. Schutt, J. Steiner, S. Esener, Journal of controlled release, 155 (2011) 358. [14] H. Wu, N.G. Rognin, T.M. Krupka, L. Solorio, H. Yoshiara, G. Guenette, C. Sanders, N. Kamiyama, A.A. Exner, Ultrasound in medicine & biology, 39 (2013) 2137. [15] T. Yin, P. Wang, J. Li, R. Zheng, B. Zheng, D. Cheng, R. Li, J. Lai, X. Shuai, Biomaterials, 34 (2013) 4532. [16] P.S. Sheeran, P.A. Dayton, Curr Pharm Design, 18 (2012) 2152. [17] N. Rapoport, Wiley Interdisciplinary Reviews: Nanomedicine and Nanobiotechnology, 4 (2012) 492. [18] N.Y. Rapoport, A.M. Kennedy, J.E. Shea, C.L. Scaife, K.-H. Nam, Journal of Controlled Release, 138 (2009) 268. [19] Z. Gao, A.M. Kennedy, D.A. Christensen, N.Y. Rapoport, Ultrasonics, 48 (2008) 260. [20] N. Rapoport, Z. Gao, A. Kennedy, Journal of the National Cancer Institute, 99 (2007) 1095. [21] D. Gao, M. Xu, Z. Cao, J. Gao, Y. Chen, Y. Li, Z. Yang, X. Xie, Q. Jiang, W. Wang, J. Liu, ACS applied materials & interfaces, 7 (2015) 13524. [22] X. Guan, Y. Li, Z. Jiao, J. Chen, Z. Guo, H. Tian, X. Chen, Acta biomaterialia, 9 (2013) 7672. [23] S.A. Senevirathne, K.E. Washington, M.C. Biewer, M.C. Stefan, Journal of Materials Chemistry B, 4 (2016) 360. [24] E. Tahover, Y.P. Patil, A.A. Gabizon, Anti-cancer drugs, 26 (2015) 241. [25] H.S. Yoo, E.A. Lee, T.G. Park, Journal of Controlled Release, 82 (2002) 17. [26] Y. Bae, N. Nishiyama, S. Fukushima, H. Koyama, M. Yasuhiro, K. Kataoka, Bioconjugate chemistry, 16 (2005) 122.
A
CC
EP
TE
D
M
A
N
U
SC RI PT
[27] Z. Yang, Z. Jiang, Z. Cao, C. Zhang, D. Gao, X. Luo, X. Zhang, H. Luo, Q. Jiang, J. Liu, Nanoscale, 6 (2014) 10193. [28] Z.X. Liao, S.F. Peng, Y.C. Ho, F.L. Mi, B. Maiti, H.W. Sung, Biomaterials, 33 (2012) 3306. [29] T. Akagi, M. Higashi, T. Kaneko, T. Kida, M. Akashi, Macromolecular bioscience, 5 (2005) 598. [30] J.-C. Zhang, J.-X. Ding, C.-S. Xiao, C.-L. He, X.-L. Zhuang, Y.-N. Yang, X.-S. Chen, Chem. J. Chin. Univ., Chin, 33 (2012) 2809. [31] C. Du, D. Deng, L. Shan, S. Wan, J. Cao, J. Tian, S. Achilefu, Y. Gu, Biomaterials, 34 (2013) 3087. [32] D. Gao, J. Gao, M. Xu, Z. Cao, L. Zhou, Y. Li, X. Xie, Q. Jiang, W. Wang, J. Liu, Molecular pharmaceutics, 14 (2017) 984. [33] K. Knop, R. Hoogenboom, D. Fischer, U.S. Schubert, Angewandte chemie international edition, 49 (2010) 6288. [34] J. Qin, T.-Y. Wang, J.K. Willmann (Eds.), Therapeutic Ultrasound; Springer, 2016, p. 263. [35] W. Fan, B. Yung, P. Huang, X. Chen, Chemical reviews, 117 (2017) 13566.
SC RI PT U N A
A
CC
EP
TE
D
M
Fig.1. TEM images and particle size distribution of PFP-BNDs (A, C) and PFP-TNDs (B, D).
SC RI PT U N A M
Fig.2. (A) Changes of particle average size by time for PFP-NDs, PFP-BNDs, PFP-TNDs incubated
A
CC
EP
TE
D
in PBS (0.01 M, pH=7.4) containing 10% FBS. (B) Cell viability of HepG2, CT-26 and HUVEC cells incubated with PFP-NDs/PGA-g-mPEG of different polymer concentration for 24 h. (C) Microscope images of erythrocyte incubated with PFP-BNDs, PFP-TNDs, PBS (negative control), Triton X-100 (positive control). (D) Hemolysis degree relative to polymer concentration of PFP-BNDs and PFPTNDs, as well as hemolysis photos of nanodroplets with a polymer concentration of 80 μg/mL.
SC RI PT
Fig.3. Contrast images of diluted PFP-TNDs at 25 ℃ and 37 ℃ compared with degassed water (A)
A
CC
EP
TE
D
M
A
N
U
and a graph of quantified gray-scale intensities (B).
SC RI PT U N A M D TE
Fig.4. Contrast images of diluted PFP-TNDs treated with ultrasound (3.5 MHz) in different MI (A) at
A
CC
EP
37 ̊C and graph of quantified gray-scale intensities (B). Ultrasound images of diluted PFP-TNDs treated with ultrasound (MI=0.08) in different frequencies (C) and graph of quantified gray-scale intensities (D). Ultrasound images of PFP-TNDs treated with ultrasound (3.5 MHz, MI=0.08) at different time points (E).
SC RI PT U N A M
Fig.5. Cell uptake profile of CT-26 cells and HepG2 cells at 2 h (A, B) and 4 h (C, D) after ultrasound
A
CC
EP
TE
D
irradiation. The mean fluorescence intensity (MFI) of each group was calculated from the uptake profile.
SC RI PT U N A M
A
CC
EP
TE
D
Fig. 6 Cell viability of CT-26 cells (A, B) and HepG2 cells (C, D) incubated with PFP-TNDs, free CAD, free DOX with and without ultrasound irradiation. Groups are divided as (A, C) groups with medium refreshed at 6 h following ultrasound irradiation, and (B, D) groups with medium unchanged for 24 h (*p≤0.05, **p≤0.01, results were taken from three measurements).
SC RI PT U N A M Fabrication of PFP-TNDs and the mechanisms of ultrasound-triggered imaging and
D
Scheme 1.
A
CC
EP
TE
delivery of CAD.
Table 1 Size, PDI and zeta potential of PFP-NDs. optimized PFP-BNDs and PFP-TNDs. Sample PFP-NDs
PDI
Zeta Potential (mV)
340±23
0.133
72.4±8.6
a
351±10
0.166
45.9±2.4
b
432±19
0.164
18.1±3.8
PFP-BNDs PFP-TNDs
The optimized mass ratio of C9F17-PAsp(DET) to CAD of the PFP-BNDs is 5/1.
SC RI PT
a.
Size (nm)
A
CC
EP
TE
D
M
A
N
U
b. The optimized C/N ratio of the PFP-TNDs is 0.2.
S0927-7765(18)30829-4 https://doi.org/10.1016/j.colsurfb.2018.11.046 COLSUB 9811
To appear in:
Colloids and Surfaces B: Biointerfaces
Received date: Revised date: Accepted date:
10 June 2018 24 October 2018 19 November 2018
Please cite this article as: Gao J, Yu B, Li C, Xu M, Cao Z, Xie X, Wang W, Liu J, Ultrasound triggered phase-change nanodroplets for doxorubicin prodrug delivery and ultrasound diagnosis: an in vitro study, Colloids and Surfaces B: Biointerfaces (2018), https://doi.org/10.1016/j.colsurfb.2018.11.046 This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting proof before it is published in its final form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.
Ultrasound triggered phase-change nanodroplets for doxorubicin prodrug delivery and ultrasound diagnosis: an in
SC RI PT
vitro study
Jinbiao Gao a,†, Baiqing Yu a,†, Chao Li a,†, Ming Xub, Zhong Caoa, Xiaoyan Xieb, Wei Wangb,*, Jie Liua,* a
School of Biomedical Engineering, Sun Yat-sen University, Guangzhou, Guangdong 510006,
China
Department of Medical Ultrasonics, Institute of Diagnostic and Interventional Ultrasound,
U
b
A
N
The First Affiliated Hospital of Sun Yat-sen University, Guangzhou, Guangdong, 510080, China
These authors contribute equally to this work.
*
E-mail: [email protected]
M
†
A
CC
EP
TE
D
[email protected]
Graphical Abstract
Highlights Phase-change contrast agent and acid-cleavable prodrug were co-loaded in a single nanodroplet system
Ultrasound triggered phase-change nanodroplets for enhanced ultrasound imaging
Enhanced uptake of drug and therapeutic effect were achieved with ultrasound irradiation
SC RI PT
ABSTRACT
U
Ultrasound-triggered delivery system is among the various multifunctional and stimuli-
N
responsive strategies that hold great potential as a robust solution to the challenges of
A
localized drug delivery and gene therapy. In this work, we developed an ultrasoundtriggered delivery system PFP/C9F17-PAsp(DET)/CAD/PGA-g-mPEG nanodroplet, which
M
combined ultrasound responsive phase-change contrast agent, acid-cleavable doxorubicin prodrug and cationic amphiphilic fluorinated polymer carrier, aiming to achieve both high
D
imaging contrast and preferable ultrasound-triggered anti-cancer therapeutic effect. The
TE
optimized nanodroplets were characterized as monodispersed particles with a diameter of about 400 nm, slightly positive surface charge and high drug-loading efficiency. The functional augmenter PGA-g-mPEG provided the nanodroplets good sustainability, low
EP
cytotoxicity and good serum compatibility, as confirmed by stability and biocompatibility tests. In ultrasound imaging study, the nanodroplets produced significant contrast with
CC
ultrasound irradiation (3.5 MHz, MI=0.08) at 37 ℃. Cell uptake and cytotoxicity studies in HepG2 and CT-26 cells showed the enhanced drug uptake and therapeutic effect with the
A
combination of nanodroplets and ultrasound irradiation. These results suggest that the PFP/CAD-loaded phase change nano-emulsion can be utilized as an efficient theranostic agent for both ultrasound contrast imaging and drug delivery.
KEYWORDS: ultrasound; perfluorocarbon; phase-change contrast agent; nanodroplets; doxorubicin;
1. Introduction
SC RI PT
Ultrasound-triggered delivery system is among the various multifunctional and stimuliresponsive strategies that have been extensively investigated and employed for localized
drug delivery and gene therapy[1]. As an external stimulus, ultrasound outperforms many other physical and biochemical stimuli with its low cast, noninvasiveness, temporal and
spatial control, depth of stimulation and good patient acceptance[2]. The robust properties of an ideal ultrasound-triggered drug delivery system include in vivo stability of drug
U
carrier prior to ultrasound irradiation, enhanced localized therapeutic effect under the
N
acoustic field, and contrast effect for imaging[3]. These advantages of the ultrasound-
A
triggered drug delivery vehicles are attributed to the unique mechanics of ultrasound irradiation that mainly include 1) cavitation caused by ultrasound shockwave that disrupts
M
the drug carriers and nearby tissue membrane to promote release and extravasation, 2) acoustic streaming that drives particles into the target site and 3) hyperthermia that can act
D
as an additional stimulus[4, 5].
TE
The most explored ultrasound-trigger drug delivery vehicles include ultrasoundresponsive polymer or lipid nanocarriers and gas-filled microbubbles, which actively
EP
respond to cavitation and enhance signal echoes for imaging[5]. Early studies exploited the ultrasound responsiveness of microbubbles by co-administrating microbubbles with
CC
therapeutic agents or nanocarriers to facilitate the delivery and uptake[6, 7]. The application of microbubbles was taken one step forward with the development of functionalized shelled microbubbles, including lipid-microbubbles and polymer-
A
microbubbles that carry drug or gene payload[8-10]. However, the large-size microbubbles have limited permeability, short circulation time and relatively low payload[11-13]. Nanobubbles emerged recently with refined fabrication strategies, aiming to strengthen passive targeting and increase the half-life of the system, and have achieved positive results[14, 15].
As an alternative to gas bubbles, phase-change agent has gained inflated attention with its combined strength: the high tissue permeability due to the small size and high acoustic responsiveness due to the phase-changing attribute[16, 17]. Phase-change agents can be encapsulated and stabilized by polymer micelles in the form of nanodroplets in normal physiological environment. When exposed to ultrasound, acoustic droplet vaporization
SC RI PT
(ADV) of the phase-change agent is triggered, producing contrast effect and promoting
nanodroplet extravasation[18]. Perfluoropentane (PFP) is one of the pioneering phase-
change agents applied in ultrasound-triggered drug delivery vehicles. In depth reports
indicate that PFP encapsulated in drug-loaded nanodroplets has a phase-changing temperature as high as 37 ℃, providing contrast effect, enhancing the uptake of therapeutic
agents in vivo and thus showing prospective value in drug delivery and gene therapy[19,
U
20]. In our previous works, we developed an amphiphilic cationic polymer C 11F17-
N
PAsp(DET), which has a fluorinated end group affinitive to PFP and cationic side chains for gene loading. The PFP-containing nanocarrier showed encouraging enhancement of
A
gene transfection efficacy with insonation and promising potential as an efficient drug
M
delivery vehicle[21].
Doxorubicin(DOX), an anthracycline topoisomerase II inhibitor, has been long applied
D
as an effective chemotherapy agent in cancer treatments. Free DOX is not tumor-specific,
TE
and causes damage to healthy tissues[22], requiring the assistance of efficient drug delivery systems. The engineering of DOX delivery systems is inspired by one of the most explored
EP
site-specific stimulus: the acidic environment of tumor tissue (pH≈6.5) and intracellular endosomes (pH=5.0-6.0). A promising approach in the design of pH-responsive systems
CC
involves conjugating the drug molecule with acid-sensitive linkages like hydrazone, carbamate or cis-aconityl bond. Drug conjugated remains stable at neutral pH during transportation and is triggered to release after arriving at tumor site or intracell, thus
A
promoting target-specificity, cell uptake and reducing toxicity in irrelevant sites[1, 23]. Prodrugs with conjugated DOX are reported to have high loading efficiency, longer circulation half-life and decent pH-responsiveness[24-26]. It is worth noting that besides acid-sensibility, cis-aconityl doxorubicin (CAD) is highly soluble in water and negatively charged, and is thus a superior choice of drug payload for the cationic amphiphilic carrier system.
In this study, we developed a PFP-containing and CAD-loaded ultrasound-triggered drug delivery system to achieve optimized imaging, uptake and therapeutic effect. As shown schematically in Scheme 1, The phase-change agent PFP is encapsulated in the amphiphilic cationic polymer C9F17-PAsp(DET) in the form of nanodroplets. The pretreated DOX, cis-aconityl doxorubicin (CAD) is absorbed onto the nanodroplet
SC RI PT
electrostatically. The nanodroplet is further coated with PGA-g-mPEG, an augmenting coating used in our previous study[27]. This coating material consists of polyethene glycol
(PEG) chains to stabilize the nanodroplet and gamma-polyglutamic acid (γ-PGA), an anionic polymer known to specifically interact with gamma-glutamyl transpeptidase
(GGT) in mammalian cell lines to enhance cell uptake via a GGT-mediated pathway[28,
29]. The obtained ternary nanodroplets, PFP/C9F17-PAsp(DET)/CAD/PGA-g-mPEG
U
(PFP-TNDs), was measured with dynamic light scattering (DLS) and observed with
N
transmission electron microscopy (TEM) to determine the physical characteristics. Biocompatibility tests and competitive inhibition experiments were conducted to
A
investigate the stability of the TNDs and the effect of γ-PGA, respectively. Furthermore,
M
the ultrasound responsiveness and therapeutic efficacy of the PFP-containing, CAD-loaded
2. Methods
EP
2.1. Materials
TE
cytotoxicity assays.
D
TNDs against cancer cells were evaluated by in vitro ultrasound imaging, cell uptake and
β-benzyl-L-aspartate N-carboxyanhydride (BLA-NCA) was purchased from Beijing
CC
HWRK Chem (Beijing, China) and used as received. 2,2,3,3,4,4,5,5,6,6,7,7,8,8,9,9,9heptadecafluorononylamine (C9F17-NH2) and fluorescence isothiocyanate (FITC) were purchased from Sigma-Aldrich. N-methyl-2-pyrrolidone (NMP) and diethylenetriamine
A
(DET) were purchased from the Aladdin Industrial (Shanghai, China) and used after distillation. Perfluoro-n-pentane (PFP) was purchased from Strem Chemicals. Doxorubicin hydrochloride (DOX·HCl) was purchased from Beijing HVSF United Chemical Materials (Beijing, China). Cis-aconitic anhydride (CA) was purchased from Alfa Aesar China (Tianjin, China). Dimethyl sulfoxide-d6 (DMSO-d6) and deuterium oxide (D2O) were purchased from Adamas Reagent. Poly(γ-glutamic acid) (γ-PGA; 5000 Da) was purchased
from Nanjing Sai Taisi Biotechnology (Nanjing, China). HepG2 cells, CT-26 cells and HUVEC cells were purchased from the Shanghai Cell Bank of the Chinese Academy of Sciences (Shanghai, China). HepG2 and HUVEC cells were cultured in Dulbecco’s modified Eagle’s medium (DMEM, Hyclone) and CT-26 cells were cultured in RPMI Medium 1640 (Hyclone), supplemented with 1% (w/v) penicillin-streptomycin and 10%
SC RI PT
(v/v) fetal bovine serum (FBS) and in a 5% CO2 environment at 37 ℃.
2.2. Synthesis of C9F17-poly{N-[N’-(2-aminoethyl)]-aspartamide} [C9F17-PAsp(DET)]
C9F17-PAsp(DET) was prepared stepwise according to Scheme S1. C9F17-PBLA was
synthesized with the BLA-NCA ring opening polymerization initiated by C9F17- NH2. BLA-NCA (0.5 g) and C9F17- NH2 (37.5 mg) was dissolved in 5 mL of anhydrous dimethyl
U
formamide. Dichloromethane (5 mL) was added dropwise into the above mixture solution.
N
The polymerization was kept in a dry N2 atmosphere for 3 days at 25 ℃. Then, the solution was dialyzed against dimethyl sulfoxide for 1 day and against deionized water for 2 days.
A
C9F17-PBLA (white powder) was obtained by lyophilization. C9F17-PAsp(DET) was
M
synthesized with aminolysis of the side chains. Briefly, C9F17-PBLA (150 mg) and was dissolved in NMP (5 mL) and DET (1 mL, mixed with 4 mL of NMP) was added dropwise
D
into the solution. 1 M HCl (9.7 mL) was injected to the mixture in an ice bath after 6 h of
TE
stirring, followed by dialysis against 0.01 M HCl for 1 day and against water for 2 days. C9F17-PAsp(DET) was obtained as a yellowish solid after lyophilization. C9F17PAsp(DET)-FITC was synthesized by adding FITC (12.3 mg dispersed in 4 mL deionized
EP
water) into C9F17-PAsp(DET) solution (50 mg in 6 mL deionized water) and stirring for 12 h, followed by dialysis against water and lyophilization. All products in the workflow were
CC
stored in -20 ℃.
A
2.3. Synthesis of cis-aconityl-doxorubicin (CAD) The CAD was synthesized according to the previous study [30]. Briefly, Doxorubicin
hydrochloride (DOX·HCl, 50 mg) was dissolved in 6 mL of 0.1M carbonate buffer (pH=9.4) and cooled in an ice bath. CA (70 mg) in 0.6 mL absolute 1,4-dioxane solution was added dropwise with stirring, and the pH was kept at 9 by careful addition of cold 0.5 M NaOH and the following 15 min of stirring. Insoluble contents were removed by
centrifugation of the mixture (13000 rpm, 5 min, 4 ℃). CAD was precipitated by addition of cooled 1 M HCl to the mixture (the pH was reduced to 1) and was collected with centrifugation. The collected CAD was re-dissolved in deionized water with pH of the solution adjusted to neutral by the addition of 0.2 M NaOH. To further purify the product, the solution was washed with 4 mL chloroform/methanol (95:5, v/v) for five times. The
SC RI PT
organic solvent was removed on the rotary evaporator and the solution was lyophilized. CAD was obtained as a red powder and was stored in dark environment at -20 ℃. 2.4. Characterization of the synthetic products
C9F17-PBLA (in DMSO-d6) and C9F17-PAsp(DET) (in D2O) were characterized by proton nuclear magnetic resonance spectroscopy (1H-NMR) using a Bruker AvanceIII
U
spectrometer (400 MHz). Furthermore, the molecular weight and of the polymers was
N
calculated from the spectra.
A
The DOX and CAD were characterized by 1H-NMR spectroscopy (both in DMSO-d6), Fourier transform infrared spectroscopy (FTIR) and electrospray ionization mass
M
spectrometry (ESI-MS). The FTIR spectra were obtained with the potassium bromide method. The ESI-MS was detected with a Thermo Finnigan TSQ Quantum Ultra AM EMR
D
Mass Spectrometer. The mass value of CAD was determined by negative ion mode (ESI-)
was 0.2 mL/min.
TE
and DOX was determined by positive ion mode (ESI+). The flow rate of the mobile phase
EP
2.5. Preparation of nanodroplets The PFP/C9F17-PAsp(DET) nanodroplets (PFP-NDs) were fabricated with oil/water
CC
emulsification. Briefly, 5 mg of C9F17-PAsp(DET) was dissolved in 1 mL of deionized water and cooled at 4 ℃ for 5 min. PFP (60 μL) was slowly added to the solution, followed
A
by sonication with a Misonix Sonicator (S-4000, with probe) for 90 s in a cycle of 1 s sonication at 22% amplitude and 1 s standby in an ice bath. The resulting PFP-NDs were kept at 4 ℃ and used in no more than 1 h. The PFP/C9F17-PAsp(DET)/CAD binary nanodroplets (PFP-BNDs) and modified PFP/C9F17-PAsp(DET)/CAD/PGA-g-mPEG ternary nanodroplets (PFP-TNDs) were prepared stepwise as shown in Scheme 1. The PFP-BNDs were prepared in different C9F17-
PAsp(DET)/CAD mass ratios by gently mixing the PFP-NDs aqueous suspension with the CAD aqueous solution. After incubation at 4 ℃ for 30 min, a certain amount of PGA-gmPEG (prepared according to the method previously reported by Yang et al) aqueous solution was added to obtain the ternary nanodroplets PFP-TNDs in various C/N ratios (the molar ratio of the dissociative carboxyl in PGA-g-mPEG to the primary amine groups in
SC RI PT
the N-substituted poly-aspartamides), followed by incubation at 4 ℃ for 30 min. Finally,
the mixture was dialyzed against 5% glucose solution for 12 h at 4 ℃. For biocompatibility test and uptake experiments, we fabricated PFP/C9F17-PAsp(DET)/PGA-g-mPEG and PFP
C9F17-PAsp(DET)-FITC/PGA-g-mPEG nanodroplets, respectively, in which CAD solution was replaced by the same amount of deionized water. Nanodroplets were used
U
upon dialysis.
N
2.6. Characterization of nanodroplets
The particle size of PFP-NDs, PFP-BNDs, and PFP-TNDs were determined with
A
dynamic light scattering (DLS, Malvern Zetasizer). The zeta potential of the nanodroplets
M
were determined by phase analysis light scattering (Malvern Zetasizer). The morphology of the nanodroplets (stained with phosphotungstic acid) was observed by a transmission
D
electron microscope (TEM, JEOLJEM 1400).
TE
For quantification of CAD in the following tests, the UV-vis absorbance spectra of PFPTNDs, PFP-NDs/PGA-g-mPEG and free CAD was measured by the UV-vis
EP
spectrophometer (Beckman DU730) and the fluorescence spectra with the excitation wavelength of 480 nm was measured by the spectrofluorometer (Horiba FluoroMax 4). For calculation of drug loading content (DLC) and drug loading efficiency (DLE) of the
CC
fabricated PFP-TNDs, 300 μL of acetonitrile was added to 100 μL of the dialyzed PFP TNDs solution to destroy the nanodroplets. The light absorbance of the mixture at 480 nm
A
(the characteristic absorption peak of CAD[31]) was measured by the microplate reader (BioTek Synergy 4) and a standard curve was applied to calculate the concentration of CAD in the samples. DLC and DLE were calculated as: DLC% =
WCAD loaded WPFP−TNDs
× 100% and DLE% =
WCAD loaded WCAD added
× 100%
(WCAD loaded : the weight of the loaded CAD. WCAD added : the weight of CAD added in fabrication of PFP-TNDs. WPFP−TNDs : the weight of PFP-TNDs.). 2.7. Stability and biocompatibility study The stability of the PFP-BNDs and PFP-TNDs was evaluated by monitoring the change
SC RI PT
of particle size in a PBS solution (0.01 M, pH = 7.4, 10% FBS) at 25 ℃ and 37 ℃ on a shaker (80 rpm). Samples were collected at various time points and the size distribution of the nanodroplets was analyzed by DLS. To evaluate the stability related to vaporization,
PFP-TNDs was incubated in PBS at various temperatures and the microscopic pictures of bubble generation were captured. The stability of CAD loading was evaluated by
measuring the accumulative release of CAD from PFP-TNDs over time in stirring PBS
U
(0.01 M, pH = 7.4 and 5.0) at 37 ℃ with samples treated with ultrasound (frequency 3.5
N
MHz, MI=1.5, 60 s of irradiation) in 30 min.
A
To investigate the biocompatibility of PFP-TNDs, HepG2, CT-26 cells (tumor cell lines) and HUVEC (normal cell line) were seeded in 96-well plates at a density of 5×105
M
cells/well and incubated at 37 ℃ in a 5% CO2 atmosphere for 12 h. PFP-NDs/PGA-gmPEG was added in various concentrations of C9F17-PAsp(DET) and the cells were
D
incubated for 24 h. The cell viability was determined by MTT assay. Briefly, the cells were
TE
rinsed twice with PBS (0.01 M, pH = 7.4), followed by addition of 100 μL fresh medium containing MTT (1 mg/mL) and 4 h incubation. The medium was replaced and the
EP
formazan salt crystals were dissolved by 200 μL of DMSO. After incubation for 5 min, the light absorbance of the samples was determined using a microplate reader (BioTek Synergy A
CC
4) at 570 nm. The cell viability was calculated as: Viability% = Asample × 100% (Asample control
and Acontrol represent the absorbance of the sample and the untreated control, respectively).
A
2.8. Erythrocyte agglutination and hemolysis The serum compatibility of the PFP-TNDs was evaluated with erythrocyte agglutination
and hemolysis tests. Erythrocytes were prepared by centrifuging healthy human blood (1500 rpm, 10 min) and washing the depositing erythrocytes with PBS thrice. Erythrocytes were resuspended with PBS at a density of 1×105 cells/mL and various concentrations (5-
80 μg/mL) of the PFP-BNDs and PFP-TNDs were mixed with the erythrocyte suspension for 2 h at 37 ℃. For the agglutination assay, 10 μL of the mixture was observed with a microscope. For hemolysis assay, the mixture was centrifuged (1500 rpm for 5 min) and the light absorbance of supernatant was measured with a microplate reader at 413 nm. Photographs of hemolysis in the samples were taken. Erythrocytes were also treated with
SC RI PT
Triton X-100 (0.1%, w/v) as a positive control or PBS as a negative control. 2.9. In vitro ultrasound imaging
PFP-TNDs were diluted in degassed water (500 μg/mL), loaded in a plastic pipette with
its head submerged in degassed water. A commercial ultrasound platform (Aplio500, Toshiba Medical Systems) with a transducer head (375BT, 3 cm × 1 cm) was used to
U
produce focused ultrasound. Contrast-enhanced ultrasonography imaging was generated
N
with pulsed ultrasound (pulse length 250 ms and mean frequency 4 kHz) at room temperature (25 ℃) and body temperature (37 ℃) with different resonance frequencies and
A
mechanical index (MI) values. For the samples measured at body temperature, the imaging
M
was performed after heating for 20 s. Ultrasound contrast was quantified digitally by
D
measuring the grey-scale intensities on the images captured.
TE
2.10. Cellular uptake with ultrasound Flow cytometry was performed to study the cellular uptake profile of PFP-TNDs triggered by ultrasound irradiation. Briefly, HepG2 cells and CT-26 cells were seeded in
EP
6-well plates at a density of 3×105 cells/well for 12h. The CAD or PFP-TNDs solution was added to each well (net DOX concentration 4 μg/mL) for 15 min, followed by ultrasound
CC
irradiation (frequency 3.5 MHz, MI=1.5, 60 s of irradiation) from an experimental ultrasound platform. Ultrasonic couplant was used to fill the gap between the bottom of
A
plastic plates and the diagnostic probe to facilitate the transmission of sound energy. The uptake of CAD in both cells at 2 h and 4 h following ultrasound irradiation was analyzed with flow cytometry. Briefly, the cells were washed thrice with PBS (0.01 M, pH = 7.4), digested with trypsin, centrifugated (1500 rpm, 5 min) and resuspended in cooled PBS (0.5 mL). The harvested cells were analyzed by flow cytometry (FACS Callibur, Becton Dickinson) at the wavelength of 480 nm. To investigate the cellular uptake of PFP-TNDs
themselves, the PFP-TNDs conjugated with FITC were prepared and used for flow cytometry study using the same method described before. 2.11. Inhibition study The targeting efficacy of γ-PGA was investigated with competitive inhibition cell
SC RI PT
uptake and cytotoxicity study on HepG2 cells (GGT-positive cell line). In the cell uptake
study, HepG2 cells was seeded in a 6-well plate at a density of 2×105 cells/well and
incubated for 12 h. Serum-free culture medium dissolved with various concentrations of free γ-PGA were added to each well for 1 h, followed by replacement with the medium of
10% FBS containing the same amount γ-PGA. PFP-TNDs (4 μg/mL net concentration of
DOX) were then added and the cells were incubated for another 4 h. The uptake of the
U
CAD in HepG2 cells was analyzed using flow cytometry with the method described above.
N
In the cytotoxicity study, HepG2 cells were seeded in 96-well plates at a density of 3×103 cells per well and pre-incubated with serum-free culture medium containing various
A
concentrations of free γ-PGA. After 1 h incubation, the serum-free medium was replaced
M
with the medium containing 10% FBS and the same amount of γ-PGA. Then, the PFPTNDs (4 μg/mL net concentration of DOX) were added to each well, followed by 24 h
D
incubation. The cell viability was measured with MTT assay and calculated with the same
TE
method as mentioned above.
2.12. Cytotoxicity with ultrasound
EP
The cytotoxicity of the PFP-TNDs with the effect of ultrasound irradiation was evaluated by MTT assay. Briefly, HepG2 and CT-26 cells were seeded in 96-well plates at
CC
a density of 3×103 cells per well and then incubated for 24 h. the cells were incubated in medium (DMEM for HepG2 cells and 1640 for CT-26 cells) with free CAD, free
A
DOX·HCl or PFP-TNDs in various net DOX doses for 15 min at 37°C, followed by US irradiation produced by an ultrasonic probe beneath the bottom of wells (frequency 3.5 MHz, MI=1.5, 60 s of irradiation). To ensure accurate transmission of ultrasound energy, all wells adjacent to the sample wells were added with PBS and the gap between the bottom of plastic plates and the diagnostic probe was filled with ultrasonic couplant. The control groups were added with fresh medium instead and were not treated with US. After 6h
following ultrasound exposure, the medium was replaced with 200 μL of fresh medium for the groups that required medium change. And all the cells were incubated for a total of 24 h. The cell viability was measured with MTT assay and calculated according to the methods described above.
SC RI PT
2.13. Statistical Analysis All data are presented as the mean ± standard error of the mean of experiments repeated no less than three times. Comparisons between two groups were statistically analyzed using
the Student’s t test. When the P value was <0.05, the difference was considered statistically
U
significant.
N
3. Results
A
3.1. Polymers and prodrug Synthesis
M
The synthesis strategy of amphiphilic fluorinated copolymer with high affinity to perfluorocarbon and capacity for drug loading was designed and developed based on our previous work[21, 32]. In this work, C9F17-NH2 was used as the ROP initiator to synthesize
D
fluorinated polymer precursor C9F17-PBLA, followed by aminolysis to produce the
TE
cationic C9F17-PAsp(DET). From the 1H-NMR spectrum of C9F17-PBLA, a successful ROP was confirmed by the appearance of aromatic peak (Fig. S1A, 7.34 [COOCH2Ph]).
EP
The degree of polymerization of the benzyl-L-aspartate unit was calculated to be 40.38 and the molecular weight to be 8,776 Da. From the spectrum of C9F17-PAsp(DET), complete
CC
conversion of the benzyl esters to DET moieties in the polymer side chains was confirmed by the disappearance of aromatic peak and the appearance of the DET peaks (Fig.S1B, 2.75
[NHCOCH2CH],
3.20
[NHCH2CH2NH],
3.30
[NH2CH2CH2NH,
A
NHCH2CH2NH], 3.49 [NH2CH2CH2]). DOX was modified by cis-aconityl linkage to form acid-sensitive CAD. The anthracene
protons (Fig. S2A ) was detected in both the spectrums of DOX and CAD, while an additional peak (Fig. S2B ) appeared in the spectrum of CAD, likely corresponding to the proton of -COCH=. The synthesized CAD was further characterized with FTIR (Fig.
S3). Spectrum of DOX marks the light absorbance of the -NH2 group at 3525 and 3329 cm-1, while spectrum of CAD showed an absorbance of -NH- at 3440 cm-1 only. The mass values of CAD and DOX were investigated by ESI- and ESI+, respectively (Fig. S4). Four signals were detected in the spectrum of CAD: (mass to charge ratio, m/z) 676.1, 698.2, 720.1 and 721.1. The theoretical molecular weight of CAD is 699.6 g/mol. Apparently, the
SC RI PT
signal at m/z=676.1 was for [M-CO2-H]-, m/z=698.2 was for [M-H]- and m/z=720.1 and 721.1 were caused by isotope molecules. In the spectrum of DOX, the signal at m/z=544.2
was for [M+H]+ and the rests were caused by isotope molecules. The aforementioned data indicated the successful synthesis of CAD.
Synthesis of PGA-g-mPEG was confirmed by 1H-NMR spectrum (Fig. S5), which showed the characteristic group -OCH3 () in the PGA main chain and -
U
CH2CH2CONH- () in the PEG side chain. The grafting ratio of PEG is calculated to
N
be 14%, which is close to the balanced ratio 10% [21] that offers decent stability and
A
biocompatibility with minimized the hindrance effect of the interaction of nanodroplets to
M
cell membranes.
D
3.2. Fabrication and Optimization of Nanodroplets
TE
The PFP loaded NDs (PFP/ C9F17-PAsp(DET), PFP-NDs) were fabricated using ultrasonic emulsification. The emulsification was clearly observed as the immiscible
EP
mixture of PFP and polymer solution changed into white milk-like dispersion. The dispersion was maintained for at least 12 h, indicating that the PFP was successfully
CC
stabilized by the fluorinated amphiphilic polymer. DLS analysis of the PFP-NDs sample showed a uniformed distribution of particle size that was 340 ± 23 nm and the zeta potential
A
that was 68.5 ± 1.84 mV. CAD was loaded onto the PFP-NDs electrostatically, forming CAD/PFP/ C9F17-
PAsp(DET) binary nanodroplets (PFP-BNDs). For optimization of drug loading efficacy, PFP-BNDs with various mass ratios of C9F17-PAsp(DET) to CAD were fabricated. Measurement results of the PFP-BNDs (Error! Reference source not found.) showed a reduce of zeta potential compare with PFP-NDs at the mass ratio of 20/1 (49.1 ± 2.1 mV
compared with 68.5 ± 1.84 mV), indicating the presence of negatively charged CAD absorbed on the positively charged NDs. Zeta potential of the PFP-BNDs further reduced with the increase of the content of CAD. The size of the nanodroplets became less uniform with the decrease of the mass ratio. Judging from the dispersity index (PDI), the polydispersity increased with the increase of the content of CAD. The sample became non-
SC RI PT
uniform when the mass ratio is below 1/1 and seemingly precipitated. This change of the
size distribution of the nanodroplets is probably attributed to the electrostatic force by the
loaded and oppositely charged CAD that compresses and destabilizes of the particle surface. This is also justified by the reduce of average diameter of the PFP-BNDs when the mass ratio decreased from 20/1 to 5/1. According to the results described above, mass ratio of
C9F17-PAsp(DET) to CAD of 5/1 was selected for all subsequent experiments to maximize
U
the loading of CAD and ensure the integrity of nanodroplets.
N
For further optimization of the nanodroplets, PGA-g-mPEG was applied in this study
A
as a stabilizing coating. PGA-g-mPEG offers superior chemo-physical properties with two functional segments: 1) the γ-PGA segment that mitigates the strong surface positive
M
charge of the PFP-BNDs, which is known to attribute to in vivo aggregation and nonspecific protein binding, and 2) the PEG segments that reduces the clearance in vivo by
D
avoiding opsonization and degradation[33]. The amount of PGA-g-mPEG added was
TE
relative to the amount of the cationic C9F17-PAsp(DET), simplified as the C/N ratio (the total mass of carboxyl groups on PGA chains vs. the total mass of primary amine groups
EP
on PAsp(DET) chains). In order to fabricate the most stable and effective nanodroplet, multiple C/N ratios were tested and the resulting PFP-TNDs were evaluated with DLS
CC
(Error! Reference source not found.). The results showed that at C/N=0.5, the PFPTNDs has a slightly positive surface potential (18.1 ± 3.78 mV), which is both electrostatically stable and preferable for interaction with negatively charged cell
A
membranes. Therefore, the C/N ratio of 0.5 was applied in the fabrication of nanodroplets for the rest of the study. Standardized characterization of the CAD content (UV-vis spectra and fluorescence spectra shown in Fig. S6) showed that the optimized PFP-TNDs had a DLC of 7.56 ± 1.65% and a DLE of 45.13 ± 9.79%. The PFP-BNDs and PFP-TNDs appeared to be spherical in TEM images (0). The size distribution of both nanodroplets appeared were
consistent to the DLS results (Table 1). It is noteworthy that PFP-BNDs tend to aggregate, possibly due to strong surface charge, while PFP-TNDs tend to disperse, indicating the stabilizing effect of PGA-g-mPEG coating.
SC RI PT
3.3. Stability and biocompatibility study The time-dependent stability of the PFP-TNDs was evaluated by in vitro serum test (0A). PFP-TNDs maintained an average size below 550 nm in FBS/PBS solution for at least 48 h at 25℃ or at least 8 h at 37℃. The temporal size increase of PFP-TNDs within 1 h at 37℃ was likely due to partial vaporization as visualized in Fig. S7. By contrast, the size of the PFP-NDs and PFP- BNDs increased dramatically in 5 h and seemingly collapsed
U
afterwards even at lower temperature 25℃. From the unshown DLS data, in addition to the
N
increase of size, PDI of the PFP-NDs and PFP-BNDs rapidly increased from less than 0.2 at 15 min to 1 at 5 h, while the count rate of the samples decreased from above 200 to less
A
than 100. The increase of PDI and decrease of count rate were signs of particle aggregation
M
and vaporization, suggesting the instability of PFP-NDs and PFP-BNDs. In contrast, PFPTNDs are protected from serum protein adsorption by the PGA-g-mPEG coating. Results
D
from CAD release studies (Fig. S8) suggested the stability of CAD loading prior to
PFP-TNDs.
TE
ultrasound irradiation, the ultrasound-triggered CAD release and acid-sensitiveness of
The cytotoxicity of PFP-TNDs was evaluated in cell viability test with HepG2, CT-26
EP
and HUVEC cells incubated with non-loaded nanodroplets (contained no CAD) for 24 h. PFP-NDs/PGA-g-mPEG showed low cytotoxicity (cell viability higher than 80%) in both
CC
cells with concentration as high as 80 µg/ml (0B). In erythrocyte agglutination test, erythrocytes incubated with PFP-TNDs maintained normal morphology comparable to that
A
of the negative control (PBS). In contrast, PFP-BNDs group showed significant erythrocyte agglutination similar to the positive control (Triton X) (0C). The serum biocompatibility of PFP-TNDs was quantified in polymer concentration-dependent hemolysis analysis, where hemolysis activity stayed low in PFP-TNDs group for concentrations as high as 80 µg/ml but skyrocketed in PFP-BNDs group with varying polymer concentration (0D).
3.4. Ultrasonic imaging study The ultrasound contrast capability of PFP-TNDs was tested using a clinical ultrasound imaging device. For an ultrasound theranostic system, a balanced setting is necessary to achieve desirable imaging and therapeutic effects. Diluted PFP-TNDs suspension was first
SC RI PT
examined at 25 ℃ and at 37 ℃ (0). PFP-TNDs noticeably vaporized and produced significant ultrasound contrast at 37 ℃, indicating that the body temperature is preferable for imaging.
We further explored several ultrasound parameters, including resonance frequency and
mechanical index (MI), to acquire the optimal setting that triggers phase transition of PFP,
produces the best backscatter for imaging and is clinically preferable. PFP-TNDs were
U
irradiated with ultrasound at frequencies of 2.5 MHz to 8.0 MHz and a set mechanical
index at 0.08. As visualized by ultrasound imaging (0A), the contrast generated in PFP-
N
TNDs samples enhanced with the increase of frequency, peaked at 3.5 MHz, and depleted
A
at higher frequencies. The results are consistent to the quantification test that showed
M
highest grayscale intensity at the frequency of 3.5 MHz (0B). In the imaging test with different MI, enhanced contrast was clearly observed at an MI of 0.08 (0C, D), signaling
D
the occurrence of microbubble cavitation. The ultrasound contrast depleted with MI above 0.4, which can be explained by bubble coalescence and escape due to high MI [4]. From
TE
the results shown above, we concluded that medium frequency (3.5 MHz) and low MI (0.08) is preferable for imaging upon the administration of PFP-TNDs. This optimized
EP
ultrasound setting, according to previous studies, facilitates ADV and microbubble cavitation[4] and also ensures clinical safety[34]. Time-dependent imaging test was
CC
performed using the optimal setting concluded above (3.5 MHz, MI = 0.08). Strong contrast persisted for at least 40 min (0E), confirming the capability of PFP-TNDs to
A
provide sustained ultrasound imaging.
3.5. Cell uptake study The uptake performance of CAD-loaded PFP-TNDs was examined in cell uptake experiment with HepG2 cells (GGT-positive) and CT-26 cells (GGT-negative), and the
uptake of CAD in the cells was quantified utilizing the characteristic light absorption of CAD (at 480 nm). In addition to the uptake profile, the difference of CAD uptake level was emphasized by comparisons of the mean fluorescence intensity (MFI) as shown in 0. HepG2 cells presented greater sensitivity to ultrasound with PFP-TNDs and overall enhanced uptake compared with CT-26 cells, as the MFI of groups with PFP-TNDs were
SC RI PT
apparently higher than groups with free CAD, and the group treated with ultrasound
showed enhanced uptake both at 2 h and 4 h upon irradiation (0B and 5D). The CT-26 cells even showed slight enhancement effect of ultrasound, however, there was no significant increase of uptake with PFP-TNDs (0A and 5C). The difference in the uptake profile
between HepG2 and CT-26 cells is possibly attributed to the specific nature of HepG2 cell,
which is γ-glutamyl transpeptidase (GGT) positive. As highlighted by previous studies, γ-
U
PGA specifically interacts with GGT receptor, resulting in a significant increase in the
N
uptake of the PFP-TNDs [28]. Furthermore, in order to confirm the intracellular fluorescence signal came from CAD-loaded nanodroplets, the FITC conjugated PFP-TNDs
A
(C9F17-PAsp(DET)(FITC)) was used for additional cell uptake study. Fig. S9 showed
M
similar intracellular fluorescence trend compared with the intracellular CAD fluorescence, which suggested the PFP-nanodroplets could be efficiently uptaked by tumor cells. The
D
above results indicated PFP-TNDs combined with ultrasound irradiation enhances the
TE
uptake of CAD loaded, especially with GGT-positive HepG2 cells.
EP
3.6. Competitive inhibition study To evaluate this enhancement effect of γ-PGA cell targeting delivery, PFP-TNDs were
CC
investigated with receptor competition experiments in GGT-positive HepG2 cells. Flow cytometry results showed a decreased uptake of CAD in samples pretreated with high
A
concentration (30 µM) of the competitive reagent γ-PGA. (Error! Reference source not found.A). Further cytotoxicity study showed that the cytotoxicity of PFP-TNDs steadily decreased with the addition of γ-PGA (Error! Reference source not found.B), confirming the effect of PGA-g-mPEG as an augmenting coating on the uptake of the nanodroplets.
3.7. Cytotoxicity study The therapeutic effect of the ultrasound triggered PFP-TNDs was investigated with in vitro cell viability of HepG2 and CT-26 cells. To highlight the effect of ultrasound, a half of the groups had the medium refreshed at 6 h following ultrasound irradiation so that the
SC RI PT
sustained influence of PFP-TNDs remnants was removed. For comparison, the other half of the groups were incubated for 24 h straight with the medium unchanged.
As shown in 0, at high level of DOX net concentration (above 2 µg/mL) free CAD had lower cytotoxicity than free DOX in both cells, which was explained by the anionic CAD
being less interactive to cell membrane and non-cleaverable in ECM environment. The cytotoxicity of PFP-TNDs without ultrasound treatment was higher than that of free CAD
U
(highlighted in Fig. 6B, D), especially at higher DOX concentrations (above 2 µg/mL).
This result was possibly attributed to the properties of the nanocarrier, including small size,
N
positive surface charge and the PGA-g-mPEG coating, which enhanced drug uptake.
A
Ultrasound irradiation noticeably enhanced the therapeutic effect, as the PFP-TNDs treated
M
with ultrasound presented higher cytotoxicity than untreated PFP-TNDs. This effect appeared on groups with refreshed medium (0A, C) and was more significant on groups
D
with unchanged medium (0B, D). Ultrasound triggered PFP-TNDs presented enhanced cytotoxicity comparable to free DOX on HepG2 cells (0C, D), indicating the augmenting
TE
effect of PGA-g-mPEG coating. The results above show that ultrasound triggered PFPTNDs achieve decent therapeutic effect and it is contributed by factors including
EP
ultrasound cavitation, nanodroplet formulation and the functional prodrug. From the results on cellular uptake and cytotoxicity, we noticed a large room for
CC
improving the ultrasound triggered therapeutic effect of the PFP-TNDs. The intense bubble generation from PFP-TNDs in higher temperatures (Fig. S7) suggested a large window for
A
improvement in the vaporization of PFP and ultrasound cavitation. Strategies of promoting phase-change efficiency and ultrasound cavitation, for example introducing photothermal mechanisms[35], are in consideration in future study to further enhance the ultrasound trigger therapeutic effect of the PFP-TNDs.
4. Conclusions In this study, we formulated an ultrasound triggered, perfluorocarbon and drug loaded polymeric nanodroplet for ultrasound imaging and tumor treatment. The optimized nanodroplet presented good serum stability and biocompatibility. The PFP stabilized in the
SC RI PT
nanodroplets indulged ultrasound sensitivity, enhanced contrast and sustained imaging. Furthermore, the nanodroplets achieved decent cellular uptake of drug and therapeutic effect with the integration of ultrasound cavitation, functional modification of drug and augmenting coating of PGA-g-mPEG. The imaging and therapeutic capability of the ultrasound triggered nanodroplet system can be further improved by applying addition
U
strategies of promoting phase-change efficiency and ultrasound cavitation in future study.
N
5. Acknowledgement
A
This work was supported by the National Natural Science Foundation of China
M
(51773231, 81572726), the National Natural Science Foundation of Guangdong Province (2016A030313315, 2014A030312018), the Tip-top Scientific and Technical Innovative
D
Youth Talents of Guangdong special support program (2014TQ01R651), and the Project of Key Laboratory of Sensing Technology and Biomedical Instruments of Guangdong
A
CC
EP
TE
Province (2011A060901013).
References
A
CC
EP
TE
D
M
A
N
U
SC RI PT
[1] V.P. Torchilin, Nature reviews Drug discovery, 13 (2014) 813. [2] H. Xia, Y. Zhao, R. Tong (Eds.), Therapeutic Ultrasound; Springer, 2016, Chapter 21. [3] S.R. Sirsi, M.A. Borden, Advanced drug delivery reviews, 72 (2014) 3. [4] N. Rapoport, D.A. Christensen, A.M. Kennedy, K.H. Nam, Ultrasound in medicine & biology, 36 (2010) 419. [5] K.W. Ferrara, Advanced drug delivery reviews, 60 (2008) 1097. [6] R.J. Price, D.M. Skyba, S. Kaul, T.C. Skalak, Circulation, 98 (1998) 1264. [7] Q.L. Lu, H.D. Liang, T. Partridge, M.J. Blomley, Gene therapy, 10 (2003) 396. [8] A. Kheirolomoom, P.A. Dayton, A.F. Lum, E. Little, E.E. Paoli, H. Zheng, K.W. Ferrara, Journal of controlled release, 118 (2007) 275. [9] P. Li, Y. Zheng, H. Ran, J. Tan, Y. Lin, Q. Zhang, J. Ren, Z. Wang, Journal of controlled release, 162 (2012) 349. [10] C. Niu, Z. Wang, G. Lu, T.M. Krupka, Y. Sun, Y. You, W. Song, H. Ran, P. Li, Y. Zheng, Biomaterials, 34 (2013) 2307. [11] K. Ferrara, R. Pollard, M. Borden, Annual review of biomedical engineering, 9 (2007) 415. [12] T.M. Krupka, L. Solorio, R.E. Wilson, H. Wu, N. Azar, A.A. Exner, Molecular pharmaceutics, 7 (2010) 49. [13] S. Ibsen, M. Benchimol, D. Simberg, C. Schutt, J. Steiner, S. Esener, Journal of controlled release, 155 (2011) 358. [14] H. Wu, N.G. Rognin, T.M. Krupka, L. Solorio, H. Yoshiara, G. Guenette, C. Sanders, N. Kamiyama, A.A. Exner, Ultrasound in medicine & biology, 39 (2013) 2137. [15] T. Yin, P. Wang, J. Li, R. Zheng, B. Zheng, D. Cheng, R. Li, J. Lai, X. Shuai, Biomaterials, 34 (2013) 4532. [16] P.S. Sheeran, P.A. Dayton, Curr Pharm Design, 18 (2012) 2152. [17] N. Rapoport, Wiley Interdisciplinary Reviews: Nanomedicine and Nanobiotechnology, 4 (2012) 492. [18] N.Y. Rapoport, A.M. Kennedy, J.E. Shea, C.L. Scaife, K.-H. Nam, Journal of Controlled Release, 138 (2009) 268. [19] Z. Gao, A.M. Kennedy, D.A. Christensen, N.Y. Rapoport, Ultrasonics, 48 (2008) 260. [20] N. Rapoport, Z. Gao, A. Kennedy, Journal of the National Cancer Institute, 99 (2007) 1095. [21] D. Gao, M. Xu, Z. Cao, J. Gao, Y. Chen, Y. Li, Z. Yang, X. Xie, Q. Jiang, W. Wang, J. Liu, ACS applied materials & interfaces, 7 (2015) 13524. [22] X. Guan, Y. Li, Z. Jiao, J. Chen, Z. Guo, H. Tian, X. Chen, Acta biomaterialia, 9 (2013) 7672. [23] S.A. Senevirathne, K.E. Washington, M.C. Biewer, M.C. Stefan, Journal of Materials Chemistry B, 4 (2016) 360. [24] E. Tahover, Y.P. Patil, A.A. Gabizon, Anti-cancer drugs, 26 (2015) 241. [25] H.S. Yoo, E.A. Lee, T.G. Park, Journal of Controlled Release, 82 (2002) 17. [26] Y. Bae, N. Nishiyama, S. Fukushima, H. Koyama, M. Yasuhiro, K. Kataoka, Bioconjugate chemistry, 16 (2005) 122.
A
CC
EP
TE
D
M
A
N
U
SC RI PT
[27] Z. Yang, Z. Jiang, Z. Cao, C. Zhang, D. Gao, X. Luo, X. Zhang, H. Luo, Q. Jiang, J. Liu, Nanoscale, 6 (2014) 10193. [28] Z.X. Liao, S.F. Peng, Y.C. Ho, F.L. Mi, B. Maiti, H.W. Sung, Biomaterials, 33 (2012) 3306. [29] T. Akagi, M. Higashi, T. Kaneko, T. Kida, M. Akashi, Macromolecular bioscience, 5 (2005) 598. [30] J.-C. Zhang, J.-X. Ding, C.-S. Xiao, C.-L. He, X.-L. Zhuang, Y.-N. Yang, X.-S. Chen, Chem. J. Chin. Univ., Chin, 33 (2012) 2809. [31] C. Du, D. Deng, L. Shan, S. Wan, J. Cao, J. Tian, S. Achilefu, Y. Gu, Biomaterials, 34 (2013) 3087. [32] D. Gao, J. Gao, M. Xu, Z. Cao, L. Zhou, Y. Li, X. Xie, Q. Jiang, W. Wang, J. Liu, Molecular pharmaceutics, 14 (2017) 984. [33] K. Knop, R. Hoogenboom, D. Fischer, U.S. Schubert, Angewandte chemie international edition, 49 (2010) 6288. [34] J. Qin, T.-Y. Wang, J.K. Willmann (Eds.), Therapeutic Ultrasound; Springer, 2016, p. 263. [35] W. Fan, B. Yung, P. Huang, X. Chen, Chemical reviews, 117 (2017) 13566.
SC RI PT U N A
A
CC
EP
TE
D
M
Fig.1. TEM images and particle size distribution of PFP-BNDs (A, C) and PFP-TNDs (B, D).
SC RI PT U N A M
Fig.2. (A) Changes of particle average size by time for PFP-NDs, PFP-BNDs, PFP-TNDs incubated
A
CC
EP
TE
D
in PBS (0.01 M, pH=7.4) containing 10% FBS. (B) Cell viability of HepG2, CT-26 and HUVEC cells incubated with PFP-NDs/PGA-g-mPEG of different polymer concentration for 24 h. (C) Microscope images of erythrocyte incubated with PFP-BNDs, PFP-TNDs, PBS (negative control), Triton X-100 (positive control). (D) Hemolysis degree relative to polymer concentration of PFP-BNDs and PFPTNDs, as well as hemolysis photos of nanodroplets with a polymer concentration of 80 μg/mL.
SC RI PT
Fig.3. Contrast images of diluted PFP-TNDs at 25 ℃ and 37 ℃ compared with degassed water (A)
A
CC
EP
TE
D
M
A
N
U
and a graph of quantified gray-scale intensities (B).
SC RI PT U N A M D TE
Fig.4. Contrast images of diluted PFP-TNDs treated with ultrasound (3.5 MHz) in different MI (A) at
A
CC
EP
37 ̊C and graph of quantified gray-scale intensities (B). Ultrasound images of diluted PFP-TNDs treated with ultrasound (MI=0.08) in different frequencies (C) and graph of quantified gray-scale intensities (D). Ultrasound images of PFP-TNDs treated with ultrasound (3.5 MHz, MI=0.08) at different time points (E).
SC RI PT U N A M
Fig.5. Cell uptake profile of CT-26 cells and HepG2 cells at 2 h (A, B) and 4 h (C, D) after ultrasound
A
CC
EP
TE
D
irradiation. The mean fluorescence intensity (MFI) of each group was calculated from the uptake profile.
SC RI PT U N A M
A
CC
EP
TE
D
Fig. 6 Cell viability of CT-26 cells (A, B) and HepG2 cells (C, D) incubated with PFP-TNDs, free CAD, free DOX with and without ultrasound irradiation. Groups are divided as (A, C) groups with medium refreshed at 6 h following ultrasound irradiation, and (B, D) groups with medium unchanged for 24 h (*p≤0.05, **p≤0.01, results were taken from three measurements).
SC RI PT U N A M Fabrication of PFP-TNDs and the mechanisms of ultrasound-triggered imaging and
D
Scheme 1.
A
CC
EP
TE
delivery of CAD.
Table 1 Size, PDI and zeta potential of PFP-NDs. optimized PFP-BNDs and PFP-TNDs. Sample PFP-NDs
PDI
Zeta Potential (mV)
340±23
0.133
72.4±8.6
a
351±10
0.166
45.9±2.4
b
432±19
0.164
18.1±3.8
PFP-BNDs PFP-TNDs
The optimized mass ratio of C9F17-PAsp(DET) to CAD of the PFP-BNDs is 5/1.
SC RI PT
a.
Size (nm)
A
CC
EP
TE
D
M
A
N
U
b. The optimized C/N ratio of the PFP-TNDs is 0.2.