High-amplitude ultrasound waves have been reported to be capable of cavitation-induced soft tissue destruction (Barnard et al. 1955; Cathignol et al. 1998; Dunn and Fry 1971; Fry and Dunn 1956; Fry et al. 1970; Tavakkoli et al. 1997). Histotripsy uses short (<20 cycles), high-pressure (>10 MPa) ultrasound pulses to generate contained dense bubble clouds and produce well-demarcated tissue fractionation (Lake et al. 2008; Parsons et al. 2006a; Roberts 2005; Vlaisavljevich et al. 2013). When these energetic bubble clouds are targeted at a fluid–tissue interface, controlled tissue erosion can also be created (Miller et al. 2013; Owens et al. 2011; Xu et al. 2004, 2010). Additionally, histotripsy can induce controlled comminution of model renal calculi at a fluid–calculus interface (Duryea et al. 2011a, 2011b).
Maxwell et al. (2013) found that when histotripsy is applied with pulses shorter than 2 cycles, the formation of a dense bubble cloud depends only on the applied peak negative pressure (p−) exceeding the “intrinsic threshold” of the medium (absolute value of 26–30 MPa in most soft nitric oxide synthase inhibitor with high water content). With an applied p− not significantly higher than this threshold, a very precise, sub-wavelength lesion could consistently be generated (“microtripsy”) (Lin et al. 2014b).
Our recent study (Lin et al. 2014a) reported that a sub-threshold high-frequency probe pulse (3 MHz, <2 cycles) can be enabled by a sub-threshold low-frequency pump pulse (500 kHz, <2 cycles) to exceed the intrinsic threshold. This pump–probe method of controlling a supra-threshold volume is called dual-beam histotripsy. Because the low-frequency pulse experiences less attenuation/aberration, and the high-frequency pulse can provide precision in lesion formation, this dual-beam histotripsy approach can be very useful in situations where precise lesion formation is required through a highly attenuative/aberrative medium, especially if a small acoustic window is available for the high-frequency pulse (Lin et al. 2014a).
Conventionally, the transmission pulse amplitude of a diagnostic ultrasound transducer does not exceed defined limits to avoid inducing possible harmful bio-effects. Thermal index (TI) and mechanical index (MI) are the two primary metrics that the U.S. Food and Drug Administration (FDA) used to regulate the acoustic output of a diagnostic ultrasound system. However, for therapeutic ultrasound systems, these restrictions no longer apply, and some studies have investigated using diagnostic ultrasound transducers to perform therapeutic procedures. Specifically, Bailey et al. used acoustic radiation forces generated by a diagnostic transducer and a Verasonics system to displace kidney stones, to expel small stones or relocate an obstructing stone to a non-obstructing location (Bailey et al. 2013; Harper et al. 2013; Sorensen et al. 2013).
In this study, a 20-element 345-kHz array transducer was used to provide the low-frequency pump pulses, whereas an ATL L7-4 imaging transducer (Philips Healthcare, Andover, MA, USA) pulsed by a Verasonics ultrasound system was used to generate the high-frequency probe pulses. The feasibility of this dual-beam histotripsy approach using an imaging transducer was tested with red blood cell (RBC) tissue-mimicking phantoms and validated in ex vivo porcine liver tissue. The capability of steering bubble clouds and lesions by steering the imaging transducer was also investigated. In the ex vivo porcine liver experiments, the L7-4 imaging transducer was used together with the Verasonics system to provide image feedback for treatment monitoring as well as to form lesion-producing bubble clouds.
We have illustrated that a sub-threshold high-frequency probe pulse provided by an imaging transducer can create lesion-producing bubble clouds when this probe pulse is “enabled” by a sub-threshold low-frequency pump pulse to exceed the intrinsic threshold (dual-beam histotripsy [Lin et al. 2014a] using an “imaging transducer”).