- Email: [email protected]
WFUMB safety symposium on echo-contrast agents: Exposure from diagnostic ultrasound equipment relating to cavitation risk
Ultrasound in Med. & Biol., Vol. 33, No. 2, pp. 214 –223, 2007 Copyright © 2007 World Federation for Ultrasound in Medicine & Biology Printed in the USA. All rights reserved 0301-5629/07/$–see front matter
doi:10.1016/j.ultrasmedbio.2006.07.011
● Safety of Ultrasound Contrast Agents WFUMB SAFETY SYMPOSIUM ON ECHO-CONTRAST AGENTS: EXPOSURE FROM DIAGNOSTIC ULTRASOUND EQUIPMENT RELATING TO CAVITATION RISK T. ANTHONY WHITTINGHAM Regional Medical Physics Department, Newcastle General Hospital, Newcastle upon Tyne Hospitals NHS Trust, Newcastle upon Tyne, UK
CHAPTER 5
1981), possibly due to partly developed bubbles or cavitation products from one cycle or pulse persisting to seed the following one. Low viscosity favours a lower pressure threshold, as does increased temperature (which might be caused by the ultrasonic exposure itself), partly due to reduced gas solubility, surface tension and viscosity. In tissues, there are additional inhibiting factors such as the tissue binding forces that will oppose bubble expansion, and the barriers to gas diffusion presented by cells surrounding a bubble. Theoretical (Holland and Apfel 1989) and experimental work (Apfel and Holland 1991) has indicated that the likelihood of inertial cavitation occurring, given a population of bubble nuclei of all sizes, is given by the ratio of the square of the peak negative in situ (at the nucleus) pressure to the pulse frequency. They concluded that this ratio must exceed 0.5 (assuming units of MPa for pressure and MHz for frequency) if inertial cavitation is to be possible for a single cycle pulse in water. This has led to the adoption of a “mechanical index (MI)” as an on-screen hazard indicator for nonthermal bioeffects:
Exposure parameters of relevance to cavitation hazard The two most important factors determining the likelihood of cavitation of un-encapsulated gas or vapour bubbles in liquids are the peak negative pressure amplitude and the duration of the negative pressure half cycle. The exposure parameter representing the former is the spatial-peak temporal-peak negative pressure (p⫺). This is the largest negative pressure excursion to be found anywhere in the beam (or the scan slice in the case of scanned modes, such as B-mode and colour Doppler Imaging), at any time in the scan sequence. The exposure parameter usually quoted in relation to the duration of negative pressure half cycle is the centre frequency fc of the pulse energy spectrum. Peak negative pressure p⫺ may not be the most appropriate pressure parameter in all circumstances. For example, pulses with sufficient amplitude to cause cavitation may have severely distorted asymmetric waveforms (Fig. 1) when measured in water, leading to underestimated p⫺ values. A better predictor may then be the amplitude of the fundamental frequency of the pulse spectrum (Morton et al. 1982). In tissue containing preexisting gas pockets, it has been suggested (Hartman at al. 1990) that peak positive pressure may be a better predictor, since the tissue would restrain gas pocket expansion, making the event principally one of collapse. However, in general, p⫺ remains the pressure parameter that is most often used. Pressure thresholds are lower, and cavitation effects enhanced, for longer pulses, longer exposure times and higher pulse repetition frequencies (Ciaravino et al.
MI ⫽ p0.3 ⁄ 兹fc where p0.3 is the peak “derated” (see below) negative pressure measured in MPa at a specified position in front of the probe, and fc is the centre frequency of the transmitted pulse measured in MHz. As the MI is the square root of the threshold ratio found by Apfel and Holland 1991, the MI threshold for inertial cavitation, given nuclei of all sizes and a single cycle pulse, is approximately 0.7. The value of MI is displayed on the equipment screen during scanning, and is automatically updated as the machine controls (output power, focus, zoom, etc.) are adjusted. It is not always displayed on permanent
Address correspondence to: Regional Medical Physics Department, Newcastle General Hospital, Newcastle upon Tyne, NE4 6BE, UK. E-mail: [email protected] 214
Exposure from diagnostic ultrasound equipment ● T. A. WHITTINGHAM
Fig. 1. Typical pressure waveforms in water. The upper waveforms are for large pulse amplitude settings. They show marked distortion due to nonlinear propagation in water. The lower waveforms are the result of reducing the output by 15 dB; the pulse shapes are more representative of those that would exist in soft tissue.
hard copy or videos of scans, however, as some manufacturers arrange for it to be removed from the screen during such recordings. The FDA requires that this index be less than 1.9, but it should be noted that displayed values of MI can be in error because of machine to machine variations for a given manufacturer and model. Although essentially a safety index, MI is commonly used to describe exposure conditions when using ultrasound contrast agents, and it has proved to be useful as a rough predictor of microbubble behaviour for a given machine and contrast agent. It should be noted, however, that MI is based on the modelled growth and behaviour of vapour filled bubbles exposed to single cycle pulses, not microbubbles with stabilising shells exposed to the sometimes long pulses used in contrast agent scanning. It cannot be assumed, therefore, that MI values give an absolute indication of the likelihood of the rupture or cracking of microbubble shells, or of associated bioeffects, that is universally valid for all machines, scanning modes and ultrasound contrast agents. It is shown in Chap. 3 that the threshold pressure value for in vitro bioeffects in the presence of shelled microbubbles appears to be proportional to frequency, rather than the square root of frequency, as inherent in the MI definition. Above the threshold pressure, bioeffects may increase in proportion to a power higher than the square of pressure. Until an index more appropriate to encapsulated microbubbles is introduced, display of the basic acoustic parameters p⫺ and fc might be helpful.
215
Exposure parameters of relevance to tissue heating Heating in tissue exposed to diagnostic ultrasound is due to two separate causes: direct absorption of ultrasound wave energy in the tissue and heat conducted through superficial tissues from a self-heated ultrasound probe. The precise temperature rise produced by absorption in any particular situation is very difficult to predict as it depends on many factors: the intensity distribution in the tissue, the ultrasound absorption coefficients of the various exposed tissues, the dimensions, specific heat and thermal conduction of the various tissues and the degree to which blood flow provides heat clearance. Most modern machines display a number (thermal index TI) on the screen that gives a very rough indication (within a factor of two) of the worst-case temperature rise that might be expected anywhere in the exposed tissues, according to very simple models of the tissue and the intensity distribution. As for the MI, the TI value is constantly updated according to the settings of the controls that influence the transmitted intensity and its spatial distribution, notably the choice of operating mode (B-mode, spectral Doppler, low MI or high MI contrast specific modes, etc.), output power and the depth at which the operator sets the transmission focus. Considerations relating to measurements of pressure in water: derating and nonlinearity Measurements are normally made in water for convenience but, ideally, values close to those found at the depth of a particular target in tissue (in situ values) would be preferred. Some standards and regulations make reference to “derated” values as a practical rough substitute for actual in situ values. Derated values are “in-water” values, reduced by ␣.fc.z decibels, representing the attenuation that would be introduced by a medium with an attenuation coefficient of (␣ dB cm⫺1 MHz⫺1), where z (cm) is the distance of the measurement site from the transducer, and fc (MHz) is the pulse centre frequency. In the definition of MI given in the previous section, the subscript 0.3 specifies a derating factor of 0.3 dB cm⫺1 MHz⫺1. This is somewhat less than the attenuation coefficients typical of soft tissue, but it allows for occasional low attenuation liquid regions in the sound path in general abdominal and obstetric studies. In formulating a more appropriate index for contrast agent work, as mentioned earlier, consideration might be given to a larger derating factor in view of the increased attenuation introduced by ultrasound contrast agents, and the greater tissue attenuation typical of common contrast applications such as transthoracic imaging of the heart or liver. A further consideration is that, at diagnostic frequencies and pressures, severe nonlinear propagation effects lead to pulse wave distortion, generation of har-
216
Ultrasound in Medicine and Biology
Volume 33, Number 2, 2007
monics, excess energy absorption and acoustic saturation (Duck 2002). This is demonstrated in the pulse waveforms in the upper part of Fig. 1. As the voltage drive applied to the transducer increases, the magnitude of the peak negative pressure becomes less than it would be in the absence of nonlinear distortion and the location of peak negative pressure moves toward the transducer (Duck and Starritt 1986). These effects appear more readily in water than in tissue, and are greater for greater transmitted pulse amplitudes. As a general result of these effects, the in situ peak negative pressure may be underestimated by up to a factor of two, being greater for larger depths (Duck and Bacon 1988). The effect of nonlinear propagation on in-water measurements of MI is less, since the formulation of the displayed MI requires p0.3 to be measured at a specific depth, which is usually comparatively close to the transducer. Using present protocols, it seems safe to assume that an in-water measurement of MI will underestimate a true in situ value by no more that 20% (Christopher 1999; Cahill and Humphrey 2000). One solution that has been proposed to the problem is that in situ peak negative pressures should be estimated by linear extrapolation from in-water measurements made at sufficiently low amplitude to assure quasilinear propagation. Spatial distribution of the acoustic field Normally, in soft tissue, negligible acoustic pressure will be produced at depths greater than those that can be usefully imaged. However, this may not be true in the rare event that extremely high reject settings or very low dynamic range or gain settings are selected. Where it is necessary to increase the transmitted amplitude to image targets behind strongly attenuating structures, for example when imaging intracranial targets, the pressure amplitude at the target may be much higher than when imaging through soft tissue. Consider a particular type of target imaged via a soft tissue path that returns an echo of sufficient amplitude to be just detected by a given probe. If the soft tissue path is replaced by a barrier introducing K times more attenuation (e.g., the skull wall), the transmission amplitude will have to be increased by a factor of K2 to compensate for the increased two-way attenuation. More significantly, from an exposure point of view, the amplitude at the target will now be K times greater than for the soft tissue path. In nonscanned modes (e.g., A-mode, M-mode and spectral Doppler mode) the beam cross-section from an array probe is rectangular close to the probe, becoming elliptical in cross section at depth. This is illustrated in Fig. 2, where the horizontal plane is the scan plane and the vertical plane is the elevation (slice thickness) plane. In the scan plane dimension, the beam is narrowest at the range of the “focus” as set by the operator, while in the
Fig. 2. Simplified representations of the acoustic field in different modes. Top: Nonscanned beam with the scan plane focus coincident with the slice thickness focus, together with typical pulse pressure waveforms measured in water at different ranges in Doppler mode. Centre: As above, with the scan-plane focus positioned deeper than the slice thickness focus. Bottom: Scanned mode with weak focusing in the scan plane.
slice thickness dimension it is narrowest at the fixed focus of the cylindrical lens at the front of the probe. Pulses with large peak pressures will be produced when the scan plan focus is set to equal the range of minimum slice thickness, as this concentrates the pulse power into the smallest beam cross-section. In scanning modes, such as in B-mode and Doppler Imaging modes, the acoustic field is in the form of a slice, with a width (scan plane) and thickness (elevation plane) that vary with range. Generally, the field width at all ranges is approximately equal to that of the probe face, but if a narrow “zoom box” (see section “Machine control settings. . .”) is selected, the strong scan-plane focusing of individual beams means that the field width will be less at the range of the box. The slice thickness (elevation plane) varies with range in a manner determined by the probe’s fixed cylindrical lens. It is equal to the height of the lens close to the probe, reducing to a minimum at the fixed range of the focus of the lens, and then increasing with range.
Exposure from diagnostic ultrasound equipment ● T. A. WHITTINGHAM
217
ducer further from the scan plane. The resulting acoustic field is more uniform in pressure amplitude versus range, but the pulse centre frequency reduces as the range increases.
Fig. 3. Example of the variation of on-axis pulse peak negative pressure with depth in B-mode. (a) In-water; (b) Derated by 0.3 dB MHz⫺1 cm-⫺; (c) Derated by 0.6 dB MHz⫺1 cm⫺1.
In all modes, the variation of in situ peak pressure with depth will be influenced by the attenuation coefficients of the particular tissues involved. Figure 3 shows the effect of applying two different attenuation coefficients to an axial profile of in-water peak pulse pressure versus depth, to estimate the in situ profiles. The range of maximum in situ peak pressure will always be closer to the probe than the maximum in-water peak pressure. This may not be a big effect if the focusing action is pronounced and the attenuation is low [curve (b)], but attenuations typical of soft tissues [curve (c)] can have a large effect. The centre frequency of the in situ pulse will also be less than that measured in water, reducing with increasing depth, due to the greater absorption by tissues of the higher frequencies in the pulse spectrum. A complication of estimating in situ pressures when using contrast agents is that the microbubbles themselves increase attenuation by absorption. This extra absorption may lead to underestimation of thermal hazard by the Thermal Index. Some probes (so-called 1.5 D arrays) have several rows of elements to provide a variable focus in the slice thickness dimension. When a single (scan-plane) focus is selected by the operator, the machine can then automatically set the slice thickness focus to be at the same depth. This increases the peak pressure at the focus, as shown in Fig. 2. If several foci are selected at different ranges, the overall transmission slice thickness and peak pressure can be made more uniform with depth. A more uniform distribution of pressure amplitude with depth is also achieved by those probes that use an Hanafi lens. This type of probe has a piezoelectric transducer element with a cylindrically concave transmitting face, resulting in the transducer being thinnest in the scan plane. Higher frequencies are transmitted from the thinner central part of the transducer to ranges close to the probe, while principally lower frequencies are transmitted to greater depths from the thicker parts of the trans-
Temporal variation of pressure in different modes The waveform of a single pulse and the temporal variation of peak pressure amplitude over a complete frame depend on the mode selected. A simplified representation of the pressure variation versus time at one point in front of the probe is shown in Fig. 4. In practice, more pulses than those shown may be experienced at a particular field point, particularly at positions between the probe and the transmission focus. This is due to the practice of transmitting pulses along other scan lines while waiting for echoes to return from a scan line (beam multiplexing). In A-mode or M-mode, pulses typically have two to three cycles. Since the beam is not scanned, there is no variation from pulse to pulse. In B-mode, the pulses are similar in shape and amplitude to those used in A- and M-modes, but as the beam is swept across the field of
Fig. 4. Simplified representation of the variation of peak pulse amplitude with time over a complete frame, at one point in front of the probe. Each pulse is shown by a single line, the height of which is proportional to the peak amplitude of the pulse. Expanded pulse waveforms are shown to the left of each frame representation.
218
Ultrasound in Medicine and Biology
view, a given field point experiences almost zero exposure for most of each frame, and a transient rise and fall of pulse amplitude as the beam passes over. If a narrow write zoom box is selected (see the following section), the temporal average intensity is likely to be increased at all depths, since the transmitting aperture is likely to be increased and hence a given field point will be in a beam (insonated) for a greater proportion of each frame. In tissue harmonic imaging (THI) mode, the transmitted pulses are likely to have a lower centre frequency than typical B-mode pulses, since the transducer has to be able to operate at both the transmission centre frequency (fundamental) and the second harmonic (twice the transmission frequency). Pressure amplitudes are normally at the high end of the range used for B-mode, to promote second harmonic generation when propagating in tissue. If a filter is used to suppress the fundamental echo component, the need to avoid overlap between the spectra of the transmitted pulse and that of the echoes means that narrower bandwidth and, hence, longer pulses are transmitted. If the pulse inversion technique is used, two pulses are transmitted per scan line, one being an inverted version of the other. These transmitted pulses can have normal bandwidth and, hence, pulse length, although some effort is often made to reduce their second harmonic content. In Spectral Doppler, pulses are longer (typically 3 to 20 cycles, depending on the width of the range gate), at a higher pulse repetition frequency (PRF) to reduce aliasing, and generally with a lower centre frequency. To keep time-averaged intensities within FDA regulatory limits despite the higher PRFs, pressure amplitudes are usually lower than in A-, M- or B-mode, becoming less as range-gate width increases. In colour Doppler imaging modes, typically 8 to 20 pulses are transmitted along each scan line before the beam is advanced (the larger number being for a high “Colour Quality” control setting). A sequence of fourteen pulses per Doppler line is shown in Fig. 4. The individual pulses are similar to those used in spectral Doppler for a short range-gate setting. As the beam is advanced across a given field point, the peak pulse amplitude builds up and then dies away, being almost zero for most of the frame. Note that in Fig. 4 one frame of colour Doppler imaging/LOC pulses is shown occupying the same space as for other modes. In practice, the need to transmit 8 to 20 pulses down each scan line means that the time taken to complete a colour Doppler frame is likely to be several times longer than for other modes. Contrast-specific modes for use with encapsulated microbubbles are commonly categorised as either “low MI” or “high MI”, although, as mentioned earlier, the appropriateness of the MI safety index is questionable. For a review of the various methods see, for example,
Volume 33, Number 2, 2007
Whittingham (2005). The pulses transmitted in low MI modes have low pressure amplitudes (around 0.1 MPa or less, with typical MI values of around 0.1 or less), specifically to reduce the likelihood of microbubble rupture. High MI modes, however, rely on the rupture of microbubbles using sensitive detection methods, which essentially compare echo sequences before and after microbubble rupture. The likelihood of cavitation with high MI contrastspecific techniques is higher than with noncontrast techniques due to the high MI values (around 1 or more) and the presence of cavitation nuclei from cracked or ruptured microbubbles. The likelihood is much less for low MI contrast-specific techniques, of course, but low MI techniques can involve more and longer pulses, and often involve MI values that are above the thresholds reported for in vitro bioeffects and cavitation in the presence of contrast agents. For example, the threshold of p–/f ⫽ 0.06 MPa/ MHz reported in Chap. 3 for in vitro bioeffects implies a MI threshold of only 0.06 at 1 MHz and 0.12 at 4 MHz. Also, Chen at al (2003) have detected white noise emission from insonated contrast agents, consistent with inertial cavitation, at MI values lower than 0.2. One common low MI contrast-specific method uses the pulse inversion technique of second harmonic detection mentioned above for Tissue Harmonic Imaging. As with THI, two pulses are transmitted along each line, but the peak negative pressures are much lower than in THI. Some low MI techniques involve the transmission of more than two pulses along each line (Table 4). These give greater sensitivity and improved discrimination between microbubble and tissue echoes. Such multiple transmissions can give rise to greater temporal-average intensities and powers than in normal B-mode and, hence, the potential for tissue heating can be greater. This will be indicated by a higher TI value on the screen. A selection of different low MI mode transmissions are described in Table 4. Loss of correlation (LOC) imaging is a common high MI method that uses similar transmission sequences to those used in colour Doppler imaging, although with higher pulse amplitudes (MI ⬃1). Recent versions use fewer (2– 6) and shorter (broad bandwidth) pulses and higher pulse repetition frequencies. “Pulse subtraction” is a high MI method in which only two short but high amplitude pulses are transmitted along each line. Since high MI imaging methods destroy the microbubbles, the probe is swept by hand across the patient, resulting in exposure times of only a few seconds at any one site. When used in “flash” or “burst” mode to clear microbubbles from a field of view, long sequences of large amplitude pulses are transmitted along each scan line in every frame. This is normally repeated for just a few frames, but in some instances may continue as long
Exposure from diagnostic ultrasound equipment ● T. A. WHITTINGHAM
as the mode is selected. In Fig. 4 a sequence of seven large amplitude pulses is shown being sent down each scan line. As the beam passes over the measuring hydrophone, the amplitude of the detected sequence rises and falls. Machine control settings that give the highest exposures The nature and range of controls is constantly changing with the evolution of new scanning techniques, so universal rules are difficult to formulate. Controls on some of the newer machines can have quite unexpected effects, as manufacturers often arrange for drive voltages or pulse repetition frequencies to change automatically when controls are set in a way which would otherwise cause a particular safety parameter, such as intensity, MI or TI, to exceed a regulatory limit. Apart from the obvious example of increasing the setting of the output power control, there are a few general observations that can be made about producing pulses with high pressure amplitudes. Some general observations on control settings that affect pulse repetition frequency (PRF) are also possible. In general, selecting a deep transmission focus commonly involves an increase in p⫺ and heating, whichever mode is chosen. This is because manufacturers often arrange for the transmission aperture to be increased if a deep transmission focus is selected, to maintain a narrow beam width and good sensitivity at depth. Apart from an increase in pressure amplitude due to more elements transmitting, the drive voltage applied to each element may also be increased to compensate for the greater attenuation anticipated for deep targets. Activation of a write-zoom box is another way by which exposure can be increased, particularly if the box is narrow. Unlike read-zoom, which simply magnifies part of the stored image, write-zoom involves a selected smaller area being re-scanned at a higher line density and/or frame rate. This leads to a given field point being insonated for a larger fraction of the frame time and thus being insonated by more pulses per second. Temporal average intensities and hence tissue heating will also increase. Write zoom may also lead to a higher pulse repetition frequency, since echoes from beyond the box are not wanted. If the zoom box is deep, there may also be an increase in p⫺, since the machine may automatically advance the focus to match the depth of the box, leading to an increase in p⫺, as mentioned above. A large p⫺ is usually produced if the operatorcontrolled focus (which sets the transmission focus in the scan plane) is set close to the (fixed) slice thickness focus, since this increases the strength of focusing in a 3-D sense (see Fig. 1). However, as discussed above, setting the focus (or range-gate in the case of spectral Doppler) to a greater depth may well increase the trans-
219
mission aperture and drive voltage, and hence produce even greater pressures and intensities near the scan-plane focus. Only practical measurements can establish where the greatest pulse pressure will be actually produced. In spectral Doppler mode, the only control that is likely to change the pulse length is the gate width (sample volume width) control. A high Doppler frequency scale setting, or the selection of “High PRF” mode, is likely to produce lower pressure amplitudes. In themselves, these controls would not be expected to affect p⫺ values, but manufacturers sometimes arrange for drive voltages and, hence, pressure amplitudes, to be reduced if a high PRF is selected, to keep intensities within prescribed limits. For the same reason, drive voltages are usually reduced as gate width (and hence pulse length) increases, so that higher p⫺ values are likely to be found for narrow gate settings. Controls that are unlikely to directly change the transmitted acoustic field include: Overall gain (2D Gain), TGC, dynamic range, reject, grey scale preprocessing or postprocessing (transfer characteristic), frame averaging (persistence), read-zoom (magnification), Doppler gain, Doppler filter, Doppler baseline, cursor angle (angle correction), spectrum inversion, colour gain, colour priority. Tables 1, 2, 3 and 4 summarise the function and effect on exposure of a selection of machine controls. Survey of worst-case values and trends When considering the hazard potential of an ultrasound field, it is of interest to try to find the “worstcase” values of the measured quantities. This involves repeating the measurements for the many different combinations of machine control settings that are likely to give a high output, until the combination that gives the largest value for the particular quantity is found. Contrast-specific modes have only recently been
Fig. 5. Worst-case in-water p⫺ values measured between 1995 and 1999.
220
Ultrasound in Medicine and Biology
Volume 33, Number 2, 2007
Table 1. General controls and their effect upon exposure Control name
Function
Effect on exposure
Frequency (where multi-frequency probes are selected)
Allows an optimum balance of spatial resolution and penetration to be achieved. Higher frequencies give better resolution but less penetration.
Output power
One way of controlling the overall sensitivity and penetration. (The other way is to use Overall Gain). Allows lateral resolution and sensitivity to be increased at a preferred depth by placing the transmission focus there.
Changes the centre frequency of the transmitted pulse. Lower frequencies often mean an increase in in-situ pulse amplitude (less attenuation). However, if the focus and aperture remain constant, changing to a higher frequency may cause the pulse pressure amplitude near the focus to be increased due to the narrower beam width. The higher the power setting, the greater the pressure amplitude of every transmitted pulse.
Focus depth
Number of foci
As above but for several different focus positions (at expense of frame rate)
Displayed depth
Sets the maximum depth of the field of view
Displayed width or sector width
Sets the width of the field of view
Frame rate (some machines)
Allows trade-off between image quality and frame rate in 2D or Colour Imaging modes Write Zoom allows greater spatial resolution and/or frame rate to be achieved by restricting scanning to a more limited area, as defined by the zoom box on the image. Max Depth sets the deeper margin of the zoom box
Write-zoom box - maximum depth
Write-zoom box-width Frame freeze or cine loop (cine review) Compounding THI mode
Sets the horizontal width of the zoom box Allows an image frame or sequence of frames to be reviewed for as long as required Improves delineation of interfaces; reduces speckle and noise Reduces noise artefacts due to sidelobes and reverberations
introduced, and there have been no published surveys of worst-case pressure or intensity values for these modes. Measurements and reviews of worst-case inwater values of p⫺, carried out for conventional modes over the last three decades (Whittingham 2000), show that those measured in the half decade 1995 to 1999 were higher than ever before, with evidence of a slow continuing upward trend. The increased likelihood of cavitation that this increase in p⫺ might imply is partly balanced for conventional modes by the fact that pulse centre frequencies have
Increasing the depth of the transmission focus will move the point of greatest pulse amplitude to a greater depth (not necessarily at the focus itself). It is also likely to increase the pulse pressure amplitude at all depths. Some machines may also lower the centre frequency of the pulse at the same time. Setting the focus to the same depth as the lens elevation plane focus (slice thickness minimum) produces larger peak pressure amplitude at that depth. Several pulses transmitted along each scan line, each with a focus at a different depth. Effect for each pulse / focus as above. Larger depth settings cause the PRF to be reduced. On some machines the pulse centre frequency may be reduced and the pressure amplitude may be increased. Smaller widths can result in more closely spaced scan lines (each point then experiences more pulses as it lies in more beams) and higher frame rates (less time before the beam returns to insonate each point). Results in more pulses at any given point and greater temporal average intensities. Low frame rates result in a greater line density. Each point then experiences a longer sequence of pulses. Selecting a deep box will create a matching deep focus - increasing the pulse amplitude as described previously for Focus Depth. A shallow maximum depth of box may cause the PRF to be increased and possibly the pulse centre frequency to be increased. As for Displayed Width or Sector Width. Normally stops transmissions, but not on every machine. The number of pulses insonating each point in the field may increase. Reduces the centre frequency of the pulse. If pulse inversion is used, two pulses are transmitted down each scan line. If a filter is used instead, longer pulse lengths may be used.
generally been increasing in the quest for greater spatial resolution. The measurements for the 1995 to 1999 period are presented graphically in Fig. 5. Little difference was found between M- and B-mode values, so they are grouped together. Overall, worst-case values are distributed about mean values of 2.4 to 2.8 MPa, depending on the mode. Median worst-case values are in the range 2.4 to 2.6 MPa. There is very little difference among modes, although perhaps it is worth noting that, in this survey, the most extreme p⫺ values (exceeding 5 MPa) occurred in spectral Doppler mode. Surveys of MI values
Exposure from diagnostic ultrasound equipment ● T. A. WHITTINGHAM
221
Table 2. Spectral Doppler mode controls and their effect upon exposure. A continuous stream of pulses is transmitted down a stationary beam. PRFs are higher than in A-, M- and B-modes and peak pressures are generally lower Control name
Function
Effect on exposure
Gate width (sample volume)
Increases the length of the sample volume, i.e., the range of depths from which Doppler signal is obtained.
Scale (PRF on some machines)
Allows a larger range of Doppler frequencies to be displayed on the Doppler spectrum.
High PRF
Allows aliasing to be eliminated in exchange for ambiguity in the depth of the sample volume (More than one range gate displayed) Controls the angle between the ultrasound beam and the blood flow.
Larger gate width settings increase the number of cycles in each pulse. Widening the gate commonly cause the pulse amplitude to be reduced. Increasing the scale (larger Doppler, frequency range) causes the PRF to be increased in proportion. Increases the PRF by a factor equal to the number of range gates displayed.
Doppler line steering
have not been reported. However, instances of erroneous MI values have been found. Estimation of worst-case in situ pressure amplitudes Although contrast agents are not currently licensed for obstetric applications, the importance of the fetus from a hazard perspective promotes the consideration of exposures in such applications. Estimates have been reported of either the typical or minimum attenuations introduced by tissue between the probe and the fetus. A figure of 0.5 dB MHz⫺1 has been proposed for minimum attenuation in the third trimester (Carson et al. 1989), while a figure of 3.6 dB MHz⫺1 has been reported for typical attenuation in the first two trimesters (Siddiqi et al. 1995). Table 5 shows the result of applying these attenuating factors to the in-water values of worst-case p⫺ upon which Fig. 5 is based, to give rough estimates of worst-case in situ p⫺ values at the fetus (Whittingham 2001). The minimum attenuation path third trimester model predicts that worst-case values in excess of 3 MPa are possible in all conventional modes, but in the “typical” model the upper limit is reduced to around 1 MPa. It should be noted that the models take no account of the reduced pulse distortion that would be associated with propagation in tissues, compared with the water in which the measurements were made. The in situ negative pressure values could, therefore, be greater than these estimates. CONCLUSIONS Conventional modes (i.e., non-contrast-specific modes) p⫺ values in conventional modes have increased several-fold since 1970 and have shown a slower, but still rising, trend over the last decade.
Increasing the deflection from the “straight ahead” direction can result in reduced pressure amplitude and broader beams (hence more pulses insonating each point).
The range of reported worst-case p⫺ values shows little difference between A-mode, M-mode, B-mode, spectral Doppler, CFM and power Doppler. However, typical pulse pressure amplitude values are generally higher in A-mode, M-mode and B-Mode than in Doppler modes. There are large differences in reported worst-case p⫺ values between machines from different manufacturers. Worst-case in-water p⫺ values have been reported to be as high as 4.6 MPa in M- and B-mode, 4.9 MPa in colour Doppler and 5.5 MPa in spectral Doppler.
Table 3. Colour Doppler imaging controls and their effect upon exposure. Several pulses are transmitted down each Doppler line of the colour box. PRF and pulse centre frequency are likely to be similar to those in spectral Doppler Control name Colour quality
Colour box steering
Function
Effect on exposure
High quality settings increase the consistency and accuracy of the colour, and allow lower flow velocities to be shown in colour. Controls the angle between the ultrasound beams in the colour box and the blood flow.
High quality settings increase the number of pulses transmitted along each scan line.
Increasing the deflection from the “straight ahead” direction can result in reduced pressure amplitude and broader beams (hence more pulses insonating each point).
222
Ultrasound in Medicine and Biology
Volume 33, Number 2, 2007
Table 4. Contrast-specific modes and their effect upon exposure. Commonly divided into High MI and Low MI modes. High MI techniques are associated with microbubble destruction and hence intermittent scanning. Low MI techniques try to prolong the lifetime of bubbles modes in order to allow extended real-time imaging Mode
Function
Effect on exposure
High MI LOC scanning
Images the distribution of contrast agent microbubbles by destroying the microbubbles and comparing the echo sequences before and after.
Flash or burst
Deliberately clears contrast agent from the field of view for reperfusion studies by rupturing the microbubbles. Discriminates microbubble echoes from tissue by their larger second harmonic content.
High MI (usually ⬎ 0.7) pulses transmitted which rupture the contrast microbubbles. Older systems transmit long sequences (⬃10) of longer pulses (3–5 cycles) along each line, as for Colour Doppler Imaging. Some recent methods use fewer (2–4) short (2–3 cycles) pulses and higher PRFs. Transmits a long sequence (⬃10) of large amplitude pulses along each scan line.
Low MI harmonic imaging by high-pass filtering
Low MI pulse inversion
As above.
Low MI coherent pulse inversion
As above.
Low MI coded excitation
As above
Low MI chirp coding
As above
Low MI contrast pulse sequence imaging (CPS)
Discriminates microbubble echoes from tissue by the non-linear signal in the fundamental echo component or by detecting a harmonic echo component.
Estimates of in situ pressure values at the fetus for obstetric models indicate typical p⫺ values of less than 0.4 MPa, but worst-case values as high as 2 MPa. There is no simple rule for predicting the depth at which the spatial maximum of in situ negative pressure will occur. The controls that most directly influence p⫺ are the output power and focus depth, but other controls can have an indirect influence due to the need to keep MI and ISPTA within regulatory limits. Spectral Doppler mode and Doppler imaging modes generally involve the highest pulse repetition frequencies and pulse lengths, and the lowest centre frequencies. Contrast-specific modes These modes are customarily divided into “high MI” and “low MI” techniques according to whether or not they are likely to cause rapid destruction of the contrast microbubbles. In high MI modes, two or more large amplitude pulses (MI ⬃1) are transmitted along each scan line. The large amplitude and the presence of transient small bub-
Longer transmitted pulses than in B-mode, with low amplitudes; sometimes with waveforms (e.g. Gaussian) designed to have low harmonic content. Centre frequency of the pulses may be reduced to facilitate detection of second harmonic. Two short pulses (more typical of B-mode), but with low amplitudes and the second pulse inverted, are transmitted along each line; Sometimes low harmonic transmission waveforms and reduced centre frequency, as above. One short low amplitude pulse transmitted along each line, inverted on alternate lines. Centre frequency may be reduced as above. Two dissimilar, long (8 or more cycles), low amplitude pulses transmitted along each line. Centre frequency may be reduced as above. One long (⬎20 cycles) low amplitude chirp (swept frequency) pulse transmitted along each line. Three or more low amplitude pulses transmitted along each line, sometimes with different amplitudes and/or with some pulses inverted. Centre frequencies may be reduced where harmonics are to be detected.
bles of free gas due to the fragmentation of contrast microbubbles means that the likelihood of cavitation is greater than in conventional imaging modes or low MI contrast modes. Low MI modes of operation generally involve the transmission of two or more low amplitude pulses along each scan line, sometimes with different amplitudes and polarities. Some techniques involve the transmission of a single long (up to 30 cycles) low amplitude chirp pulse along each line. Centre frequencies may be lower than for conventional B-mode. In low MI contrast modes, p⫺ values are likely to be low, so the potential for other nonthermal effects is also likely to be lower, although this may be offset to some extent by the fact that each field point is likely to experience more pulses per frame. Temporal average intensities and powers may be higher when long sequences of pulses are transmitted. The potential for thermal effects may then be increased. Low MI techniques can involve MI values that are above the thresholds reported for in vitro bioeffects and cavitation in the presence of contrast agents.
Exposure from diagnostic ultrasound equipment ● T. A. WHITTINGHAM
223
Table 5. Median (and maximum) values of worst-case in-water p- from diagnostic equipment, together with values obtained by applying attenuation factors of 0.5 dB MHz⫺1 and 3.6 dB MHz⫺1 respectively M-mode & Bmode (MPa)
Colour Doppler (MPa)
Spectral Doppler (MPa)
Attenuation dB MHz⫺1
Median
(Max)
Median
(max)
Median
(max)
0.5
2.4 1.8
(4.6) (3.1)
2.6 2.0
(4.9) (3.9)
2.4 1.7
(5.5) (3.2)
3.6
0.4
(1.0)
0.4
(1.1)
0.3
(1.1)
In water Minimum attenuation - 3rd trimester “Typical” attenuation - 1st an 3rd trimesters
No surveys have yet been published that give acoustic output data for contrast-specific modes. General Although widely available on display screens as a guide to the likelihood of nonthermal bioeffects, the mechanical index (MI) is not necessarily an adequate guide to the likelihood of microbubble rupture or of bioeffects associated with ultrasound contrast agents. Contrast agent may enhance absorption and, hence, heating. This may have implications for the proper choice of derating factor and absorption coefficient for use in thermal and mechanical indices when using contrast agents. In situ peak negative pressure can be significantly underestimated if account is not taken of nonlinear propagation effects. Displayed MI may be underestimated by up to 20% because linear assumptions are made in the calculation of in situ peak negative pressure. REFERENCES Apfel RE, Holland KH. Gauging the likelihood of cavitation from short pulse low duty cycle diagnostic ultrasound. Ultrasound Med Biol 1991;17:179 –185. Cahill MD, Humphrey VF. A theoretical investigation of the effect of nonlinear propagation on measurements of mechanical index. Ultrasound Med Biol 2000;26:433– 440. Carson PL, Rubin JM, Chiang EH. Fetal depth and ultrasound path lengths through overlying tissues. Ultrasound Med Biol 1989;15:629 – 639.
Chen WS et al. A comparison of the fragmentation thresholds and inertial cavitation doses of different ultrasound contrast agents. J Acoust Soc Am 2003;113:643– 651. Christopher T. Computing the mechanical index. J Ultrasound Med 1999;18:63– 68. Ciaravino V, Flynn HG, Miller MW. Pulsed enhancement of acoustic cavitation: A postulated model. Ultrasound Med Biol 1981;7:159 – 166. Duck FA. Nonlinear acoustics in diagnostic ultrasound. Ultrasound Med Biol 2002;28:1–18. Duck FA, Bacon DR. A fundamental criticism of hydrophone-in-water measurement. Ultrasound Med Biol 1988;14:305–307. Duck FA, Starritt HC. The locations of peak pressures and peak intensities in finite amplitude beams from a pulsed focused transducer. Ultrasound Med Biol 1986;12(5):403– 409. Hartman C, Child SZ, Macer R, et al. Lung damage from exposure to the fields of an electro-hydraulic lithotripter. Ultrasound Med Biol 1990;16:675– 679. Holland KH, Apfel RE. An improved theory for the prediction of micro-cavitation thresholds. IEEE. Trans Ultrason Ferroelectrics Fre Qcontr UFFI- 1989;36(2):139. Morton KI, ter Haar G, Stratford IJ, Hill CR. The role of cavitation in the interaction of ultrasound with V79 Chinese hamster cells in vitro. Br J Cancer 1982;45(Suppl V):147. Siddiqi TA, O’Brian, WD. Jr, Meyer RA, Sullivan JM, Miodovnik M. In situ human obstetrical ultrasound exposimetry: Estimates of derating factors for each of three different tissue models. Ultrasound Med Biol 1995;21:379 –391. Whittingham TA. Acoustic outputs of diagnostic machines. In: ter Haar G, Duck FA, eds. Safety of medical diagnostic ultrasound. London: British Institute of Radiology, 2000;16 –31. Whittingham TA. Estimated cerebral ultrasound exposures from clinical examinations. Ultrasound Med Biol 2001;27(7):877– 882. Whittingham TA. Contrast-specific imaging techniques: Technical perspective. In: Quaia E, ed. Contrast media in ultrasonography. Berlin: Springer, 2005; 43–70.
doi:10.1016/j.ultrasmedbio.2006.07.011
● Safety of Ultrasound Contrast Agents WFUMB SAFETY SYMPOSIUM ON ECHO-CONTRAST AGENTS: EXPOSURE FROM DIAGNOSTIC ULTRASOUND EQUIPMENT RELATING TO CAVITATION RISK T. ANTHONY WHITTINGHAM Regional Medical Physics Department, Newcastle General Hospital, Newcastle upon Tyne Hospitals NHS Trust, Newcastle upon Tyne, UK
CHAPTER 5
1981), possibly due to partly developed bubbles or cavitation products from one cycle or pulse persisting to seed the following one. Low viscosity favours a lower pressure threshold, as does increased temperature (which might be caused by the ultrasonic exposure itself), partly due to reduced gas solubility, surface tension and viscosity. In tissues, there are additional inhibiting factors such as the tissue binding forces that will oppose bubble expansion, and the barriers to gas diffusion presented by cells surrounding a bubble. Theoretical (Holland and Apfel 1989) and experimental work (Apfel and Holland 1991) has indicated that the likelihood of inertial cavitation occurring, given a population of bubble nuclei of all sizes, is given by the ratio of the square of the peak negative in situ (at the nucleus) pressure to the pulse frequency. They concluded that this ratio must exceed 0.5 (assuming units of MPa for pressure and MHz for frequency) if inertial cavitation is to be possible for a single cycle pulse in water. This has led to the adoption of a “mechanical index (MI)” as an on-screen hazard indicator for nonthermal bioeffects:
Exposure parameters of relevance to cavitation hazard The two most important factors determining the likelihood of cavitation of un-encapsulated gas or vapour bubbles in liquids are the peak negative pressure amplitude and the duration of the negative pressure half cycle. The exposure parameter representing the former is the spatial-peak temporal-peak negative pressure (p⫺). This is the largest negative pressure excursion to be found anywhere in the beam (or the scan slice in the case of scanned modes, such as B-mode and colour Doppler Imaging), at any time in the scan sequence. The exposure parameter usually quoted in relation to the duration of negative pressure half cycle is the centre frequency fc of the pulse energy spectrum. Peak negative pressure p⫺ may not be the most appropriate pressure parameter in all circumstances. For example, pulses with sufficient amplitude to cause cavitation may have severely distorted asymmetric waveforms (Fig. 1) when measured in water, leading to underestimated p⫺ values. A better predictor may then be the amplitude of the fundamental frequency of the pulse spectrum (Morton et al. 1982). In tissue containing preexisting gas pockets, it has been suggested (Hartman at al. 1990) that peak positive pressure may be a better predictor, since the tissue would restrain gas pocket expansion, making the event principally one of collapse. However, in general, p⫺ remains the pressure parameter that is most often used. Pressure thresholds are lower, and cavitation effects enhanced, for longer pulses, longer exposure times and higher pulse repetition frequencies (Ciaravino et al.
MI ⫽ p0.3 ⁄ 兹fc where p0.3 is the peak “derated” (see below) negative pressure measured in MPa at a specified position in front of the probe, and fc is the centre frequency of the transmitted pulse measured in MHz. As the MI is the square root of the threshold ratio found by Apfel and Holland 1991, the MI threshold for inertial cavitation, given nuclei of all sizes and a single cycle pulse, is approximately 0.7. The value of MI is displayed on the equipment screen during scanning, and is automatically updated as the machine controls (output power, focus, zoom, etc.) are adjusted. It is not always displayed on permanent
Address correspondence to: Regional Medical Physics Department, Newcastle General Hospital, Newcastle upon Tyne, NE4 6BE, UK. E-mail: [email protected] 214
Exposure from diagnostic ultrasound equipment ● T. A. WHITTINGHAM
Fig. 1. Typical pressure waveforms in water. The upper waveforms are for large pulse amplitude settings. They show marked distortion due to nonlinear propagation in water. The lower waveforms are the result of reducing the output by 15 dB; the pulse shapes are more representative of those that would exist in soft tissue.
hard copy or videos of scans, however, as some manufacturers arrange for it to be removed from the screen during such recordings. The FDA requires that this index be less than 1.9, but it should be noted that displayed values of MI can be in error because of machine to machine variations for a given manufacturer and model. Although essentially a safety index, MI is commonly used to describe exposure conditions when using ultrasound contrast agents, and it has proved to be useful as a rough predictor of microbubble behaviour for a given machine and contrast agent. It should be noted, however, that MI is based on the modelled growth and behaviour of vapour filled bubbles exposed to single cycle pulses, not microbubbles with stabilising shells exposed to the sometimes long pulses used in contrast agent scanning. It cannot be assumed, therefore, that MI values give an absolute indication of the likelihood of the rupture or cracking of microbubble shells, or of associated bioeffects, that is universally valid for all machines, scanning modes and ultrasound contrast agents. It is shown in Chap. 3 that the threshold pressure value for in vitro bioeffects in the presence of shelled microbubbles appears to be proportional to frequency, rather than the square root of frequency, as inherent in the MI definition. Above the threshold pressure, bioeffects may increase in proportion to a power higher than the square of pressure. Until an index more appropriate to encapsulated microbubbles is introduced, display of the basic acoustic parameters p⫺ and fc might be helpful.
215
Exposure parameters of relevance to tissue heating Heating in tissue exposed to diagnostic ultrasound is due to two separate causes: direct absorption of ultrasound wave energy in the tissue and heat conducted through superficial tissues from a self-heated ultrasound probe. The precise temperature rise produced by absorption in any particular situation is very difficult to predict as it depends on many factors: the intensity distribution in the tissue, the ultrasound absorption coefficients of the various exposed tissues, the dimensions, specific heat and thermal conduction of the various tissues and the degree to which blood flow provides heat clearance. Most modern machines display a number (thermal index TI) on the screen that gives a very rough indication (within a factor of two) of the worst-case temperature rise that might be expected anywhere in the exposed tissues, according to very simple models of the tissue and the intensity distribution. As for the MI, the TI value is constantly updated according to the settings of the controls that influence the transmitted intensity and its spatial distribution, notably the choice of operating mode (B-mode, spectral Doppler, low MI or high MI contrast specific modes, etc.), output power and the depth at which the operator sets the transmission focus. Considerations relating to measurements of pressure in water: derating and nonlinearity Measurements are normally made in water for convenience but, ideally, values close to those found at the depth of a particular target in tissue (in situ values) would be preferred. Some standards and regulations make reference to “derated” values as a practical rough substitute for actual in situ values. Derated values are “in-water” values, reduced by ␣.fc.z decibels, representing the attenuation that would be introduced by a medium with an attenuation coefficient of (␣ dB cm⫺1 MHz⫺1), where z (cm) is the distance of the measurement site from the transducer, and fc (MHz) is the pulse centre frequency. In the definition of MI given in the previous section, the subscript 0.3 specifies a derating factor of 0.3 dB cm⫺1 MHz⫺1. This is somewhat less than the attenuation coefficients typical of soft tissue, but it allows for occasional low attenuation liquid regions in the sound path in general abdominal and obstetric studies. In formulating a more appropriate index for contrast agent work, as mentioned earlier, consideration might be given to a larger derating factor in view of the increased attenuation introduced by ultrasound contrast agents, and the greater tissue attenuation typical of common contrast applications such as transthoracic imaging of the heart or liver. A further consideration is that, at diagnostic frequencies and pressures, severe nonlinear propagation effects lead to pulse wave distortion, generation of har-
216
Ultrasound in Medicine and Biology
Volume 33, Number 2, 2007
monics, excess energy absorption and acoustic saturation (Duck 2002). This is demonstrated in the pulse waveforms in the upper part of Fig. 1. As the voltage drive applied to the transducer increases, the magnitude of the peak negative pressure becomes less than it would be in the absence of nonlinear distortion and the location of peak negative pressure moves toward the transducer (Duck and Starritt 1986). These effects appear more readily in water than in tissue, and are greater for greater transmitted pulse amplitudes. As a general result of these effects, the in situ peak negative pressure may be underestimated by up to a factor of two, being greater for larger depths (Duck and Bacon 1988). The effect of nonlinear propagation on in-water measurements of MI is less, since the formulation of the displayed MI requires p0.3 to be measured at a specific depth, which is usually comparatively close to the transducer. Using present protocols, it seems safe to assume that an in-water measurement of MI will underestimate a true in situ value by no more that 20% (Christopher 1999; Cahill and Humphrey 2000). One solution that has been proposed to the problem is that in situ peak negative pressures should be estimated by linear extrapolation from in-water measurements made at sufficiently low amplitude to assure quasilinear propagation. Spatial distribution of the acoustic field Normally, in soft tissue, negligible acoustic pressure will be produced at depths greater than those that can be usefully imaged. However, this may not be true in the rare event that extremely high reject settings or very low dynamic range or gain settings are selected. Where it is necessary to increase the transmitted amplitude to image targets behind strongly attenuating structures, for example when imaging intracranial targets, the pressure amplitude at the target may be much higher than when imaging through soft tissue. Consider a particular type of target imaged via a soft tissue path that returns an echo of sufficient amplitude to be just detected by a given probe. If the soft tissue path is replaced by a barrier introducing K times more attenuation (e.g., the skull wall), the transmission amplitude will have to be increased by a factor of K2 to compensate for the increased two-way attenuation. More significantly, from an exposure point of view, the amplitude at the target will now be K times greater than for the soft tissue path. In nonscanned modes (e.g., A-mode, M-mode and spectral Doppler mode) the beam cross-section from an array probe is rectangular close to the probe, becoming elliptical in cross section at depth. This is illustrated in Fig. 2, where the horizontal plane is the scan plane and the vertical plane is the elevation (slice thickness) plane. In the scan plane dimension, the beam is narrowest at the range of the “focus” as set by the operator, while in the
Fig. 2. Simplified representations of the acoustic field in different modes. Top: Nonscanned beam with the scan plane focus coincident with the slice thickness focus, together with typical pulse pressure waveforms measured in water at different ranges in Doppler mode. Centre: As above, with the scan-plane focus positioned deeper than the slice thickness focus. Bottom: Scanned mode with weak focusing in the scan plane.
slice thickness dimension it is narrowest at the fixed focus of the cylindrical lens at the front of the probe. Pulses with large peak pressures will be produced when the scan plan focus is set to equal the range of minimum slice thickness, as this concentrates the pulse power into the smallest beam cross-section. In scanning modes, such as in B-mode and Doppler Imaging modes, the acoustic field is in the form of a slice, with a width (scan plane) and thickness (elevation plane) that vary with range. Generally, the field width at all ranges is approximately equal to that of the probe face, but if a narrow “zoom box” (see section “Machine control settings. . .”) is selected, the strong scan-plane focusing of individual beams means that the field width will be less at the range of the box. The slice thickness (elevation plane) varies with range in a manner determined by the probe’s fixed cylindrical lens. It is equal to the height of the lens close to the probe, reducing to a minimum at the fixed range of the focus of the lens, and then increasing with range.
Exposure from diagnostic ultrasound equipment ● T. A. WHITTINGHAM
217
ducer further from the scan plane. The resulting acoustic field is more uniform in pressure amplitude versus range, but the pulse centre frequency reduces as the range increases.
Fig. 3. Example of the variation of on-axis pulse peak negative pressure with depth in B-mode. (a) In-water; (b) Derated by 0.3 dB MHz⫺1 cm-⫺; (c) Derated by 0.6 dB MHz⫺1 cm⫺1.
In all modes, the variation of in situ peak pressure with depth will be influenced by the attenuation coefficients of the particular tissues involved. Figure 3 shows the effect of applying two different attenuation coefficients to an axial profile of in-water peak pulse pressure versus depth, to estimate the in situ profiles. The range of maximum in situ peak pressure will always be closer to the probe than the maximum in-water peak pressure. This may not be a big effect if the focusing action is pronounced and the attenuation is low [curve (b)], but attenuations typical of soft tissues [curve (c)] can have a large effect. The centre frequency of the in situ pulse will also be less than that measured in water, reducing with increasing depth, due to the greater absorption by tissues of the higher frequencies in the pulse spectrum. A complication of estimating in situ pressures when using contrast agents is that the microbubbles themselves increase attenuation by absorption. This extra absorption may lead to underestimation of thermal hazard by the Thermal Index. Some probes (so-called 1.5 D arrays) have several rows of elements to provide a variable focus in the slice thickness dimension. When a single (scan-plane) focus is selected by the operator, the machine can then automatically set the slice thickness focus to be at the same depth. This increases the peak pressure at the focus, as shown in Fig. 2. If several foci are selected at different ranges, the overall transmission slice thickness and peak pressure can be made more uniform with depth. A more uniform distribution of pressure amplitude with depth is also achieved by those probes that use an Hanafi lens. This type of probe has a piezoelectric transducer element with a cylindrically concave transmitting face, resulting in the transducer being thinnest in the scan plane. Higher frequencies are transmitted from the thinner central part of the transducer to ranges close to the probe, while principally lower frequencies are transmitted to greater depths from the thicker parts of the trans-
Temporal variation of pressure in different modes The waveform of a single pulse and the temporal variation of peak pressure amplitude over a complete frame depend on the mode selected. A simplified representation of the pressure variation versus time at one point in front of the probe is shown in Fig. 4. In practice, more pulses than those shown may be experienced at a particular field point, particularly at positions between the probe and the transmission focus. This is due to the practice of transmitting pulses along other scan lines while waiting for echoes to return from a scan line (beam multiplexing). In A-mode or M-mode, pulses typically have two to three cycles. Since the beam is not scanned, there is no variation from pulse to pulse. In B-mode, the pulses are similar in shape and amplitude to those used in A- and M-modes, but as the beam is swept across the field of
Fig. 4. Simplified representation of the variation of peak pulse amplitude with time over a complete frame, at one point in front of the probe. Each pulse is shown by a single line, the height of which is proportional to the peak amplitude of the pulse. Expanded pulse waveforms are shown to the left of each frame representation.
218
Ultrasound in Medicine and Biology
view, a given field point experiences almost zero exposure for most of each frame, and a transient rise and fall of pulse amplitude as the beam passes over. If a narrow write zoom box is selected (see the following section), the temporal average intensity is likely to be increased at all depths, since the transmitting aperture is likely to be increased and hence a given field point will be in a beam (insonated) for a greater proportion of each frame. In tissue harmonic imaging (THI) mode, the transmitted pulses are likely to have a lower centre frequency than typical B-mode pulses, since the transducer has to be able to operate at both the transmission centre frequency (fundamental) and the second harmonic (twice the transmission frequency). Pressure amplitudes are normally at the high end of the range used for B-mode, to promote second harmonic generation when propagating in tissue. If a filter is used to suppress the fundamental echo component, the need to avoid overlap between the spectra of the transmitted pulse and that of the echoes means that narrower bandwidth and, hence, longer pulses are transmitted. If the pulse inversion technique is used, two pulses are transmitted per scan line, one being an inverted version of the other. These transmitted pulses can have normal bandwidth and, hence, pulse length, although some effort is often made to reduce their second harmonic content. In Spectral Doppler, pulses are longer (typically 3 to 20 cycles, depending on the width of the range gate), at a higher pulse repetition frequency (PRF) to reduce aliasing, and generally with a lower centre frequency. To keep time-averaged intensities within FDA regulatory limits despite the higher PRFs, pressure amplitudes are usually lower than in A-, M- or B-mode, becoming less as range-gate width increases. In colour Doppler imaging modes, typically 8 to 20 pulses are transmitted along each scan line before the beam is advanced (the larger number being for a high “Colour Quality” control setting). A sequence of fourteen pulses per Doppler line is shown in Fig. 4. The individual pulses are similar to those used in spectral Doppler for a short range-gate setting. As the beam is advanced across a given field point, the peak pulse amplitude builds up and then dies away, being almost zero for most of the frame. Note that in Fig. 4 one frame of colour Doppler imaging/LOC pulses is shown occupying the same space as for other modes. In practice, the need to transmit 8 to 20 pulses down each scan line means that the time taken to complete a colour Doppler frame is likely to be several times longer than for other modes. Contrast-specific modes for use with encapsulated microbubbles are commonly categorised as either “low MI” or “high MI”, although, as mentioned earlier, the appropriateness of the MI safety index is questionable. For a review of the various methods see, for example,
Volume 33, Number 2, 2007
Whittingham (2005). The pulses transmitted in low MI modes have low pressure amplitudes (around 0.1 MPa or less, with typical MI values of around 0.1 or less), specifically to reduce the likelihood of microbubble rupture. High MI modes, however, rely on the rupture of microbubbles using sensitive detection methods, which essentially compare echo sequences before and after microbubble rupture. The likelihood of cavitation with high MI contrastspecific techniques is higher than with noncontrast techniques due to the high MI values (around 1 or more) and the presence of cavitation nuclei from cracked or ruptured microbubbles. The likelihood is much less for low MI contrast-specific techniques, of course, but low MI techniques can involve more and longer pulses, and often involve MI values that are above the thresholds reported for in vitro bioeffects and cavitation in the presence of contrast agents. For example, the threshold of p–/f ⫽ 0.06 MPa/ MHz reported in Chap. 3 for in vitro bioeffects implies a MI threshold of only 0.06 at 1 MHz and 0.12 at 4 MHz. Also, Chen at al (2003) have detected white noise emission from insonated contrast agents, consistent with inertial cavitation, at MI values lower than 0.2. One common low MI contrast-specific method uses the pulse inversion technique of second harmonic detection mentioned above for Tissue Harmonic Imaging. As with THI, two pulses are transmitted along each line, but the peak negative pressures are much lower than in THI. Some low MI techniques involve the transmission of more than two pulses along each line (Table 4). These give greater sensitivity and improved discrimination between microbubble and tissue echoes. Such multiple transmissions can give rise to greater temporal-average intensities and powers than in normal B-mode and, hence, the potential for tissue heating can be greater. This will be indicated by a higher TI value on the screen. A selection of different low MI mode transmissions are described in Table 4. Loss of correlation (LOC) imaging is a common high MI method that uses similar transmission sequences to those used in colour Doppler imaging, although with higher pulse amplitudes (MI ⬃1). Recent versions use fewer (2– 6) and shorter (broad bandwidth) pulses and higher pulse repetition frequencies. “Pulse subtraction” is a high MI method in which only two short but high amplitude pulses are transmitted along each line. Since high MI imaging methods destroy the microbubbles, the probe is swept by hand across the patient, resulting in exposure times of only a few seconds at any one site. When used in “flash” or “burst” mode to clear microbubbles from a field of view, long sequences of large amplitude pulses are transmitted along each scan line in every frame. This is normally repeated for just a few frames, but in some instances may continue as long
Exposure from diagnostic ultrasound equipment ● T. A. WHITTINGHAM
as the mode is selected. In Fig. 4 a sequence of seven large amplitude pulses is shown being sent down each scan line. As the beam passes over the measuring hydrophone, the amplitude of the detected sequence rises and falls. Machine control settings that give the highest exposures The nature and range of controls is constantly changing with the evolution of new scanning techniques, so universal rules are difficult to formulate. Controls on some of the newer machines can have quite unexpected effects, as manufacturers often arrange for drive voltages or pulse repetition frequencies to change automatically when controls are set in a way which would otherwise cause a particular safety parameter, such as intensity, MI or TI, to exceed a regulatory limit. Apart from the obvious example of increasing the setting of the output power control, there are a few general observations that can be made about producing pulses with high pressure amplitudes. Some general observations on control settings that affect pulse repetition frequency (PRF) are also possible. In general, selecting a deep transmission focus commonly involves an increase in p⫺ and heating, whichever mode is chosen. This is because manufacturers often arrange for the transmission aperture to be increased if a deep transmission focus is selected, to maintain a narrow beam width and good sensitivity at depth. Apart from an increase in pressure amplitude due to more elements transmitting, the drive voltage applied to each element may also be increased to compensate for the greater attenuation anticipated for deep targets. Activation of a write-zoom box is another way by which exposure can be increased, particularly if the box is narrow. Unlike read-zoom, which simply magnifies part of the stored image, write-zoom involves a selected smaller area being re-scanned at a higher line density and/or frame rate. This leads to a given field point being insonated for a larger fraction of the frame time and thus being insonated by more pulses per second. Temporal average intensities and hence tissue heating will also increase. Write zoom may also lead to a higher pulse repetition frequency, since echoes from beyond the box are not wanted. If the zoom box is deep, there may also be an increase in p⫺, since the machine may automatically advance the focus to match the depth of the box, leading to an increase in p⫺, as mentioned above. A large p⫺ is usually produced if the operatorcontrolled focus (which sets the transmission focus in the scan plane) is set close to the (fixed) slice thickness focus, since this increases the strength of focusing in a 3-D sense (see Fig. 1). However, as discussed above, setting the focus (or range-gate in the case of spectral Doppler) to a greater depth may well increase the trans-
219
mission aperture and drive voltage, and hence produce even greater pressures and intensities near the scan-plane focus. Only practical measurements can establish where the greatest pulse pressure will be actually produced. In spectral Doppler mode, the only control that is likely to change the pulse length is the gate width (sample volume width) control. A high Doppler frequency scale setting, or the selection of “High PRF” mode, is likely to produce lower pressure amplitudes. In themselves, these controls would not be expected to affect p⫺ values, but manufacturers sometimes arrange for drive voltages and, hence, pressure amplitudes, to be reduced if a high PRF is selected, to keep intensities within prescribed limits. For the same reason, drive voltages are usually reduced as gate width (and hence pulse length) increases, so that higher p⫺ values are likely to be found for narrow gate settings. Controls that are unlikely to directly change the transmitted acoustic field include: Overall gain (2D Gain), TGC, dynamic range, reject, grey scale preprocessing or postprocessing (transfer characteristic), frame averaging (persistence), read-zoom (magnification), Doppler gain, Doppler filter, Doppler baseline, cursor angle (angle correction), spectrum inversion, colour gain, colour priority. Tables 1, 2, 3 and 4 summarise the function and effect on exposure of a selection of machine controls. Survey of worst-case values and trends When considering the hazard potential of an ultrasound field, it is of interest to try to find the “worstcase” values of the measured quantities. This involves repeating the measurements for the many different combinations of machine control settings that are likely to give a high output, until the combination that gives the largest value for the particular quantity is found. Contrast-specific modes have only recently been
Fig. 5. Worst-case in-water p⫺ values measured between 1995 and 1999.
220
Ultrasound in Medicine and Biology
Volume 33, Number 2, 2007
Table 1. General controls and their effect upon exposure Control name
Function
Effect on exposure
Frequency (where multi-frequency probes are selected)
Allows an optimum balance of spatial resolution and penetration to be achieved. Higher frequencies give better resolution but less penetration.
Output power
One way of controlling the overall sensitivity and penetration. (The other way is to use Overall Gain). Allows lateral resolution and sensitivity to be increased at a preferred depth by placing the transmission focus there.
Changes the centre frequency of the transmitted pulse. Lower frequencies often mean an increase in in-situ pulse amplitude (less attenuation). However, if the focus and aperture remain constant, changing to a higher frequency may cause the pulse pressure amplitude near the focus to be increased due to the narrower beam width. The higher the power setting, the greater the pressure amplitude of every transmitted pulse.
Focus depth
Number of foci
As above but for several different focus positions (at expense of frame rate)
Displayed depth
Sets the maximum depth of the field of view
Displayed width or sector width
Sets the width of the field of view
Frame rate (some machines)
Allows trade-off between image quality and frame rate in 2D or Colour Imaging modes Write Zoom allows greater spatial resolution and/or frame rate to be achieved by restricting scanning to a more limited area, as defined by the zoom box on the image. Max Depth sets the deeper margin of the zoom box
Write-zoom box - maximum depth
Write-zoom box-width Frame freeze or cine loop (cine review) Compounding THI mode
Sets the horizontal width of the zoom box Allows an image frame or sequence of frames to be reviewed for as long as required Improves delineation of interfaces; reduces speckle and noise Reduces noise artefacts due to sidelobes and reverberations
introduced, and there have been no published surveys of worst-case pressure or intensity values for these modes. Measurements and reviews of worst-case inwater values of p⫺, carried out for conventional modes over the last three decades (Whittingham 2000), show that those measured in the half decade 1995 to 1999 were higher than ever before, with evidence of a slow continuing upward trend. The increased likelihood of cavitation that this increase in p⫺ might imply is partly balanced for conventional modes by the fact that pulse centre frequencies have
Increasing the depth of the transmission focus will move the point of greatest pulse amplitude to a greater depth (not necessarily at the focus itself). It is also likely to increase the pulse pressure amplitude at all depths. Some machines may also lower the centre frequency of the pulse at the same time. Setting the focus to the same depth as the lens elevation plane focus (slice thickness minimum) produces larger peak pressure amplitude at that depth. Several pulses transmitted along each scan line, each with a focus at a different depth. Effect for each pulse / focus as above. Larger depth settings cause the PRF to be reduced. On some machines the pulse centre frequency may be reduced and the pressure amplitude may be increased. Smaller widths can result in more closely spaced scan lines (each point then experiences more pulses as it lies in more beams) and higher frame rates (less time before the beam returns to insonate each point). Results in more pulses at any given point and greater temporal average intensities. Low frame rates result in a greater line density. Each point then experiences a longer sequence of pulses. Selecting a deep box will create a matching deep focus - increasing the pulse amplitude as described previously for Focus Depth. A shallow maximum depth of box may cause the PRF to be increased and possibly the pulse centre frequency to be increased. As for Displayed Width or Sector Width. Normally stops transmissions, but not on every machine. The number of pulses insonating each point in the field may increase. Reduces the centre frequency of the pulse. If pulse inversion is used, two pulses are transmitted down each scan line. If a filter is used instead, longer pulse lengths may be used.
generally been increasing in the quest for greater spatial resolution. The measurements for the 1995 to 1999 period are presented graphically in Fig. 5. Little difference was found between M- and B-mode values, so they are grouped together. Overall, worst-case values are distributed about mean values of 2.4 to 2.8 MPa, depending on the mode. Median worst-case values are in the range 2.4 to 2.6 MPa. There is very little difference among modes, although perhaps it is worth noting that, in this survey, the most extreme p⫺ values (exceeding 5 MPa) occurred in spectral Doppler mode. Surveys of MI values
Exposure from diagnostic ultrasound equipment ● T. A. WHITTINGHAM
221
Table 2. Spectral Doppler mode controls and their effect upon exposure. A continuous stream of pulses is transmitted down a stationary beam. PRFs are higher than in A-, M- and B-modes and peak pressures are generally lower Control name
Function
Effect on exposure
Gate width (sample volume)
Increases the length of the sample volume, i.e., the range of depths from which Doppler signal is obtained.
Scale (PRF on some machines)
Allows a larger range of Doppler frequencies to be displayed on the Doppler spectrum.
High PRF
Allows aliasing to be eliminated in exchange for ambiguity in the depth of the sample volume (More than one range gate displayed) Controls the angle between the ultrasound beam and the blood flow.
Larger gate width settings increase the number of cycles in each pulse. Widening the gate commonly cause the pulse amplitude to be reduced. Increasing the scale (larger Doppler, frequency range) causes the PRF to be increased in proportion. Increases the PRF by a factor equal to the number of range gates displayed.
Doppler line steering
have not been reported. However, instances of erroneous MI values have been found. Estimation of worst-case in situ pressure amplitudes Although contrast agents are not currently licensed for obstetric applications, the importance of the fetus from a hazard perspective promotes the consideration of exposures in such applications. Estimates have been reported of either the typical or minimum attenuations introduced by tissue between the probe and the fetus. A figure of 0.5 dB MHz⫺1 has been proposed for minimum attenuation in the third trimester (Carson et al. 1989), while a figure of 3.6 dB MHz⫺1 has been reported for typical attenuation in the first two trimesters (Siddiqi et al. 1995). Table 5 shows the result of applying these attenuating factors to the in-water values of worst-case p⫺ upon which Fig. 5 is based, to give rough estimates of worst-case in situ p⫺ values at the fetus (Whittingham 2001). The minimum attenuation path third trimester model predicts that worst-case values in excess of 3 MPa are possible in all conventional modes, but in the “typical” model the upper limit is reduced to around 1 MPa. It should be noted that the models take no account of the reduced pulse distortion that would be associated with propagation in tissues, compared with the water in which the measurements were made. The in situ negative pressure values could, therefore, be greater than these estimates. CONCLUSIONS Conventional modes (i.e., non-contrast-specific modes) p⫺ values in conventional modes have increased several-fold since 1970 and have shown a slower, but still rising, trend over the last decade.
Increasing the deflection from the “straight ahead” direction can result in reduced pressure amplitude and broader beams (hence more pulses insonating each point).
The range of reported worst-case p⫺ values shows little difference between A-mode, M-mode, B-mode, spectral Doppler, CFM and power Doppler. However, typical pulse pressure amplitude values are generally higher in A-mode, M-mode and B-Mode than in Doppler modes. There are large differences in reported worst-case p⫺ values between machines from different manufacturers. Worst-case in-water p⫺ values have been reported to be as high as 4.6 MPa in M- and B-mode, 4.9 MPa in colour Doppler and 5.5 MPa in spectral Doppler.
Table 3. Colour Doppler imaging controls and their effect upon exposure. Several pulses are transmitted down each Doppler line of the colour box. PRF and pulse centre frequency are likely to be similar to those in spectral Doppler Control name Colour quality
Colour box steering
Function
Effect on exposure
High quality settings increase the consistency and accuracy of the colour, and allow lower flow velocities to be shown in colour. Controls the angle between the ultrasound beams in the colour box and the blood flow.
High quality settings increase the number of pulses transmitted along each scan line.
Increasing the deflection from the “straight ahead” direction can result in reduced pressure amplitude and broader beams (hence more pulses insonating each point).
222
Ultrasound in Medicine and Biology
Volume 33, Number 2, 2007
Table 4. Contrast-specific modes and their effect upon exposure. Commonly divided into High MI and Low MI modes. High MI techniques are associated with microbubble destruction and hence intermittent scanning. Low MI techniques try to prolong the lifetime of bubbles modes in order to allow extended real-time imaging Mode
Function
Effect on exposure
High MI LOC scanning
Images the distribution of contrast agent microbubbles by destroying the microbubbles and comparing the echo sequences before and after.
Flash or burst
Deliberately clears contrast agent from the field of view for reperfusion studies by rupturing the microbubbles. Discriminates microbubble echoes from tissue by their larger second harmonic content.
High MI (usually ⬎ 0.7) pulses transmitted which rupture the contrast microbubbles. Older systems transmit long sequences (⬃10) of longer pulses (3–5 cycles) along each line, as for Colour Doppler Imaging. Some recent methods use fewer (2–4) short (2–3 cycles) pulses and higher PRFs. Transmits a long sequence (⬃10) of large amplitude pulses along each scan line.
Low MI harmonic imaging by high-pass filtering
Low MI pulse inversion
As above.
Low MI coherent pulse inversion
As above.
Low MI coded excitation
As above
Low MI chirp coding
As above
Low MI contrast pulse sequence imaging (CPS)
Discriminates microbubble echoes from tissue by the non-linear signal in the fundamental echo component or by detecting a harmonic echo component.
Estimates of in situ pressure values at the fetus for obstetric models indicate typical p⫺ values of less than 0.4 MPa, but worst-case values as high as 2 MPa. There is no simple rule for predicting the depth at which the spatial maximum of in situ negative pressure will occur. The controls that most directly influence p⫺ are the output power and focus depth, but other controls can have an indirect influence due to the need to keep MI and ISPTA within regulatory limits. Spectral Doppler mode and Doppler imaging modes generally involve the highest pulse repetition frequencies and pulse lengths, and the lowest centre frequencies. Contrast-specific modes These modes are customarily divided into “high MI” and “low MI” techniques according to whether or not they are likely to cause rapid destruction of the contrast microbubbles. In high MI modes, two or more large amplitude pulses (MI ⬃1) are transmitted along each scan line. The large amplitude and the presence of transient small bub-
Longer transmitted pulses than in B-mode, with low amplitudes; sometimes with waveforms (e.g. Gaussian) designed to have low harmonic content. Centre frequency of the pulses may be reduced to facilitate detection of second harmonic. Two short pulses (more typical of B-mode), but with low amplitudes and the second pulse inverted, are transmitted along each line; Sometimes low harmonic transmission waveforms and reduced centre frequency, as above. One short low amplitude pulse transmitted along each line, inverted on alternate lines. Centre frequency may be reduced as above. Two dissimilar, long (8 or more cycles), low amplitude pulses transmitted along each line. Centre frequency may be reduced as above. One long (⬎20 cycles) low amplitude chirp (swept frequency) pulse transmitted along each line. Three or more low amplitude pulses transmitted along each line, sometimes with different amplitudes and/or with some pulses inverted. Centre frequencies may be reduced where harmonics are to be detected.
bles of free gas due to the fragmentation of contrast microbubbles means that the likelihood of cavitation is greater than in conventional imaging modes or low MI contrast modes. Low MI modes of operation generally involve the transmission of two or more low amplitude pulses along each scan line, sometimes with different amplitudes and polarities. Some techniques involve the transmission of a single long (up to 30 cycles) low amplitude chirp pulse along each line. Centre frequencies may be lower than for conventional B-mode. In low MI contrast modes, p⫺ values are likely to be low, so the potential for other nonthermal effects is also likely to be lower, although this may be offset to some extent by the fact that each field point is likely to experience more pulses per frame. Temporal average intensities and powers may be higher when long sequences of pulses are transmitted. The potential for thermal effects may then be increased. Low MI techniques can involve MI values that are above the thresholds reported for in vitro bioeffects and cavitation in the presence of contrast agents.
Exposure from diagnostic ultrasound equipment ● T. A. WHITTINGHAM
223
Table 5. Median (and maximum) values of worst-case in-water p- from diagnostic equipment, together with values obtained by applying attenuation factors of 0.5 dB MHz⫺1 and 3.6 dB MHz⫺1 respectively M-mode & Bmode (MPa)
Colour Doppler (MPa)
Spectral Doppler (MPa)
Attenuation dB MHz⫺1
Median
(Max)
Median
(max)
Median
(max)
0.5
2.4 1.8
(4.6) (3.1)
2.6 2.0
(4.9) (3.9)
2.4 1.7
(5.5) (3.2)
3.6
0.4
(1.0)
0.4
(1.1)
0.3
(1.1)
In water Minimum attenuation - 3rd trimester “Typical” attenuation - 1st an 3rd trimesters
No surveys have yet been published that give acoustic output data for contrast-specific modes. General Although widely available on display screens as a guide to the likelihood of nonthermal bioeffects, the mechanical index (MI) is not necessarily an adequate guide to the likelihood of microbubble rupture or of bioeffects associated with ultrasound contrast agents. Contrast agent may enhance absorption and, hence, heating. This may have implications for the proper choice of derating factor and absorption coefficient for use in thermal and mechanical indices when using contrast agents. In situ peak negative pressure can be significantly underestimated if account is not taken of nonlinear propagation effects. Displayed MI may be underestimated by up to 20% because linear assumptions are made in the calculation of in situ peak negative pressure. REFERENCES Apfel RE, Holland KH. Gauging the likelihood of cavitation from short pulse low duty cycle diagnostic ultrasound. Ultrasound Med Biol 1991;17:179 –185. Cahill MD, Humphrey VF. A theoretical investigation of the effect of nonlinear propagation on measurements of mechanical index. Ultrasound Med Biol 2000;26:433– 440. Carson PL, Rubin JM, Chiang EH. Fetal depth and ultrasound path lengths through overlying tissues. Ultrasound Med Biol 1989;15:629 – 639.
Chen WS et al. A comparison of the fragmentation thresholds and inertial cavitation doses of different ultrasound contrast agents. J Acoust Soc Am 2003;113:643– 651. Christopher T. Computing the mechanical index. J Ultrasound Med 1999;18:63– 68. Ciaravino V, Flynn HG, Miller MW. Pulsed enhancement of acoustic cavitation: A postulated model. Ultrasound Med Biol 1981;7:159 – 166. Duck FA. Nonlinear acoustics in diagnostic ultrasound. Ultrasound Med Biol 2002;28:1–18. Duck FA, Bacon DR. A fundamental criticism of hydrophone-in-water measurement. Ultrasound Med Biol 1988;14:305–307. Duck FA, Starritt HC. The locations of peak pressures and peak intensities in finite amplitude beams from a pulsed focused transducer. Ultrasound Med Biol 1986;12(5):403– 409. Hartman C, Child SZ, Macer R, et al. Lung damage from exposure to the fields of an electro-hydraulic lithotripter. Ultrasound Med Biol 1990;16:675– 679. Holland KH, Apfel RE. An improved theory for the prediction of micro-cavitation thresholds. IEEE. Trans Ultrason Ferroelectrics Fre Qcontr UFFI- 1989;36(2):139. Morton KI, ter Haar G, Stratford IJ, Hill CR. The role of cavitation in the interaction of ultrasound with V79 Chinese hamster cells in vitro. Br J Cancer 1982;45(Suppl V):147. Siddiqi TA, O’Brian, WD. Jr, Meyer RA, Sullivan JM, Miodovnik M. In situ human obstetrical ultrasound exposimetry: Estimates of derating factors for each of three different tissue models. Ultrasound Med Biol 1995;21:379 –391. Whittingham TA. Acoustic outputs of diagnostic machines. In: ter Haar G, Duck FA, eds. Safety of medical diagnostic ultrasound. London: British Institute of Radiology, 2000;16 –31. Whittingham TA. Estimated cerebral ultrasound exposures from clinical examinations. Ultrasound Med Biol 2001;27(7):877– 882. Whittingham TA. Contrast-specific imaging techniques: Technical perspective. In: Quaia E, ed. Contrast media in ultrasonography. Berlin: Springer, 2005; 43–70.