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Pulse sequences for uniform perfluorocarbon droplet vaporization and ultrasound imaging
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Joint Department of Biomedical Engineering, University of North Carolina at Chapel Hill and North Carolina State University, CB 7575, Chapel Hill, NC, USA
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C. Puett, P.S. Sheeran, J.D. Rojas, P.A. Dayton ⇑
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Article history: Received 22 April 2014 Accepted 19 May 2014 Available online xxxx
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Keywords: Phase-change Ultrasound contrast agent Perfluorocarbon Vaporization Decafluorobutane Octafluoropropane
a b s t r a c t Phase-change contrast agents (PCCAs) consist of liquid perfluorocarbon droplets that can be vaporized into gas-filled microbubbles by pulsed ultrasound waves at diagnostic pressures and frequencies. These activatable contrast agents provide benefits of longer circulating times and smaller sizes relative to conventional microbubble contrast agents. However, optimizing ultrasound-induced activation of these agents requires coordinated pulse sequences not found on current clinical systems, in order to both initiate droplet vaporization and image the resulting microbubble population. Specifically, the activation process must provide a spatially uniform distribution of microbubbles and needs to occur quickly enough to image the vaporized agents before they migrate out of the imaging field of view. The development and evaluation of protocols for PCCA-enhanced ultrasound imaging using a commercial array transducer are described. The developed pulse sequences consist of three states: (1) initial imaging at sub-activation pressures, (2) activating droplets within a selected region of interest, and (3) imaging the resulting microbubbles. Bubble clouds produced by the vaporization of decafluorobutane and octafluoropropane droplets were characterized as a function of focused pulse parameters and acoustic field location. Pulse sequences were designed to manipulate the geometries of discrete microbubble clouds using electronic steering, and cloud spacing was tailored to build a uniform vaporization field. The complete pulse sequence was demonstrated in the water bath and then in vivo in a rodent kidney. The resulting contrast provided a significant increase (>15 dB) in signal intensity. Ó 2014 Elsevier B.V. All rights reserved.
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1. Introduction
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Micron-sized encapsulated bubbles of gas (microbubbles) are used clinically as ultrasound contrast agents (UCAs). OctafluoroÒ propane (OFP) microbubble contrast agents having lipid (Definity ) Ò and protein (Optison ) shells are approved by the Food and Drug Administration (FDA) in the United States as a diagnostic tool for opacifying the left ventricle and improving endocardial border delineation during echocardiography [1,2]. Additional diagnostic applications have been demonstrated utilizing the unique nonlinear acoustic signatures of oscillating microbubbles [3,4]. Experimental therapeutic applications of microbubbles include enhancement of tissue ablation due to microbubble cavitation in high intensity focused ultrasound (HIFU) [5–7], increased apoptosis following radiation therapy [8], and utilization in drug delivery [9]. However, there are limitations to the direct application of currently available UCAs. Microbubbles survive for just minutes
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⇑ Corresponding author. Tel.: +1 919 843 9521. E-mail address: [email protected] (P.A. Dayton).
in vivo [10], evaporating as they circulate through the lungs and being cleared by the reticuloendothelial system [11]. Additionally, their large size (mean diameter >1 lm) constrains them to the vasculature [12]. Sub-micron liquid perfluorocarbon (PFC) agents provide an attractive tool to introduce microbubbles into a region of interest. Droplets can be manufactured small enough to extravasate through the leaky vasculature found in malignant and inflamed tissues and can therefore enter diseased interstitial spaces [13–17]. Also, droplets can be designed to have much longer half-lives than microbubbles in vivo and yet can still be vaporized back into microbubbles by the temperature and pressure changes achievable with ultrasound [14,16,18] As such, they are referred to as phase-change contrast agents (PCCAs), and their vaporization (termed ‘‘activation’’) allows microbubbles to be generated within the acoustic field [19]. Manipulating the critical activation step from droplet to microbubble is an area of active research. Vaporization is a complex phenomenon [20] depending on many factors, including the properties of the ultrasound wave as well as the size and composition of the
http://dx.doi.org/10.1016/j.ultras.2014.05.013 0041-624X/Ó 2014 Elsevier B.V. All rights reserved.
Please cite this article in press as: C. Puett et al., Pulse sequences for uniform perfluorocarbon droplet vaporization and ultrasound imaging, Ultrasonics (2014), http://dx.doi.org/10.1016/j.ultras.2014.05.013
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PFC droplets [21–24]. As expected, droplets containing more volatile PFCs, such as decafluorobutane (DFB, C4F10, boiling point: 2 °C) and octafluoropropane (OFP, C3F8, boiling point: 37 °C), are easier to vaporize than those containing PFCs with higher boiling points [23]. These volatile PFC agents are well suited for diagnostic applications, since they can be vaporized to microbubbles with acoustic pressures used for imaging. However, PCCA-enhanced diagnostic ultrasound is still in its infancy, typically requiring separate transducers for activation and imaging due to the demanding pulse sequence requirements [17,25]. Herein, we present a unique ‘‘image-activate-image’’ protocol that demonstrates the feasibility of performing initial sub-activation threshold imaging, activation, and post-activation imaging with a single transducer. In this protocol, the initial plane-wave imaging step (both anatomical B-mode and pulse inversion for baseline bubble assessment) is used to provide the operator with information to determine the desired spatial activation pattern and is followed by a user-initiated activation step, during which electronically-steered, focused-wave pulses vaporize PFC droplets with high spatial specificity. A final plane-wave pulse inversion imaging step visualizes the vaporization field of microbubbles. Although the imaging sequences used in this study to maximize bubble detection and minimize artifacts (pulse inversion and coherent plane-wave compounding) are well described [26–30], much of the challenge of PCCA-enhanced diagnostic ultrasound lies in optimizing the activation step of the protocol. The goal is to achieve a relatively uniform vaporization field of microbubbles throughout the selected target area, utilizing acoustic energy exposures that carry a low risk of mechanical or thermal tissue injury. Since the geometry of the focal zone and the intensity achieved within the zone vary as a function of the location of the focus relative to the transducer, vaporization responses also vary. Therefore, uniform vaporization requires anticipating differences in the sizes and shapes of microbubble clouds, adjusting acoustic parameters to account for these differences, and then spacing individual clouds appropriately. In this study, testing and development were done using sub-micron DFB and OFP droplets dispersed in a degassed water bath, and the completed pulse sequence was then applied in vivo to image the rat kidney following the intravenous administration of these PCCAs. Image-activate-image protocols that direct the generation of uniform vaporization fields are useful tools for the ongoing study of PCCAs and will be critical in the eventual clinical application of these agents.
2. Methods
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2.1. Preparation of the phase-change contrast agents
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PCCAs used in this work consist of droplets of liquid PFC [either decafluorobutane (DFB, C4F10) or octafluoropropane (OFP, C3F8)] contained in a phospholipid and polyethylene glycol shell. They were prepared by the condensation procedure described previously [22]. Briefly, lipids [1,2-distearoyl-sn-glycero-3-phosphocholine (DSPC) and 1,2-distearoyl-sn-glycero-3-phosphoethanolamine-N-methoxy (polyethylene–glycol)-2000 (DSPE-PEG2000), purchased from Avanti Polar Lipids (Alabaster, AL) in a 9:1 M ratio and a total lipid concentration of 1.0 mg/mL] were emulsified in a solution of propylene glycol 15% (v/v), glycerol 5% (v/v), and phosphate-buffered saline (PBS) 80% (v/v). 1.5 mL of this emulsion was placed in a sealed 3 mL glass vial and the air headspace exchanged with PFC gas purchased from Fluoromed (Round Rock, TX, USA). Microbubbles with PFC cores and phospholipid shells formed spontaneously when agitated in a Vialmix Shaker (Bristol–Myers– Squibb, New York, NY, USA). The concentration of microbubbles in this emulsion was measured via optical techniques (Accusizer 780, Particle Sizing Systems, Santa Barbara, CA, USA) and was found to be 1010 microbubbles/mL. The microbubbles were condensed to droplets by cooling the emulsion under an adjustable high-pressure air source in an isopropanol bath maintained between 6 and 10 °C using dry ice. Dynamic light scattering (Malvern Nano ZS, Malvern Instruments, Ltd., Westborough, MA, USA) demonstrated a polydisperse population with a mean diameter of 240 ± 65 nm. Assuming a 1:1 conversion from microbubble to droplet, a stock solution containing 1010 PFC droplets/mL was used immediately or stored frozen as a stock solution for use within 24 h.
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2.2. In Vitro setup
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In vitro testing was done in an acrylic-walled tank containing 4.7 L of degassed water. DFB or OFP droplets were added to the water to achieve a final concentration of approximately 106 droplets/mL. For studies using DFB droplets, the water was maintained at 37 °C by circulating warm water through copper tubing in the bath. Because OFP droplets are relatively unstable at 37 °C [23], testing with OFP was done at room temperature (22 °C) to ensure an adequate half-life for experimentation. The water was stirred between experimental runs to redistribute PFC droplets, and the
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Fig. 1. The image-activate-image protocol for phase-change contrast agent (PCCA)-enhanced diagnostic ultrasound. This flow diagram identifies the tasks accomplished through each step of the protocol as well as the time required to complete the pulse sequences for each step.
Please cite this article in press as: C. Puett et al., Pulse sequences for uniform perfluorocarbon droplet vaporization and ultrasound imaging, Ultrasonics (2014), http://dx.doi.org/10.1016/j.ultras.2014.05.013
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transducer was angled to reduce direct ultrasound reflections off the tank walls.
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2.3. Pulse sequence design
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Our approach to PCCA-enhanced diagnostic ultrasound involves a unique ‘‘image-activate-image’’ pulse sequence (Fig. 1). A programmable research ultrasound system (Verasonics, Redmond, WA) driving an ATL L12-5 38 mm 192-element linear array transducer was used to implement this pulse sequence. The initial imaging step included seven transmissions of steered plane waves (angles: 18°, 12°, 6°, 0°, 6°, 12°, 18°). 128 elements (the maximum number of connector channels) were used to transmit single-cycle pulses at 9 MHz. Echoes were received at the same frequency, bandpass filtered using a 21 tap finite impulse response filter with a 4–14 MHz bandwidth, and averaged using built-in Verasonics reconstruction processes. Pixel-based algorithms displayed a B-mode image of the anatomy [31]. This image allowed for selection of a region of interest. Next, successive pulse inversion plane-wave transmissions were delivered at seven angles as previously described. At each angle, two 1-cycle pulses at 4.5 MHz were delivered 180° out of phase and the received signals were summed to isolate the nonlinear microbubble oscillation. The echoes received from different transmission angles were averaged to obtain a coherent plane-wave compounded image, providing a background (baseline) microbubble assessment. Electronically focused ultrasound pulses were transmitted during the activation step to vaporize droplets and generate microbubble clouds. In order to image activated bubbles before significant motion occurs, the entire activation process needs to be accomplished quickly. Therefore, this activation sequence directed changes to transducer function that require little time, including sub-aperture adjustments (the position and number of active elements) and the number of cycles in each activation pulse. To manipulate individual microbubble cloud geometries, focused pulses were delivered at 5 MHz using 32, 64, 96, or 128 elements and included 2, 5, 10, or 15 cycles per pulse, while the sub-aperture was shifted to minimize the steering required to target focal sites. The final imaging step visualized the microbubble clouds and involved repeating the baseline seven-angle pulse inversion transmissions at 4.5 MHz and averaging the received echoes to form a coherent plane wave-compounded image. These reconstructed
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Fig. 2. Measurement of droplet activation (vaporization) as a function of the voltage applied during the imaging steps of the image-activate-image protocol. Insonation involved seven plane-wave transmissions using coherent plane wave compounding with pulse inversion applied over a voltage range from 10 to 30 V. This voltage range corresponds to free-field peak negative pressures (PNPs) from 0.38 to 1.13 MPa. Non-targeted activation was measured as the difference in the average background brightness before and after the imaging step being tested. A significant (p < 0.0001) increase in non-targeted decafluorobutane (DFB) droplet vaporization occurred when the imaging step used >25 V, which equated to a PNP of 0.96 MPa.
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images were compared to those obtained prior to activation. The custom image-activate-image sequence was implemented in Matlab (The Mathworks, Natick, MA) on the scanning system. Changing voltage in real time during the image-activate-image protocol is problematic, since it is a relatively slow process, requiring a delay of at least 1 ms due to system limitations. Therefore, testing was done to determine the optimum voltage for use throughout the protocol. Selecting this voltage required finding a balance between pressures high enough for focused-wave activa-
Fig. 3. Energy deposition as a function of depth in the acoustic field. Free-field pressures during plane-wave imaging were measured at depths from 0.5 to 1.5 cm, and then derated to calculate the mechanical index (MI) (A), spatial peak-pulse average intensity (ISPPA.3) (B), and spatial peak-temporal average intensity (ISPTA.3) (C) resulting from 20 V (gray square), used for OFP studies, and 25 V (black diamond), used for DFB studies. The derated values were compared to the FDA thresholds (dashed lines) for small organ imaging (MI: 1.9, ISPPA.3: 190 W/cm2, ISPTA.3: 94 mW/cm2). The Definity suggested MI threshold of 0.8 [Definity Imaging] when bubbles are present in the acoustic field (dotted line) is indicated in (A). The depth-dependent ISPTA.3 during focused-wave activation (D) is also shown.
Please cite this article in press as: C. Puett et al., Pulse sequences for uniform perfluorocarbon droplet vaporization and ultrasound imaging, Ultrasonics (2014), http://dx.doi.org/10.1016/j.ultras.2014.05.013
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tion, but low enough to avoid droplet vaporization during planewave imaging. In order to determine the maximum allowable pressures, the imaging protocol described above was carried out at transducer voltages of 10, 15, 20, 25, and 30 V (the operating range specified in code for this ultrasound system) in the water bath containing DFB or OFP droplets. Average background brightness was measured before (baseline bubble presence) and after using low pressures (peak negative pressure: 0.38 MPa). The difference in the average brightness reflected the amount of vaporization triggered by the imaging step. Once established, the maximum voltages allowable for DFB and OFP were applied throughout each PCCA-specific image-activate-image protocol. Acoustic pressures associated with the selected voltages were measured in degassed water using a needle hydrophone (Onda HNA-0400, Sunnyvale, CA, USA). Since pressures generated in the acoustic field in response to a fixed voltage vary with depth, the values reported in this study are based on the depth of measurement. When insonation covered an area at multiple depths, a range of pressures is reported. The maximum free-field rarefactional or peak negative pressure (PNP) is located at the transducer’s elevational focus, a depth of 1.25 cm in these studies. However, in the animal, acoustic pressures diminish as a result of tissue attenuation. To determine the distribution of depth dependent acoustic pressures generated in the rat kidney, pressures were measured
at depths from 0.5 to 1.5 cm in 0.25 increments in the water bath. The free-field pressure waveforms (P) were derated (P0.3) by 0.3 dB/cm/MHz to correct for tissue attenuation and used to calculate the derated spatial peak intensity [ISP.3 = (P0.3)2/impedance] [32], using a standard renal tissue impedance of 1.65 MRayls [33]. Changes in the ISP.3 as a function of time through the duration of the pulse were averaged to find the derated spatial peak-temporal average (ISPTA.3) and spatial peak-pulse average (ISPPA.3) intensities. Similarly, the free-field PNP was derated (PNP0.3), and the derated mechanical index [MI = (PNP0.3 in MPa)/(frequency in MHz)1/2] calculated. The values for plane-wave insonation (the imaging steps of the protocol) were compared to the FDA guidelines for imaging small organs. The depth-dependent ISPTA.3 produced during focused-wave insonation (the activation step of the protocol) was also calculated, although it should be noted that the insonation time at each focal site was quite brief, lasting at most 3 ls. Previous in vitro work has demonstrated that droplet expansion takes approximately 2 ls, followed by oscillations as the bubble settles to its final resting size [28].
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2.4. Measuring focused ultrasound beam profiles
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The Onda HNA-0400 hydrophone was used to map pressures and determine focal zones in the water bath when focused ultra-
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Fig. 4. The effect of changing the sub-aperture position on focal pressures and vaporization responses (microbubble cloud size and shape) at different locations in the acoustic field. The free-field beam maps (A) and corresponding pulse inversion images of decafluorobutane (DFB) (B) and octafluoropropane (OFP) (C) microbubble clouds showed that shifting the sub-aperture corrected the lateral steering effects, although depth dependent changes in pressures and cloud geometry remained.
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sion imaging data and then counted all pixels with a brightness of at least 1% of this maximum. Pixel dimensions were set as 0.5 0.25 wavelengths of the receive frequency, resulting in a pixel area of 0.0035 mm2 at 9 MHz. The number of pixels was multiplied by the pixel size to calculate the total area. Differences in individual microbubble cloud geometries were assessed as a function of the location of the focus in the acoustic field, the size and position of the sub-aperture, and the number of cycles in the activation pulse.
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sound pulses were directed at different depths and lateral locations within a 4-cm2 area. A 3-axis motion stage (Newport XPS-RC, Irvine, CA, USA) stepped the transducer laterally and axially at 0.1 mm intervals to measure the PNP distributions around selected acoustic foci. The findings were combined to form composite beam maps and analyzed in Matlab to determine the sizes and shapes of the surrounding 3 dB to 12 dB focal zones. Beam maps were also obtained at the transducer’s elevational focus using 32, 64, 96, or 128 elements.
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2.5. Quantifying individual microbubble cloud sizes
2.6. Spacing between individual microbubble clouds
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The image-activate-image protocol was used to vaporize PFC droplets and visualize the resulting discrete microbubble clouds. The size of each cloud was defined as the area in the acoustic plane that contained echoes indicating the presence of bubbles. These areas were quantified using Matlab code that identified the brightest pixel in a selected region using the uncompressed pulse inver-
To determine the appropriate spacing between individual microbubble clouds, line testing was performed in which three vertical or two horizontal parallel lines of microbubble clouds were generated. The separation between these rows was gradually decreased from 2 mm in 0.25 mm increments. The optimum spacing between individual microbubble clouds was defined as the
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Fig. 5. Focal pressures and microbubble cloud geometries as a function of the sub-aperture size. Focused-wave ultrasound was applied using 32, 64, 96, or 128 center elements. The free-field beam maps (A) and focal zone plots (C) were compared to the corresponding pulse inversion decafluorobutane (DFB) microbubble cloud images (B). Note that the lower pressures reached using fewer elements were spread over a larger area, such that microbubble cloud geometries remained relatively stable when at least 64 elements were used.
Please cite this article in press as: C. Puett et al., Pulse sequences for uniform perfluorocarbon droplet vaporization and ultrasound imaging, Ultrasonics (2014), http://dx.doi.org/10.1016/j.ultras.2014.05.013
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The image-activate-image protocol was tested in the water bath by recording the time required to fill areas of 1, 2, 3, and 4 cm2 with bubbles.
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2.8. Contrast-enhanced renal imaging in the rat following the intravenous administration of phase-change contrast agents
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greatest distance at which the brightness plot profiles (determined by pixel assessment in Matlab as described above) for adjacent clouds showed overlap of their full-width half-maxima. Since focal pressures and vaporization responses are depth dependent, horizontal and vertical line tests were performed at three different depths within the acoustic field: 0.75, 1.25, and 1.75 cm.
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2.7. Generating uniform vaporization fields of microbubbles in vitro
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The results of measuring the geometry of individual microbubble clouds as a function of the location of the focus in the acoustic field, transducer sub-aperture position and size, and number of cycles during the vaporization of DFB and OFP droplets, as well as the results of line tests, were used to develop PCCA-specific activating pulse sequences. Focused pulse characteristics were chosen to maximize the size of individual microbubble clouds and space them appropriately in the selected region of interest as efficiently as possible. Activating pulses were separated by 300 ls and delivered left-to-right and deepest-to-most shallow to minimize shadowing as bubble clouds accumulate. The activation pulse sequences were linked to pre- and post-vaporization imaging sequences designed to maximize bubble contrast. Custom programs on the scanning system directed the function of a single linear array transducer through these image-activate-image steps.
The image-activate-image protocol was applied to the rat for PCCA-enhanced renal imaging. All in vivo studies were approved by the Institutional Animal Care and Use Committee at UNC Chapel Hill. Fischer rats weighing approximately 150 g were anesthetized using inhaled isoflurane vapor at 3% with oxygen. 60 lL of DFB or OFP stock solution (1010 droplets/mL) was diluted in 60 lL of saline and injected intravenously into the tail vein. The image-activate-image protocol was initiated 1.5 min after injection. The increase in signal intensity (in decibels) provided by the generated bubble field was calculated by measuring the difference in average pixel intensity within the user-selected region of interest before and after droplet activation. 2.9. Statistical analysis
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Reported means and corresponding standard deviations were based on sample sizes (N) of 3 animals and at least 5 in vitro acquisitions. One way analysis of variance (ANOVA) was used to compare the means when more than two sample groups were present in the data set. P values of 25 V were applied in the presence of DFB droplets (Fig. 2) and >20 V with OFP. These voltages corresponded with maximum free-field PNPs of 0.96 MPa and 0.78 MPa respectively, during plane-wave imaging. The rat kidney is located at a depth between 0.5 and 1.5 cm, and over this range, tissue attenuation resulted in
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Fig. 6. Microbubble cloud size as a function of the sub-aperture size employed during the activation pulse. Peak negative pressures (PNPs) were dependent on the number of active elements. Activation was achieved using 3.7–8.7 MPa for decafluorobutane (DFB) and 3.1–7.4 MPa for octafluoropropane (OFP). Microbubble clouds were significantly larger (p < 0.0001) when at least 64 elements were used in the presence of both DFB (A) and OFP (B), compared to 32 elements. Note that OFP clouds were more than twice as large as DFB clouds, even though the acoustic pressure was lower.
Fig. 7. Microbubble cloud size as a function of the number of cycles during the activation pulse. Focused ultrasound was applied at 2, 5, 10, or 15 cycles for decafluorobutane (DFB) (A) and octafluoropropane (OFP) (B). The peak negative pressures (PNPs) generated by 15 cycles were 8.7 MPa for DFB and 7.4 MPa for OFP. There was a statistically significant (p < 0.0001 by ANOVA) dose-responsive relationship between the cycle number and microbubble cloud size for DFB. In contrast, the largest OFP clouds resulted from 10 cycles, although sizes were similar over the range from 5 to 15 cycles.
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Fig. 8. Line testing to determine the optimal axial and lateral spacing between individual microbubble clouds. These representative pulse inversion images of horizontal decafluorobutane (DFB) (A) and vertical octafluoropropane (OFP) (C) cloud rows and their corresponding brightness profiles (B and D) were generated when focused ultrasound was delivered to a depth of 0.75 cm. Optimum spacing was considered the separation distance at which the brightness profile full-width half-maxima overlap. At this depth, the optimum axial and lateral spacing were 0.75 and 0.5 mm.
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depth-dependent derated mechanical indices and averaged spatial peak intensities (MI, ISPTA.3, and ISPPA.3) that met FDA guidelines during the plane-wave imaging steps of the protocol (Fig. 3). Although no guidelines exist for acceptable exposures during droplet activation, the maximum ISPTA.3 reached during the focused pulse sequences in these experiments [15 cycles using 128 elements at a PNP0.3 of 6.9 MPa (depth 1 cm)] was 28 W/cm2.
ture. Significant differences in the sizes of microbubble clouds also occurred as a function of the depth of the focus. For DFB, a microbubble cloud formed at a depth of 5 mm was approximately one quarter the size of a cloud formed at 1.5 cm, and this difference was more pronounced with OFP. With regard to shape, DFB microbubble clouds tended to be more compact and have more regular borders, compared to OFP clouds.
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3.2. Microbubble cloud geometry as a function of the location of the focal zone in the acoustic field
3.3. Microbubble cloud geometry as a function of the sub-aperture size employed during the activation pulse
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There were significant differences in the pressures achieved during focused ultrasound, depending on the location of the focus in relation to the transducer aperture. These differences resulted in focal zones of variable shape and size (Fig. 4A). As expected when using a linear array transducer, the highest pressures were achieved in the region near the elevational focus of the transducer at a depth of 1.25 cm. Steering the focus to target more lateral sites produced lower pressures and asymmetric focal zones. These pressure differences affected the size, shape, and brightness of the microbubble clouds generated at these sites (Fig. 4B and C). However, this variability could be minimized by shifting the sub-aper-
There were significant differences in the sizes of the acoustic focal zones and the pressures achieved within these zones when the size of the sub-aperture was changed (Fig. 5A and C). As expected, insonation using fewer elements resulted in the production of larger focal zones containing lower pressures. Interestingly, these differences tended to balance, so that the geometries of the microbubble clouds formed using 64, 96, and 128 elements varied only slightly. Fig. 5B includes representative pulse inversion images of microbubble clouds formed in the presence of DFB. The ability to generate microbubble clouds of relatively similar size using fewer elements (Fig. 6) suggested the opportunity to lower
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Fig. 9. Pulse inversion images of decafluorobutane (DFB, top row) and octafluoropropane (OFP, bottom row) vaporization fields of microbubbles throughout rectangular regions of interest in the water bath. The image-activate-image protocol generated these vaporization fields measuring 1, 2, 3, and 4 cm2 in 370, 440, 500, and 540 ms for DFB and 370, 430, 480, and 510 ms for OFP. 393 394 395 396 397 398 399 400 401
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the intensities (required acoustic energy) in the focal zones while maintaining microbubble cloud size. However, this trade-off had a lower limit. If only 32 elements were used, the OFP microbubble cloud was less than half the size of those formed in response to more elements (Fig. 6B), and almost no vaporization could be achieved in the presence of DFB with this low element number (Fig. 6A). In general, microbubble clouds formed from more volatile OFP droplets were approximately twice as large as DFB clouds, despite the use of a lower voltage (acoustic pressure).
3.4. Microbubble cloud size as a function of the number of cycles used for an activation pulse The size of the microbubble cloud tended to increase steadily (ANOVA p < 0.001) as a function of the number of cycles per pulse in the presence of DFB (Fig. 7A), increasing from 2.5 to 4.4 mm2 following the application 2–15 cycles. In the presence of OFP, the largest microbubble clouds (averaging 8 mm2) were measured following the application of 10 cycles per pulse (Fig. 7B). However, the cloud sizes were similar throughout the range from 5 to 15 cycles. Again, it was noted that microbubble clouds formed with more volatile OFP were larger than DFB clouds and tended to have more irregular borders.
3.5. Optimum spacing between microbubble clouds
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Line tests were used to determine the optimum axial and lateral spacing between adjacent microbubble clouds. Fig. 8 provides representative pulse inversion images of horizontal DFB (Fig. 8A) and vertical OFP (Fig. 8C) microbubble cloud lines and their corresponding brightness plot profiles (Fig. 8B and D) collected during testing at a depth of 0.75 cm. Spacing was found to depend on the depth of the focus in relation to the transducer aperture. Different spacings were needed to achieve full-width half-maxima overlap at depths of 0.75, 1.25, and 1.5 cm. Separating the depth of focus into three zones (0.5–1.0 cm, 1.0–1.5 cm, and 1.5–2.0 cm) was adequate to achieve relatively uniform spacing between microbubble cloud rows, and the imaging sequence therefore adjusted axial and lateral spacing based on the depth zone within which the focus was located. Sub-aperture shifting minimized the need to adjust spacing in response to lateral changes in the focal location at a given depth.
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3.6. Demonstration of the image-activate-image protocol in a water bath and its application in the rat for contrast-enhanced renal imaging
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The application of this image-activate-image pulse sequence was demonstrated in the water bath by activating droplets within
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Fig. 10. Representative images through the pulse sequence as applied in vivo to the rat kidney following the intravenous injection of decafluorobutane (DFB, top row) and octafluoropropane (OFP, bottom row) droplets. The green dots identify the calculated target locations for focused activation pulses. (For interpretation of the references to colour in this figure legend, the reader is referred to the web version of this article.)
Please cite this article in press as: C. Puett et al., Pulse sequences for uniform perfluorocarbon droplet vaporization and ultrasound imaging, Ultrasonics (2014), http://dx.doi.org/10.1016/j.ultras.2014.05.013
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selected regions of interest measuring 1, 2, 3, and 4 cm2. In their entirety, these image-activate-image protocols required 370, 440, 500, and 540 ms for DFB and 370, 430, 480, and 510 ms for OFP. The activation times were 310 ms and 280 ms respectively to fill a 4 cm2 region, with OFP requiring less time, since the individual OFP microbubble clouds were larger and therefore fewer activation sites were required. Fig. 9 includes representative pulse inversion images of vaporization fields generated from DFB and OFP droplets. Activation in these examples was accomplished using settings that produced the largest individual microbubble clouds for each agent. The pulse sequences were also successfully used in the rat to direct contrast imaging of the kidney after the intravenous administration of these PCCAs (Fig. 10). The kidney boundaries were identified by baseline imaging and resulted in the selection of regions of interest averaging 40 mm2. On average, the pulse sequences calculated 105 focal sites and filled these areas with bubbles in 130 ms (activation time). The average increase in signal intensity over baseline was 16 ± 3 dB for DFB and 20 ± 2 dB for OFP.
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Ongoing research in medical ultrasound includes advancing transducer technology, developing new transmit-receive protocols, and exploring new ultrasound contrast agents (UCAs). Currently available FDA-approved UCAs [octafluoropropane (OFP) microbubbles >1 lm in diameter] are limited by their short half-lives and inability to escape the vasculature. However, the discovery that more stable, sub-micron, liquid perfluorocarbon droplets can be vaporized by ultrasound to acoustically-active microbubbles (which also allows for targeted applications) has greatly expanded the potential for UCAs to enhance both diagnostic and therapeutic ultrasound. This study developed an efficient pulse sequence for PCCA experimentation using a single transducer. Our approach to applying PCCA enhancement over an extended spatial region involves an image-activate-image protocol, with each step requiring an adjustment in transducer operation. Although the imaging steps use standard techniques, many questions remain regarding the optimum transducer settings to accomplish the activation step. In this setting, optimum refers to trade-offs in the amount of delivered energy and time required to generate the bubble field, so that adequate contrast is achieved both safely and efficiently. Establishing a vaporization field quickly can be important for imaging. This is especially true if the goal is to collect information about the anatomy of the vascular network or perfusion through a circulatory bed where blood flow is present. The pulse sequence used in this study adjusts system settings that can be changed quickly, including the selection of sub-aperture size and position as well as pulse cycle number. These selections manipulate the geometries of individual microbubble clouds and space them appropriately. Adjustments can be made in real time, and testing has shown that the pulse sequence through this image-activateimage protocol generated a relatively uniform bubble field in the rat kidney in 130 ms, when the goal was to maximize contrast. In these experiments, an eightfold increase in contrast was achieved on average. However, for some applications, such as measuring vascular perfusion, minimizing the activation time may be more important than maximizing contrast. Modifications that shorten the activation sequence time (fewer pulse cycles and increased spacing between focal sites) can easily be made. As such, image-activate-image protocols with PCCAs could potentially provide information similar to the UCA bubble destruction-reperfusion protocols currently being tested to visualize circulation [34]. This is an area for future research.
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The acoustic energy exposures delivered by the imaging pulse sequences in this study meet FDA limits on exposure during diagnostic ultrasound for small organ imaging. In considering the potential risk for tissue injury during the activation step (focused ultrasound) of the protocol, it is important to note that each focal site is exposed to peak intensities for less than 3 ls, making risk of thermal injury low. Future experiments to measure tissue temperature changes throughout the image-activate-image protocol will be needed, perhaps using magnetic resonance thermometry. Additionally, given the risk of possible mechanical bioeffects, especially as the pulse cycle number increases, acoustic parameter optimization studies with post-imaging histology are planned. Although these pulse sequences are specific to this ultrasound system and to these PCCAs, the goals of this study are applicable to the use of PCCA-enhanced ultrasound in general. These goals included the ability to select a region of interest in real time, deliver focused waves to fill the targeted region with bubbles quickly, and link the activation step with transmit-receive imaging protocols that achieve maximum contrast from the bubble field, using a single commercial linear array transducer. Our laboratory is studying its application for imaging both healthy and diseased tissues, using different types and combinations of PCCAs. By increasing the amount of energy delivered during the focused activation step, this imaging sequence is easily converted to an imageablate-image protocol. As such, it would direct a PCCA-enhanced high intensity focused ultrasound (HIFU) sequence for therapeutic applications. Combining diagnostic and therapeutic steps in real time, using an agent that allows targeted intervention, is the goal of ‘‘theranostics,’’ and these protocols may provide useful tools for future theranostic experimentation.
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5. Conclusions
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At the current time, microbubble-based ultrasound contrast agents (UCAs) have a relatively limited role in clinical medicine. However, advances in transducer design, transmit and receive protocols, as well as the UCAs themselves are demonstrating a wide range of potential diagnostic and therapeutic applications for contrast-enhanced ultrasound. The discovery that acoustically-active microbubbles can be generated by vaporizing stable liquid droplets has opened the possibility that UCAs can be targeted. Targeting droplet vaporization (activation) and visualizing the resulting bubble field requires a combination of imaging and activating steps. This study examined factors influencing droplet activation and developed a pulse sequence that directs a single transducer through an image-activate-image protocol. Such pulse sequences are useful tools for the ongoing study of phase-change contrast agent-enhanced ultrasound and will be essential for the clinical application of this technology.
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Acknowledgements
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These studies were supported in part by a Summer Undergraduate Research Fellowship from UNC Chapel Hill (Connor Puett), the Focused Ultrasound Foundation, pilot funds from the National Science Foundation (DMR#1122483), and a Graduate Research Fellowship from the National Science Foundation (Paul S. Sheeran).
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References
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[18] L.C. Phillips, P.S. Sheeran, C. Puett, K.F. Timbie, R.J. Price, G.W. Miller, P.A. Dayton, Dual perfluorocarbon nanodroplets enhance high intensity focused ultrasound heating and extend therapeutic window in vivo, J. Acoust. Soc. Am. 134 (2013) 4049. [19] P.S. Sheeran, P.A. Dayton, Phase-change contrast agents for imaging and therapy, Curr. Pharm. Des. 18 (2012) 2152–2165. [20] O. Shpak, M. Verweij, H.J. Vos, N. de Jong, D. Lohse, M. Versluis, Acoustic droplet vaporization is initiated by superharmonic focusing, Proc. Natl. Acad. Sci. USA 111 (2014) 1697–1702. [21] K.C. Schad, K. Hynynen, In vitro characterization of perfluorocarbon droplets for focused ultrasound therapy, Phys. Med. Biol. 55 (2010) 4933–4947. [22] P.S. Sheeran, S. Luois, P.A. Dayton, T.O. Matsunaga, Formulation and acoustic studies of a new phase-shift agent for diagnostic and therapeutic ultrasound, Langmuir 27 (2011) 10412–10420. [23] P.S. Sheeran, S.H. Luois, L.B. Mullin, T.O. Matsunaga, P.A. Dayton, Design of ultrasonically-activatable nanoparticles using low boiling point perfluorocarbons, Biomaterials 33 (2012) 3262–3269. [24] R. Singh, G.A. Husseini, W.G. Pitt, Phase transitions of nanoemulsions using ultrasound: experimental observations, Ultrason. Sonochem. 19 (2012) 1120– 1125. [25] O.D. Kripfgans, J.B. Fowlkes, D.L. Miller, O.P. Eldevik, P.L. Carson, Acoustic droplet vaporization for therapeutic and diagnostic applications, Ultrasound Med. Biol. 26 (2000) 1177–1189. [26] A.E. Forbrich, T.J. Harrison, R. Paproski, R.J. Zemp. Realtime flash difference ultrasound imaging of phase-change perfluorocarbon nanodroplet activation. Ultrasonics Symposium, IEEE International 2012, pp. 687–690. [27] G. Montaldo, M. Tanter, J. Bercoff, N. Benech, M. Fink, Coherent plane-wave compounding for very high frame rate ultrasonography and transient elastography, IEEE Trans. Ultrason. Ferroelectr. Freq. Control 56 (2009) 489– 506. [28] P.S. Sheeran, T.O. Matsunaga, P.A. Dayton, Phase change events of volatile liquid perfluorocarbon contrast agents produce unique acoustic signatures, Phys. Med. Biol. 59 (2013) 379–401. [29] C.C. Shen, Y.H. Chou, P.C. Li, Pulse inversion techniques in ultrasonic nonlinear imaging, J. Med. Ultrasound 13 (2005) 3–17. [30] D.H. Simpson, C.T. Chin, P.N. Burns, Pulse inversion Doppler: a new method for detecting nonlinear echoes from microbubble contrast agents, IEEE Trans. Ultrason. Ferroelectr. Freq. Control 46 (1999) 372–382. [31] R.E. Daigle, U.S. Patent No. 8,287,456. U.S. Patent and Trademark Office, Washington, DC, 2012. [32] V. Bull, G.R. ter Haar, The physics of ultrasound energy sources, in: E.G. Moros (Ed.), Physics of Thermal Therapy: Fundamentals and Clinical Applications, CRC Press, Boca Raton, FL, 2013, pp. 75–94. [33] H. Azhari, Basics of Biomedical Ultrasound for Engineers, Wiley-IEEE Press, New Jersey, 2010. [34] P. Kogan, K.A. Johnson, S. Feingold, N. Garrett, I. Guracar, W.J. Arendshorst, P.A. Dayton, Validation of dynamic contrast enhanced ultrasound in rodent kidneys as an absolute quantitative method for measuring blood perfusion, Ultrasound Med. Biol. 37 (2011) 900–908.
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Pulse sequences for uniform perfluorocarbon droplet vaporization and ultrasound imaging
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Joint Department of Biomedical Engineering, University of North Carolina at Chapel Hill and North Carolina State University, CB 7575, Chapel Hill, NC, USA
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C. Puett, P.S. Sheeran, J.D. Rojas, P.A. Dayton ⇑
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Article history: Received 22 April 2014 Accepted 19 May 2014 Available online xxxx
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Keywords: Phase-change Ultrasound contrast agent Perfluorocarbon Vaporization Decafluorobutane Octafluoropropane
a b s t r a c t Phase-change contrast agents (PCCAs) consist of liquid perfluorocarbon droplets that can be vaporized into gas-filled microbubbles by pulsed ultrasound waves at diagnostic pressures and frequencies. These activatable contrast agents provide benefits of longer circulating times and smaller sizes relative to conventional microbubble contrast agents. However, optimizing ultrasound-induced activation of these agents requires coordinated pulse sequences not found on current clinical systems, in order to both initiate droplet vaporization and image the resulting microbubble population. Specifically, the activation process must provide a spatially uniform distribution of microbubbles and needs to occur quickly enough to image the vaporized agents before they migrate out of the imaging field of view. The development and evaluation of protocols for PCCA-enhanced ultrasound imaging using a commercial array transducer are described. The developed pulse sequences consist of three states: (1) initial imaging at sub-activation pressures, (2) activating droplets within a selected region of interest, and (3) imaging the resulting microbubbles. Bubble clouds produced by the vaporization of decafluorobutane and octafluoropropane droplets were characterized as a function of focused pulse parameters and acoustic field location. Pulse sequences were designed to manipulate the geometries of discrete microbubble clouds using electronic steering, and cloud spacing was tailored to build a uniform vaporization field. The complete pulse sequence was demonstrated in the water bath and then in vivo in a rodent kidney. The resulting contrast provided a significant increase (>15 dB) in signal intensity. Ó 2014 Elsevier B.V. All rights reserved.
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1. Introduction
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Micron-sized encapsulated bubbles of gas (microbubbles) are used clinically as ultrasound contrast agents (UCAs). OctafluoroÒ propane (OFP) microbubble contrast agents having lipid (Definity ) Ò and protein (Optison ) shells are approved by the Food and Drug Administration (FDA) in the United States as a diagnostic tool for opacifying the left ventricle and improving endocardial border delineation during echocardiography [1,2]. Additional diagnostic applications have been demonstrated utilizing the unique nonlinear acoustic signatures of oscillating microbubbles [3,4]. Experimental therapeutic applications of microbubbles include enhancement of tissue ablation due to microbubble cavitation in high intensity focused ultrasound (HIFU) [5–7], increased apoptosis following radiation therapy [8], and utilization in drug delivery [9]. However, there are limitations to the direct application of currently available UCAs. Microbubbles survive for just minutes
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⇑ Corresponding author. Tel.: +1 919 843 9521. E-mail address: [email protected] (P.A. Dayton).
in vivo [10], evaporating as they circulate through the lungs and being cleared by the reticuloendothelial system [11]. Additionally, their large size (mean diameter >1 lm) constrains them to the vasculature [12]. Sub-micron liquid perfluorocarbon (PFC) agents provide an attractive tool to introduce microbubbles into a region of interest. Droplets can be manufactured small enough to extravasate through the leaky vasculature found in malignant and inflamed tissues and can therefore enter diseased interstitial spaces [13–17]. Also, droplets can be designed to have much longer half-lives than microbubbles in vivo and yet can still be vaporized back into microbubbles by the temperature and pressure changes achievable with ultrasound [14,16,18] As such, they are referred to as phase-change contrast agents (PCCAs), and their vaporization (termed ‘‘activation’’) allows microbubbles to be generated within the acoustic field [19]. Manipulating the critical activation step from droplet to microbubble is an area of active research. Vaporization is a complex phenomenon [20] depending on many factors, including the properties of the ultrasound wave as well as the size and composition of the
http://dx.doi.org/10.1016/j.ultras.2014.05.013 0041-624X/Ó 2014 Elsevier B.V. All rights reserved.
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PFC droplets [21–24]. As expected, droplets containing more volatile PFCs, such as decafluorobutane (DFB, C4F10, boiling point: 2 °C) and octafluoropropane (OFP, C3F8, boiling point: 37 °C), are easier to vaporize than those containing PFCs with higher boiling points [23]. These volatile PFC agents are well suited for diagnostic applications, since they can be vaporized to microbubbles with acoustic pressures used for imaging. However, PCCA-enhanced diagnostic ultrasound is still in its infancy, typically requiring separate transducers for activation and imaging due to the demanding pulse sequence requirements [17,25]. Herein, we present a unique ‘‘image-activate-image’’ protocol that demonstrates the feasibility of performing initial sub-activation threshold imaging, activation, and post-activation imaging with a single transducer. In this protocol, the initial plane-wave imaging step (both anatomical B-mode and pulse inversion for baseline bubble assessment) is used to provide the operator with information to determine the desired spatial activation pattern and is followed by a user-initiated activation step, during which electronically-steered, focused-wave pulses vaporize PFC droplets with high spatial specificity. A final plane-wave pulse inversion imaging step visualizes the vaporization field of microbubbles. Although the imaging sequences used in this study to maximize bubble detection and minimize artifacts (pulse inversion and coherent plane-wave compounding) are well described [26–30], much of the challenge of PCCA-enhanced diagnostic ultrasound lies in optimizing the activation step of the protocol. The goal is to achieve a relatively uniform vaporization field of microbubbles throughout the selected target area, utilizing acoustic energy exposures that carry a low risk of mechanical or thermal tissue injury. Since the geometry of the focal zone and the intensity achieved within the zone vary as a function of the location of the focus relative to the transducer, vaporization responses also vary. Therefore, uniform vaporization requires anticipating differences in the sizes and shapes of microbubble clouds, adjusting acoustic parameters to account for these differences, and then spacing individual clouds appropriately. In this study, testing and development were done using sub-micron DFB and OFP droplets dispersed in a degassed water bath, and the completed pulse sequence was then applied in vivo to image the rat kidney following the intravenous administration of these PCCAs. Image-activate-image protocols that direct the generation of uniform vaporization fields are useful tools for the ongoing study of PCCAs and will be critical in the eventual clinical application of these agents.
2. Methods
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2.1. Preparation of the phase-change contrast agents
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PCCAs used in this work consist of droplets of liquid PFC [either decafluorobutane (DFB, C4F10) or octafluoropropane (OFP, C3F8)] contained in a phospholipid and polyethylene glycol shell. They were prepared by the condensation procedure described previously [22]. Briefly, lipids [1,2-distearoyl-sn-glycero-3-phosphocholine (DSPC) and 1,2-distearoyl-sn-glycero-3-phosphoethanolamine-N-methoxy (polyethylene–glycol)-2000 (DSPE-PEG2000), purchased from Avanti Polar Lipids (Alabaster, AL) in a 9:1 M ratio and a total lipid concentration of 1.0 mg/mL] were emulsified in a solution of propylene glycol 15% (v/v), glycerol 5% (v/v), and phosphate-buffered saline (PBS) 80% (v/v). 1.5 mL of this emulsion was placed in a sealed 3 mL glass vial and the air headspace exchanged with PFC gas purchased from Fluoromed (Round Rock, TX, USA). Microbubbles with PFC cores and phospholipid shells formed spontaneously when agitated in a Vialmix Shaker (Bristol–Myers– Squibb, New York, NY, USA). The concentration of microbubbles in this emulsion was measured via optical techniques (Accusizer 780, Particle Sizing Systems, Santa Barbara, CA, USA) and was found to be 1010 microbubbles/mL. The microbubbles were condensed to droplets by cooling the emulsion under an adjustable high-pressure air source in an isopropanol bath maintained between 6 and 10 °C using dry ice. Dynamic light scattering (Malvern Nano ZS, Malvern Instruments, Ltd., Westborough, MA, USA) demonstrated a polydisperse population with a mean diameter of 240 ± 65 nm. Assuming a 1:1 conversion from microbubble to droplet, a stock solution containing 1010 PFC droplets/mL was used immediately or stored frozen as a stock solution for use within 24 h.
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2.2. In Vitro setup
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In vitro testing was done in an acrylic-walled tank containing 4.7 L of degassed water. DFB or OFP droplets were added to the water to achieve a final concentration of approximately 106 droplets/mL. For studies using DFB droplets, the water was maintained at 37 °C by circulating warm water through copper tubing in the bath. Because OFP droplets are relatively unstable at 37 °C [23], testing with OFP was done at room temperature (22 °C) to ensure an adequate half-life for experimentation. The water was stirred between experimental runs to redistribute PFC droplets, and the
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Fig. 1. The image-activate-image protocol for phase-change contrast agent (PCCA)-enhanced diagnostic ultrasound. This flow diagram identifies the tasks accomplished through each step of the protocol as well as the time required to complete the pulse sequences for each step.
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transducer was angled to reduce direct ultrasound reflections off the tank walls.
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2.3. Pulse sequence design
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Our approach to PCCA-enhanced diagnostic ultrasound involves a unique ‘‘image-activate-image’’ pulse sequence (Fig. 1). A programmable research ultrasound system (Verasonics, Redmond, WA) driving an ATL L12-5 38 mm 192-element linear array transducer was used to implement this pulse sequence. The initial imaging step included seven transmissions of steered plane waves (angles: 18°, 12°, 6°, 0°, 6°, 12°, 18°). 128 elements (the maximum number of connector channels) were used to transmit single-cycle pulses at 9 MHz. Echoes were received at the same frequency, bandpass filtered using a 21 tap finite impulse response filter with a 4–14 MHz bandwidth, and averaged using built-in Verasonics reconstruction processes. Pixel-based algorithms displayed a B-mode image of the anatomy [31]. This image allowed for selection of a region of interest. Next, successive pulse inversion plane-wave transmissions were delivered at seven angles as previously described. At each angle, two 1-cycle pulses at 4.5 MHz were delivered 180° out of phase and the received signals were summed to isolate the nonlinear microbubble oscillation. The echoes received from different transmission angles were averaged to obtain a coherent plane-wave compounded image, providing a background (baseline) microbubble assessment. Electronically focused ultrasound pulses were transmitted during the activation step to vaporize droplets and generate microbubble clouds. In order to image activated bubbles before significant motion occurs, the entire activation process needs to be accomplished quickly. Therefore, this activation sequence directed changes to transducer function that require little time, including sub-aperture adjustments (the position and number of active elements) and the number of cycles in each activation pulse. To manipulate individual microbubble cloud geometries, focused pulses were delivered at 5 MHz using 32, 64, 96, or 128 elements and included 2, 5, 10, or 15 cycles per pulse, while the sub-aperture was shifted to minimize the steering required to target focal sites. The final imaging step visualized the microbubble clouds and involved repeating the baseline seven-angle pulse inversion transmissions at 4.5 MHz and averaging the received echoes to form a coherent plane wave-compounded image. These reconstructed
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Fig. 2. Measurement of droplet activation (vaporization) as a function of the voltage applied during the imaging steps of the image-activate-image protocol. Insonation involved seven plane-wave transmissions using coherent plane wave compounding with pulse inversion applied over a voltage range from 10 to 30 V. This voltage range corresponds to free-field peak negative pressures (PNPs) from 0.38 to 1.13 MPa. Non-targeted activation was measured as the difference in the average background brightness before and after the imaging step being tested. A significant (p < 0.0001) increase in non-targeted decafluorobutane (DFB) droplet vaporization occurred when the imaging step used >25 V, which equated to a PNP of 0.96 MPa.
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images were compared to those obtained prior to activation. The custom image-activate-image sequence was implemented in Matlab (The Mathworks, Natick, MA) on the scanning system. Changing voltage in real time during the image-activate-image protocol is problematic, since it is a relatively slow process, requiring a delay of at least 1 ms due to system limitations. Therefore, testing was done to determine the optimum voltage for use throughout the protocol. Selecting this voltage required finding a balance between pressures high enough for focused-wave activa-
Fig. 3. Energy deposition as a function of depth in the acoustic field. Free-field pressures during plane-wave imaging were measured at depths from 0.5 to 1.5 cm, and then derated to calculate the mechanical index (MI) (A), spatial peak-pulse average intensity (ISPPA.3) (B), and spatial peak-temporal average intensity (ISPTA.3) (C) resulting from 20 V (gray square), used for OFP studies, and 25 V (black diamond), used for DFB studies. The derated values were compared to the FDA thresholds (dashed lines) for small organ imaging (MI: 1.9, ISPPA.3: 190 W/cm2, ISPTA.3: 94 mW/cm2). The Definity suggested MI threshold of 0.8 [Definity Imaging] when bubbles are present in the acoustic field (dotted line) is indicated in (A). The depth-dependent ISPTA.3 during focused-wave activation (D) is also shown.
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tion, but low enough to avoid droplet vaporization during planewave imaging. In order to determine the maximum allowable pressures, the imaging protocol described above was carried out at transducer voltages of 10, 15, 20, 25, and 30 V (the operating range specified in code for this ultrasound system) in the water bath containing DFB or OFP droplets. Average background brightness was measured before (baseline bubble presence) and after using low pressures (peak negative pressure: 0.38 MPa). The difference in the average brightness reflected the amount of vaporization triggered by the imaging step. Once established, the maximum voltages allowable for DFB and OFP were applied throughout each PCCA-specific image-activate-image protocol. Acoustic pressures associated with the selected voltages were measured in degassed water using a needle hydrophone (Onda HNA-0400, Sunnyvale, CA, USA). Since pressures generated in the acoustic field in response to a fixed voltage vary with depth, the values reported in this study are based on the depth of measurement. When insonation covered an area at multiple depths, a range of pressures is reported. The maximum free-field rarefactional or peak negative pressure (PNP) is located at the transducer’s elevational focus, a depth of 1.25 cm in these studies. However, in the animal, acoustic pressures diminish as a result of tissue attenuation. To determine the distribution of depth dependent acoustic pressures generated in the rat kidney, pressures were measured
at depths from 0.5 to 1.5 cm in 0.25 increments in the water bath. The free-field pressure waveforms (P) were derated (P0.3) by 0.3 dB/cm/MHz to correct for tissue attenuation and used to calculate the derated spatial peak intensity [ISP.3 = (P0.3)2/impedance] [32], using a standard renal tissue impedance of 1.65 MRayls [33]. Changes in the ISP.3 as a function of time through the duration of the pulse were averaged to find the derated spatial peak-temporal average (ISPTA.3) and spatial peak-pulse average (ISPPA.3) intensities. Similarly, the free-field PNP was derated (PNP0.3), and the derated mechanical index [MI = (PNP0.3 in MPa)/(frequency in MHz)1/2] calculated. The values for plane-wave insonation (the imaging steps of the protocol) were compared to the FDA guidelines for imaging small organs. The depth-dependent ISPTA.3 produced during focused-wave insonation (the activation step of the protocol) was also calculated, although it should be noted that the insonation time at each focal site was quite brief, lasting at most 3 ls. Previous in vitro work has demonstrated that droplet expansion takes approximately 2 ls, followed by oscillations as the bubble settles to its final resting size [28].
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The Onda HNA-0400 hydrophone was used to map pressures and determine focal zones in the water bath when focused ultra-
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Fig. 4. The effect of changing the sub-aperture position on focal pressures and vaporization responses (microbubble cloud size and shape) at different locations in the acoustic field. The free-field beam maps (A) and corresponding pulse inversion images of decafluorobutane (DFB) (B) and octafluoropropane (OFP) (C) microbubble clouds showed that shifting the sub-aperture corrected the lateral steering effects, although depth dependent changes in pressures and cloud geometry remained.
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sion imaging data and then counted all pixels with a brightness of at least 1% of this maximum. Pixel dimensions were set as 0.5 0.25 wavelengths of the receive frequency, resulting in a pixel area of 0.0035 mm2 at 9 MHz. The number of pixels was multiplied by the pixel size to calculate the total area. Differences in individual microbubble cloud geometries were assessed as a function of the location of the focus in the acoustic field, the size and position of the sub-aperture, and the number of cycles in the activation pulse.
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sound pulses were directed at different depths and lateral locations within a 4-cm2 area. A 3-axis motion stage (Newport XPS-RC, Irvine, CA, USA) stepped the transducer laterally and axially at 0.1 mm intervals to measure the PNP distributions around selected acoustic foci. The findings were combined to form composite beam maps and analyzed in Matlab to determine the sizes and shapes of the surrounding 3 dB to 12 dB focal zones. Beam maps were also obtained at the transducer’s elevational focus using 32, 64, 96, or 128 elements.
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2.6. Spacing between individual microbubble clouds
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The image-activate-image protocol was used to vaporize PFC droplets and visualize the resulting discrete microbubble clouds. The size of each cloud was defined as the area in the acoustic plane that contained echoes indicating the presence of bubbles. These areas were quantified using Matlab code that identified the brightest pixel in a selected region using the uncompressed pulse inver-
To determine the appropriate spacing between individual microbubble clouds, line testing was performed in which three vertical or two horizontal parallel lines of microbubble clouds were generated. The separation between these rows was gradually decreased from 2 mm in 0.25 mm increments. The optimum spacing between individual microbubble clouds was defined as the
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Fig. 5. Focal pressures and microbubble cloud geometries as a function of the sub-aperture size. Focused-wave ultrasound was applied using 32, 64, 96, or 128 center elements. The free-field beam maps (A) and focal zone plots (C) were compared to the corresponding pulse inversion decafluorobutane (DFB) microbubble cloud images (B). Note that the lower pressures reached using fewer elements were spread over a larger area, such that microbubble cloud geometries remained relatively stable when at least 64 elements were used.
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The image-activate-image protocol was tested in the water bath by recording the time required to fill areas of 1, 2, 3, and 4 cm2 with bubbles.
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2.8. Contrast-enhanced renal imaging in the rat following the intravenous administration of phase-change contrast agents
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greatest distance at which the brightness plot profiles (determined by pixel assessment in Matlab as described above) for adjacent clouds showed overlap of their full-width half-maxima. Since focal pressures and vaporization responses are depth dependent, horizontal and vertical line tests were performed at three different depths within the acoustic field: 0.75, 1.25, and 1.75 cm.
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The results of measuring the geometry of individual microbubble clouds as a function of the location of the focus in the acoustic field, transducer sub-aperture position and size, and number of cycles during the vaporization of DFB and OFP droplets, as well as the results of line tests, were used to develop PCCA-specific activating pulse sequences. Focused pulse characteristics were chosen to maximize the size of individual microbubble clouds and space them appropriately in the selected region of interest as efficiently as possible. Activating pulses were separated by 300 ls and delivered left-to-right and deepest-to-most shallow to minimize shadowing as bubble clouds accumulate. The activation pulse sequences were linked to pre- and post-vaporization imaging sequences designed to maximize bubble contrast. Custom programs on the scanning system directed the function of a single linear array transducer through these image-activate-image steps.
The image-activate-image protocol was applied to the rat for PCCA-enhanced renal imaging. All in vivo studies were approved by the Institutional Animal Care and Use Committee at UNC Chapel Hill. Fischer rats weighing approximately 150 g were anesthetized using inhaled isoflurane vapor at 3% with oxygen. 60 lL of DFB or OFP stock solution (1010 droplets/mL) was diluted in 60 lL of saline and injected intravenously into the tail vein. The image-activate-image protocol was initiated 1.5 min after injection. The increase in signal intensity (in decibels) provided by the generated bubble field was calculated by measuring the difference in average pixel intensity within the user-selected region of interest before and after droplet activation. 2.9. Statistical analysis
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Reported means and corresponding standard deviations were based on sample sizes (N) of 3 animals and at least 5 in vitro acquisitions. One way analysis of variance (ANOVA) was used to compare the means when more than two sample groups were present in the data set. P values of 25 V were applied in the presence of DFB droplets (Fig. 2) and >20 V with OFP. These voltages corresponded with maximum free-field PNPs of 0.96 MPa and 0.78 MPa respectively, during plane-wave imaging. The rat kidney is located at a depth between 0.5 and 1.5 cm, and over this range, tissue attenuation resulted in
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Fig. 6. Microbubble cloud size as a function of the sub-aperture size employed during the activation pulse. Peak negative pressures (PNPs) were dependent on the number of active elements. Activation was achieved using 3.7–8.7 MPa for decafluorobutane (DFB) and 3.1–7.4 MPa for octafluoropropane (OFP). Microbubble clouds were significantly larger (p < 0.0001) when at least 64 elements were used in the presence of both DFB (A) and OFP (B), compared to 32 elements. Note that OFP clouds were more than twice as large as DFB clouds, even though the acoustic pressure was lower.
Fig. 7. Microbubble cloud size as a function of the number of cycles during the activation pulse. Focused ultrasound was applied at 2, 5, 10, or 15 cycles for decafluorobutane (DFB) (A) and octafluoropropane (OFP) (B). The peak negative pressures (PNPs) generated by 15 cycles were 8.7 MPa for DFB and 7.4 MPa for OFP. There was a statistically significant (p < 0.0001 by ANOVA) dose-responsive relationship between the cycle number and microbubble cloud size for DFB. In contrast, the largest OFP clouds resulted from 10 cycles, although sizes were similar over the range from 5 to 15 cycles.
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Fig. 8. Line testing to determine the optimal axial and lateral spacing between individual microbubble clouds. These representative pulse inversion images of horizontal decafluorobutane (DFB) (A) and vertical octafluoropropane (OFP) (C) cloud rows and their corresponding brightness profiles (B and D) were generated when focused ultrasound was delivered to a depth of 0.75 cm. Optimum spacing was considered the separation distance at which the brightness profile full-width half-maxima overlap. At this depth, the optimum axial and lateral spacing were 0.75 and 0.5 mm.
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depth-dependent derated mechanical indices and averaged spatial peak intensities (MI, ISPTA.3, and ISPPA.3) that met FDA guidelines during the plane-wave imaging steps of the protocol (Fig. 3). Although no guidelines exist for acceptable exposures during droplet activation, the maximum ISPTA.3 reached during the focused pulse sequences in these experiments [15 cycles using 128 elements at a PNP0.3 of 6.9 MPa (depth 1 cm)] was 28 W/cm2.
ture. Significant differences in the sizes of microbubble clouds also occurred as a function of the depth of the focus. For DFB, a microbubble cloud formed at a depth of 5 mm was approximately one quarter the size of a cloud formed at 1.5 cm, and this difference was more pronounced with OFP. With regard to shape, DFB microbubble clouds tended to be more compact and have more regular borders, compared to OFP clouds.
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3.2. Microbubble cloud geometry as a function of the location of the focal zone in the acoustic field
3.3. Microbubble cloud geometry as a function of the sub-aperture size employed during the activation pulse
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There were significant differences in the pressures achieved during focused ultrasound, depending on the location of the focus in relation to the transducer aperture. These differences resulted in focal zones of variable shape and size (Fig. 4A). As expected when using a linear array transducer, the highest pressures were achieved in the region near the elevational focus of the transducer at a depth of 1.25 cm. Steering the focus to target more lateral sites produced lower pressures and asymmetric focal zones. These pressure differences affected the size, shape, and brightness of the microbubble clouds generated at these sites (Fig. 4B and C). However, this variability could be minimized by shifting the sub-aper-
There were significant differences in the sizes of the acoustic focal zones and the pressures achieved within these zones when the size of the sub-aperture was changed (Fig. 5A and C). As expected, insonation using fewer elements resulted in the production of larger focal zones containing lower pressures. Interestingly, these differences tended to balance, so that the geometries of the microbubble clouds formed using 64, 96, and 128 elements varied only slightly. Fig. 5B includes representative pulse inversion images of microbubble clouds formed in the presence of DFB. The ability to generate microbubble clouds of relatively similar size using fewer elements (Fig. 6) suggested the opportunity to lower
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Fig. 9. Pulse inversion images of decafluorobutane (DFB, top row) and octafluoropropane (OFP, bottom row) vaporization fields of microbubbles throughout rectangular regions of interest in the water bath. The image-activate-image protocol generated these vaporization fields measuring 1, 2, 3, and 4 cm2 in 370, 440, 500, and 540 ms for DFB and 370, 430, 480, and 510 ms for OFP. 393 394 395 396 397 398 399 400 401
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the intensities (required acoustic energy) in the focal zones while maintaining microbubble cloud size. However, this trade-off had a lower limit. If only 32 elements were used, the OFP microbubble cloud was less than half the size of those formed in response to more elements (Fig. 6B), and almost no vaporization could be achieved in the presence of DFB with this low element number (Fig. 6A). In general, microbubble clouds formed from more volatile OFP droplets were approximately twice as large as DFB clouds, despite the use of a lower voltage (acoustic pressure).
3.4. Microbubble cloud size as a function of the number of cycles used for an activation pulse The size of the microbubble cloud tended to increase steadily (ANOVA p < 0.001) as a function of the number of cycles per pulse in the presence of DFB (Fig. 7A), increasing from 2.5 to 4.4 mm2 following the application 2–15 cycles. In the presence of OFP, the largest microbubble clouds (averaging 8 mm2) were measured following the application of 10 cycles per pulse (Fig. 7B). However, the cloud sizes were similar throughout the range from 5 to 15 cycles. Again, it was noted that microbubble clouds formed with more volatile OFP were larger than DFB clouds and tended to have more irregular borders.
3.5. Optimum spacing between microbubble clouds
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Line tests were used to determine the optimum axial and lateral spacing between adjacent microbubble clouds. Fig. 8 provides representative pulse inversion images of horizontal DFB (Fig. 8A) and vertical OFP (Fig. 8C) microbubble cloud lines and their corresponding brightness plot profiles (Fig. 8B and D) collected during testing at a depth of 0.75 cm. Spacing was found to depend on the depth of the focus in relation to the transducer aperture. Different spacings were needed to achieve full-width half-maxima overlap at depths of 0.75, 1.25, and 1.5 cm. Separating the depth of focus into three zones (0.5–1.0 cm, 1.0–1.5 cm, and 1.5–2.0 cm) was adequate to achieve relatively uniform spacing between microbubble cloud rows, and the imaging sequence therefore adjusted axial and lateral spacing based on the depth zone within which the focus was located. Sub-aperture shifting minimized the need to adjust spacing in response to lateral changes in the focal location at a given depth.
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3.6. Demonstration of the image-activate-image protocol in a water bath and its application in the rat for contrast-enhanced renal imaging
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The application of this image-activate-image pulse sequence was demonstrated in the water bath by activating droplets within
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Fig. 10. Representative images through the pulse sequence as applied in vivo to the rat kidney following the intravenous injection of decafluorobutane (DFB, top row) and octafluoropropane (OFP, bottom row) droplets. The green dots identify the calculated target locations for focused activation pulses. (For interpretation of the references to colour in this figure legend, the reader is referred to the web version of this article.)
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selected regions of interest measuring 1, 2, 3, and 4 cm2. In their entirety, these image-activate-image protocols required 370, 440, 500, and 540 ms for DFB and 370, 430, 480, and 510 ms for OFP. The activation times were 310 ms and 280 ms respectively to fill a 4 cm2 region, with OFP requiring less time, since the individual OFP microbubble clouds were larger and therefore fewer activation sites were required. Fig. 9 includes representative pulse inversion images of vaporization fields generated from DFB and OFP droplets. Activation in these examples was accomplished using settings that produced the largest individual microbubble clouds for each agent. The pulse sequences were also successfully used in the rat to direct contrast imaging of the kidney after the intravenous administration of these PCCAs (Fig. 10). The kidney boundaries were identified by baseline imaging and resulted in the selection of regions of interest averaging 40 mm2. On average, the pulse sequences calculated 105 focal sites and filled these areas with bubbles in 130 ms (activation time). The average increase in signal intensity over baseline was 16 ± 3 dB for DFB and 20 ± 2 dB for OFP.
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Ongoing research in medical ultrasound includes advancing transducer technology, developing new transmit-receive protocols, and exploring new ultrasound contrast agents (UCAs). Currently available FDA-approved UCAs [octafluoropropane (OFP) microbubbles >1 lm in diameter] are limited by their short half-lives and inability to escape the vasculature. However, the discovery that more stable, sub-micron, liquid perfluorocarbon droplets can be vaporized by ultrasound to acoustically-active microbubbles (which also allows for targeted applications) has greatly expanded the potential for UCAs to enhance both diagnostic and therapeutic ultrasound. This study developed an efficient pulse sequence for PCCA experimentation using a single transducer. Our approach to applying PCCA enhancement over an extended spatial region involves an image-activate-image protocol, with each step requiring an adjustment in transducer operation. Although the imaging steps use standard techniques, many questions remain regarding the optimum transducer settings to accomplish the activation step. In this setting, optimum refers to trade-offs in the amount of delivered energy and time required to generate the bubble field, so that adequate contrast is achieved both safely and efficiently. Establishing a vaporization field quickly can be important for imaging. This is especially true if the goal is to collect information about the anatomy of the vascular network or perfusion through a circulatory bed where blood flow is present. The pulse sequence used in this study adjusts system settings that can be changed quickly, including the selection of sub-aperture size and position as well as pulse cycle number. These selections manipulate the geometries of individual microbubble clouds and space them appropriately. Adjustments can be made in real time, and testing has shown that the pulse sequence through this image-activateimage protocol generated a relatively uniform bubble field in the rat kidney in 130 ms, when the goal was to maximize contrast. In these experiments, an eightfold increase in contrast was achieved on average. However, for some applications, such as measuring vascular perfusion, minimizing the activation time may be more important than maximizing contrast. Modifications that shorten the activation sequence time (fewer pulse cycles and increased spacing between focal sites) can easily be made. As such, image-activate-image protocols with PCCAs could potentially provide information similar to the UCA bubble destruction-reperfusion protocols currently being tested to visualize circulation [34]. This is an area for future research.
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The acoustic energy exposures delivered by the imaging pulse sequences in this study meet FDA limits on exposure during diagnostic ultrasound for small organ imaging. In considering the potential risk for tissue injury during the activation step (focused ultrasound) of the protocol, it is important to note that each focal site is exposed to peak intensities for less than 3 ls, making risk of thermal injury low. Future experiments to measure tissue temperature changes throughout the image-activate-image protocol will be needed, perhaps using magnetic resonance thermometry. Additionally, given the risk of possible mechanical bioeffects, especially as the pulse cycle number increases, acoustic parameter optimization studies with post-imaging histology are planned. Although these pulse sequences are specific to this ultrasound system and to these PCCAs, the goals of this study are applicable to the use of PCCA-enhanced ultrasound in general. These goals included the ability to select a region of interest in real time, deliver focused waves to fill the targeted region with bubbles quickly, and link the activation step with transmit-receive imaging protocols that achieve maximum contrast from the bubble field, using a single commercial linear array transducer. Our laboratory is studying its application for imaging both healthy and diseased tissues, using different types and combinations of PCCAs. By increasing the amount of energy delivered during the focused activation step, this imaging sequence is easily converted to an imageablate-image protocol. As such, it would direct a PCCA-enhanced high intensity focused ultrasound (HIFU) sequence for therapeutic applications. Combining diagnostic and therapeutic steps in real time, using an agent that allows targeted intervention, is the goal of ‘‘theranostics,’’ and these protocols may provide useful tools for future theranostic experimentation.
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At the current time, microbubble-based ultrasound contrast agents (UCAs) have a relatively limited role in clinical medicine. However, advances in transducer design, transmit and receive protocols, as well as the UCAs themselves are demonstrating a wide range of potential diagnostic and therapeutic applications for contrast-enhanced ultrasound. The discovery that acoustically-active microbubbles can be generated by vaporizing stable liquid droplets has opened the possibility that UCAs can be targeted. Targeting droplet vaporization (activation) and visualizing the resulting bubble field requires a combination of imaging and activating steps. This study examined factors influencing droplet activation and developed a pulse sequence that directs a single transducer through an image-activate-image protocol. Such pulse sequences are useful tools for the ongoing study of phase-change contrast agent-enhanced ultrasound and will be essential for the clinical application of this technology.
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Acknowledgements
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These studies were supported in part by a Summer Undergraduate Research Fellowship from UNC Chapel Hill (Connor Puett), the Focused Ultrasound Foundation, pilot funds from the National Science Foundation (DMR#1122483), and a Graduate Research Fellowship from the National Science Foundation (Paul S. Sheeran).
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References
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Please cite this article in press as: C. Puett et al., Pulse sequences for uniform perfluorocarbon droplet vaporization and ultrasound imaging, Ultrasonics (2014), http://dx.doi.org/10.1016/j.ultras.2014.05.013
