The characteristics of an ultrasound beam, the propagation of ultrasound in tissue and the design of transducers as found in B-mode imaging are all relevant for Doppler techniques.^2–6

The Doppler Effect and Its Application

For all waves such as sound or light the Doppler effect is a change in the observed frequency of the wave because of motion of the source or observer. This is due either to the source stretching or compressing the wave or the observer meeting the wave more quickly or slowly as a result of their motion. In basic medical usage of the Doppler effect, the source and observer (receiver) are a transmitting and a receiving element usually positioned next to each other in a hand-held transducer (Fig. 1-1A). A continuous cyclic electrical signal is applied to the transmitting element and therefore a corresponding CW ultrasound beam is generated. When the ultrasound is scattered or reflected at a moving structure within the body, it experiences a Doppler shift in its frequency and returns to the receiving (detecting) element. Reflected ultrasound is also detected from static surfaces within the body but it has not suffered a Doppler shift in frequency. After the reflected ultrasound is received, the Doppler instrument separates the signals from static and moving structures by exploiting their different frequency.

FIGURE 1-1 Sample volumes in Doppler techniques. (A) For dual crystal continuous wave Doppler unit. (B) For pulsed wave Doppler unit. (C) Neighbouring sample volumes along a beam for imaging Doppler units.

Motion of the reflector towards the transducer produces an increase in the reflected ultrasonic frequency, whereas motion away gives a reduction. The system electronics note whether the detected ultrasound has a higher or lower frequency than that transmitted and hence extract information on the direction of motion relative to the transducer.

When the line of movement of the reflector is at an angle θ to the transducer beam, then the Doppler shift, f_D, is given by:

where f_t is the transmitted frequency, f_r is the received frequency, c is the speed of ultrasound and u.cosθ (i.e. u × cosineθ) is the component of the velocity of the reflecting agent along the ultrasonic beam direction. For a typical case of blood flow in a superficial vessel:

Transmitted frequency, f_t = 5 MHz = 5 × 10⁶ Hz

Velocity of sound in soft tissue, c = 1540 m/s

Velocity of blood movement, u = 30 cm/s

Angle between ultrasonic beam and direction of flow, θ = 45^°

The Doppler shift is therefore:

The shift in frequency is small and within the audible range. In an ultrasonic Doppler instrument, the electronics are designed to extract the difference in frequency, f_D = f_t – f_r (the Doppler shift frequency). The instrument can therefore feed a signal of frequency f_D to some output device such as a loudspeaker or frequency analyser.

So far we have considered an ultrasound beam being reflected from a structure moving at a fixed speed and hence generating a Doppler shift of one particular frequency. In practice, there are many reflecting blood cells and their speeds are different. The ultrasound signals returned to the detector from the different cells therefore have suffered different Doppler shifts and add together to give a complex signal containing a range of frequencies. The Doppler shift frequencies are extracted from the detected complex signal and can be fed to a loudspeaker where they can be interpreted by listening. High-frequency (high-pitch) components in the audible sound are related to high speeds, whereas low-frequency components correspond to low speeds. Strong signals, namely those of loud audible volume, correspond to strong echoes that have received a Doppler shift. Strong signals could be due to the detection of many blood cells, say in a large vessel, or to echoes from tissue. An output display called a spectral display or spectrogram is often used to portray the frequency content of Doppler signals.

In the PW Doppler technique, the electrical excitation signal is applied to the transmitter element at regular intervals as pulses, each containing typically 10 cycles, and therefore a corresponding train of pulses of ultrasound are transmitted, separated by non-transmission intervals of duration around 20 times that of each pulse. Regularly spaced echoes are then received back from a reflector and they can be regarded as samples of the signal which would be received if a continuous wave had been transmitted. If the reflector is moving the system electronics can extract a Doppler shift signal from the samples. The Doppler equation again applies to this Doppler shift and can be used to calculate the speed of the reflector.⁷

A bonus of PW Doppler is that since pulsed ultrasound is employed, the range of the moving target may be measured from the echo-return time, as well as its speed from the Doppler shift. The range can be measured from one echo signal; however, the calculation of the Doppler shifts and hence speed of the reflector, typically requires 50–100 echoes. As in the CW case, a group of blood cells moving with different velocities produces a range of Doppler shift frequency components in the output signal.

It was noted above that the frequency of reflected ultrasound is shifted upward or downward depending on whether the motion of the reflector is toward or away from the transducer. A numerical example illustrates this point and emphasises the small changes in frequency that the instrument must distinguish. When 2 MHz ultrasound is reflected from an object travelling at 30 cm⁻¹ toward the transducer, it returns to the receiver with a frequency of 2.00078 MHz, a shift of + 0.00078 MHz. If the object moves at 30 cm⁻¹ away from the transducer, the ultrasound returns with a frequency of 1.99922 MHz, a shift of −0.00078 MHz. Virtually all Doppler instruments which measure velocity preserve this direction information.

Continuous and Pulsed Wave Doppler Instruments

Doppler blood flow instruments are required to be extremely sensitive and to be capable of detecting weak signals from moving blood in the presence of much stronger signals from static or moving tissues; the latter give rise to low-frequency Doppler shift ‘clutter’ signals. The magnitude of the scattered signal from blood is typically 40 dB below that received from soft tissues, i.e. the blood echo amplitude is typically one hundredth of the soft tissue echo amplitude. The dB (decibel) unit is a measure of the size of a signal relative to another signal; the second signal is often a reference signal or perhaps the input signal to an amplifier to which the output is compared. Blood flow signals may be detected even though the vessel is not clearly depicted, for instance in the fetal brain, or the renal artery of the neonate.

The transducer of a basic CW Doppler unit has two independent piezoelectric elements. Since the transmitting element is continually driven to generate a continuous wave of ultrasound, a second element is used to detect the reflected ultrasound. When a CW Doppler mode is implemented as part of an ultrasound system which uses array transducers, separate groups of array elements are used for transmission and reception. On extraction of the Doppler shift frequency a filter, the ‘wall thump’ filter, is often used to remove large, low-frequency components from the signal, such as those from slowly moving vessel walls. Typically in a Doppler unit operating at 5 MHz, Doppler shift frequencies below 100 Hz are removed by filtering. Basic CW Doppler instruments are small and inexpensive; CW Doppler mode facilities are incorporated into some array systems to allow them to detect high velocities (see section on Aliasing artefact, below).

The transmitted ultrasound field and the zone of maximum receiving sensitivity overlap for a particular range in front of the transducer (Fig. 1-1A). Any moving structure within this region of overlap will contribute a component frequency to the total Doppler signal. The shape of the region of overlap (the beam shape) can be considered as having a crude focus which depends on the field and zone shapes and on their angle of orientation to each other. In practice, the beam shapes are rarely well known for CW Doppler transducers. A 5 MHz blood flow instrument might be focused at a distance of 2 or 3 cm from the transducer and a 10 MHz device at a distance of 0.5–1 cm. CW Doppler instruments normally have ultrasonic output intensities (I_spta) of less than 10 mW cm⁻² although they may be significantly higher when used in conjunction with duplex systems to measure high velocities.

A PW Doppler instrument, operating with 5 MHz ultrasonic pulses, may have a pulse repetition frequency (PRF) of 10 000 per second, i.e. 10 kHz. The highest velocity that the instrument can measure is directly proportional to its PRF (see Aliasing artefact, below); therefore the PRF is made as high as possible while still avoiding overlap between successive echo trains. A train of echoes is produced as a transmitted pulse passes through reflecting interfaces and regions of scattering targets. After amplification, successive echo signals from a specific depth are selected by electronic gating and the Doppler shift frequency is extracted as described above.

Pulsed Doppler devices can be used on their own by altering slowly the beam direction or the gated range depth while listening to the output, for example in transcranial blood flow studies. Identification of vessels is made easier by combining the PW Doppler mode with a real-time B-scan mode to form a duplex system; however, this obviously adds to the cost and complexity.

Since the ultrasound is pulsed and the excitation time is short, a stand-alone PW unit uses a single crystal transducer for transmission and reception (Fig. 1-1B). On setting the electronic gate to select a signal from a specific range, reflectors within a volume, known as the sample volume, contribute to the signal. The shape and size of the sample volume are determined by a number of factors: the transmitted pulse length, the beam width, the gated range length, and the characteristics of the electronics and transducer. The sample volume is often described as a tear drop in shape (Fig. 1-1B). Sample volume lengths are usually altered by changing the gated range length. In a blood flow unit for superficial vessels, the sample volume length may be as short as 1 mm, whereas in a transcranial device it can be 1 or 2 cm; however, the precise lengths are rarely known.

The ultrasonic output intensity of pulsed Doppler instruments varies considerably from unit to unit. The intensity (see Safety section, below) may typically be a few hundred mW cm⁻² but can be as high as 1000 mW cm⁻², particularly when they are required to penetrate bone, as in transcranial Doppler. At present the most common use of stand-alone PW units is in transcranial examinations of cerebral vessels.

A summary of technical factors relating to the use of continuous wave and pulsed wave Doppler instruments is given in Box 1-1.

BOX 1-1 Technical Factors in the Use of CW and PW Doppler

1. Doppler beams are subject to the same physical processes in tissue as B-mode beams, i.e. attenuation, refraction, speed of sound variation, defocusing, etc.

2. Since stand-alone CW and PW units are used blind, the beam direction and also the sample volume in the PW case must be systematically moved through the region of interest to maximise both the volume and pitch of the audible Doppler signal.

3. PW Doppler is subject to the aliasing artefact in the measurement of high velocities, CW Doppler is not.

4. The sensitivity (gain, transmit power) of the Doppler unit should not be so high that noise detracts from the signal quality.

5. The instrument should be assessed on normal vessels where the blood flow pattern is known and the expected Doppler signal well understood.

6. The wall-thump filter should just be high enough to remove the strong low-frequency signal from vessel walls and any other moving tissue.

7. The final result in many cases should be a distinct display, called a ‘spectrogram’ or ‘sonogram’ (see section on the Spectrum analyser, below), with a clearly defined maximum-velocity trace.

8. Since the beam–vessel angle is unlikely to be known, the sonogram cannot be calibrated in velocity and the vertical axis remains as Doppler shift frequency.

9. Care should be taken to ensure good acoustic coupling between the transducer and the patient. Since there is no associated image it is not always apparent that a weak signal may be due to a lack of coupling agent.

10. If possible, information should be obtained on the shape of the sample volume for both CW and PW beams. The sample volume size can then be related to the size of the vessel under study. With CW Doppler there is very little depth discrimination. With PW Doppler the sample volume depth and size are set by the user.

Imaging and Doppler

There are three types of imaging used with Doppler techniques. The first, known as ‘duplex Doppler’, uses a real-time B-scanner to locate the site at which blood flow is to be examined then a Doppler beam interrogates that site. The second type creates an image from Doppler information, i.e. an image of velocities in regions of blood flow.⁸ Known as ‘colour Doppler’, ‘colour flow imaging’ or ‘colour velocity imaging’, it is normally combined with a conventional real-time B-scan so that both tissue structure and areas of flow are displayed. The third type of Doppler imaging is similar to colour Doppler, but generates an image of the power of the Doppler signal from pixel locations throughout the field of view and is known as ‘power Doppler imaging’ (power Doppler).⁹ A power Doppler image depicts the amount of blood moving in each region, i.e. an image of the detected blood pool.

DUPLEX INSTRUMENTS

Duplex systems link CW or PW Doppler features and real-time B-scanners so that the Doppler beam can interrogate specific locations in the B-scan image (Fig. 1-2). CW duplex is normally only used where very high velocities have to be measured without the aliasing artefact, for example in the estimation of the velocity of a jet through stenosed heart valves. The direction of the CW beam is shown as a line across the B-scan image. In the case of PW Doppler, markers on the beam line show the position of the sample volume. The Doppler beam is often directed across the field of view so that it does not intersect the blood flow at 90°.

FIGURE 1-2 Duplex system combining real-time B-mode (- – – – ) and a Doppler beam (—-) of variable position across the B-mode field of view.

In duplex systems, the transmitted ultrasound frequency in the Doppler mode is often lower than that for the B-mode. The low Doppler beam frequency is to enable higher velocities to be handled before aliasing occurs, while the high B-scan frequency is to optimise resolution in the image. An example could be 5 MHz for Doppler and 7 MHz for B-mode when studies of superficial vessels are undertaken.

A summary of technical factors relating the use of duplex Doppler instruments is given in Box 1-2.

BOX 1-2 Technical Factors in the Use of Duplex Doppler

1. The factors quoted above for CW and PW Doppler may also be relevant.

2. A spectrogram can be obtained from a known vessel and a known location within the vessel.

3. The Doppler beam may be refracted and not pass along the line shown in the B-mode image.

4. The beam–vessel angle may be measured manually, allowing blood velocity to be estimated.

5. Simple estimates of blood flow rate may be obtained from the measured diameter and mean velocity.

6. Spectral broadening due to the use of wide aperture transducers to give good focusing can cause large errors in the measurement of maximum velocity.

7. Since the Doppler beam is held fixed and the PRF is high, particular attention should be paid to the outputs of PW units in relation to safety.

COLOUR DOPPLER IMAGING

Pulsed Doppler techniques require between 50 and 100 ultrasonic pulses to be transmitted in each beam direction for the determination of velocities of blood in a sample volume. It is therefore not possible to move the beam rapidly through the scan plane to build up real-time Doppler images of velocity of flow. Such imaging became possible when signal processing was developed which could quickly produce a measure of mean blood velocity at each sample volume from a small number of ultrasonic echo pulses. A technique called ‘autocorrelation processing’ of the signals from blood quickly gives the mean velocity in each small sample volume along the beam (Fig. 1-1C). This real-time colour Doppler imaging processes between 2 and 16 echo signals from each sample volume. In addition, the direction of flow is obtained by examining the signals for the direction of the shift as for CW and PW Doppler devices. Each image pixel is then colour-coded for direction in relation to the transducer and mean Doppler shift (Fig. 1-3A).

FIGURE 1-3 (A) Colour flow image of flow from left to right in a straight tube. When the flow velocity component is along the beam towards the transducer it is colour-coded red, when the component along the beam is away from the transducer it is coded blue. (B) Power Doppler image of flow in a straight tube. Direction of flow is not measured so it is not colour-coded.

B-scanning and Doppler imaging are carried out with the common types of real-time transducer. Echo signals from the blood and tissues are processed along two signal paths in the system electronics (Fig. 1-4). Going along one path, the signals produce the real-time B-scan image; going along the other path, autocorrelation function processing and direction flow sensing are employed to give a colour flow image. An important exclusion circuit in the autocorrelation path separates large-amplitude signals which arise from tissue and excludes them from the blood velocity processing. The B-mode and mean velocity images are then superimposed in the final display. Strictly speaking, the flow image is of the mean Doppler shift frequency and not the mean velocity, since the beam–vessel angles throughout the field of view are not measured. Colour shades in the image can indicate the magnitude of the velocity, for example light red for high velocity and dark red for low velocity. Turbulence, related to the range of velocities in each sample volume, may be presented as a different colour or as a mosaic of colours.

FIGURE 1-4 B-mode and Doppler imaging signal processing paths in a scanning machine.

Doppler images typically contain about 64 genuine lines of information and 128 consecutive sample volumes along each line. The frame rate varies from 5 to 40 frames per second, depending on the depth of penetration and the width of the field of view. As in B-scanning, the appearance of the image is usually improved by inserting additional lines or frames whose data are calculated from the genuine lines, a process known as interpolation. Alteration in flow can occur rapidly over the cardiac cycle, therefore a cine-loop store of say the last 128 frames is of value for review purposes. Doppler spectrograms can be made by selecting the appropriate beam direction and sample volume location in the image and then switching to the PW or CW mode. PW and CW Doppler techniques provide more detailed information on blood velocities than colour Doppler, so spectral information is still of value.

A summary of technical factors relating to the use of colour Doppler imaging is given in Box 1-3.

BOX 1-3 Technical Factors in the Use of Colour Doppler Imaging

1. The mean Doppler frequency shift is the quantity which is presented in a colour-coded form in each pixel. When the colour bar is labelled in velocity the beam–vessel angle has been assumed to be zero throughout the image.

2. The velocity component along the Doppler beam is heavily dependent on the angle between the direction of flow and the beam direction (the cosine θ dependence). The colours depicted in the image are therefore heavily angle dependent.

3. The flow pattern on the colour Doppler display can be related to the structures shown in the B-mode image.

4. Colour Doppler imaging is a pulsed technique so aliasing is a problem.

5. A good colour Doppler machine is one which discriminates well between signals from tissue and those from blood.

6. The colour Doppler field of view box should be adjusted to cover only the region of interest and therefore maximise the frame rate.

7. The velocity range covered by the colour scale should be carefully matched to the velocities expected in the study.

8. A cine-loop is useful for the review of fast-changing blood flow patterns.

9. A change in the colour registered in a vessel from, say, blue to red may not mean a change in the direction of flow along a vessel. It may merely mean the beam–flow angle has changed from less than 90° to more than 90°.

POWER DOPPLER IMAGING

The power of the Doppler signal from each small sample volume in the field of view may be displayed rather than the mean frequency shift (Fig. 1-3B). The power of the signal from each point relates to the number of moving blood cells in that sample volume. The power Doppler image may be considered to be an image of the blood pool. The power mode does not measure velocity or direction and therefore the image shows little angle dependence, nor does it suffer from aliasing; however, it obviously presents less information about blood flow. The attraction of power Doppler images is that they suffer less from noise than velocity images, as the power of the background noise for any sample volume with no blood flow signal is less than the power of the background noise plus Doppler signal when blood flow is present. The background noise may be used to set a threshold above which signals are accepted for Doppler flow. Noise from sample volume regions lacking blood flow is therefore reduced in the power image by a threshold detector. However, when the same signal is used in the velocity imaging mode, the noise will produce a mean velocity value which the machine will treat as a genuine blood velocity and which will therefore appear in the image. The power Doppler mode is therefore less prone to noise and hence more sensitive and can be used to detect small vessels. Further sensitivity can be obtained by averaging power images over several frames to reduce spuriously distributed noise even more. In velocity imaging there is interest in showing quick changes in blood flow and hence less averaging is used.

Power Doppler imaging often provides a more complete image of the vasculature than velocity imaging. This has made it popular in clinical use and it is commonly used initially to locate regions of interest prior to investigation by colour Doppler or duplex methods. It is also possible to use the direction information in the signal to colour-code the corresponding power image.

A summary of technical factors relating to the use of power Doppler imaging is given in Box 1-4.

BOX 1-4 Technical Factors in the Use of Power Doppler Imaging

1. In the power Doppler image, the power of the Doppler signal at each pixel is colour-coded.

2. No velocity information is displayed.

3. The power Doppler signal is direction insensitive; so the display is the same regardless of whether the blood is flowing toward or away from the transducer. However, some systems do display different colours depending on the direction of blood flow by including some directional Doppler information.

4. The power Doppler image is insensitive to angle except near 90° where the Doppler signal may fall below the clutter filter, which cuts out low Doppler shifts, and no signal is displayed.

5. There is no aliasing, because the frequency information (i.e. velocity) is not estimated from the Doppler signal.

6. The power Doppler image is very sensitive to movement of the tissue or probe (the flash artefact). Some machines incorporate a flash filter to try to reduce this effect.

Ultrasonic Microbubble Contrast Agents

A number of agents have been considered which can enhance the scattering of ultrasound from blood and hence could be employed as echo-enhancing or contrast agents. Ophir and Parker ¹⁰ have reviewed contrast agents. From this review and more recent experience it has become obvious that agents in the form of encapsulated microbubbles are by far the most likely to be successful echo-enhancing agents in the immediate future. This is due to the large difference in acoustic impedance between the gas in the bubbles and the surrounding blood. In addition, bubbles of a few microns in diameter have a fundamental resonance frequency of a few megahertz. For example, 4 μm diameter bubbles resonate at 4 MHz, which is well within the range of medical ultrasound systems. Bubbles of these dimensions are important since, even with only very thin wall encapsulation, they are able to pass through the capillaries of the lung into the systemic circulation. An investigation by a committee of the American Society of Echocardiography concluded that contrast echocardiography carried a minimal risk for patients and that there were few residual or complicating side-effects.¹¹ Other studies have confirmed these conclusions; however, more work is required on new agents as they become available.¹²

The development around 1990 of contrast agent microbubbles that can be used by percutaneous venous injection was the breakthrough which gave rise to the current high level of activity in this field. Table 1-1 gives examples of agents which are currently under commercial development. Large gas molecules are encapsulated in some agents to reduce the rate of diffusion and so increase the lifetime of the bubbles. Typically the lifetime in blood ranges from 2–3 min up to 20–30 min. An attraction of contrast agents is the ability to increase the signal obtained from small blood vessels which are difficult to detect by conventional Doppler methods, such as cerebral or renal vessels. There is also interest in perfusion studies, for example to observe and measure the wash-in and wash-out of agent in the myocardium in a manner analogous to nuclear medicine studies.

TABLE 1-1

Properties of Some Commercially Available Ultrasound Contrast Agents

By courtesy of CM Moran, University of Edinburgh.

Enhanced scattering is obtained if a bubble is insonated with ultrasound of a frequency equal to that of the fundamental resonance frequency of the bubble. At low power (that is low ultrasonic wave pressure amplitudes) the oscillations of the bubble are about its centre and are directly in proportion to the size of the pressure fluctuations in the ultrasound wave. However, at higher powers the oscillations become distorted and ultrasound at frequencies different from that of the incident wave is generated by the bubbles. These frequencies are known as harmonics and are simply related to the fundamental resonance frequency of the bubble, so that the second harmonic frequency is twice the fundamental frequency. There is considerable interest in detecting and using the second harmonic, since tissue does not produce this effect to any great extent and the second harmonic signal comes predominantly from the echo-enhancing agent in the blood vessels. Both pulse–echo and Doppler systems have been designed to pick out the second harmonic component in the ultrasound returned to the transducer and use it to enhance the signal from the agent in blood, possibly by as much as 20–30 dB. These systems are being evaluated in clinical practice.¹³

The scattering from contrast agents can also be enhanced if the acoustic pressure fluctuations in the beam are large enough to damage the microbubbles, causing them to leak. An unencapsulated gas bubble then forms next to the original one; however, since it has no outer shell, the scattering from it is undamped and can be around 1000 times higher than that from the encapsulated bubble. This effect has been exploited in a technique known as ‘intermittent’ or ‘transient’ imaging, which allows time for the damaged bubbles to be replaced between sweeps of the ultrasound beam.¹⁴

A summary of technical factors relating to the use of microbubble contrast agents is given in Box 1-5.

BOX 1-5 Technical Factors in the Use of Microbubble Contrast Agents

Contrast agents are still being developed but some technical points of importance include:

1. Check the age and shelf life of the agent.

2. Microbubble contrast agents are fairly fragile. The handling instructions should be carefully followed.

3. The ultrasonic beam is capable of destroying some microbubbles; higher transmit power will therefore not necessarily produce a stronger echo signal.

4. Several contrast agents can be insonated at high acoustic pressures to give enhanced backscatter due to bubble destruction. This phenomenon is also known as ‘acoustically stimulated emission’. Aspects of its safety still need to be fully evaluated.

5. Determine the physical phenomena which are expected to be present under the operating conditions and which will enhance the echo signal, e.g. harmonic scattering, free bubble creation and bubble destruction.

Information from Doppler Signals

When the Doppler signal with the information on blood velocities has been obtained, it has to be interpreted. It is essential to obtain good-quality signals in order to be able to detect disease.

THE SPECTRUM ANALYSER

A Doppler signal may be analysed into its frequency components in order to give a display of the velocities of the blood cells at each instant (Fig. 1-5).² Short time intervals of the Doppler signal are analysed, for example a segment of 5 ms duration. This produces an instantaneous spectrum of the frequencies in the sample volume for that time period. If an angle correction is then applied, this spectrum will represent the range of velocities in the sample volume. The frequencies in each spectrum are displayed along a vertical line on which the power of each frequency component is presented as a shade of grey. The consecutive velocity spectra are then displayed as side-by-side grey-shade vertical lines. In this way a spectral display, or spectrogram, is built up (Fig. 1-6). Note the difference between the instantaneous velocity spectrum which tells us about the pattern of velocities in the sample volume at that instant and the spectral display, or spectrogram, which shows how the velocity pattern varies with time. Spectrograms are generated in real time during the clinical examination and it is usually possible to store a few seconds of the trace in an analyser for subsequent review.

FIGURE 1-5 Schematic representation of the analysis of a Doppler signal to form a sonogram (spectrogram). (From McDicken², Copyright Elsevier 1991)

FIGURE 1-6 A normal sonogram (A); (B) a magnified view of the segment between the arrows illustrating the pixel structure and speckle pattern.

The temporal resolution in a spectrogram, that is the smallest discernible time interval, equals the length of the portion of Doppler signal used to produce each instantaneous spectrum and is typically 5–10 ms. The frequency (or velocity) axis of the spectrogram usually has about 100 scale intervals each of 100 Hz. The Doppler signal is therefore resolved into frequency components separated by 100 Hz, the frequency resolution of the spectrogram.

It is often desirable to make measurements on a spectrogram, for example an estimate of the time between two events or of the maximum velocity during systole or diastole. Measurement is performed by placing a cursor on the relevant points of interest then a variety of calculations can be performed by the system computer. Indices related to the spectrogram shape and hence to normality or abnormality of flow velocities can be calculated within the analyser and displayed on its screen. These are discussed later in this chapter. The oscillatory shape of a spectrum or a trace derived from it is often referred to as a waveform.

A summary of technical factors relating to spectral analysis is given in Box 1-6.

BOX 1-6 Technical Factors in Relation to Spectral Analysis

1. Arrange the frequency and time scales to best display the detailed information in the sonogram by adjusting velocity scale, baseline and sweep speed.

2. Arrange the grey tones, or colours, of the spectral display to give the best ‘image’ quality by adjusting the Doppler gain.

3. Treat with caution information from weak signals in a noisy background.

4. A clear spectrogram is required before the maximum frequency can be traced.

5. It is usually best to make measurements over at least five heart cycles. However, if this is not possible, perhaps due to respiratory movement, useful information can be obtained from one or two cardiac cycles.

6. Make use of the total storage capacity of the analyser and use the scroll facility to select the most suitable part of the spectrogram to make relevant measurements.

7. Check thoroughly the reliability and accuracy of automatic tracing techniques and calculations derived by the ultrasound system.

8. Note whether the vertical axis has been calibrated in velocity by allowing for the beam–vessel angle.

9. To obtain reproducible results, try to use the same beam–vessel angle for all examinations, e.g. 60°.

Spectrograms (Sonograms) and Indices

A good-quality audible Doppler signal will result in a good-quality sonogram. Degradation of a sonogram is due to electronic and acoustic noise which should only be a problem when the gain of the Doppler unit or analyser is very high for the detection of weak signals. Noise in sonograms creates considerable difficulties for the automatic calculation of quantities and they should therefore be used with caution. If the automatic mode cannot deal with the noise, then the traces should be drawn manually making use of the eye’s ability to distinguish the true signal from the noise. It should also be remembered that the vertical axis of a sonogram can only be labelled as velocity after the beam–vessel angle has been measured.

WAVEFORM INDICES

Waveform indices are derived from a combination of a few dominant features of the waveform.³ Indices that have the same or similar names in the literature are occasionally defined differently, so a first step is to check the definition of any index to be used. In practice only two classes of index are used to any great extent, those related to the degree of diastolic flow and others related to spectral broadening. The variation in time of the maximum velocity displayed in a spectrogram is commonly used as a source of data for the derivation of an index (Fig. 1-7). Since the maximum velocity is not always clearly apparent in a spectrogram, some analysers produce a trace which is closely related to the maximum velocity trace. One example is a trace showing the upper velocity boundary below which the velocity components contain seven-eighths of the power of the Doppler signal.

FIGURE 1-7 A spectral waveform showing maximum velocity trace (white) and mean velocity trace (blue).

The mean velocity waveform (average velocity waveform) is also employed (Fig. 1-7). To calculate the mean velocity at each instant, the values of velocity and the intensities of the signal for each velocity component in the instantaneous spectrum are used. The mean velocity is used together with the vessel cross-sectional area to calculate blood flow rate. However, it is difficult to measure mean velocity accurately and there are several other problems associated with calculating flow rates; these are discussed further in Chapter 2.

Since the beam–vessel angle may not always be known, the waveforms or spectrograms will not be corrected for angle. Indices are therefore defined involving ratios of velocities. In such a ratio, the angle factors appear on both the top and bottom and hence cancel each other out, so that the index is independent of beam angle. Errors are also reduced by averaging the calculated indices over several heart cycles.

A number of the most commonly encountered indices are briefly discussed below:

A/B Ratio

The A/B ratio is defined as the ratio of two specified velocities, e.g. maximum velocities, at two points in the cardiac cycle (Fig. 1-8). It is usually employed where there is no reverse flow in the waveform.

FIGURE 1-8 Waveform indices, normally calculated from a maximum velocity waveform, but the mean velocity waveform can also be used. (From McDicken², Copyright Elsevier 1991)

Resistance Index (RI)

In Fig. 1-8:

High resistance in the distal vessels produces low diastolic flow in the supplying artery and results in a high value for this index; a low resistance results in a low value as there is higher diastolic flow. It is also known as the Pourcelot index.

Pulsatility Index (PI)

The pulsatility index (PI) is defined as:

This ratio is used in vessels where reverse flow may occur, for example in the lower limbs (Fig. 1-8). The PI may typically have a value of 10 for the normal common femoral artery but be around 2 when proximal disease severely dampens the waveform.

As defined above, the PI index is heart rate dependent. To avoid this, PI can be calculated over a specified time from the start of systole, e.g. for the first 500 ms. The pulsatility index is then labelled ‘PI(500)’.

Damping Factor

The damping factor is defined as the ratio of pulsatility indices at two sites along an artery. It quantifies the damping of the waveform downstream along a diseased vessel.

The numerical value of this index increases as disease becomes more severe, a value of 2 being typical of a high degree of damping. This index is mentioned for completeness, it is not widely used.

Spectral Broadening

Turbulence increases the range of blood cell velocities in a vessel. One index to quantify this broadening of the spectrum of velocities is:

Conclusions with regard to the presence of turbulence should be made with caution and only after familiarity has been gained with the patterns for laminar and plug flow for the particular instrument being used. Other technical factors can cause spectral broadening: for example ‘geometrical spectral broadening’, which results from the range of Doppler angles which any given blood corpuscle subtends to different points on the face of the transducer.

Transit Time

The time for the pulse pressure wave to travel along a length of artery can be measured by placing a Doppler probe at either end of it. From this time and knowing the length, the pulse wave velocity (PWV) is calculated. Alternatively, using the electrocardiogram (ECG) and one Doppler unit the transit time can also be measured, the QRS of the ECG giving the time when the pressure pulse leaves the heart. The time is first measured for the pulse to travel from the heart to the proximal artery site; a second time is then measured for the pulse to go from the heart to the distal site. Subtracting these two times gives the transit time and, if the length of the artery is known, the PWV can be calculated. Any error associated with the assumption that the QRS represents the time at which the pulse leaves the heart is removed by the subtraction.

A normal aorta has a PWV of around 10 m/s. The transit time along 0.5 m is then 50 ms. Pulse wave velocity depends on disease states of the artery wall and blood pressure. This index is also mentioned for completeness, it is not widely used.

A summary of technical factors relating to waveform analysis is given in Box 1-7.

BOX 1-7 Technical Factors in the Application of Waveform Indices

1. Use good-quality spectral data to calculate indices.

2. Most indices are calculated using the maximum Doppler frequency as this is relatively insensitive to beam–vessel alignment and sample volume size.

3. Mean Doppler frequency shift is very sensitive to alignment and sample volume size and is not widely used to estimate indices.

4. Indices of waveform shape are insensitive to the beam–vessel angle except near 90° where the end-diastolic component may disappear below the high-pass filter.

5. For obstetrics use, the wall-thump filter must be set low, typically 50–100 Hz, in order that absent end-diastolic flow may be ascertained unambiguously.

6. The most suitable index or measurement for each application can be determined from the literature. Indices are useful in that they allow velocity patterns to be classified and related to disease state. However, they do not extract any more information than can be deduced by direct observation of the whole sonogram.

Artefacts in Doppler Techniques

The most important artefacts are mentioned here and methods of dealing with them are suggested. Further details can be found in other texts.^2,⁵ Artefacts are usually dealt with by explaining their origin or by recognising that they occur fairly frequently and are not of significance.

OVERESTIMATION OF MAXIMUM VELOCITY

The maximum blood velocity is often of clinical interest since it is of physiological significance and often easy to identify on a spectrogram. However large errors can occur in its measurement if steps are not taken to avoid them. If maximum velocity is measured in the laboratory using a string phantom, which simulates a line of blood cells passing through a sample volume, it is found that the velocity is overestimated (Fig. 1-9A) and the error increases markedly as the angle approaches 90°. As an active array segment has a finite length, typically 20–40% of the full array, so different elements will register different beam–vessel angles and therefore Doppler shifts, as shown in Fig. 1-9B; the elements with the higher angles will also be more affected by the angle error shown in Fig. 1-9A. The highest Doppler shift and hence the highest measured velocity will be registered by the elements with the smallest beam–vessel angles. Strictly speaking a correction should be applied to allow for the fact that the maximum velocity is not measured by ultrasound travelling along the central axis of the beam, as assumed in the basic Doppler equation, but manufacturers have not adopted this approach. The errors in maximum velocity estimation shown in Fig. 1-9 are typical of modern linear arrays. The errors for phased arrays, which are used in cardiology, are smaller due to their smaller Doppler apertures. In clinical practice with linear arrays, errors are reduced by the use of small angles (typically 45–60°) and by the use of indices employing velocity ratios as described in Fig. 1-8.^15,¹⁶

FIGURE 1-9 Overestimation of maximum velocity: (A) Velocity errors measured using a string phantom. The true string velocity is 0.61 m/s but the velocity measured using Doppler ultrasound is overestimated, with the degree of overestimation increasing as the beam-filament angle approaches 90^°. (B) Geometrical spectral broadening due to the finite size of the active area of an array. Higher Doppler shifts are received at one end of the Doppler aperture, smaller shifts at the other end. The Doppler system will usually perform angle correction with respect to the centre of the aperture.

SPECKLE AND THE SPECTRAL DISPLAY

The speckled appearance of a sonogram results from fluctuations in the power levels of the velocity components in neighbouring pixels (Fig. 1-6). These fluctuations are due to variations in the ultrasonic signal received from the random distribution of blood cells. Due to this speckle noise, the power level in a pixel cannot be directly related to the number of cells moving with a particular velocity. Averaging the power levels in neighbouring pixels gives a more accurate measure of the number of cells moving with each velocity.

INADEQUATE COUPLING

Weak Doppler signals can be attributed to a lack of coupling medium between the transducer and the skin. In CW and PW Doppler applications this artefact is not as evident as in imaging techniques, where poor penetration is readily seen.

ELECTRICAL PICK-UP

Doppler devices are required to be highly sensitive and are therefore prone to electrical pick-up of stray signals. Some such signals can be recognised by their pattern in the sonogram. The spectrum analyser unit may allow the operator to attempt to clean up the sonogram by deleting spurious signals.

COMPRESSION OF THE SPECTRAL DISPLAY

Information is lost when a sonogram is compressed, either in its grey shades, or in its scale of presentation. A sonogram should be treated as a type of image and therefore presented with optimum grey-shade contrast and spatial detail. The latter means that the velocity and time axis scales should be selected to show detailed structure in the sonogram. The accuracy of measurements from the sonogram will also be affected by poor presentation.

ERRONEOUS DIRECTION SENSING

The direction-sensing circuitry does not always function correctly since its design and implementation are difficult. Flow will then be presented in the wrong direction, either in a sonogram or in a colour Doppler image; if this is suspected then the set-up of the system can be checked by examining a normal artery in which the direction of flow is known.

FILTERING

Filters are used to reduce low frequencies, such as those obtained from arterial walls. Filters also remove information on slow-moving blood but this is not usually a serious problem unless it is desired to measure mean velocity accurately or slow flow specifically (Fig. 1-10).

FIGURE 1-10 Filtering. (A) Standard sonogram. Raising the high pass (wall thump) filter (B) removes low-velocity signals and information.

HARMONIC GENERATION BY LARGE SIGNAL DISTORTION

The harmonics of a frequency are higher multiples of that frequency; for example, harmonics of 100 Hz are 200 Hz, 300 Hz, etc. If a signal is too large to be handled by the electronics it becomes distorted and then contains additional harmonic frequency components. When such a distorted signal is analysed, the harmonics appear at regular frequency intervals in the spectral display. Strong blood flow signals exhibit harmonic components as a higher frequency part of the sonogram above that which would probably be obtained if the gain were reduced (Fig. 1-11).

FIGURE 1-11 Increasing gain from (A) to (C). Very high gain distorts the signal and introduces harmonic frequencies.

HIGH OR LOW SENSITIVITY

In colour Doppler and power Doppler imaging, setting up the system with too low a sensitivity causes blood flow signal to be lost. Too high a sensitivity causes spurious echoes to be colour-coded as blood (Fig. 1-12).

FIGURE 1-12 Variation of gain in a colour Doppler image. The image content is very sensitive to gain setting: (A) shows too much gain; (B) shows a satisfactory gain; (C) shows too little gain.

ALIASING

Pulsed Doppler and colour Doppler units have to reconstruct the Doppler shift signal from regularly timed samples of information, rather than the complete signal as used in CW units. The sampling rate is equal to the PRF (pulse repetition frequency) of the Doppler unit. If the sampling rate is too low, less than half the Doppler shift frequency, then the frequency of the reconstructed Doppler signal is in error and the direction of flow is presented wrongly. In a spectral display or flow image, this is known as an ‘aliasing’ artefact (Figs 1-13 and 1-14).

FIGURE 1-13 Aliasing in a sonogram. (A) Satisfactory PRF and spectral display. (B) PRF is too low and the highest velocity signals appear in the reverse channel (arrows).

FIGURE 1-14 Increase in aliasing in a colour Doppler velocity image. Note that there is no black space between the red and blue pixels in the lateral colour displays (arrows), indicating that the pixel colour-coding has gone off the top of the red colours and wrapped round to the blue colours. However, there is a black space between the colour displays centrally (arrowhead), indicating reversal of flow direction in relation to the transducer.

The aliasing artefact is encountered when high-frequency Doppler shift signals are produced, usually by high-velocity flow. It also occurs when sampling deeper vessels, as the PRF is reduced to allow time for the echoes from a pulse to return before the next pulse is transmitted. If the PRF is too low for the Doppler shift frequencies from the blood in the vessel, aliasing will occur. An approach to raising the velocity level at which aliasing becomes a problem is to use a high PRF (the high-PRF mode), even although the echoes from deep structures have not died out before the next pulse is transmitted. If the deep echoes are strong enough, Doppler signals from deep vessels may then be superimposed on those from a more superficial site. This uncertainty with regard to the source of the signal is referred to as ‘range ambiguity’. The ultrasound display will show two or more sample volumes from which data are being analysed and the examination is best performed if only one sample volume is positioned in the vessel, with the others overlying non-vascular structures so that any Doppler shift detected is most likely to come from the vessel. Although this mode can be useful, it also increases the intensity of the beam, which is another reason for using it only when necessary. Aliasing can be of value in colour Doppler imaging since it allows high-velocity jets to be identified. Power Doppler does not suffer from aliasing.

EFFECT OF BEAM ANGLE TO FLOW DIRECTION

The quality of a Doppler signal depends on beam–vessel angle and above 70° it degrades quite rapidly (Fig. 1-15). If the direction of an ultrasound beam is at 90° to the direction of the flowing blood, no Doppler signal is expected since, in the Doppler equation, cos90° = 0. However, a poor-quality Doppler signal is usually obtained for two reasons. First, the ultrasound beam may converge or diverge slightly from the beam axis, so all of it is never at 90° to the flow. Second, there may be some turbulence in the flow, in which case the blood cells are not all travelling in parallel paths at 90° to the beam.

FIGURE 1-15 Variation of quality of signal with beam–vessel angle: (A) at 60°; (B) at 70°; (C) at 80°.

Colour Doppler images obtained at 90° to the direction of flow appear dark or noisy, corresponding to anabsent or small Doppler shift (Fig. 1-3). Power Doppler images are relatively insensitive to angle, except near 90°, where the low Doppler frequencies may fall below the clutter filter and no power signal is displayed (Fig. 1-3).

In a vessel in which the direction of flow alters with respect to the ultrasound beam, different regions of the vessel will be colour-coded differently. Note that this may be due to a genuine change in flow direction as seen in the normal carotid bulb, or merely due to the changing beam angle (see Fig. 4-7) which is particularly common in sector scan imaging.

EFFECT OF VELOCITY SCALE

The choice of velocity scale can dramatically change the appearance of a colour Doppler image (Fig. 1-16). The scale should be chosen to accommodate the range of velocities thought to be present. Too low a scale will cause aliasing and too high a scale results in the flow being depicted as a few dark colours in colour Doppler imaging.

FIGURE 1-16 Effect of changing velocity scale (PRF): (A) scale set too high for flow velocity; (B) satisfactory setting; (C) scale set too low for flow velocity, resulting in aliasing.

UNEXPECTED MACHINE ARTEFACTS

Doppler technology is developing rapidly and can still have gremlins in it. The operator must therefore check the performance and calibration of the instrument. This is most readily done in a situation where the flow pattern is considered to be well understood, such as in a clearly seen normal blood vessel or a flow test-object. Figure 1-17 shows an unexpected artefact in which the maximum velocity measured varies with the beam position in the field of view. This has arisen because the transducer aperture used by the system is a different size in the different positions, resulting in a different amount of spectral broadening.

FIGURE 1-17 Variation of sonogram for different beam points of origin. A wider aperture is used towards the centre of an array for better focusing. This increases the range of angles of insonation, producing more spectral broadening (A) compared to a beam from the edge of the array (B).

INTERFERENCE FROM NEIGHBOURING VESSELS

If part or all of a neighbouring vessel in addition to the vessel of interest is within the sample volume of a CW or PW instrument, the Doppler signal will contain a contribution from the extra vessel. Moving the sample volume or redirecting the ultrasound beam to try to interrogate only the vessel of interest may reduce this artefact.

VESSEL COMPRESSION

It is easy to compress superficial vessels by transducer pressure. Increased velocity of flow through the restriction in the compressed vessel results in a higher-pitched Doppler sound or a colour change in an image.

FACTORS AFFECTING THE PATIENT

It is necessary to have as complete a knowledge as possible of the patient’s physiological status when undertaking blood flow studies, since many factors affect the cardiovascular system. Examples of these factors are exercise, heart rate, temperature, anxiety, posture, food, smoking and other drugs.

PATIENT OR VESSEL MOVEMENT

If movement causes the sample volume of a CW or PW beam to interrogate a different region, the blood flow signal will obviously be altered. It can be difficult to eliminate this factor, especially in abdominal examinations, and it is not always clear whether respiration has actually affected the flow or just moved the vessel.

FLASH ARTEFACT

Movement of the patient, an organ, or the transducer during Doppler imaging gives tissues a velocity relative to the transducer and hence scattered ultrasound is Doppler shifted; a large area of the image is therefore colour-coded for the duration of the movement. This artefact is more severe with power Doppler imaging due to its increased sensitivity.

TWINKLE ARTEFACT

This artefact is seen as an intense, rapidly changing mixture of colour signals behind strongly reflective stationary structures such as renal calculi (see Figs 9-7, 9-8 and 13-24). This artefact is thought to be produced by factors relating to the irregularity of the reflecting media, multiple reflections and tiny random variations in the clock that synchronises pulse transmissions (clock jitter). These factors result in fluctuating returned signals which the electronics present on the display as similar to colour-coded Doppler signals. The effect is highly machine and control setting dependent.^17,18

BEAM POSITION WITHIN THE VESSEL LUMEN

The Doppler signal obtained from a vessel depends on how the beam insonates the vessel. The effect is less marked if the beam is wider than the vessel. A narrow beam through the centre of a vessel, however, overemphasises the high velocities while a beam through the side of the vessel detects lower velocities.

ONE-DIMENSIONAL SCAN

Blood flow often occurs in an unknown direction in space in relation to the transducer, for example in the heart or at a vessel bifurcation. Detecting the flow from one direction only measures the velocity component along that beam direction. Measurement of the actual velocity in these situations requires components to be measured in three directions not in the same plane. However, when laminar flow occurs in a vessel lying in an imaging scan plane, measurement of one velocity component and the beam–vessel angle permits measurement of the actual velocity.

Safety and Prudent Use of Doppler Instruments

Ultrasound beams transmit energy into tissue so the possibility of hazard has to be considered. The most likely mechanisms for harmful effects are thought to be tissue heating as the ultrasound energy is absorbed, or cavitation in which gas bubbles in the tissue react violently under the influence of the pressure fluctuations of the ultrasound field. The most sensitive structures are considered to be the developing fetus, the brain, the eye, the lung and bone–tissue interfaces.

There is a considerable amount of literature on bioeffects and safety of diagnostic ultrasound.¹⁹ The literature is scrutinised by several national and international bodies who produce statements on safety and the prudent use of ultrasound. Organisations actively monitoring the safety of ultrasound are the World Federation for Ultrasound in Medicine and Biology (WFUMB),²⁰ the European Federation of Societies for Ultrasound in Medicine and Biology (EFSUMB),²¹ the British Medical Ultrasound Society²² and the American Institute of Ultrasound in Medicine (AIUM). It is still true to say that there are no confirmed harmful effects of diagnostic ultrasound. Often the possibility of an effect is reported but it is not confirmed by further work. There is a need for well-controlled studies but these are increasingly difficult to conduct since unscanned control populations are almost non-existent in the developed world. Although no harmful effects have been confirmed, there is some concern that the outputs of machines have been increasing by factors of as much as 3 or 5 since 1991, as manufacturers seek to produce better B-mode images and more sensitive Doppler units.²³ The situation is summed up in a commentary by ter Haar.²⁴

Until the early 1990s, attempts were made to specify the maximum intensities permissible for different clinical applications. This proved to be both limiting and impractical, so the approach now is to use the ALARA principle (As Low As Reasonably Achievable) borrowed from the field of ionising radiations. The user is now informed of the output of the machine and has the responsibility to keep the exposure to a low value which will still give a diagnosis. Many systems now display the output on the screen in terms of a thermal index (TI), related to tissue heating, and a mechanical index (MI), related to the possibility of producing cavitation. These indices are defined in the Output Display Standard (ODS) developed jointly by the AIUM and the National Electrical Manufacturers Association (NEMA) in the USA.²⁵ The Food and Drug Administration (FDA) in the USA requires adherence to this standard and this is followed by many countries throughout the world. MI is also used to provide a measure of output during contrast agent studies and hence helps to describe the technique. When a new technique is to be implemented clinically, attention should be paid to the TI and MI values used by its developers. Transducer heating was a problem with some devices in the past due to inefficient conversion of electrical energy into acoustic energy. It is worth checking that the transducer face is not hot.

EFSUMB puts out regular statements on safety and currently it says that the use of B-mode is not contraindicated in routine scanning during pregnancy. However, it is more cautious with regard to pulsed Doppler, saying ‘routine examination of the developing embryo during the particularly sensitive period of organogenesis using pulsed Doppler devices is considered to be inadvisable at present’.

A summary of technical factors relating the prudent use and safety of ultrasound is given in Box 1-8.

BOX 1-8 Technical Factors Affecting Prudent Use and Safety

1. Use the lowest transmit power which will give a diagnostic result. This involves keeping the MI and TI less than 1 if possible.

2. Use high receiver gain rather than high output power to achieve high sensitivity.

3. Use the minimum scan time possible.

4. Check that the transducer ceases to transmit when the imaging mode is frozen.

5. Take particular care when a fixed beam direction PW Doppler mode is being used near sensitive tissues.

6. Compare the maximum output values (intensity, power, pressure amplitudes) for your machine to those quoted in the published data from equipment surveys.

7. Safety considerations related to contrast agents need to be studied at frequent intervals during their developmental stage.

Future Instrumentation

The technical performance of Doppler ultrasound systems is regularly improved by the introduction of new transducers and computer processing algorithms for ultrasonic echo signals. Clinical performance is also improved by increased operator experience and the identification of new applications. The introduction of completely new technology typically occurs at the rate of one or two new instruments per decade.

The production of colour Doppler images by sweeping a single pulsed beam across a scan plane has already been discussed. With this technique the frame rate might typically be 5 to 40 frames s⁻¹. New techniques are being explored to increase frame rates and to overcome angle dependence in the measurement of velocity. Computer control of the excitation of the individual transducer elements and the reception of individual echo signals introduces great flexibility in the production of images. These new techniques allow the generation of frame rates in excess of 100 s⁻¹ making it possible to image rapidly changing flow patterns in 2D and to have more useful volume scan rates in 3D.

At an interrogated blood flow site, techniques are being developed to measure velocity components in more than one direction (Fig. 1-18). Beam direction and interrogation sites can be altered very rapidly so that it is possible to interrogate flow sites across the scan plane and measure velocity components in quick succession. From these components the speed and true direction of flow at sites in the scan plane can be calculated, i.e. the velocity vector in the scan plane is obtained for each site. A quantity is called a vector in mathematics when its magnitude (speed in the present case) and direction are known. The velocity vector at each site in a blood vessel is then presented on the display as an arrow whose length is the magnitude of the velocity and whose direction is the direction of flow (Fig. 1-18). High frame rate vector Doppler images are possible because of the speed with which modern electronics can excite transducer elements and computers can process received echoes. High frame rates are very desirable in situations where flow patterns change quickly throughout the cardiac cycle, indeed low frame rates can be very misleading in such situations.