Therapeutic Ultrasound Parameter Optimization for Drug Delivery Applied to a Murine Model of Hepatocellular Carcinoma

Abstract

Ultrasound and microbubble (USMB)-mediated drug delivery is a valuable tool for increasing the efficiency of the delivery of therapeutic agents to cancer while maintaining low systemic toxicity. Typically, selection of USMB drug delivery parameters used in current research settings are either based on previous studies described in the literature or optimized using tissue-mimicking phantoms. However, phantoms rarely mimic in vivo tumor environments, and the selection of parameters should be based on the application or experiment. In the following study, we optimized the therapeutic parameters of the ultrasound drug delivery system to achieve the most efficient in vivo drug delivery using fluorescent semiconducting polymer nanoparticles as a model nanocarrier. We illustrate that voltage, pulse repetition frequency and treatment time (i.e., number of ultrasound pulses per therapy area) delivered to the tumor can successfully be optimized in vivo to ensure effective delivery of the semiconducting polymer nanoparticles to models of hepatocellular carcinoma. The optimal in vivo parameters for USMB drug delivery in this study were 70 V (peak negative pressure = 3.4 MPa, mechanical index = 1.22), 1-Hz pulse repetition frequency and 100-s therapy time. USMB-mediated drug delivery using in vivo optimized ultrasound parameters caused an up to 2.2-fold (p < 0.01) increase in drug delivery to solid tumors compared with that using phantom-optimized ultrasound parameters.

Keywords: Drug delivery; Hepatocellular carcinoma; Microbubbles; Sonoporation; Therapy; Ultrasound.

Figure 1:

US-mediated drug-delivery system: (a) dual-transducer design used in (Chowdhury et al., 2016, 2018), and (b) corresponding B-mode image. (c) Single transducer design, used in this study, and (d) corresponding B-mode image.

Figure 2:

Schematic view of the in vivo US parameter optimization experiments. Top row: PRF and voltage parameter optimization experiments. CEUS image intensity was used to determine the optimal PRF and voltage. Middle row: Treatment time optimization experiments. Confocal immunofluorescence imaging was used to determine the optimal number of pulses per focal region. Bottom row: Comparison of in vivo-optimized and phantom-optimized therapy parameters experiments. Confocal immunofluorescence imaging used to determine the therapy resulting in higher SPN delivery to the HCC tumor.

Figure 3:

Measured therapy beam’s profile. a) Lateral scan; b) axial scan; c) elevation scan; d) 2D lateral/elevation scan with −6 dB area contour shown in white.

Figure 4:

Peak negative pressure in water as a function of applied voltage for the focused excitation from the L11-5 transducer.

Figure 5:

CEUS images of HCC tumor in mice injected with different microbubbles: a) MicroMarker; b) BR38; c) Definity d) Sonovue. White arrows indicate the location of visible microbubbles. All images show 50 dB of dynamic range.

Figure 6:

Experimentally obtained (blue) and fitted (red) time-intensity curves after applying a therapeutic US pulse of 90 V. The time-intensity curve measured here was t_p = 1.51 ± 0.11 sec, yielding a PRF of ~0.7 Hz.

Figure 7:

Measured time-intensity curves (blue) and their running average (black) in HCC tumors after: (a) applying a single therapeutic US pulse (red dashed line) at 70 V; (b) average curve for all experiments at 70 V; (c) applying a single therapeutic US pulse at 90 V; (d) average curve for all experiments at 90 V; (e) first pulse at 90 V; (f) 100th pulse at 90 V applied to the same focal region as in (e).

Figure 8:

Representative immunofluorescence images of SPN model drug (red) and endothelial marker CD31 (green) for blood vessel visualization in subcutaneous HCC tumors. The tumors were treated with in vivo-optimized PRF (1 Hz) and voltage (70 V, MI = 1.22), while the therapy time was varied: (a) 10 sec; (b) 30 sec; (c) 50 sec; (d) 75 sec; (e) 100 sec. Scale bars are 75 μm.

Figure 9:

Amount of SPN (per unit vessel area, normalized to 10 pulses) delivered to the HCC tumor cells after applying US therapy with in vivo-optimized PRF_opt (1 Hz) and voltage (70 V, MI = 1.22) and varying treatment time. Error bars represent the standard deviation from the mean value.

Figure 10:

Representative immunofluorescence images of SPN model drug (red) and endothelial marker CD31 (green) for blood vessel visualization in subcutaneous HCC tumors. (a) treated region with in vivo-optimized US parameters; (b) treated region with phantom-optimized US parameters; (c) magnification of a region inside the square in (a); (d) magnification of a region inside the square in (b); (e) untreated region in the control tumor, same mouse as in (a); (f) untreated region in the control tumor, same mouse as in (b). Scale bars are 100 μm in (a), (b), (d), (e), and 20 μm in (c), (d).

Figure 11:

Comparison of therapy using phantom-optimized and in vivo-optimized parameters. Left: SPN penetration depth from blood vessels into tumor. The central mark indicates the median, and the edges of the box indicate interquartile range. There was no significant difference between the control tumors or those treated with in vivo optimized US parameters. Right: Quantification of the control-normalized increase in the SPN intensity for therapies using in vivo-optimized and phantom-optimized parameters. Error bars represent the standard deviation from the mean value.

Figure 12:

Representative H&E-stained tumor sections in treated groups with phantom-optimized or in vivo-optimized US acoustic parameters show no histological damage compared to the control tumors without US treatment. Scale bar, 100 μm.