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. Author manuscript; available in PMC: 2018 Jan 1.
Published in final edited form as: Nanomedicine. 2016 Aug 23;13(1):59–67. doi: 10.1016/j.nano.2016.08.020

Improving Performance of Nano-Scale Ultrasound Contrast Agents Using N,N diethylacrylamide Stabilization

Reshani H Perera 1, Hanping Wu 1, Pubudu Peiris 1, Christopher Hernandez 1, Alan Burke 1, Helen Zhang 1, Agata A Exner 1
PMCID: PMC5237399  NIHMSID: NIHMS831257  PMID: 27565686

Abstract

The design of nanoscale yet highly echogenic agents for imaging outside of the vasculature and for ultrasound-mediated drug delivery remains a formidable challenge. We have previously reported on formulation of echogenic perfluoropropane gas nanobubbles stabilized by a lipid-Pluronic surfactant shell. In the current work we describe the development of a new generation of these nanoparticles which consist of perfluoropropane gas stabilized by a surfactant and lipid membrane and a crosslinked network of N, N-diethylacrylamide. The resulting crosslinked nanobubbles (CL-PEG-NB) were 95.2 ± 25.2 nm in diameter and showed significant improvement in stability and retention of echogenic signal over 24 h. In vivo analysis via ultrasound and fluorescence mediated tomography showed greater tumor extravasation and accumulation with CL-PEG-NB compared to microbubbles. Together these results demonstrate the capabilities and advantages of a new, more stable, nanometer-scale ultrasound contrast agent that can be utilized in future work for diagnostic scans and molecular imaging.

Keywords: ultrasound contrast agent, nanobubble, microbubble, extravasation, cancer

Graphical abstract

Traditional ultrasound contrast agents are too large to extravasate beyond the vasculature, making them ineffective for imaging of target biomarkers located on the cancer cell surface. This work describes the formulation and comprehensive characterization of nanoparticle-sized (~100 nm) ultrasound contrast agents which are based on lipid/surfactant stabilized perfluorocarbon gas. To add stability and improve longevity of ultrasound signal, the nanobubbles are further stabilized by a crosslinked polymer component. Such agents can be inherently be imaged at ultrasound frequencies that are clinically relevant with little post formulation modification have not been previously described.

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BACKGROUND

Clinically utilized ultrasound contrast agents (UCA) are comprised of gas bubbles stabilized by lipid, polymer or protein membranes. Their use capitalizes on the outstanding safety, relatively low cost and wide accessibility of ultrasound to augment diagnostic and therapeutic capabilities of this imaging modality. Recently, UCAs have gained more attention as theranostic agents and are currently being explored as an image-guided, ultrasound-triggered drug delivery system in various applications1–4. Clinically available UCAs are FDA approved in the United States for use in echocardiography studies, and are in clinical trials for other applications such as liver tumor diagnosis5–7. However, existing UCAs have a relatively large hydrodynamic diameter (1–8 μm) confining them to the vasculature and a short circulation half-life stemming from relatively limited structural stability8. Reducing the size of UCAs to near 100 nm addresses the first of these concerns. The submicron contrast agents can exploit the leaky vasculature and hampered lymphatic drainage of many tumors via the enhanced permeability and retention effect (EPR)9, giving them the ability to penetrate into tumor parenchyma and target cell surface markers. Another attractive feature of these agents is the relative ease with which payload delivery can be enhanced within the tumor following triggered on-demand release of therapeutic using external low frequency ultrasound stimulation10, 11.

A number of existing methods of producing nanometer-sized nanobubbles exist currently, yet most utilize post-formulation manipulations, such as centrifugation or filtration, which reduce both yield and stability12–16. Others utilize external forces and changes in temperature to achieve the desired sizes and echogenicity17. In addition, decreasing particle radius using the same material will often lead to stiffening of the membrane and increasing resonant frequency, thus leading to a loss of signal at clinically relevant ultrasound transducer frequencies18, 19, which ultimately harms image quality. Our group has previously reported on a simple formulation strategy of echogenic nano-sized perfluoropropane gas bubbles using the surfactant Pluronic to stabilize the phospholipid membrane and reduce size of the bubbles to as low as 100 nm in both ambient and physiological conditions20–22. Pluronics (also known as poloxamers) are a family of nonionic tri-block copolymers assembled with polyethylene oxide (PEO) and polypropylene oxide (PPO), following the general structure of PEOx-PPOy-PEOx. While the mechanism of action is not fully understood, it is known that Pluronic interacts with a lipids in the shell and reduces the size of bubbles to nanometer range20. Pluronic-stabilized nanobubbles are self-assembled, have a low tendency to coalesce into larger bubbles and are believed to be the first lipid perfluorocarbon formulation that does not require post-formulation manipulation to achieve consistent sub-micron size distribution. Both stability and echogenicity are believed to be related to some degree to lipid shell fluidity, which can be altered by addition of Pluronic20, 21, 23.

Although this agent has echogenicity and stability comparable to clinically-used UCAs20, 21, bubble stability especially in applications of molecular imaging and intratumoral drug delivery requires improvement20, 21. In the current approach, we report on the formulation and characterization of second generation Pluronic-stabilized nanobubbles. These are created by adding interpenetrating crosslinking biodegradable polymer N, N-diethyl acrylamide (NNDEA) and N,N-bis(acryoyl) cystamine (BAC) (Fig.1) and integrating additional polyethylene glycol (PEG) groups above and beyond those of the Pluronic PEO subunits on the surface of the nanoparticles24, 25. Incorporation of crosslinking agents has been shown previously to increase stability of Pluronic polymeric micelles below their critical micelle concentration (CMC)26–28. In these crosslinked Pluronic-lipid-perfluorocarbon bubbles (CL-PEG-NB) the hydrophobic network is non-covalently integrated into the inner ring or the hydrophobic domain of thebubble, which should improve structural stability while retaining membrane flexibility and reduce diffusion of hydrophobic perfluorocarbon gas out of the core. In this report, the agents were characterized in vitro and their biodistribution and extravasation were examined in a LS174T colorectal tumor xenograft in live mice.

Figure 1.

Figure 1

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(A) Schematic diagram of lipid and Pluronic-stabilized nanobubbles with interpenetrating crosslinking biodegradable polymer N, N-diethyl acrylamide (NNDEA) and N,N-bis(acryoyl) cystamine (BAC) crosslinking network.

MATERIALS AND METHODS

Formulation of Nanobubbles

To formulate stabilized CL-PEG-NB, 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) (Avanti Polar Lipids, Pelham, AL), and mPEG-DSPE (1,2-Distearoyl-phosphatidylethanol amine-methyl-poly ethylene glycol conjugate-2000) (Laysan Lipids, Arab, AL) were dissolved in chloroform in a 4:1:1:1 mass ratio. The solvent was then removed by evaporation, leaving a lipid film. The film was hydrated by adding 1ml of 0.6 mg/ml Pluronic solution (Sigma Aldrich, Milwaukee, WI) in 0.5% Irgacure 2959 (Fisher Scientific; Pittsburgh, PA) in PBS in the presence of glycerol (50 μl) at 75° C for 30 min. Next, NNDEA (Polysciences, Warrington, PA) and BAC (Sigma Aldrich, Milwaukee, WI) (2:1 weight ratio) were added, and air was removed from the sealed vials and replaced with octafluoropropane until the vial pressure equalized. Finally, the vial was shaken on a VialMix shaker (Bristol-Myers Squibb Medical Imaging, Inc., N. Billerica, MA) for 45s, and the bubble vials were irradiated at 254 nm using a UV lamp (Spectronics Co. Westbury, NY) for 30 min. If not used immediately, CL-PEG-NB were stored at 4°C.

To formulate Pluronic PEG-NB, the lipid film was prepared using above lipids and hydrated by adding 1ml of 0.6 mg/ml Pluronic L10 solution in the presence of glycerol (50 μl). After keeping the solution at 75 °C for 30 min, air was removed from the sealed vials and octafluoropropane (C3H8) was added to the vials until vial pressure equalized. Then, the vial was shaken on a VialMix shaker for 45 s. Bubble samples were stored at 4°C until use.

Nanobubble Characterization

The mean diameter and polydispersity of Pluronic nanobubbles were measured using dynamic light scattering (DLS) (90 Plus, Brookhaven Instruments Corp). Measurements were performed at 25 °C, with a laser wavelength of 660 nm at an angle of 90°. Bubble size was measured by diluting a sample 1:1000 with PBS at pH 7.4 (n =3). Bubble size is reported as a number average. Nanobubble morphology was imaged using scanning electron microscopy (SEM) and gated Stimulated Emission Depletion (STED) imaging. In order to prepare samples for SEM imaging, a drop of freshly prepared bubble solution was placed on dust-free foil and kept in a desiccator to evaporate the solvent. Then the samples were sputter coated with palladium and images were obtained using scanning electron microscope (Hitachi S4500) with a gun acceleration voltage of 3.0 kV and 7mm working distance15.

The qNano platform from Izon Science (Izon Science, Christchurch, New Zealand) was used to analyze the absolute diameter and the concentration of CL-PEG-NB. The qNano was equipped with the NP150 nanopore for detection of particles in the range of 70–200 nm. The optimized parameters of the qNano were adjusted (45 mm stretch, 0.7 V current) to obtain a stable baseline current with PBS buffer. Then, 40 μl of diluted CL-PEG-NB solution was loaded to the upper fluid cell and pressure (0.5kPa) applied to obtain the desired 500 particle count. The size and concentration distribution plots were created by Izon Control Suite software, version 2.1.

In vitro Ultrasound Stability Characterization

A custom tissue mimicking phantom (ATS Laboratories, Bridgeport, CT) was immersed in a 1L glass beaker containing 700 ml of PBS (Fisher Scientific; Pittsburgh, PA) at 37° C. Nanobubble solution (700 μl) was injected into the PBS solution, and the contents of the beaker were continuously stirred at 150 rpm to agitate the bubbles (Fig. 3a). The change in contrast as a function of time was measured using an AplioXG SSA-790A clinical ultrasound imaging system (Toshiba Medical Imaging Systems, Otawara-Shi, Japan) equipped with a 12 MHz linear array transducer. System acquisition parameters were set to Contrast Harmonic Imaging (CHI) with 8.0 MHz harmonic frequency, 0.1 mechanical index (MI), 1 Hz imaging frame rate, 65 dB dynamic range, and 80 dB gain.

Figure 3.

Figure 3

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(A) Representative ultrasound images of nanobubbles in vitro, showing both custom-made tissue phantom area and the contrast agent area (B) Enhanced in vitro stability of CL-PEG-NB compared to original PEG-NB over 1h (inset) and 24 h as measured as a function of time.

Images were obtained over a 60 min period using a low imaging frame rate (0.1 Hz). Regions of interest (ROI) of the same size were drawn in the tissue phantom area and in the contrast agent vicinity. The data were exported to Excel (Microsoft, Redmond, WA) and normalized by the contrast of tissue phantom. The signal intensity (logarithmic scale of the normalized data) as a function of time was plotted to obtain the decay rate of nanobubbles. To acquire the rate of nanobubble dissolution over 24 h period, at different time points (t = 0, 1, 3, 6, 24 h) of post injection of bubbles, 10 images were obtained and signal intensity was plotted as a function of time.

Cell Culture

LS174T human colorectal adenocarcinoma cells (ATCC, Manassas, VA) were cultured in complete MEM medium (10% fetal bovine serum, 1% penicillin-streptomycin; Invitrogen, Carlsbad, CA) and placed in a humidified atmosphere at 37 °C and 5% CO2. At 90% confluence, cells were detached using 0.25% trypsin-EDTA (Invitrogen, Carlsbad, CA) for passaging.

Animal Preparation and Tumor Inoculation

Animals were handled according to a protocol approved by the Institutional Animal Care and Use Committee at Case Western Reserve University and were in accordance with all applicable protocols and guidelines in regards to animal use. In all procedures, the animals were anesthetized with 3% isoflurane with 1L/min oxygen. LS174T cells (1×107 cells) suspended in 200 μl of growth medium were injected subcutaneously into the flank of each mouse.

Ultrasound Biodistribution of Nanobubbles In Vivo

Two weeks after inoculation, mice with tumor diameter of 0.8 cm or larger, were selected for biodistribution studies (n=3 per group). To visualize ultrasound images of liver, kidney, and right subcutaneous tumor, the animal was imaged through the sagittal plane. After injection of nanobubbles, the change of tissue contrast of liver, kidney, and tumor was measured using Contrast Harmonic Imaging (CHI, frequency 8.0 MHz; MI, 0.08; dynamic range, 65 dB; gain, 80 dB; imaging frame rate, 1 frame/s). The images were acquired in raw data format as a function of time. Fifteen seconds after raw data acquisition started, nanobubble solution was administrated and continuous image acquisition continued for 8 min. At this point, image acquisition continued for 22 min. with an imaging frame rate of 0.2 frames/s.

The raw data were processed with software provided by the scanner manufacturer22. The liver, kidney, and tumor areas were delineated by drawing regions of interest (ROIs) and the signal intensity in each ROI as a function of time (Time Intensity Curve -TIC) was calculated. The data were exported to Excel, the baseline was subtracted from TIC, and the calculated peak value of TIC was used to normalize the data to obtain the decay of signal at each time point. The log of the data was fitted by simple linear regression to derive the decay slope.

In Vivo Biodistribution of Nanobubbles Measured Using Fluorescence Molecular Tomography (FMT)

In vivo biodistribution was studied using FMT (FMT 2500, PerkinElmer Inc., Boston, MA). One week before FMT imaging, tumor bearing mice were placed on a special diet to reduce gut autofluorescence. Before injection, mice were anesthetized and imaged on the FMT imaging system to rule out auto fluorescence. Animals were divided into 3 groups. For the first group, mice were injected with 200 μl of 1:3 diluted CL-PEG-NB-Vivotag 680 and FMT imaging was performed at multiple time points post-injection (t= 5, 15, 30 min and 3 h, n=5). PEG-microbubbles (PEG-MB n=3) were injected as a clinically-relevant control and saline (n=2) was injected as a negative control. FMT was carried out using the 680 nm channel and body scanning 3D reconstructions of the imaging data were performed with TrueQuant software (version 3.1). ROI’s were selected for each organ, including tumor, in all 3 imaging planes (X, Y, Z) and total fluorochrome concentration was determined per ROI. At t= 3 h animals were euthanized and liver, kidney, heart and tumor were harvested and weighed. Uptake of CL-PEG-NB-Vivo 680 in tumor and other organs will be calculated as the percentage of the injected dose per gram of tissue (% ID/g).

Histological analysis

Animals were divided into three groups: CL-PEG-NB (n=5), MB (n=3), and no contrast control (n=2). Mice received either 200 μL of contrast material diluted (1:4) with PBS or PBS alone via the tail vein. Three hours after bubble injection, animals were anesthetized again and Texas red tagged lectin solution (Vector Laboratories, Burlingame, CA) (0.1mL of 1mg/mL) was inject through a tail vein over a period of 20–30s. Five minutes after lectin injection, PBS perfusion was performed with 50 ml PBS through left ventricle. Organs (lung, liver, kidney) and tumors were excised and embedded in optimal cutting temperature compound (OCT Sakura Finetek USA, Inc., Torrance, CA) and frozen on dry ice. The tissues were cut into 8 μm slices using Leica CM1850 cryostat (Leica, Germany). The fluorescence images were obtained and analyzed using AxioVision V 4.8.1.0, Carl Zeiss software (Thornwood, NY).

Statistical analysis

All data are presented as mean ± standard deviation, unless otherwise noted. Statistical significance of differences between experimental groups was derived using one-way ANOVA model. Two-tailed unpaired student’s t-test with unequal variants were used to determine the significance of the outcome. Data analysis was performed with Microsoft Excel and ANOVA.

RESULTS

Nanobubble Characterization

DLS and qNano measurements showed CL-PEG-NB diameters of 95.2 ± 25.2 nm (n=10) and 90 ± 3 nm, respectively (Fig. 2 a, b). Gated Stimulated Emission Depletion (STED) confocal images confirmed the size and spherical shape of CL-PEG-NB (Fig 2d). Surface morphology and size distribution of CL-PEG-NB were obtained using SEM. Further, SEM images (Fig. 2e) confirmed that the nanobubbles were 100 ±70 nm sized, spherical and non-aggregated. A donut shape was seen in this formulation but not seen with the original nanobubbles suggesting that crosslinks are located in an annular fashion within the hydrophobic lipid and Pluronic portions surrounding the hydrophobic gas core. Concentration of CL-PEG-NB obtained from qNano was 7×1012 particles per ml (Fig. 2c), and was 2.5×1012 particles per ml for the PEG-NB, (data not shown).

Figure 2.

Figure 2

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Size distribution of the CL-PEG-NB obtained by (A) DLS and (B) qNano; (C) Concentration distribution of CL-PEG-NB obtained from qNano; (D) STED image of Alexa Fluor 488 tagged CL-EG-NB; (E) Surface morphology of the CL-PEG-NB and normal PEG-NB visualized using scanning electron microscopy (SEM).

In Vitro Echogenicity Characterization

In vitro stability was analyzed by imaging the nanobubbles at 37°C under constant agitation and calculating the relative signal loss over time. Fig. 3a shows representative gray scale ultrasound images of PEG-NB and CL-PEG-NB in vitro in the tissue mimicking phantom. The ratios of mean echo-power values of contrast agents (both PEG-NB and the CL-PEG-NB) and the tissue were plotted over time period of 60 min (Fig. 3b inset). The initial contrast at time t=0 was significantly higher with CL-PEG-NB (P<0.01) than the initial contrast of regular PEG-NB. The CL-PEG-NB showed a 2.8 times lower decay rate compared to the decay rate of non-crosslinked PEG-NB (−0.266 db/min compared to − 0.742 db/min) over 1 h (Fig. 3 insert). As shown in Fig. 3b the signal intensities of CL-PEG-NB at 6 and 24 h are significantly higher than PEG-NB (P < 0.05).

In Vivo Biodistribution of Nanobubbles Visualized with Ultrasound

In vivo ultrasound biodistribution of nanobubbles was examined using ultrasound. Fig. 4 shows representative contrast images of selected ROIs (tumor, kidney, liver). The mean echo-power value in the ROIs as a function of time or time-intensity curve (TIC) is shown in Fig. 5a and 5b. Subsequently administrated of either normal PEG-NB or the CL-PEG-NB via the tail vein of LS174T tumor model bearing mice, the three tested regions (kidney, liver and tumor) showed rapid contrast enhancement. The peak enhancement occurred in both cases 21s after injection, which was followed by gradual wash-out in all tested regions. The PEG-NB accumulation in the kidney and liver was 1.4 times higher than the CL-PEG-NB accumulation in the same organs. CL-PEG-NB showed nearly 2 fold higher enhancement in the tumor area compared to PEG-NB.

Figure 4.

Figure 4

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Ultrasound images showing the contrast in each organ and the tumor after the bubble injection as a function of time. Regions of interests (ROIs) delineate the liver, kidney, and subcutaneous tumor.

Figure 5.

Figure 5

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Ultrasound intensity of each organ and the tumor after injecting (A) PEG-NB; (B) CL-PEG-NB; (C) Normalized ultrasound intensity of each organ at the initial phase after peak enhancement (n=3).

After peak enhancement the bubble signal dissipation occurred in two phases. Fig. 5 shows the percent of peak signal intensity of tested regions as a function of time after the peak enhancement that corresponds to the initial phase of bubble dissipation. The decay slopes of CL-PEG-NB in tumor and kidney were significantly slower than those of normal PEG-NB in the initial phase (P <0.05; 15.63 ± 4.06 compared to the 25.03 ± 3.22 in kidney and 13.05 ± 4.19 compared to the 21.96 ± 3.03 in tumor). There were no significant differences observed in the decay slope in the liver during the initial phase.

In Vivo Biodistribution Studies Measured with FMT

The FMT signal was initially validated using a phantom experiment which showed linear correlation between signal and the various concentrations of bubbles tagged with the fluorescent probe. To quantify the percentage of initial dose of bubbles accumulated per gram of tissue (%ID/g), the fluorescence signal of each organ was analyzed using FMT scan and normalized to the mass of the organ recorded after organ excision. Fig. 6a shows the representative FMT images that were taken 1 h after IV injection of the CL-PEG-NB, PEG-MB and saline control. As shown in Fig. 6b, the intratumoral accumulation of fluorescent signal in tumor from PEG-MB was significantly lower compared to the CL-PEG-NB at each time point post injection (p< 0.01).

Figure 6.

Figure 6

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Tumor nanobubble and microbubble kinetics measured with FMT. (A) FMT images show the accumulation of Vivo680-tagged CL-PEG-NB and Vivo680 tagged PEG-MB in tumor at 1 h post injection; (B) Fluorescence intensities of tumor after application of above treatments as a function of time. CL-PEG-NB perform better than the MB control.

Fig. 7 shows biodistribution of CL-PEG-NB compared to PEG-MB. We did not observe any statistically significant differences in bubble accumulation in liver immediately after or 1 h post-injection. However, in the kidney, the PEG-MB accumulation was significantly increased in all time points compared to CL-PEG-NB.

Figure 7.

Figure 7

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Fluorescence intensities of each organ (liver, kidney, heart) 5 and 60 min of post application of Vivo680-CL-PEG-NB and Vivo680-PEG-MB.

Histological analysis

Results are shown in Fig. 8. Here, the red color indicates the vessels stained with Texas red and the green color indicates presence of nanobubbles or nanobubble components. In the CL-PEG-NB group, bubble signal appeared inside the tumor, outside the vessels, which provides strong evidence of bubble extravasation and subsequent interstitial penetration. In the PEG-MB group the bubbles remained primarily within the vasculature (Fig. 8a, 8b). In both cases signal from tumor vasculature appeared to be 21–23 % of the total tissue (Fig. 8c). The nanobubble signal intensity in tumor tissue was 12 times higher (36±19 % of the total tumor tissue) compared to the microbubble signal (3±9 %) and was significantly different (p<0.05).

Figure 8.

Figure 8

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(A) Representative fluorescence images showing the bubble distribution in tumor (green). Alexa 488-tagged CL-PEG-NB showed higher extravasation compared to the Alexa 488-tagged PEG-MB. MB confined to the tumor vasculature (red); (B) Representative image showing the extravasation of CL-PEG-NB into the tumor beyond the tumor vasculature; (c) The signal intensities of bubble and vessel expressed as the percentage of total cells of tumor tissues.

DISCUSSION

The goal of this study was to formulate a stabilized ultrasound-visible UCA in the size range of nanoparticles typically utilized for molecular imaging and drug delivery in cancer applications. DLS, SEM and STED data provide evidence of the size and morphology of the crosslinker stabilized Pluronic bubbles which are suitable for imaging and delivery applications. Importantly, the small size does not affect the echogenicity of the new formulation. The distinct donut shape seen in SEM images with the CL-PEG-NB formulation suggests stable crosslinks arranged in an annular fashion in the hydrophobic core of the bubble. The persistence of this structure throughout sample preparation also may suggest that these networks play a key role in stabilizing the nanobubble shape. The “deflated” look presumably results from gas diffusion out of the constructs. In vitro echogenicity of regular PEG-NB and the CL-PEG-NB was compared next. Immediately after injection of nanobubbles, the signal was significantly higher (p<0.01) for the CL-PEG-NB, compared to PEG-NB. Based on qNano data, while the concentrations for both formulations were in the 1012 range, there were indeed more CL-PEG-NB, which could account for the higher signal. It is also possible that higher signal generated by CL-PEG-NB could be due, in part, to more gas entrapped trapped in the high surface area of the crosslinked NNDEA network29, 30. The elasticity of the crosslinked polymer could also account for the enhanced echogenicity31, 32. More investigation is needed to confirm these suppositions. CL-PEG-NB also showed a decreased rate of decay in ultrasound signal compared to PEG-NB. This could also be a result of the crosslinked polymer network, which may reduce diffusion of gas from the construct.

In vivo ultrasound biodistribution studies data demonstrate better tumor enhancement with CL-PEG-NB compared to PEG-NB. The enhanced signal is likely due to persistent echogenicity and stability leading to the greater accumulation and retention of CL-PEG-NBs in the tumor. With the CL-PEG-NB, we observed the contrast not only in the tumor periphery, but also inside the tumor matrix, which suggests improved nanobubble penetration into the tumor in contrast to the regular PEG-NB and other reported nanobubbles32.

The liver and kidney contrast enhancement was higher with PEG-NB than CL-PEG-NB. In the same organs, PEG-NB had a higher initial decay rate in the liver and kidney compared to the CL-PEG-NB. In contrast, the CL-PEG-NB cleared from the tumor much more slowly than the non-crosslinked NBs, supporting the notion that CL-PEG-NB may accumulate in tumors more extensively than the PEG-NB24, 25. In vivo FMT results show that CL-PEG-NB accumulate in the tumor over time at a much greater rate than microbubbles. A limitation of FMT is the difficulty of assigning the ROI for the selected organ in the reconstructed image, especially in the abdomen. The gut signal may overlap with the liver and kidney signals and may account for the inaccuracy and the high standard deviation of the quantitative assessment of the signal33–36. The signal in the kidney after injection of MB was elevated compared to the CL-PEG-NB injected kidney. This may be due to the renal retention of microbubble37, 38. Continued investigation into bubble biodistribution is necessary to better understand this behavior.

Accumulation of nanobubbles in tumors relies on extravasation into the tumor parenchyma. Upon histological analysis, CL-PEG-NB showed more extensive extravasation beyond the tumor vessel area, while large PEG-MB showed limited extravasation. However, EPR-driven extravasation depends greatly on the type, stage and morphology of the tumor among many other parameters. Thus these data are only supportive of the process in our LS174T colorectal tumors and additional studies using different tumor models will be necessary to gather more broadly applicable data.

We have demonstrated the development of stable ultrasound-sensitive nanobubbles by incorporating a interpenetrating crosslinked network of N, N-diethylacrylamide and N, N-bis(acryoyl) cystamine into Pluronic-lipid-perfluorocarbon bubbles. The CL-PEG-NB showed enhanced ultrasound signal and lower decay rate compared to PEG-NB without the crosslink network. The new formulation shows promising performance in terms of in vitro and in vivo stability, echogenicity and extravasation into tumor. Future investigation will explore the feasibility of the nanobubbles as a targeted molecular imaging agent for tumor detection and as a carrier vehicle for image guided chemotherapy since the crosslinked acrylamide mesh augments the cargo capacity within the bubble core.

Acknowledgments

This work was supported by the National Cancer Institute of the National Institutes of Health and the Department of Defense Ovarian Cancer Research Program (R01CA136857 and OC110149 to AAE). The content is solely the responsibility of the authors and does not necessarily represent the official views of the National Institutes of Health or of the Department of Defense. We thank Bernadette O. Erokwu, PhD for the assistance with animal studies and Luis Solorio, PhD for support in obtaining SEM images.

Footnotes

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