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. Author manuscript; available in PMC: 2020 Sep 1.
Published in final edited form as: J Pharm Sci. 2019 May 13;108(9):3091–3098. doi: 10.1016/j.xphs.2019.05.004

Enhancing Tumor Drug Distribution with Ultrasound-Triggered Nanobubbles

Pinunta Nittayacharn 1,*, Hai-Xia Yuan 2,3,*, Christopher Hernandez 1, Peter Bielecki 1, Haoyan Zhou 4, Agata A Exne 1,3
PMCID: PMC6708467  NIHMSID: NIHMS1529247  PMID: 31095958

Abstract

Issues with limited intratumoral drug penetration and heterogeneous drug distribution continue to impede the therapeutic efficacy of nanomedicine-based delivery systems. Ultrasound-enhanced drug delivery has emerged as one effective means of overcoming these challenges. Acoustic cavitation in the presence of nanoparticles has shown to increase the cellular uptake and distribution of chemotherapeutic agents in vivo. In this study, we investigated the potential of a drug-loaded echogenic nanoscale bubbles in combination with low frequency (3 MHz), high energy (2 W/cm2) ultrasound for antitumor therapy. The doxorubicin-loaded nanobubbles (Dox-NBs) stabilized with an interpenetrating polymer mesh were 171.5±20.9 nm in diameter. When used in combination with therapeutic ultrasound (TUS), Dox-NBs combined with free drug showed significantly higher (*p< 0.05) intracellular uptake and therapeutic efficacy compared to free drug. When injected intravenously in vivo, Dox-NBs+TUS showed significantly higher (*p<0.05) accumulation and better distribution of Dox in tumors when compared to free drug. This strategy provides an effective and simple method to increase the local dose and distribution of otherwise systemically toxic chemotherapeutic agents for cancer therapies.

Keywords: cavitation, sonoporation, ultrasound, nanobubbles, doxorubicin, drug delivery

Introduction

Despite the availability of effective chemotherapy drugs, limitations in drug transport to solid tumors still accounts for many failures in clinical applications1,2. Lipid-based particles have emerged as widely researched vehicles for the intravascular delivery of insoluble chemotherapeutic drugs and genes in cancer therapy3. A majority of these formulations overcome initial protection mechanisms, such as kidney filtration and the reticuloendothelial system (RES), designed to clear toxins from the circulation in order to achieve a higher tumor drug concentration. However, cytotoxic effects are reduced but not eliminated as drug delivery vehicles accumulate in off-site organs4. Further, these approaches have problems achieving adequate penetration of drugs away from the tumor vasculature, and some, like liposomes, do not release the drug easily once in the cell. To circumvent these problems, methods for triggered release of drug from particles using ultraviolet and near-infrared light, temperature, magnetic resonance imaging (MRI) and ultrasound (US) have been developed to deliver controlled local therapeutic doses from an external source5–9. Of these, US serves as a widely available, portable, inexpensive, and non-invasive method for contrast imaging and triggered release of particles in real time10. The drug delivery vehicles used in US-mediated drug delivery are US-visible microbubbles (1–10 μm in diameter and remain in blood pool) in combination with therapeutics11,12, liquid perfluorocarbon emulsions13,14 or US-sensitive drug loaded nanoparticles (such as liposomes or micelles)15,16,17,18 which are not directly visible on US at clinically-utilized frequencies19,20.

Microbubbles (MBs) provide a means by which low-energy acoustic waves can be focused at the vasculature for reversible endothelium permeability, local drug release and penetration21. The inertial cavitation of gas bubbles triggers payload release from the carrier20 while increasing vascular permeability through mild hyperthermia and disruption of cell-cell junctions and cell membranes22. Because MBs localize acoustic energy to the vasculature, they circumvent problems of thermal tissue damage associated with focused US (FUS) without compromising therapeutic efficacy23. However, effective tumor drug delivery using MBs continues to be challenging. The heterogeneous tumor environment contains regions of high and low perfusion, hypoxic domains, pockets of necrosis, mixed cell populations, sporadically permeable blood vessels and high interstitial pressure24,25. Yet, due to their larger size (1–10 μ m) MBs are unable to take advantage of passive targeting of the tumor via the enhanced permeability and retention effect (EPR), which requires a drug delivery vehicle diameter less than 400–800 nm26–28.

Our group has previously developed sub-micron echogenic ultrasound contrast agents, by the incorporation of the surfactant, Pluronic, into the lipid shell of perfluoropropane (C3F8) gas bubbles29–32 We have hypothesized that the incorporation of Pluronic leads to a reduction in surface tension of these nanobubbles (NBs), as seen in lipid monolayers33,34, thereby increasing their stability and potentially decreasing their resonant frequency. Though the EPR can be itself quite heterogeneous and variable in primary versus metastatic tumors and between tumors of different origins2,35, we have shown that these NBs can utilize the limited EPR available, and serve as a diagnostic contrast agent for solid tumors31,32. For this study, we have expanded the application of our previously developed diagnostic NBs, synthesized with an interpenetrating polymer mesh15,32, to include the chemotherapeutic drug doxorubicin (Dox) and serve as a potential theranostic agent. We have also used this method previously to solubilize porphyrin, which was unable to be solubilized in NBs without the NNDEA36. This polymer mesh has been shown to increase the stability and circulation time of NBs and can also increase Dox loading.

The Dox carrying NBs in combination with the sonoporating effects of their inertial cavitation, penetrate physiological and cell barriers that may limit typical nano-based device drug release to intended target sites. To our knowledge, there are no successful applications of inherently gas-core NBs to carry drug and image and deploy in real time with therapeutic US (TUS). To demonstrate that NBs can cause bioeffects and increase free Dox penetration in tumors, in this study, the effectiveness of Dox carrying NBs in synergy with free drug and TUS in improving tumor drug accumulation was assessed in a mouse model of LS-174T colorectal cancer.

Materials and Methods

Materials

The lipids DPPC (1,2-Dipalmitoyl-sn-Glycero-3-Phosphocholine), DPPA (1,2 Dipalmitoyl-sn-Glycero-3-Phosphate), DPPE (1,2-dipalmitoyl-sn-glycero-3-phosphoethanolamine) were obtained from Avanti Polar Lipids (Pelham, AL), and mPEG-DSPE (1,2-Distearoyl-phosphatidylethanol amine-methyl-poly ethylene glycol conjugate-2000) was obtained from Laysan Lipids (Arab, AL). N, N-diethyl acrylamide (NNDEA), 2-Hydroxy-4(2-hydroxyethoxy)-2-methylpropiophenone (Irgacure 2959), cell proliferation reagent WST-1, (N, Nbis(acryoyl) cystamine (BAC), and Doxorubicin (Dox) were purchased from Sigma Aldrich (Milwaukee, WI). Pluronic L10 was donated by BASF (Shreveport, LA). LS174-T human colorectal adenocarcinoma cells were purchased from ATCC (Manassas, VA). Minimum essential medium (MEM) with 10% fetal bovine serum and 1% penicillin-streptomycin, trypsin-EDTA were purchased from Invitrogen (Carlsbad, CA).

Formulation of Dox-loaded NBs

To formulate Dox-loaded NBs, Dox was encapsulated in our previous crosslinked NBs37. Briefly, the lipids DPPC, DPPA, DPPE, mPEG-DSPE were dissolved in chloroform in a 4:1:1:1 mass ratio. The solvent was then removed by evaporation and the lipids were hydrated in a solution containing 50 μ L of glycerol and 1 ml of PBS containing 0.06 wt. % Pluronic L10, 0.5 wt. % Irgacure 2959, and 0.2 wt. % Dox at 70°C for 30 min. Following the hydration step, NNDEA and BAC were added and dissolved into the lipid solution and the vial was resealed before air inside the vial was replaced with octafluoropropane (C3F8). Finally, the vial was shaken on a VialMix shaker (Bristol-Myers Squibb Medical Imaging, Inc., N. Billerica, MA) for 45s, and were irradiated at 254 nm using a UV lamp (Spectronics Co. Westbury, NY) for 30 min. The bubble vial was stored at 4°C for 1h to allow for separation of the larger bubbles. After that, the lower part solution (NB population) was drawn and used immediately. This fraction will be referred to as combination therapy (Dox-NBs) for the remainder of this paper as the final solution was shown to contain Dox loaded on the NBs, Dox encapsulated in non-buoyant particles (micelles/liposome), and residual free Dox.

Characterization of Dox-loaded NBs

The mean diameter and polydispersity of the Dox-NBs (n=3) were measured using dynamic light scattering (DLS) (90 Plus, Brookhaven Instruments Corp). Bubble size was measured by diluting a sample 1:1000 with PBS at pH 7.4. Measurements were performed at 25°C, with a laser wavelength of 660 nm at an angle of 90° and bubble size was reported as a number average. In addition, the size and concentration of buoyant and non-buoyant particles (particle/mL) was also measured using resonant mass measurement (RMM) (Archimedes, Malvern Instruments)38–41 with a nanosensor which provides measurement of particles ranging from 100 nm to 2 microns. Samples (n=3) were diluted 1:100 with PBS at pH 7.4 for measurement. For ultrasound imaging, Dox-NBs were diluted with PBS and imaged with a linear ultrasound probe (12 MHz transducer, MI:0.1) in an acrylamide gel phantom. The bubbles were imaged again immediately after application of TUS with an Omnisound® 3000 device (Accelerated Care Plus Corp., Reno, NV) at 3 MHz, 2 W/cm2 power density, and 20% duty cycle for 1 min. Dox-NBs were imaged under fluorescent microscopy (Zeiss Axio Observer Z1 motorized FL inverted microscope) using a 63x objective.

Determination of Dox loading content and encapsulation efficiency

Free Dox was removed from the Dox-NBs by centrifugation using centrifugal filter with a molecular weight cut-off of 50,000 Da (Vivaspin20, GE) at 4000 rpm for 50 min. To determine drug loading and encapsulation efficiency, lyophilized Dox loaded nanoparticles were weighed and dissolved in the mixture of methanol and PBS in 1:1, v/v, causing complete dissolution of the nanoparticles and release of the encapsulated Dox. The amount of the encapsulated Dox was determined by measuring the fluorescence at an excitation of 495 nm and an emission of 595 nm. Dox content was expressed as the drug loading content (DLC) which is weight ratio between loaded Dox and total weight of materials including (lipids, NNDEA, BAC, Glycerol), and encapsulation efficiency (%EE) as the percent of encapsulated Dox to initial feeding Dox. All loading measurements were performed in triplicate.

In vitro Dox release from the NBs

Dox-NBs were loaded with 2 mg Dox. The entire contents of the bubble vial were diluted with PBS, pH 7.4, to a total of 10 mL and transferred into 300 kDa Spectra/Por® dialysis tubing. The Dox-NBs were then placed in a beaker filled with 1L of PBS at 37°C to begin dialysis under continuous stirring. Samples were taken of the bath side solution at predetermined time points up to 2 hrs. For Dox-NBs exposed to TUS, Dox-NBs in dialysis were transferred to the 1L PBS beaker at T=0 min and immediately exposed to 1 min of TUS (4 cm2 surface probe, 3 MHz frequency, 2.0 W/cm2 power, and 20% duty cycle). Samples were read on a TECAN plate reader (Infinite M200, San Jose, CA) at an excitation of 495 nm and an emission of 595 nm to quantify the amount of Dox released with time (n=3).

In vitro therapeutic efficacy

LS-174T human colorectal adenocarcinoma cells were cultured in complete MEM medium (10% fetal bovine serum, 1% penicillin-streptomycin) and placed in a humidified atmosphere at 37°C and 5% CO2. Cells were passaged until they were 90% confluent, then detached with 0.25% trypsin-EDTA. Dox intracellular uptake was evaluated using fluorescence microscopy. Briefly, cells were plated on 12-well plates (2.5×104 cells/mL) one day before treatment. After 24 h of incubation, the surface medium was aspirated and the cells were treated under the following conditions: (a) MEM medium only (no treatment control); (b) 1 μg/mL Dox (Dox); (c) Dox-NBs; (d) 1 μg/mL Dox plus therapeutic ultrasound (Dox+TUS); (e) Dox-NBs plus therapeutic ultrasound (Dox-NBs+TUS); (f) mixed solution of Dox-NBs without any gas (Burst Dox-NBs); (g) mixed solution of Dox-NBs without any gas plus therapeutic ultrasound (Burst Dox-NBs+TUS). In the last two groups (f and g) Dox-NBs were exposed repeatedly to high intensity ultrasound (US, MI1.5) to destroy the bubble portion prior to addition into cells. The 1 μg/mL concentration of Dox was determined to be total amount of the drug in the injected bubble dose since encapsulated drug was not separated from free Dox in these experiments. The treatment solution was aspirated after 2 h and cells were washed with PBS. The attached cells were then fixed with 3% formaldehyde for 10 min and were imaged using a fluorescence microscope (Zeiss Axio Observer Z1). Average increase in Dox fluorescence intensity (%) for treatment group was measured using ImageJ (National Institutes of Health).

For in vitro cell toxicity studies, LS-174T cells were plated (2.5×104 cells/ well) in a flat bottom 96-well plate 24 h prior to treatment. Following 24 h, the media in each well was replaced with 400 μ L of the same treatment solutions as described above (group a-e). The wells were then sealed with Parafilm® (Fisher Scientific; Pittsburgh, PA), and inverted for 20 min to ensure close contact between cells and Dox-NBs. For groups with US treatment, cells were exposed to TUS at 3 MHz, 2 W/cm2 power density and 20% duty cycle for 1 min. After a 3hour incubation at 37°C, the treatment solution was removed, cells were washed and then cultured in complete media for 3 days. Following the three-day incubation, cell viability was measured using the WST-1 cell viability assay. Assay absorbance was measured using a TECAN plate reader at 450 nm.

Animal preparation and tumor model

All animals were handled according to a protocol approved by the Institutional Animal Care and Use Committee at Case Western Reserve University in accordance with all applicable protocols and guidelines in regards to animal use. Athymic nude mice (NCR nu/nu) were purchased from the Athymic Animal and Xenograft Core Facility of Case Western Reserve University. In all procedures, mice were anesthetized with 3% isoflurane with 1L/min oxygen. Then, 200 μ L LS-174T cells (1×106 cells/mL) suspended in incomplete MEM were injected subcutaneously into the flank of each mouse.

In vivo drug distribution analysis

After tumor inoculation, growth was measured every other day. When tumor diameter reached 8–10 mm, the mice were randomly divided into 4 groups (10 mice/group) as follows: no treatment, Dox+TUS, Dox-NBs and Dox-NBs+TUS. Each animal received only one treatment via intravenously tail vein injection (200 μ L injection volume) and the amount of Dox was held constant for all treatment groups at 184 nmol (5mg/kg). Bubble accumulation in tumors was monitored with a linear ultrasound probe (8 MHz transducer, MI: 0.08). After 1-minute post-intravenous injection, mice in Dox+TUS and Dox-NBs+TUS groups received ultrasound-mediated locoregional treatment at the tumor site via the same protocol as above (3 MHz at 2 W/cm2 power density and 20% duty cycle for 1 min). Animals were euthanized 3 hours post treatment, and tissues were collected for analysis.

To determine Dox accumulation in each tumor (n=5 mice/group), tissues were homogenized in water (10 mL/g) after washing in PBS. 200 μ L of the tissue homogenate was mixed with 100 μ L of PBS buffer, 200 μ L of water, and 1.5 mL of acidified isopropanol (0.75 N HCl). Solutions were mixed thoroughly, and incubated overnight at −20 °C. On the next day, the tubes were warmed to room temperature, vortexed, and centrifuged at 15,000g for 20 minutes42. The samples were plated into 96-well plate to measure the fluorescence intensity using a plate reader and used to calculate the Dox concentration remaining in tissue.

For histological imaging, tumors (n=5 mice/group) were excised and mounted in optimal cutting temperature compound (OCT, Sakura Finetek USA, Inc., Torrance, CA) and frozen at −80°C. The tissues were cut into 10 μ m slices using Leica CM1850 cryostat (Leica, Germany) and stained with DAPI using standard techniques. The tissue sections were imaged at 5x and 20x on the fluorescence inverted microscope (Zeiss Axio Observer Z1).

Statistical analysis

All data are presented as mean ± SD (standard deviation) unless otherwise noted. All statistical analyses were performed using Graphpad Prism. Statistical significance of differences between experimental groups was derived using a one-way ANOVA test with Tukey Test or a student T-test with Welch’s correction. P<0.05 was considered to be significant.

Results

Dox-NB characterization

Dox-NBs had an average diameter of 171.5 ± 20.9 nm with an average polydispersity of 0.15 ± 0.13 measured by DLS (n=3) (Figure. 2A). This diameter was similar to other crosslinked stabilized Pluronic-lipid bubbles tested in our previous studies43. Bubble concentration was measured using RMM, which works by observing the change in the resonant frequency of a vibrating micro-cantilever produced by the particle and measuring the buoyant mass of particles44. RMM results show that buoyant particle (Dox loaded-NBs) and non-buoyant particle (Dox-micelles/liposome, lipid aggregation) concentrations were present in solution at concentrations of 9.76 μ 109 and 2.44 μ 1010 particles/mL, respectively (Figure.2B). It is possible that not all of the non-buoyant particles are entirely free of gas, but are indeed denser than the carrier fluid. The microscopic images of large bubbles were provided to clearly show the co-localization of Dox signal in the bubble shell (Figure.2C). Separation of the NB population was not performed for this experiment, in order to accurately visualize loading in larger particles and as most of them would be below the light diffraction limit. Diagnostic US images of Dox-NBs before and after application of TUS reveals a complete loss of contrast (Figure.3A), suggesting that the high acoustic amplitude of the TUS was sufficient for inertial cavitation and destruction of the bubbles. When injected intravenously, Dox-NBs provided an enhancement in tumor contrast (Figure. 3B)

Figure 2.

Figure 2.

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Dox-NB Characterization. (A) Representative size distribution of Dox-NBs obtained by DLS measurement. (B) Size distribution of buoyant and non-buoyant particles measured by RMM. (C) Representative fluorescence (left) and bright field (middle) image of Dox-NBs White arrows show co-localization of Dox signal in the bubble shell.

Figure 3.

Figure 3.

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(A) Ultrasound image of Dox-NBs before and after application of TUS. (B) Representative US image showing Dox-NB contrast in tumor before and 1.5 min after injection.

Loading and release of Dox-NBs

Encapsulated Dox content in particles was determined using centrifuge filtration technique45–48. Dox loading content was 14.35 ± 2.24 μg of Dox per mg of materials including lipids and others reagents. The encapsulation efficiency was 22.8 ±1.73 %. The estimated total amount of Dox in bubbles was 129.16±0.08 μg which was calculated from the amount of Dox in the total number of buoyant particles (per mL)49. The release study of free Dox, Dox-NBs with/without TUS was performed in dialysis with a starting Dox concentration of 0.2 mg/mL. Due to the release of Dox from NBs with and without TUS within 60 min is similar, and a significant release (*, p <0.05) in both groups was observed only at the 90 and 120min, there is no clear effect of ultrasound on the drug release. The drug seems to be re-sequestered by a function of polymer mesh network into micelles as RMM result shows that we have predominately solid particles. However, as shown next, in cells and in vivo there is still quite a significant difference in drug uptake and distribution, indicating that ultrasound-generated effects and cavitation of the gas bubbles contribute more to these effects rather than just to triggering the drug release on demand.

Enhancement of Dox uptake and cytotoxicity by TUS in cells

In order to determine if the application TUS would affect the uptake of Dox, LS-174T cells were exposed with TUS in the presence of free Dox, Dox-NBs, or the mixed solution of Dox-NBs without gas and immediately imaged under fluorescent microscopy. The average increase in Dox fluorescence intensity of each treatment group is presented in Figure. 5A and confirms a significant enhancement (*p< 0.0001) of intracellular Dox delivery with the Dox-NBs+TUS in comparison to all other control groups (no treatment, free Dox, free Dox+TUS, and Dox-NB). We have previously demonstrated that exposure to only the TUS used in this study does not induce any cytotoxic effect on LS-174T cells and therefore was not tested here29. As shown in Figure. 5B, 72 h post treatment, the application of TUS decreased the cell viability of cells treated with free Dox from 30.6 ± 8.2% to 23.8 ± 4.78%. Cells treated with Dox-NBs+TUS had the lowest viability of all treatment groups (*p<0.05) with 15.9 ± 5.39%.

Figure 5.

Figure 5.

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Enhancement of Dox uptake into LS-174T cells after treatment with Dox-NBs and TUS. (A) Average increase in Dox fluorescence intensity for each treatment group. (B) Cell viability of LS-174T cells for different treatments normalized to the untreated control (n=6 ± STDEV).

In vivo tumor drug accumulation analysis

Figure. 6A shows the typical fluorescence images of sectioned tumors from the following treatment groups: no treatment control, free Dox +TUS, Dox-NBs, and Dox-NBs+TUS. Cell nuclei were stained with DAPI (blue). The average normalized fluorescence signal (Dox/DAPI) for all treatment groups is shown in Figure.6B and shows a significantly higher (***p< 0.05) accumulation of Dox inside tumors treated with Dox-NBs+TUS than all other treatment groups. There was no statistical difference in the relative signal intensity from free Dox+TUS and Dox-NBs, however both were statistically higher (p<0.05) than the no treatment control. Extraction of Dox from tissues 3 h after treatment (Figure. 6C) demonstrated a significantly higher (*p<0.05) dose of drug in tumors treated with Dox-NBs+TUS (1.16 ± 0.43 μg/g) compared to Dox-NBs (0.67 ± 0.23 μg/g) and free Dox+TUS (0.75 ± 0.17 μg/g).

Figure 6.

Figure 6.

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(A) Representative fluorescence images of showing the distribution of Dox in tumors following each treatment. (B) The average Dox signal in each image was normalized by its corresponding DAPI signal (n=5 ± STDEV). (C) Extracted mass of Dox in tumors 3 hours after treatment (n=5 ± STDEV). Scale bar represents 50 μm.

Discussion

Contrast agent-assisted ultrasonic cavitation has been investigated as a means to increase the delivery of therapeutic agents such as genes and small molecules11,19,20,50. However, the vast majority of these therapies involve the delivery of drug, either free or encapsulated in a nanoparticle, in the presence of MBs51–53. In the case of cancer therapy, MBs are limited to the blood pool because of their size, thereby limiting their therapeutic benefits to areas immediately adjacent to the tumor vasculature. The purpose of this study was to determine whether, when combined with ultrasound, a drug-loaded, acoustically responsive sub-micron bubble is capable of improving tumor drug accumulation and distribution following IV injection. In order to achieve this, we have expanded the application of our previously developed, polymer-stabilized NB formulation32 and incorporated the chemotherapeutic agent, Dox. This method has been used previously to solubilize porphyrin, which was unable to be solubilized in NBs without the NNDEA36.

DLS characterization points to a relatively uniform particle with a mean diameter <200 nm and a small sub-population of particles around 600 nm. Examining the formulation in more detail with RMM reveals a combination of buoyant (Dox-loaded NBs) and non-buoyant particles (Dox-micelles/liposome). Heterogeneous nanoparticle composition may be a possibility in many other drug-loaded formulations. However, because RMM measurement is not widely used, this disparity is typically not observed or explored. Theoretically it is possible that, given the relatively neutral buoyancy of lipid and polymer drug-loaded nanoparticles and higher bubble susceptibility to shear stress which can lead to gas dissipation from bubbles (including drug loaded MBs), that there may exist a mixture of gas filled and non-buoyant particles in most such formulations. However, the control experiments in this study suggest that the bubbles in the population play the major role in the observed beneficial effects.

While the release of Dox from NBs with and without TUS is similar, there is likely a more complex mechanism at play where TUS and loaded NBs closely interact with the cell to improve drug uptake. The lack of increased drug release from the constructs upon intonation is somewhat surprising. It is possible, as others have shown, that the release of Dox from the particles is rapid, and the drug is re-captured into the lipid vesicles much more rapidly than our sampling can show. This has been demonstrated in the prior work54–56. A great deal of earlier research has shown that US combined with bubble cavitation events can introduce defects into the cell membrane and can drive drug into cells. However, this has not been shown in the case of the lipid-polyacrylamide formulation of NBs used in this work37. It is thus possible that the bubbles combined with drug have a significant effect on drug distribution in tumors and uptake into cells without having a drastic effect on immediate drug release. It is also possible that once the gas is removed from the bubbles following US application, the bubbles turn into lipid-polyacrylamide micelles which sequester the drug, are taken up into tumor cells and remain there for an extended time. We have shown that the particle structure remains intact with these constructs in prior work57. An in-depth investigation of this complex mechanism will be carried in the future work.

In cell experiments, the combination therapy of Dox-NBs in the presence of low frequency, high intensity ultrasound lead to the improved in intracellular uptake of Dox in comparison to all control groups (no treatment, free Dox, free Dox+TUS, Dox-NBs, Burst Dox-NBs, and Burst Dox-NBs+TUS) in vitro. The burst Dox-NBs with and without TUS were carried out in order to support our hypothesis that Dox carrying NBs in synergy with US are the responsible for the observed effects in vitro and in vivo. We found that the mixed solution of Dox-NBs without any gas (Burst Dox-NBs) with and without US showed lower fluorescence signal as shown in Figure 5A, compared to the solution with intact NBs. An increase in Dox fluorescence signal could be from the acute toxicity from cavitation after expose bubbles to US which causes a mechanical damage to the cell58. However, since the goal of this work was to see if the combination of NBs and US could increase penetration of the drug in cells and, more importantly, in tumors, we will examine the mechanism in future work. Moreover, a significant reduction in LS-174T cell viability was found when cells were treated with Dox-NBs+TUS compared to all other treatment groups (Figure 5B). These provide evidence that while the sonication alone has some benefit in improving cell uptake from any nanoparticle, it is the presence of the gas-filled, drug loaded NBs that contributes a much more significant benefit.

In vivo, a significant increase in Dox accumulation in tumors was seen via two methods. First significantly increased Dox signal was observed in histological slices of mice treated with Dox-NBs+TUS compared to either free Dox plus US or Dox-NBs which did not receive US. Second, extraction of Dox from excised tumor tissue showed significant increases in tumor drug accumulation in the Dox-NBs+TUS compared to all other groups. We hypothesize that the enhanced local accumulation and cellular uptake of Dox are due to both the passive accumulation of drug-loaded NBs, as well as the membrane permeating effects of cavitation combined with free Dox in the cell area. Overall, these results support the use of our combination therapy +TUS for effective treatment to limit potential off-site cytotoxicity in healthy tissue.

This study has several limitations. First, the mechanism of strong activity from stabilized gas NBs at 12 MHz remains unclear, although reduction of surface tension by Pluronic33,34, buckling of the shell in response to the ultrasound field59–61, and a contribution from larger sub-micron bubbles which are not reflected in the mean hydrodynamic diameter measurements from DLS could all be contributing factors. A more in depth examination is required to fully understand the underlying principles; these studies are ongoing. Another limitation is the use of a mixed formulation. Although typically RMM measurements were not performed in related studies, the free drug was removed from the nano-sized drug carriers by centrifugation prior attaching to the surface of the bubble in some studies62,63 but not in others which did not purify their drug-loaded particles66,67 However, unlike other non-buoyant nanoparticles such as micelles, bubbles are fragile and can be easily destroyed by shear stress and by rapid gas diffusion over time. Thus separation of the Dox-loaded NBs from free drug solution by standard centrifugation or filtration techniques is challenging. Accordingly, the NBs were isolated from the larger MBs using their difference inherent buoyancy by placing the NB vial up side-down in the fridge resulting in an average size of 171.5±20.9 nm in diameter. This method allowed more buoyant of larger MBs rise up and separate from submicron bubbles and other non-buoyant particles including drug loaded-micelles/liposome and free drug. However, it did not enable removal of free drug. Affinity column separation of free drug, while ideal, also resulted in significantly reduced bubble yield (not shown). Thus to maintain bubble integrity, we opted to inject in vivo a formulation that is a combination of free and encapsulated drug in micelles and loaded on NBs. With more robust nanobubble formulations able to better withstand shear stress64,65, the separation of free drug will be feasible and likely to reduce off target effects. Finally, the technique was applied in a flank model of colorectal cancer in immunodeficient mice. While this is appropriate for elucidating relative difference between treatment groups, the application of this technique to tumors located either in the colorectal region or metastasized to the liver and the effect of the immune system on therapeutic outcomes will require additional testing.

Conclusions

We demonstrated that the combination therapy of Dox-loaded NBs, Dox bearing non-buoyant particles and residual free drug in the presence of therapeutic ultrasound provide significant improvement in intracellular uptake and tumor drug accumulation. Due to the co-localization of drug and gas within the NB core, drug release from Dox-NBs is likely governed by similar cavitation principles as MBs but with increased tumor penetration associated with nanoparticles. Future studies include evaluating the cytotoxicity of polymer-stabilized NBs, optimizing the Dox-NB formulation for improved therapeutic efficacy and stability and assessing the benefits of purified NBs in this application.

Figure 4.

Figure 4.

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Drug release profiles of Dox-loaded formulations with and without therapeutic ultrasound.

Acknowledgements

This work was supported by the National Institute of Biomedical Imaging and Bioengineering [R01EB016960]; the National Cancer Institute of the National Institutes of Health [F31CA200373]; the Department of Defense [W81XWH-16-1-0372]. Pinunta Nittayacharn is supported by the Royal Thai Government, Thailand. Dr. Hai-Xia Yuan was supported by Shanghai Municipal Commission of Health and Family Planning, China [201640263]. Views and opinions of, and endorsements by the author(s) do not reflect those of the National Institutes of Health or the Department of Defense. Illustration credits (Figure. 1) to Tiffany Yang, tyangdesign.com.

Figure 1.

Figure 1.

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Schematic diagram of Dox-loaded polymer/lipid NBs with interpenetrating crosslinking biodegradable polymer N,N-diethyl acrylamide (NNDEA) and N,N-bis(acryoyl) cystamine (BAC) crosslinking network.

Footnotes

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References

  • 1.Sriraman SK, Aryasomayajula B, Torchilin VP 2014. Barriers to drug delivery in solid tumors. Tissue Barriers 2:e29528. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 2.Prabhakar U, Maeda H, Jain RK, Sevick-Muraca EM, Zamboni W, Farokhzad OC, Barry ST, Gabizon A, Grodzinski P, Blakey DC 2013. Challenges and key considerations of the enhanced permeability and retention (EPR) effect for nanomedicine drug delivery in oncology. Cancer research 73(8):2412–2417. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 3.Jin S, Ye KM 2007. Nanoparticle-mediated drug delivery and gene therapy. Biotechnol Progr 23(1):32–41. [DOI] [PubMed] [Google Scholar]
  • 4.Svenson S 2012. Clinical translation of nanomedicines. Curr Opin Solid St M 16(6):287–294. [Google Scholar]
  • 5.Fomina N, McFearin C, Sermsakdi M, Edigin O, Almutairi A 2010. UV and Near-IR Triggered Release from Polymeric Nanoparticles. J Am Chem Soc 132(28):9540–9542. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 6.Derfus AM, von Maltzahn G, Harris TJ, Duza T, Vecchio KS, Ruoslahti E, Bhatia SN 2007. Remotely triggered release from magnetic nanoparticles. Adv Mater 19(22):3932–+. [Google Scholar]
  • 7.Manzoor AA, Lindner LH, Landon CD, Park JY, Simnick AJ, Dreher MR, Das S, Hanna G, Park W, Chilkoti A, Koning GA, ten Hagen TLM, Needham D, Dewhirst MW 2012. Overcoming Limitations in Nanoparticle Drug Delivery: Triggered, Intravascular Release to Improve Drug Penetration into Tumors. Cancer Research 72(21):5566–5575. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 8.Husseini GA, Pitt WG 2008. Micelles and nanoparticles for ultrasonic drug and gene delivery. Adv Drug Deliver Rev 60(10):1137–1152. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 9.Chilkoti A, Dreher MR, Meyer DE, Raucher D 2002. Targeted drug delivery by thermally responsive polymers. Adv Drug Deliver Rev 54(5):613–630. [DOI] [PubMed] [Google Scholar]
  • 10.Bohmer MR, Klibanov AL, Tiemann K, Hall CS, Gruell H, Steinbach OC 2009. Ultrasound triggered image-guided drug delivery. Eur J Radiol 70(2):242–253. [DOI] [PubMed] [Google Scholar]
  • 11.Rychak JJ, Klibanov AL 2014. Nucleic acid delivery with microbubbles and ultrasound. Adv Drug Deliver Rev 72:82–93. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 12.McDannold N, Arvanitis CD, Vykhodtseva N, Livingstone MS 2012. Temporary disruption of the blood–brain barrier by use of ultrasound and microbubbles: safety and efficacy evaluation in rhesus macaques. Cancer research 72(14):3652–3663. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 13.Sheeran PS, Dayton PA 2012. Phase-change contrast agents for imaging and therapy. Current pharmaceutical design 18(15):2152–2165. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 14.Sheeran PS, Streeter JE, Mullin LB, Matsunaga TO, Dayton PA 2013. Toward ultrasound molecular imaging with phase-change contrast agents: an in vitro proof of principle. Ultrasound in medicine & biology 39(5):893–902. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 15.Husseini GA, Christensen DA, Rapoport NY, Pitt WG 2002. Ultrasonic release of doxorubicin from Pluronic P105 micelles stabilized with an interpenetrating network of N,N-diethylacrylamide. Journal of controlled release : official journal of the Controlled Release Society 83(2):303–305. [DOI] [PubMed] [Google Scholar]
  • 16.Pruitt JD, Pitt WG 2002. Sequestration and ultrasound-induced release of doxorubicin from stabilized Pluronic P105 micelles. Drug delivery 9(4):253–258. [DOI] [PubMed] [Google Scholar]
  • 17.Kim H, Britton GL, Peng T, Holland CK, McPherson DD, Huang S-L 2014. Nitric oxide-loaded echogenic liposomes for treatment of vasospasm following subarachnoid hemorrhage. International journal of nanomedicine 9:155. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 18.Kheirolomoom A, Lai C-Y, Even LM, Ingham ES, Fite BZ, Watson KD, Ferrara KW 2013. Achieving complete response to locoregional disease without toxicity using temperature-sensitive liposomes and ultrasound-mediated hyperthermia. Cancer Research 73(8 Supplement):2155–2155. [Google Scholar]
  • 19.Klibanov AL 2006. Microbubble contrast agents: targeted ultrasound imaging and ultrasound-assisted drug-delivery applications. Investigative radiology 41(3):354–362. [DOI] [PubMed] [Google Scholar]
  • 20.Sirsi SR, Borden MA 2014. State-of-the-art materials for ultrasound-triggered drug delivery. Adv Drug Deliver Rev 72(0):3–14. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 21.Martin KH, Dayton PA 2013. Current status and prospects for microbubbles in ultrasound theranostics. Wires Nanomed Nanobi 5(4):329–345. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 22.Pitt WG, Husseini GA, Staples BJ 2004. Ultrasonic drug delivery--a general review. Expert Opin Drug Deliv 1(1):37–56. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 23.Fan C-H, Yeh C-K 2014. Microbubble-enhanced Focused Ultrasound-induced Blood–brain Barrier Opening for Local and Transient Drug Delivery in Central Nervous System Disease. Journal of Medical Ultrasound 22(4):183–193. [Google Scholar]
  • 24.Maeda H 2015. Toward a full understanding of the EPR effect in primary and metastatic tumors as well as issues related to its heterogeneity. Adv Drug Deliver Rev 91:3–6. [DOI] [PubMed] [Google Scholar]
  • 25.Prabhakar U, Maeda H, Jain RK, Sevick-Muraca EM, Zamboni W, Farokhzad OC, Barry ST, Gabizon A, Grodzinski P, Blakey DC 2013. Challenges and Key Considerations of the Enhanced Permeability and Retention Effect for Nanomedicine Drug Delivery in Oncology. Cancer Research 73(8):2412–2417. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 26.Greish K 2010. Enhanced permeability and retention (EPR) effect for anticancer nanomedicine drug targeting. Methods in molecular biology 624:25–37. [DOI] [PubMed] [Google Scholar]
  • 27.Maeda H 2001. The enhanced permeability and retention (EPR) effect in tumor vasculature: the key role of tumor-selective macromolecular drug targeting. Advances in enzyme regulation 41:189–207. [DOI] [PubMed] [Google Scholar]
  • 28.Tanaka T, Shiramoto S, Miyashita M, Fujishima Y, Kaneo Y 2004. Tumor targeting based on the effect of enhanced permeability and retention (EPR) and the mechanism of receptor-mediated endocytosis (RME). International journal of pharmaceutics 277(1–2):39–61. [DOI] [PubMed] [Google Scholar]
  • 29.Perera RH, Solorio L, Wu H, Gangolli M, Silverman E, Hernandez C, Peiris PM, Broome AM, Exner AA 2014. Nanobubble ultrasound contrast agents for enhanced delivery of thermal sensitizer to tumors undergoing radiofrequency ablation. Pharm Res 31(6):1407–1417. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 30.Wu H, Rognin NG, Krupka TM, Solorio L, Yoshiara H, Guenette G, Sanders C, Kamiyama N, Exner AA 2013. Acoustic characterization and pharmacokinetic analyses of new nanobubble ultrasound contrast agents. Ultrasound Med Biol 39(11):2137–2146. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 31.Krupka TM, Solorio L, Wilson RE, Wu HP, Azar N, Exner AA 2010. Formulation and Characterization of Echogenic Lipid-Pluronic Nanobubbles. Mol Pharmaceut 7(1):49–59. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 32.Perera RH, Wu H, Peiris P, Hernandez C, Burke A, Zhang H, Exner AA 2017. Improving performance of nanoscale ultrasound contrast agents using N,N-diethylacrylamide stabilization. Nanomedicine: Nanotechnology, Biology and Medicine 13(1):59–67. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 33.Maskarinec SA, Lee KYC 2003. Comparative study of poloxamer insertion into lipid monolayers. Langmuir 19(5):1809–1815. [Google Scholar]
  • 34.Hadicke A, Blume A 2013. Interactions of Pluronic block copolymers with lipid monolayers studied by epi-fluorescence microscopy and by adsorption experiments. J Colloid Interf Sci 407:327–338. [DOI] [PubMed] [Google Scholar]
  • 35.Maeda H Toward a full understanding of the EPR effect in primary and metastatic tumors as well as issues related to its heterogeneity. Adv Drug Deliver Rev (0). [DOI] [PubMed] [Google Scholar]
  • 36.Bosca F, Bielecki PA, Exner AA, Barge A 2018. Porphyrin-Loaded Pluronic Nanobubbles: A New US-Activated Agent for Future Theranostic Applications. Bioconjug Chem 29(2):234–240. [DOI] [PubMed] [Google Scholar]
  • 37.Perera RH, Wu HP, Peiris P, Hernandez C, Burke A, Zhang H, Exner AA 2017. Improving performance of nanoscale ultrasound contrast agents using N,N-diethylacrylamide stabilization. Nanomed-Nanotechnol 13(1):59–67. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 38.Godin M, Bryan AK, Burg TP, Babcock K, Manalis SR 2007. Measuring the mass, density, and size of particles and cells using a suspended microchannel resonator. Appl Phys Lett 91(12). [Google Scholar]
  • 39.Burg TP, Godin M, Knudsen SM, Shen W, Carlson G, Foster JS, Babcock K, Manalis SR 2007. Weighing of biomolecules, single cells and single nanoparticles in fluid. Nature 446(7139):1066–1069. [DOI] [PubMed] [Google Scholar]
  • 40.Olcum S, Cermak N, Wasserman SC, Christine KS, Atsumi H, Payer KR, Shen WJ, Lee JC, Belcher AM, Bhatia SN, Manalis SR 2014. Weighing nanoparticles in solution at the attogram scale. P Natl Acad Sci USA 111(4):1310–1315. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 41.Lee J, Shen WJ, Payer K, Burg TP, Manalis SR 2010. Toward Attogram Mass Measurements in Solution with Suspended Nanochannel Resonators. Nano Lett 10(7):2537–2542. [DOI] [PubMed] [Google Scholar]
  • 42.Laginha KM, Verwoert S, Charrois GJ, Allen TM 2005. Determination of doxorubicin levels in whole tumor and tumor nuclei in murine breast cancer tumors. Clinical cancer research : an official journal of the American Association for Cancer Research 11(19 Pt 1):6944–6949. [DOI] [PubMed] [Google Scholar]
  • 43.Perera RH, Wu H, Peiris P, Hernandez C, Burke A, Zhang H, Exner AA 2017. Improving performance of nanoscale ultrasound contrast agents using N,N-diethylacrylamide stabilization. Nanomedicine 13(1):59–67. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 44.Nieves LM, Hernandez C, Nittayacharn P, Hadley J, Coyne R, Mangadlao J, Advincula R, Exner A 2017. Effect of the Surfactant Pluronic on the Stability of Lipid-Stabilized Perfluorocarbon Nanobubbles. Ieee Int Ultra Sym. [Google Scholar]
  • 45.Nasongkla N, Shuai X, Ai H, Weinberg BD, Pink J, Boothman DA, Gao JM 2004. cRGD-functionalized polymer micelles for targeted doxorubicin delivery. Angew Chem Int Edit 43(46):6323–6327. [DOI] [PubMed] [Google Scholar]
  • 46.Puntawee S, Theerasilp M, Reabroi S, Saeeng R, Piyachaturawat P, Chairoungdua A, Nasongkla N 2016. Solubility enhancement and in vitro evaluation of PEG-b-PLA micelles as nanocarrier of semi-synthetic andrographolide analogue for cholangiocarcinoma chemotherapy. Pharm Dev Technol 21(4):437–444. [DOI] [PubMed] [Google Scholar]
  • 47.Tambunlertchai S, Srisang S, Nasongkla N 2017. Development of antimicrobial coating by later-by-layer dip coating of chlorhexidine-loaded micelles. J Mater Sci-Mater M 28(6). [DOI] [PubMed] [Google Scholar]
  • 48.Eawsakul K, Chinavinijkul P, Saeeng R, Chairoungdua A, Tuchinda P, Nasongkla N 2017. Preparation and Characterizations of RSPP050-Loaded Polymeric Micelles Using Poly(ethylene glycol)-b-Poly(epsilon-caprolactone) and Poly(ethylene glycol)-b-Poly(D,L-lactide). Chem Pharm Bull 65(6):530–537. [DOI] [PubMed] [Google Scholar]
  • 49.Nittayacharn P, de Leon A, Abenojar EC, Exner AA 2018. The Effect of Lipid Solubilization on the Performance of Doxorubicin-loaded Nanobubbles. Ieee Int Ultra Sym. [Google Scholar]
  • 50.Couture O, Foley J, Kassell NF, Larrat B, Aubry J-F 2014. Review of ultrasound mediated drug delivery for cancer treatment: updates from pre-clinical studies. Translational Cancer Research 3(5):494–511. [Google Scholar]
  • 51.Escoffre JM, Piron J, Novell A, Bouakaz A 2011. Doxorubicin Delivery into Tumor Cells with Ultrasound and Microbubbles. Mol Pharmaceut 8(3):799–806. [DOI] [PubMed] [Google Scholar]
  • 52.Tinkov S, Coester C, Serba S, Geis NA, Katus HA, Winter G, Bekeredjian R 2010. New doxorubicin-loaded phospholipid microbubbles for targeted tumor therapy: In-vivo characterization. J Control Release 148(3):368–372. [DOI] [PubMed] [Google Scholar]
  • 53.Gao Z, Kennedy AM, Christensen DA, Rapoport NY 2008. Drug-loaded nano/microbubbles for combining ultrasonography and targeted chemotherapy. Ultrasonics 48(4):260–270. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 54.Husseini GA, de la Rosa MAD, Gabuji T, Zeng Y, Christensen DA, Pitt WG 2007. Release of doxorubicin from unstabilized and stabilized micelles under the action of ultrasound. J Nanosci Nanotechno 7(3):1028–1033. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 55.Marin A, Sun H, Husseini GA, Pitt WG, Christensen DA, Rapoport NY 2002. Drug delivery in pluronic micelles: effect of high-frequency ultrasound on drug release from micelles and intracellular uptake. Journal of Controlled Release 84(1–2):39–47. [DOI] [PubMed] [Google Scholar]
  • 56.Husseini GA, Myrup GD, Pitt WG, Christensen DA, Rapoport NAY 2000. Factors affecting acoustically triggered release of drugs from polymeric micelles. Journal of Controlled Release 69(1):43–52. [DOI] [PubMed] [Google Scholar]
  • 57.Hernandez C, Gulati S, Fioravanti G, Stewart PL, Exner AA 2017. Cryo-EM Visualization of Lipid and Polymer-Stabilized Perfluorocarbon Gas Nanobubbles - A Step Towards Nanobubble Mediated Drug Delivery. Sci Rep-Uk 7. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 58.Afadzi M, Strand SP, Nilssen EA, Masoy SE, Johansen TF, Hansen R, Angelsen BA, Davies CD 2013. Mechanisms of the Ultrasound-Mediated Intracellular Delivery of Liposomes and Dextrans. Ieee T Ultrason Ferr 60(1):21–33. [DOI] [PubMed] [Google Scholar]
  • 59.Kooiman K, Vos HJ, Versluis M, de Jong N 2014. Acoustic behavior of microbubbles and implications for drug delivery. Adv Drug Deliver Rev 72:28–48. [DOI] [PubMed] [Google Scholar]
  • 60.Marmottant P, Bouakaz A, de Jong N, Quilliet C 2011. Buckling resistance of solid shell bubbles under ultrasound. J Acoust Soc Am 129(3):1231–1239. [DOI] [PubMed] [Google Scholar]
  • 61.Shapiro MG, Goodwill PW, Neogy A, Yin M, Foster FS, Schaffer DV, Conolly SM 2014. Biogenic gas nanostructures as ultrasonic molecular reporters. Nat Nanotechnol 9(4):311–316. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 62.Yu FTH, Chen XC, Wang JJ, Qin B, Villanueva FS 2016. Low Intensity Ultrasound Mediated Liposomal Doxorubicin Delivery Using Polymer Microbubbles. Mol Pharmaceut 13(1):55–64. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 63.Burke CW, Alexander E, Timbie K, Kilbanov AL, Price RJ 2014. Ultrasound-activated Agents Comprised of 5FU-bearing Nanoparticles Bonded to Microbubbles Inhibit Solid Tumor Growth and Improve Survival. Mol Ther 22(2):321–328. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 64.Hernandez C, Abenojar EC, Hadley J, de Leon A, Coyne R, Perera R, Gopalakrishnan R, Basilion JP, Kolios MC, Exner AA 2019. Sink or float? Characterization of shell-stabilized bulk nanobubbles using a resonant mass measurement technique. Nanoscale 11(3):851–855. [DOI] [PMC free article] [PubMed] [Google Scholar]
  • 65.Abenojar ENP, De Leon A, Perera R, Wang Y, Bederman I, Exner A 2019. Effect of bubble concentration on the performance of highly stable lipid shell-stabilized micro- and nanoscale ultrasound contrast agents. Langmuir. [DOI] [PubMed] [Google Scholar]
  • 66.Geers B, Lentacker I, Sanders NN, Demeester J, Meairs S, De Smedt SC 2011. Self-assembled liposome-loaded microbubbles: The missing link for safe and efficient ultrasound triggered drug-delivery. J Control Release 152(2):249–256. [DOI] [PubMed] [Google Scholar]
  • 67.Lentacker I, Geers B, Demeester J, De Smedt SC, Sanders NN 2010. Design and Evaluation of Doxorubicin-containing Microbubbles for Ultrasound-triggered Doxorubicin Delivery: Cytotoxicity and Mechanisms Involved. Mol Ther 18(1):101–108. [DOI] [PMC free article] [PubMed] [Google Scholar]

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