Method, apparatus and applications for combining transmit wave functions to obtain synthetic waveform in ultrasonic imaging system

A signal processing technique for improving the axial resolution and/or the sensitivity of an ultrasonic imaging system. This technique also improves the lateral resolution and the depth of field. The technique is based on combining two or more transmit wave functions in an optimal manner to achieve a synthetic waveform which has greater bandwidth and/or energy than the individual wave functions. This scheme operates on the imaging data before the data reaches the envelope detector, while the phase information is still maintained within the signal. Using a synthetic transmit wave design approach, the effective emitted pressure waveform would have a bandwidth which would be wider than the transducer bandwidth with high sensitivity. The synthetic transmit waveform design scheme consists of firing two or more relatively long transmit waveforms for each single A-line in a given focal zone. The frequency spectrum for each of these transmit pulses is centered at a slightly different frequency. The received signals from all of these transmit waveforms are added, while maintaining their phase information, to produce a synthetic waveform having an wider bandwidth than that of an impulse excitation.

FIELD OF THE INVENTION 
This invention generally relates to ultrasound imaging, primarily clinical 
ultrasound images as well as industrial ultrasonic images. In particular, 
the invention relates to a method for improving the axial resolution, 
contrast resolution, lateral resolution, depth of field and sensitivity of 
an ultrasonic imaging system. 
BACKGROUND OF THE INVENTION 
Conventional ultrasound imaging system comprise an array of ultrasonic 
transducers which are used to transmit an ultrasound beam and then receive 
the reflected beam from the object being studied. For ultrasound imaging, 
the array typically has a multiplicity of transducers arranged in a line 
and driven with separate voltages. By selecting the time delay (or phase) 
and amplitude of the applied voltages, the individual transducers can be 
controlled to produce ultrasonic waves which combine to form a net 
ultrasonic wave that travels along a preferred beam direction and is 
focused at a selected range along the beam. Multiple firings may be used 
at various depths or different positions in the image to acquire a full 
two dimensional data set representing the desired anatomical information 
along a multiplicity of scan lines. The beamforming parameters of each of 
the firings (or transmitted beams) may be varied to provide a change in 
the position of focus, depth of field or the shading (or apodization) 
function. Similarly, the beam forming parameters can be changed for the 
received beam. A dynamic receive beam is typically used for the reception 
where the delay focus is continuously changed as different data are 
received from different depths. However, during the transmission a beam of 
ultrasound energy with a specific focal position is transmitted. Typically 
multiple beams are transmitted along the same direction with different 
focal lengths for improved resolution. Multiple transmit and reception 
beams are used in a plane to construct a two dimensional image. 
The same principles apply when the transducer is employed to receive the 
reflected sound (receiver mode). The voltages produced at the receiving 
transducer elements are summed so that the net signal is indicative of the 
ultrasound reflected from a single focal point in the object. As with the 
transmission mode, this focused reception of the ultrasonic energy is 
achieved by imparting separate time delay (and/or phase shifts) and gains 
to the signal from each receiving transducer element. The reflected 
ultrasound is sampled from the focal zones of two or more transmit beams 
each focused at different depths along the same scan line. In most recent 
ultrasound imaging systems the received signal is dynamically focused as 
signals from different depths are received. For each steering angle, the 
sampled data from contiguous focal zones is acquired and then spliced to 
make one vector or A-line. A multiplicity of transmit vectors, one beam 
for each focal point, are used, along with interpolated data values, are 
used to collect all the image information which are displayed on the 
monitor to form a full image frame. This information is displayed on a 
pixel by pixel basis. 
Such scanning comprises a series of measurements in which the steered or 
non-steered beams of ultrasonic wave is transmitted, the system switches 
to receive mode after a short time interval, and the reflected, or 
backscattered, ultrasonic wave is received and stored. Typically, 
transmission and reception are steered in the same direction during each 
measurement to acquire data from a series of points along a scan line. 
Multiple reception beams can be formed for a single transmit beam for 
improved frame rate. For example, for a single transmit beam two reception 
beams on either side of transmit beam can be formed simultaneously using 
parallel beamforming or alternatively using high-speed a multiplexed 
beamforming which would process both beams simultaneously. The receiver is 
dynamically focused at a succession of ranges or depths along the scan 
line as the reflected ultrasonic waves are received. 
Referring to FIG. 1, the ultrasonic imaging system incorporating the 
invention includes a transducer array 10 comprised of a plurality of 
separately driven transducer elements 12, each of which produces a burst 
of ultrasonic energy when energized by a pulsed waveform produced by a 
transmitter 22. The ultrasonic energy reflected back to transducer array 
10 from the object under study is converted to an electrical signal by 
each receiving transducer element 12 and applied separately to a receiver 
24 through a set of transmit/receive (T/R) switches 26. Transmitter 22, 
receiver 24 and switches 26 are operated under control of a digital 
controller 28 responsive to commands by a human operator. A complete scan 
is performed by acquiring a series of echoes in which switches 26 are set 
to their transmit position, transmitter 22 is gated ON momentarily to 
energize each transducer element 12, switches 26 are then set to their 
receive position, and the subsequent echo signals detected by each 
transducer element 12 are applied to receiver 24, which combines, or 
beamform, the separate echo signals from each transducer element to 
produce a single echo signal which is used to produce a line in an image 
on a display monitor 30. 
Transmitter 22 drives transducer array 10 such that the produced beam of 
ultrasonic energy is directed, or steered, along a specific steering 
angle. To accomplish this, transmitter 22 imparts a time delay T.sub.i to 
the respective pulsed waveforms 34 that are applied to successive 
transducer elements 12. By adjusting the time delays T.sub.i appropriately 
in a conventional manner, the ultrasonic beam can be directed away from 
the normal to the plane of transducer array 36, by an angle .theta. and/or 
focused at a fixed range R. A sector scan is performed by progressively 
changing the time delays T.sub.i in successive excitations. The angle 
.theta. is thus changed in increments to steer the transmitted beam in a 
succession of directions. 
The echo signals are produced by each burst of ultrasonic energy, reflect 
from objects located at successive ranges along the ultrasonic beam. The 
echo signals are sensed separately by each transducer element 12 and a 
sample of the magnitude of the echo signal at a particular point in time 
represents the amount of reflection occurring at a specific range. Due to 
the differences in the propagation paths between a reflecting point P and 
each transducer element 12, however, these echo signals will not be 
detected simultaneously and their amplitudes will not be equal. Receiver 
24 amplifies the separate echo signals, imparts the proper time delay to 
each, and sums them to provide a single echo signal which accurately 
indicates the total ultrasonic energy reflected from point P located at 
range R along the ultrasonic beam oriented at the angle .theta.. 
Demodulation can occur either before or after the individual received 
signals are summed together. 
To simultaneously sum the electrical signals produced by the echoes 
impinging on each transducer element 12, time delays are introduced into 
each separate transducer channel 110 of receiver 24 (see FIG. 2). The beam 
time delays for reception are delays (T.sub.i) which are applied in a 
similar manner as the transmission delays described above. However, the 
time delay of each receiver channel is continuously changing during 
reception of the echo to provide dynamic focusing of the received beam at 
the range R from which the echo signal emanates. 
Under the direction of digital controller 28, receiver 24 provides delays 
during the scan such that steering of receiver 24 tracks the direction 
.theta. of the beam steered by transmitter 22 and samples the echo signals 
at a succession of ranges R and provides the proper delays and phase 
shifts to dynamically focus at points P along the beam. Thus, each 
transmission of an ultrasonic pulse waveform results in the acquisition of 
a series of data points which represent the amount of reflected sound at 
points in the focal zone of the transmit beam. 
Referring to FIG. 1, scan converter/interpolator 32 receives the series of 
data points produced by receiver 24 and converts the data into the desired 
image for display. In particular, the scan converter converts the acoustic 
image data from polar coordinate (R.sub.---- .theta.) sector format or 
Cartesian coordinate linear array to appropriately scaled Cartesian 
coordinate display pixel data at the video rate. This scan-converted 
acoustic data is then output for display on display monitor 30, which 
images the time-varying amplitude of the envelope of the signal as a gray 
scale. 
Referring to FIG. 2, a conventional receiver 24 comprises three sections: a 
time-gain control section 100, a receive beamforming section 38 and a 
processor 102. Time-gain control (TGC) section 100 includes a respective 
amplifier 105 for each of the receiver channels 110 and a time-gain 
control circuit 106. The input of each amplifier 105 is connected to a 
respective one of transducer elements 12 to amplify the echo signal which 
it receives. The amount of amplification provided by amplifiers 105 is 
controlled through a control line 107 that is driven by TGC controller 
106. The TGC is a combination of the potentiometers 108 which are set by 
the operator together with a constant gain profile, programmed into the 
controller, which compensates for tissue attenuation and diffraction gain 
variation in the image. 
The receive beamforming section 38 of receiver 24 includes separate 
receiver channels 110. Each receiver channel 110 receives the analog echo 
signal from one of amplifiers 105 at an input 111. Each received signal is 
delayed before being summed at the summing point 114 and 115. This delay 
provides the dynamic focusing which is essential for high resolution 
imaging. The summed signals indicate the magnitude and phase of the echo 
signal reflected from a point P located at range R on the steered beam 
(.theta.). Each amplified signal is conveyed as a pair of quadrature 
signals in the respective receiver channel, where the phases of the mixing 
reference frequency differ by 90.degree.. Alternatively the quadrature 
signals can be produced using the Hilbert transform. A signal processor 
120 receives the beam samples from summing points 114 and 115 and produces 
an output 121 to scan converter 32 (see FIG. 1). The signal processor 120 
sums the square of the I and Q signals before taking the square root of 
this signal. This produces the envelope detected or demodulated image 
signal. Alternatively, the demodulation can be performed after the 
individual received signals are summed. The signal processor 120 comprises 
an envelope detector for forming the envelope of the complex signals (I 
and Q), at which point the phase information is lost. 
The axial resolution of an ultrasound imaging system of the foregoing type 
is primarily determined by the finite bandwidth of the transducer. In 
accordance with conventional ultrasound imaging methods, the highest 
possible resolution is obtained by means of an impulse excitation which 
utilizes the entire available bandwidth of the transducer. Unfortunately, 
the available energy in an impulse excitation is low, which results in 
poor sensitivity. In order to compensate for this, a larger driver pulse 
can be used. However, there are a number of factors which limit the amount 
of peak-to-peak voltage which can be applied to a transducer. These 
limitations are brought about by the finite peak-to-peak voltage available 
from the driver electronics, the breakdown voltage of the piezoceramic 
material and the possibility of depoling of the piezoceramic or 
piezoelectric material, the need for high-voltage driver stages and 
regulatory limits on the peak pressure to which a patient can be exposed. 
Furthermore, under impulse excitation the bandwidth of the emitted pulse 
is limited to the transducer bandwidth. The increased bandwidth would 
result in an improved spatial resolution, improved contrast resolution and 
an improved depth of field. 
SUMMARY OF THE INVENTION 
The present invention is a new signal processing technique for improving 
the resolution and/or the sensitivity of an ultrasonic imaging system 
together with an extended depth of field. The technique is based on 
combining two or more transmit wave functions in an optimal manner to 
achieve a synthetic waveform which has greater bandwidth and/or energy 
than the individual wave functions. Methods to improve the technique by 
improving the frame rate, reducing the motion induced errors in the 
processing, adjusting the phase of transmit waveforms for optimum 
response, combining different imaging modes by selectively using echo 
signals corresponding to individual or multiple transmit waveforms and its 
application for ultrasound imaging using contrast agents are described. 
This invention applies to all imaging modes (B. M, Color, Pulsed Doppler 
Imaging and Doppler). This scheme operates on the imaging data before the 
data reaches the envelope detector, while the phase information is still 
maintained within the signal. This is different from conventional 
frequency compounding which is generally performed after envelope 
detection, or incoherent summation, to reduce speckle. 
The concept of performing a coherent summation of backscattered narrowband 
RF signals centered at different frequencies to obtain a flat frequency 
response was previously described "Influence of heart rate, preload, 
afterload, and isotropic state on myocardial ultrasonic backscatter" by K. 
B. Sagar et al., Laboratory Investigation Ultrasound, circulation 77, No. 
2, pp. 478-483, 1988!. However, this was purely a tissue characterization 
study whereby an unfocused disc transducer was used to measure the 
backscatter coefficient of cardiac tissue in a very small range-gated 
subendocardial region. For each subject the coherent summation was 
performed off-line based on separately stored range-gated data for 
different frequency scans, and the result is a single numeric estimate of 
the backscatter coefficient of the range-gated tissue region. In contrast, 
the present invention pertains to use of synthetic transmit waveforms for 
real-time, two-dimensional imaging using a general purpose ultrasonic 
scanner with a transducer array. Specifically, this invention discloses: 
(1) The details of the transmit waveform design procedure for optimum 
synthetic waveform design. 
(2) Means to reduce and/or compensate for the motion-induced errors which 
can result in a significant degradation of the image when using synthetic 
waveform design. 
(3) The optimum detection hardware to combine the two waveforms. 
(4) The application of the synthetic waveform design to increase the 
transmit depth of field in an image. 
(5) The application of the synthetic transmit waveform design to improve 
the sensitivity while maintaining wide bandwidth. 
(6) The application of combining different imaging modes whereby the 
individual waveforms are used for color/Doppler mode imaging and when two 
or more waveforms at different frequencies are combined the B-mode or the 
M-mode images are generated. 
(7) The application of synthetic waveform imaging for ultrasonic imaging 
using contrast enhancement agents. 
(8) Means to increase the frame rate by having overlapping vectors and by 
having the two transmit focal zones at a slightly different position. 
Using a synthetic transmit wave design approach, the emitted pressure of 
the combined waveforms would have a bandwidth which would be wider than 
the transducer impulse response and/or with higher sensitivity. It also 
provides larger depth of field together with improved contrast resolution. 
The synthetic transmit waveform design scheme in accordance with the 
preferred embodiment of the invention consists of firing two or more 
relatively long transmit waveforms in succession, both transmit waveforms 
being focused at the same focal position and the frequency spectrum for 
each of these transmit pulses being centered at slightly different 
frequencies. Once the received signals from all of these transmit 
waveforms are added, while maintaining their phase information, the 
resultant synthetic waveform will have an overall bandwidth which can be 
higher than even an impulse excitation. This scheme also results in an 
improved sensitivity due to the increased energy associated with the 
longer transmit pulse length. Compared to an impulse excitation (applied 
twice), a two-waveform synthetic approach can produce a response with 
greater sensitivity and/or wider bandwidth.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS 
To understand the theory of operation of the invention, consider an 
ultrasonic transducer with a finite Gaussian impulse response as shown in 
FIG. 3B. This transducer was examined under two different conditions. 
First, a single-cycle excitation waveform as shown in FIG. 3A was applied 
to the transducer. The excitation waveform was at the resonance frequency 
of the transducer. Second, the same transducer was also excited with two 
waveforms, as shown in FIGS. 3C and 3D, each of which was a two-cycle 
waveform, one at a frequency above and the other at a frequency below the 
resonance frequency of the transducer. The overall pulse length for each 
of these two narrow bandwidth excitation waveforms was longer than the 
single-cycle excitation waveform used previously. Hence the integrated 
energy for the two-cycle waveforms, FIGS. 3C and 3D, is larger than the 
energy in the single-cycle waveform, FIG. 3A. 
The transducer responses to the two waveforms from the two narrow-bandwidth 
pulses were added while maintaining the phase information (in the complex 
domain, i.e., before envelope detection). This response was compared to 
the response to the single-cycle excitation multiplied by two in FIGS. 4A 
and 4B. The response to the synthetic waveform has a shorter ringdown time 
with a corresponding improvement in bandwidth together with a small 
increase in the sensitivity. This illustrates the advantage of synthetic 
transmit waveform design. 
Alternatively, the emphasis can be to increase the peak-to-peak pressure of 
the detected pulse. Again the same analysis was performed on the response 
using the synthetic transmit waveforms approach compared to the 
single-cycle excitation at the resonance frequency multiplied by two. The 
improved sensitivity obtained from synthetic transmit design approach is 
shown in FIGS. 5A and 5B. In this case, the two narrowband two-cycle 
excitations had frequencies closer to the resonance frequency of the 
transducer than the frequencies of the two narrowband two-cycle 
excitations depicted in FIG. 3 and used to generate FIGS. 4A and 4B. If 
the synthetic waveform is made of more than two transmit waveforms, then 
even larger improvements can be expected. However, it is desirable to keep 
the number of transmit waveforms close to two in order to maintain a high 
frame rate. 
The increased bandwidth using the synthetic transmit waveform also improves 
the depth of field. The beam shape is strongly influenced by the pulse 
shape. The near-field interference and the sidelobes in the far field will 
be reduced as the pulse length becomes shorter. The position of focus, the 
depth of field and the lateral beam profile are a function of the 
wavelength. A broadband impulse response comprises a number of frequency 
components. Hence the increased bandwidth also increases the depth of 
field together with a reduction in the sidelobe level. The effect of 
impulse response on the beam profile has been described by W. L. Beaver 
in: "Sonic Nearfields of a Pulsed Piston Radiator", J. Acoust. Soc. 
America, Vol. 56, pp. 1043-1048, 1974; J. A. Hossak in: "Extended Focal 
Depth Imaging for Medical Ultrasound", IEEE Ultrasonics Symposium, 
November 1996; and P. Fish in: "Physics and Instrumentation of Diagnostic 
Medical Ultrasound", John Wiley & Sons, 1990, pp. 37-39. 
In accordance with the present invention, the transmitter 22 (see FIG. 1) 
comprises means for transmitting a first transmit ultrasound waveform 
having a first frequency spectrum which is centered at a first frequency 
by exciting selected transducer elements 12 during a first time interval 
and means for transmitting a second transmit ultrasound waveform having a 
second frequency spectrum which is centered at a second frequency by 
exciting selected transducer elements during a second time interval 
immediately subsequent to the first time interval. The first transmit 
ultrasound waveform is focused at a first focal point and the second 
transmit ultrasound waveform is focused at a second focal point near or 
identical to the first focal point. Furthermore, the second frequency is 
different than the first frequency. 
Referring to FIG. 9, the invention further comprises conventional means, 
i.e., transducer 10 and beamformer 38, for forming a first basebanded 
complex signal pair (I1 and Q1 signals) of a first receive ultrasound 
waveform derived from the portion of the first transmit ultrasound 
waveform reflected back to selected receiving transducer elements by 
scatterers in a focal zone encompassing the first focal point and 
subsequently forming a second basebanded complex signal pair (I2 and Q2 
signals) of a second receive ultrasound waveform derived from the portion 
of the second transmit ultrasound waveform reflected back to selected 
receiving transducer elements by scatterers in a focal zone encompassing 
the second focal point. The receiver transfer function would remain the 
same for both waveforms. However, the amplitudes would be different in 
order to compensate for the tissue attenuation which would be different 
for two frequencies. 
The theory of operation in accordance with the present invention will now 
be described with reference to an ultrasound imaging system of the type 
having a complex signal detector, as shown in FIG. 9. In this system, an 
equalization filter 40 receives the complex I and Q outputs from the 
beamformer 38 and passes a bandwidth which is a function of the bandwidth 
of the input signals. In accordance with the invention, the outputs of 
equalization filter 40 are transmitted to a multiplexer 50 which 
selectively multiplexes the first and second complex signal pairs as 
follows: the I1 and Q1 signals are multiplexed to the delay circuits, or 
buffers, 52A and 52B respectively during one cycle; and the I2 and Q2 
signals are multiplexed directly (without delay) to the adders 54A and 54B 
respectively during the next cycle. The delay circuits 52A and 52B provide 
a delay of one cycle so that adder 54A receives the I1 and I2 signals and 
adder 54B receives the Q1 and Q2 signals during the second cycle. Adder 
54A forms the sum Isum=I1+I2; adder 54B forms the sum Qsum=Q1+Q2. The 
complex signals Isum and Qsum are then input into envelope detector 42, 
which calculates the function (Isum2+Qsum2)1/2. 
Thereafter the envelope is passed through a low pass filter 44 and then the 
filtered envelope undergoes logarithmic data compression (block 46). The 
logcompressed signal is output to the scan converter 32 and then displayed 
as a vector on monitor 30. 
In practice, the received signals for the different transmit frequency 
waveforms can have very different amplitudes due to frequency-dependent 
tissue attenuation. To compensate for tissue attenuation effects, 
different weighting coefficients should be applied to the different 
transmit waveforms, or to the different received signals before the 
coherent sum or a combination of both. In the preferred method the 
coefficients should vary dynamically with time since the goal is to 
generate a synthetic signal with a symmetrical spectrum at all depths. The 
coefficients can be specified in terms of the tissue absorption and 
frequency-dependent scattering characteristics. 
The weighting coefficients are realized in the time gain control section. 
Referring to FIG. 10, the time gain control section 100' comprises two 
sets of potentiometers 108a and 108b which are alternately connected to 
the time gain circuit 106 by means of a set of switches 120. The settings 
of potentiometers 108a are selected to provide a first set of weighting 
coefficients, which are used to adjust the amount of amplification 
provided by amplifiers 105 during a first receive interval; the settings 
of potentiometers 108b are selected to provide a second set of weighting 
coefficients, which are used to adjust the amount of amplification 
provided by amplifiers 105 during a second receive interval. In accordance 
with the preferred embodiment, all of the weighting coefficients of the 
first set have a first value; likewise all of the weighting coefficients 
of the second set have a second value different than the first value. The 
first receive interval transpires during reception of the received signal 
produced in response to the first transmit wave form; the second receive 
interval transpires during reception of the received signal produced in 
response to the second transmit waveform. 
The attenuation effect can be further compensated by having a greater 
number of cycles in the higher frequency transmit waveform, which 
effectively increases the energy in the higher frequency band edge of the 
synthetic transmit waveform. Alternatively, a dynamic equalization filter 
can be applied to each of the different received signals to compensate for 
tissue attenuation effects, before they are summed to produce a synthetic 
wideband signal. Such equalization filters can be bandpass or complex low 
pass filters depending on whether the beamformed data is of RF/IF or 
baseband type respectively. The advantage of using equalization filtering 
is that it can potentially realize additional bandwidth and/or sensitivity 
gain. Specifically, for narrowband excitation, a narrowband equalization 
filter can be applied to reject out-of-band noise before the different 
received signals are summed. The equalization filter can also be designed 
to"whiten" received spectra that are otherwise skewed, and/or to reject 
undesirable spectral sidelobes. 
The proposed theory of operation of the present invention was verified by a 
series of 5-MHz imaging experiments on a standard phantom. An I/Q 
beamformed data set for an optimum conventional transmit waveform design 
was recorded via a data acquisition unit. A similar data set was also 
obtained using the synthetic transmit waveform design approach of the 
invention. B-mode images were created off line by taking the polar 
magnitude of the I/Q data, and then displaying the logcompressed images on 
a Sun workstation. To produce the synthetic waveforms, two"narrowband" 
transmit waveforms were used (one of two received waveforms was actually 
equalization filtered to remove an undesirable spectral sidelobe). FIG. 6A 
shows the two corresponding received spectra which, for the purpose of 
comparison, have been normalized and shifted back to the demodulation 
frequency of 5 MHz. These spectra represent the ensemble-average of 10 
successive central scan lines, where each sample spectrum is computed by 
taking a complex Fast Fourier Transform over a 1-cm axial window centered 
at a depth of 3.5 cm depth. To create the synthetic signal as described in 
the previous section, the two signals were normalized in mean amplitude 
and summed (in baseband) before detection, and the result is shown in FIG. 
6B. There is a significant increase in bandwidth in the composite response 
relative to each of the two transmit waveforms. FIG. 7A shows the 
comparison between the synthetic spectrum (FIG. 6B) and the received 
spectrum from the optimal conventional transmit waveform. The bandwidth 
improvements at -6, -10 and -20 dB below the spectrum peak are 30%, 40% 
and 14%, respectively. FIG. 7B shows the corresponding image segments of 
the phantom, which clearly demonstrate an increased pin resolution and 
much finer speckle texture and increased depth of focus with the synthetic 
waveform approach. 
A similar experiment was performed in the body. Here a series of renal 
scans were performed using exactly the same set of transmit waveforms as 
in the phantom experiment. The received signals for the synthetic transmit 
methods (corresponding to those of FIGS. 6A and 7A) are shown in FIGS. 8A 
and 8B. As in the phantom case, the individual low- and high-frequency 
components add up to produce a synthetic signal which has even greater 
bandwidth than a conventional wideband pulse. B-mode images created 
off-line also showed a finer speckle texture and increased depth of focus 
with the synthetic waveform approach. 
As with conventional synthetic aperture methods, one requirement for the 
synthetic transmit waveform design of the present information is that the 
phase relation be maintained between two successive firings. If there is 
any motion of the object or the transducer during the interval between the 
two successive firings, the phase information will be distorted. This 
results in a degradation of the image compared to conventional techniques. 
In order to prevent this error, the maximum tissue displacement between 
successive firings must be less than or equal to one-tenth of a 
wavelength. Although this is a conservative estimate, it may require that 
successive transmit waveforms be fired over a very short time interval 
such that successive vectors have overlapping periods. For example, when 
operating at 3 MHz, the overlapping can be less than or equal to 4 cm for 
adequate temporal sampling. With such overlapping periods, having 
different transmit waveforms may help if the receiver can be uniquely set 
up for optimal detection of each of these signals. Furthermore, it is 
advantageous to fire the high-frequency signal first before firing the 
low-frequency waveform, since the high-frequency signal would be 
attenuated at a higher rate. This would reduce the amount of acoustic 
noise which remains from the previous firing. This is especially important 
when vectors with overlapping periods are fired. These are the vector sets 
whereby the second vector is fired while the data corresponding to the 
first vector are still being gathered. 
It is also possible to incorporate motion estimation and motion 
compensation algorithms, such as correlation techniques, to compensate for 
tissue motion during successive firings. This would reduce the amount of 
constraint on the firing interval between two successive firings. 
The basic concept of the invention can be extended in many directions. For 
example, the synthetic transmit waveforms can be designed such that the 
axial resolution is improved in the near field (e.g., using narrowband 
waveforms of the type used to produce the responses shown in FIGS. 4A and 
4B) and the sensitivity is improved in the far field (e.g., using 
narrowband wave-forms of the type used to produce the responses shown in 
FIGS. 5A and 5B). This method would fully utilize the advantages of the 
synthetic transmit waveform design at all depths. This could depend on the 
clinical examination application. Alternatively, two synthetic transmit 
waveforms can be used in the near field and three or more waveforms at 
greater depths. 
The two transmit waveforms can have different numbers of cycles. For 
example, it can be three cycles at the higher-frequency excitation and two 
cycles at the lower-frequency excitation. This can potentially increase 
the energy in the region of the spectrum which is going to be attenuated 
at a higher rate due to higher frequency. Hence the two transmit waveforms 
can have different bandwidths. 
Typical transmit waveform are tonebursts with a duration of one or more 
cycles. This would have a frequency spectrum which is a sinc (sin(x)/x!) 
function. The preferred mode of operation would design the two or more 
waveforms such that the frequency domain sidelobes of one waveform would 
overlap with the nulls (or zeros) of the proceeding waveform. When the two 
waveforms are combined using the synthetic waveform design, the frequency 
domain response would have reduced sidelobes with a corresponding shorter 
time domain ringdown time. 
The F number or active element spacing can be changed to maintain a 
constant beam width for the two firings. Hence, when operating at the 
lower frequency a larger aperture can be used to maintain a constant 
lateral point spread function. Note that this is not the same as the 
conventional synthetic aperture approach in which the two firings of the 
same waveform from different sub-apertures are combined to achieve a 
larger effective aperture. 
The proposed algorithm can be applied to power Doppler imaging (PDI). The 
axial resolution or sensitivity can be improved at different depths or for 
different applications. In this case for every packet of data in the PDI 
processing two or more firings must be made very rapidly. A similar scheme 
can be applied to color flow imaging. 
It is also possible to combine the color and/or Doppler mode with B-mode 
(or gray scale) imaging. The lower-frequency waveforms with the higher 
sensitivity can be used on its own for flow detection (Doppler or color 
mode) and the composite of high- and low-frequency waveforms can be used 
for the B-mode, where contrast and detail resolution are more important. 
In order to maintain a high frame rate, the algorithms can be applied to a 
restricted region of interest. Hence the number of transmit focal zones 
can be reduced together with the number of beams which are fired over the 
region of interest. 
In order to reduce the effect on frame rate, two successive transmit beams 
with the focus at slightly different positions can be used to cover two 
transmit focal zones with two firings. This would mitigate the reduction 
in the frame rate. However, the temporal sampling requirement still 
remains (i.e., rapid firing of two transmit waveforms). As an example, 
consider an imaging system in which four focal zones are employed to cover 
the region of interest, each focal zone having a depth of 3 cm. If two 
firings are focused at a respective focal point in the four focal zones, a 
total of eight firings would be required. Alternatively, the same 12 cm 
can be divided into three focal zones, each 4 cm deep. If for each of the 
three focal zones, two firings are focused at two different focal points 
separated by 1 cm, then the full 12 cm of depth can be covered by six 
firings instead eight, thereby mitigating the reduction in frame rate 
resulting from use of the invention. 
The decision on either the bandwidth or sensitivity improvement can also 
depend on the type of vector being fired. For example, when firing the 
steered beams (beams which are not normal to the transducer or the 
aperture planes) the sensitivity can be improved. The bandwidth can be 
improved for all other cases. 
Even if tissue attenuation is already compensated by other means, the 
narrowband synthetic transmit waveform together with the narrowband 
equalization filter can significantly improve the signal-to-noise ratio 
compared to conventional technologies. 
One can possibly compensate for the motion-induced errors using correlation 
techniques, logical operators, gradient methods or optical flow techniques 
to compensate for any motion before performing the weighted sum. This way 
the potential phase errors introduced by the motion can be reduced or 
removed. Hence the algorithm can be applied at slow frame rates. 
The synthetic waveform design can also be used for conventional contrast 
imaging as well as second harmonic imaging, when using various ultrasonic 
contrast agents. The longer transmit pulse in synthetic bandwidth imaging 
should enable use of lower maximum acoustic pressure or intensity. This 
extra degree of control in transmit signal level is likely to be very 
important for achieving the desirable effects in contrast imaging. 
Additionally, for second harmonic contrast imaging, the second harmonic 
frequency band in the transmit signal due to the spectral leakage and/or 
nonlinear propagation effects can be suppressed by using longer or more 
puretone signals. This may prove to be important for discriminating the 
second harmonic signals generated by the contrast agents from direct 
backscattering of the second harmonic frequency band from tissue. 
Finally, in the synthetic waveform approach of the present invention, two 
or more narrowband excitations may help reduce phase aberration effects 
caused by the fat/muscle surface layers. 
The foregoing preferred embodiments have been disclosed for the purpose of 
illustration. Variations and modifications will be readily apparent to 
those skilled in the art of baseband ultrasonic imaging systems. All such 
variations and modifications are intended to be encompassed by the claims 
set forth hereinafter.