Ultrasonic harmonic imaging system and method

Abstract

A method and system for ultrasonically imaging a target with energy spreading transmissions is provided. Simultaneous and sequential compound point, line and combination point and line focusing is used with harmonic receive processing. Ultrasonic energy focused at multiple depths is transmitted at the target. The target may or may not include contrast agents. In either case, echoes of the transmissions at one or more fundamental center frequencies are received at harmonics of the fundamental frequencies.

Description

BACKGROUND

This invention relates to ultrasound imaging systems, and in particular to improved methods for imaging nonlinear contrast agents and tissue with such systems.

Nonlinear contrast agents are described for example by V. Uhlendorf, et al., in "Nonlinear Acoustical Response of Coated Microbubbles in Diagnostic Ultrasound" (1995 Ultrasonic Symposium, pp. 1559-1562). Such agents possess a fundamental resonant frequency. When they are insonified with high intensity ultrasonic energy at this fundamental frequency, they radiate ultrasonic frequency at a harmonic of the fundamental frequency. Such contrast agents are often used to highlight regions containing blood loaded with the contrast agent. For example, in the case of a blood-filled chamber of the heart, the borders of the chamber can be distinguished more easily when contrast agent is used. Since the contrast agent generates harmonic ultrasound energy, echoes from tissue (containing no contrast agent) at the fundamental frequency may be eliminated by filtering at the receive beamformer.

Typically, such agents are used with an imaging system having a transmit beamformer that transmits ultrasonic energy at the fundamental frequency and a receive beamformer responsive to the harmonic. In order to image the contrast agent clearly, it is known to reduce energy at the harmonic in the transmit beam, and to reduce sensitivity of the receive beamformer to energy at the fundamental.

In the past, this has been done by using a burst of square or sine waves to form the transmit beam, and by using appropriate band pass or high pass filters in the receive beamformer. Though a large pulse count reduces energy at the harmonic, it reduces time resolution of the pulse, and therefore spatial resolution of the resulting image.

The present invention is directed to further improvements that enhance the imaging of such nonlinear contrast agents, as well as the nonlinear response of tissue.

SUMMARY

The present invention is defined by the following claims, and nothing in this section should be taken as a limitation on those claims. By way of introduction, the system described below ultrasonically images a target with energy spreading transmissions. Generally, ultrasonic energy corresponding to a multiple depth focus is transmitted into a target at a fundamental center frequency. The target may or may not include contrast agents. In either case, echoes are received at a harmonic of the fundamental center frequency.

In one mode of operation, ultrasonic energy centered at a fundamental frequency and focused at a first depth is transmitted along a scan line. Substantially simultaneously, ultrasonic energy centered at the fundamental frequency and focused at a second depth is transmitted along the scan line. Ultrasonic energy from the target is received at a harmonic of the fundamental frequency and is associated with at least the first and second depths. In alternative modes of operation, the first and second depths are associated with different scan lines.

In a second mode of operation, the target is free of contrast agent throughout an entire imaging session. Ultrasonic energy centered at a fundamental frequency and focused at a first depth is transmitted along a scan line. Ultrasonic energy from the target is received at a harmonic of the fundamental frequency and is associated with at least the first depth. After transmitting and receiving at the first depth, the transmit and receive steps are repeated for a second depth.

In a third mode of operation, sequential transmission and reception as discussed above is performed, regardless of the existence of contrast agent in the target. The process is repeated for a plurality of scan lines. A frame of data associated with the plurality of scan lines is generated.

In a fourth mode of operation, a first transmit beam focused at a first depth is transmitted, and a second transmit beam having a line focus is also transmitted. In other modes of operation, two transmit beams each having a line focus are transmitted at a fundamental center frequency. Echo information associated with harmonic frequencies are received.

The transducer is preferably tuned or optimized for transmission of waveforms at a fundamental frequency and receive of waveforms at a harmonic frequency is provided.

Further aspects and advantages of the invention are discussed below in conjunction with the preferred embodiments.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a block diagram of an ultrasonic imaging system suitable for use with the method of this invention.

FIG. 2 is a block diagram of a transmit beamformer suitable for use in the system of FIG. 1.

FIG. 3 is a graph of a Gaussian pulse in the frequency domain.

FIG. 4 is a graph of a waveform corresponding to the Gaussian pulse of FIG. 3 in the time domain.

FIG. 5 is a graph of a focusing profile suitable for an axicon focus.

FIG. 6 is a schematic representation of a compound focus arrangement.

FIG. 7 is a graph showing high and low frequency wavefronts.

FIG. 8 is a block diagram of a second transmit beamformer suitable for use in the system of FIG. 1.

FIG. 9 is a block diagram of a third transmit beamformer suitable for use in the system of FIG. 1.

FIGS. 10, 11, and 11a are block diagrams of portions of transmit beamformers that incorporate preferred embodiments of this invention.

FIG. 12 is a waveform diagram illustrating operation of the beamformers of FIGS. 10-11a.

FIGS. 13-16 are graphs illustrating the temporal and frequency characteristics of waveforms produced by the beamformers of FIGS. 10-11a.

FIG. 17 is a graph illustrating a filter transfer function suitable for use in the beamformers of FIGS. 10-11a.

FIGS. 18-23 are block diagrams of alternative filter circuits suitable for use with ultrasonic harmonic imaging systems.

FIG. 24 is a circuit diagram of one embodiment for tuning a transducer.

FIG. 25 is a circuit diagram of another embodiment for tuning a transducer.

FIG. 26 is a graph illustrating the frequency response of a tuned transducer.

DETAILED DESCRIPTION OF THE PRESENTLY PREFERRED EMBODIMENTS

The preferred embodiments described below are designed to reduce harmonic energy in the transmitted beam, and to provide an improved spatial distribution of fundamental energy in the transmitted beam. FIG. 1 shows generally an ultrasonic imaging system 10 which can be used to practice the method of this invention.

The system 10 includes a transmit beamformer 12 that supplies high voltage transmit waveforms via a multiplexer 14 to a transducer array 16. The transducer array 16, which can be any suitable type, generates an ultrasonic transmit beam in response to the transmit waveforms, and this transmit beam propagates outwardly through the subject 18 being imaged. In this case, the subject being imaged includes a nonlinear contrast agent 20, such as that described above. As used herein, nonlinear contrast agent refers to any contrast agent with a nonlinear response, whether the contrast agent is designed for fundamental or harmonic imaging. Any suitable contrast agent may be used, as long as it absorbs ultrasonic energy at a first frequency and radiates ultrasonic energy at a second frequency, different from the first frequency when subject to a high intensity acoustic pulse. In this example, the first frequency is referred to as the fundamental frequency, and the second frequency is a harmonic of the first frequency. As used herein, "harmonic" is intended broadly to include sub-harmonics and fractional harmonic energy (e.g. 1/2 or 3/2 of the fundamental), as well as higher harmonics (e.g. 2 or 3 times the fundamental).

Ultrasonic energy radiated by the nonlinear contrast agent 20 at the harmonic frequency is received by the transducer array 16, focused by the receive beamformer 22, and displayed as an image by a display processor (not shown). As described above, the receive beamformer 22 includes conventional filters to substantially prevent reflected ultrasonic energy at the fundamental frequency from being imaged.

Turning now to FIG. 2, this figure shows a block diagram of a first preferred embodiment 40 of the transmit beamformer of FIG. 1. As shown in FIG. 2, the transmit beamformer 40 includes N channels, one for each of the transducers of the transducer array 16 (FIG. 1). Each channel includes a delay memory 42, a waveform memory 44, and a delay counter 46 (FIG. 2). The delay memory 42 includes 256 words 48, one for each possible steering angle or ultrasound transmit scan line. Each word 48 of the delay memory 42 is set equal to a negative number equal to the number of clock cycles on the clock signal line 50 that elapse between a start of line signal on line 54 and the first non-zero value of the associated waveform. For simplicity, it is assumed that zero is defined as a word 48 having the most significant bit equal to one and all other bits equal to zero. Hence, the most significant bit becomes an enable signal for the waveform memory 44.

The waveform memory 44 in this embodiment stores a single waveform in digital form, which is used for all transmit scan lines. The waveform memory 44 can include for example 64 or 128 successive 8 bit words. The magnitude of each 8 bit word corresponds to the voltage amplitude at the respective position in the waveform. When the waveform memory 44 is read with a 40 MHz clock on the line 50, the resulting sequence of digital values defines a waveform approximately 1.6 to 3.2 microseconds in duration.

The delay memory 42 is not required, but it reduces memory requirements for the waveform memory 44. This is because the delay memory 42 eliminates the need to store a large number of leading zeros when the ultrasound line is steered at a large angle.

In use, each channel responds to a scan line selection signal on line 52 by loading the word 48 for the selected scan line into the delay counter 46. The delay counter 46 responds to a start of scan line signal on line 54 by incrementing the stored value with each cycle of the 40 MHz clock on line 50. When the counter 46 increments to zero, it enables the waveform memory 44. Subsequently generated values of the counter 46 (incrementing now from zero upwards) become address values for the waveform memory 44. As each word of the waveform memory 44 is addressed, the corresponding 8 bit word is read and applied to a digital to analog converter 56.

The analog output signal of the converter 56 is passed through a low pass filter such as a Bessel filter 58 to reduce sampling effects and then to an amplifier 60. The output of the amplifier 60 can be passed through an additional low pass filter 62 to improve harmonic rejection. The output of the low pass filter 62 is the transmit waveform discussed above that is applied to the respective transducer via the multiplexer 14 (FIG. 1). The low pass filters 58, 62 preferably provide a sharp cut-off with a low stop band level in order substantially to eliminate ultrasonic energy in the transmitted pulse at the harmonic frequency.

The transmit beamformer 40 utilizes values for the waveforms stored in the waveform memory 44 and the delays stored in the delay memory 42 that enhance insonification of the nonlinear contrast agent in the subject.

The waveform stored in the waveform memory 44 is shaped to suppress ultrasonic energy in a wide pass band centered at the harmonic frequency. For example, the spectrum of the desired pulse can be designed on a computer 64. FIG. 3 shows the frequency spectrum of one suitable pulse 70 which is centered at the fundamental frequency of 2.5 MHz and is generally Gaussian in shape. The particular Gaussian shape shown in FIG. 3 has an amplitude reduced by 71 dB at 3.71 MHz. The bandwidth of the pulse 70 is 30% of the center frequency, measured at points -6.8 dB with respect to the peak amplitude. Such a pulse will be referred to herein as a 30% BW Gaussian pulse. Note that the pulse 70 has substantially no energy at 5 MHz, the first harmonic of the fundamental frequency. This invention is not limited to use with Gaussian pulses, and a wide range of spectra can be used.

Once the desired pulse has been designed, an inverse fast Fourier transform is then performed to generate the corresponding time domain waveform. FIG. 4 shows a waveform 80 which corresponds to the pulse 70 of FIG. 3. Note that the waveform 80 includes an oscillating component having a frequency of about 2.5 MHz. This oscillating component is shaped by an envelope 82. The envelope 82 rises gradually from zero amplitude to a maximum amplitude, and then falls gradually from the maximum amplitude back to zero amplitude. Thus, the envelope 82 is quite different from the envelope for a switched pulse train, which is substantially rectangular in shape. The gradually increasing and gradually decreasing envelope 82 of FIG. 4 brings with it the advantage of reduced ultrasonic energy at harmonics of the fundamental frequency.

Once a waveform such as the waveform 80 of FIG. 4 has been designed, the waveform 80 can be coded into binary samples at a suitable sample rate and then stored in the waveform memory 44. The waveform memory 44 may be a read only memory, in which case the computer 64 may not be required to be connected to the transmit beamformer 40. Alternately, the waveform memory 44 may be a volatile memory which is programmed at power-up initialization by the computer 64. The computer 64 may perform any desired subset of the pulse designing steps described above. Typically, the desired pulse may be one of several selectable pulses included in a menu for user choice.

When the waveform in the waveform memory 44 is designed as described above, the result is a broad band waveform in the waveform memory 44 which simultaneously has substantially no radiated energy in a broad band centered on the harmonic. In the example of FIGS. 3 and 4, substantially no ultrasonic energy is radiated at frequencies above 4 MHz, or in a bandwidth of .quadrature.1 MHz with respect to the first harmonic (5 MHz). Preferably, the energy component at the harmonic is more than 30 dB reduced with respect to the magnitude of the fundamental frequency, and ideally is reduced by more than 40 dB.

Of course, it is not necessary to define the waveform 80 initially in the frequency domain. A Gaussian pulse can be defined in the time domain. Furthermore, the envelope need not be Gaussian, it may be some other window function such as a Hamming pulse, a modified Gaussian pulse, or any other suitable pulse. In some applications it may be preferable to use a narrow bandwidth pulse and thereby achieve a very high reduction of energy at the harmonic, since the harmonic of the lower bandedge is well above the upper bandedge. On other occasions it may be preferable to use a wider bandwidth pulse, for example, to obtain better axial (temporal) resolution. In this case, somewhat reduced reduction of energy at the harmonic may be accepted.

An alternative approach is to design the pulse 70 to ensure that substantially no energy is transmitted above 1.5 times the center frequency of the intended fundamental pulse (3.75 MHz in this case). It is preferable to limit low frequency energy in the transmitted pulse so that the harmonic energy associated with this frequency does not fall within the spectrum of the transmitted pulse. If the cut off frequency is 3.75 MHz, there should be very little transmitted energy below 1.875 MHz.

In designing the pulse 70 and the waveform 80, the Gaussian waveform may be passed through a high order, low pass filter to eliminate all harmonic energy. This filtering may be done off line in the computer 64.

Optimum imaging of the nonlinear contrast agent is obtained when the transmit beam insonifies the agent at power levels within a desired range. Power levels below this range may not be sufficiently high to cause the nonlinear contrast agent to radiate at the harmonic frequency. Power levels above this range may destroy the contrast agent prematurely. Additionally, since there are FDA limits on ultrasound intensity, a sharply focused transmit beam is not optimal. Such a sharply focused beam provides a high intensity (near the FDA limits) at the focus, but an intensity that is undesirably low at many other points along the associated receive scan line.

The receive beamformer 22 preferably receives samples along an entire scan line for each transmit event. For this reason, it is preferable that the region of insonification within the desired power level range be extended over a substantial portion of the length of the receive scan line. Thus, it is preferable that the intensity of the transmitted ultrasonic energy be substantially uniform and at a high level throughout the field of interest (which is typically a large fraction of the displayed depth).

The delay memory 42 preferably stores delay values to provide a controlled spread to the beam intensity in a way to optimize imaging of the contrast agent. Also, by making the intensity of harmonic energy received at the receive beamformer 22 more uniform throughout the field of interest, the levels of harmonic back scatter may be better controlled, resulting in manageable voltage swings at the receiver input.

In this embodiment, the delay values stored in the delay memory 42 are selectively chosen to spread the beam along the current ultrasound line. One way to accomplish this is to use the well-known axicon focusing arrangement, as described, for example by C. Burckhardt in "Ultrasound Axicon: A Device for Focusing over a Large Depth" (J. Acoust. Soc. of Am., 54, 6, pp. 1628-1630 (1973)). The axicon focusing arrangement may utilize a focusing profile as shown in FIG. 5. Typically, this focusing profile provides a near focal limit corresponding to a circular arc centered on the near focal limit. Typically, the delay profile extends linearly outwardly from this circular arc to some outer limit, as shown in FIG. 5.

The objective is to spread the ultrasound energy throughout a region of the target, and many different delay profiles may accomplish this result. For example, the delay profile may be slightly curved, with a nonlinear variation of focal point with respect to transducer element position. There may be an outer focal limit, in which case the delay profile can include a circular portion at the ends of the array.

In many applications, it will be desirable to select the delay values in the delay memory 42 such that at least first frequency components of the transmit beam from at least a first plurality of transducers are focused at a first, shorter range, and that at least second frequency components of the transmit beam from at least a second plurality of transducers are focused at a second, longer range. One example is shown in FIG. 6, where substantially all of the ultrasonic energy from the transducers 16a at the end portions of the transducer array 16 are focused at a single longer range X1, and substantially all of the ultrasonic energy from the transducers 16b at central portions of the array are focused at a single, shorter range X2. By properly selecting the delay values, a line focus or a multiple-point compound focus may be obtained. When a line focus is used, the line may be straight or curved.

Another approach begins with focal delays required for a chosen focal point in the conventional manner. A random delay error is superimposed on these focal delays to smear or defocus the resulting beam. The greater the degree of defocusing, the more spread out the beam is. Preferably, a user control is provided to allow the user to vary the degree of defocusing by increasing the relative level of the defocusing delays. Also, it is preferable to increase the transmitted energy level to partially compensate for the loss of peak field intensity due to defocusing, either in an automatic (internally programmed) manner or under user control. The defocusing hardware can consist of a modified version of aberration correction hardware in which the delay corrections are pre-programmed random numbers rather than values which are continuously updated, for example by means of cross-correlating the waveforms from adjacent elements.

Conventional imaging at the fundamental frequency using an axicon or defocused beam is known to produce side lobes. However, such side lobes are not anticipated to create substantial problems in this application, particularly if the side lobes are below the harmonic activation threshold intensity and they therefore do not contribute to the generation of harmonic energy. Full dynamic receive focusing is preferably performed in the receive beamformer 22 to reduce the effect of side lobes in the transmit beam further.

Regardless of the precise configuration of the pulse 70 and the waveform 80, the waveform 80 preferably provides more uniform field intensity through a usefully extended depth of field. This results in more uniform generation of harmonic energy by the contrast agent and possibly a higher overall signal to noise ratio since more of the field is being unisonified at a sufficiently high power level to cause the contrast agent to radiate harmonic energy, but at a sufficiently low power level to prevent unnecessarily rapid destruction of the contrast agent.

If desired, the axicon focus may be obtained with a lens. Also, an axicon focusing scheme may be used in the elevation direction, if it is desired to increase the dimension of the insonified region in that direction.

The transmit beamformer 40 of FIG. 2 is quite similar to the beamformer shown in FIG. 13 of U.S. Pat. No. 5,608,690, and the techniques described above relating to Gaussian waveforms modified to provide a line focus can be performed in the manner described in the above-identified application. Note in particular pages 23 and 24 of the above-identified application as filed, which expressly relate to optimized transmit beamformers for use with nonlinear contrast agents. With this approach, a plurality of transmit waveforms are provided, each for a respective one of the transducers of the transducer array. This plurality of transmit waveforms includes a central transmit waveform associated with a central one of the transducers. As explained in the above-identified patent application in detail, the central transmit waveform preferably comprises a lower frequency component and a higher frequency component, and a lower frequency component of the central transmit waveform occurs earlier in time than the higher frequency component of the central transmit waveform, as shown in FIG. 7. FIG. 7 corresponds to FIG. 11 of above-identified U.S. Pat. No. 5,608,690, and that application can be referenced for additional information regarding these figures. When the system of U.S. Pat. No. 5,608,690 is adapted for use with this invention, it is preferred that the frequencies along the line focus all be near the fundamental frequency to insonify the non-linear contrast agent effectively.

Additionally, the transmit beamformer described in Cole, et al., U.S. patent application Ser. No. 08/286,652, (abandoned) filed Aug. 5, 1994, and in U.S. patent application Ser. No. 08/432,056, filed May 2, 1995 (now U.S. Pat. No. 5,675,554), both assigned to the assignee of the present invention, can be adapted for use with this invention. Once the desired output is defined as described above in terms of very low harmonic signal, one can then define the ideal output signal in the frequency domain and then convert it to the time domain. This time domain signal can then be divided by the carrier to obtain the desired envelope using complex shapes for both the time domain signal and the carrier. This combination of envelope and carrier can then be programmed into the transmit waveformer, using the parameters of attached Appendix 1. Appendix 1 provides parameters for both B-mode and Flow Mode. Appendix 3 provides a revised set of transmit beamformer parameters, and Appendix 4 provides a preferred set of transmit beamformer parameters adapted for use with the Acuson 3V2C transducer probe.

The envelope is sampled at a relatively low frequency, and as a result of imperfections in real implementations, remnants of harmonics relating to the sampling frequency of the carrier and the base band signal may appear in the final result.

In this example, harmonic energy at the second harmonic from the contrast agent is imaged using the receive beamformer described in Wright, et al. U.S. patent application Ser. No. 08/286,658, (abandoned) filed Aug. 5, 1994, and in U.S. patent application Ser. No. 08/432,615, filed May 2, 1995 (now U.S. Pat. No. 5,685,308), both assigned to the assignee of the present invention. This receive beamformer can be programmed using the parameters of attached Appendix 2.

For example, the transmit beamformer of U.S. patent applications Ser. Nos. 08/286,652 and 08/432,056 can be operated with a nominal center frequency FO equal to 5.0 MHz, a transmit carrier frequency F, equal to 2.5 MHz, and a Gaussian envelope having a 50% bandwidth at -6.8 dB with 8/4 envelope sampling.

FIG. 8 shows another transmit beamformer 100 that can be used in the ultrasound imaging system 10. The beamformer 100 includes a pulse generator 102 which supplies a burst of pulses to a low pass filter 104. In this case, the pulse generator 102 switches between a selectable high voltage DC level and ground. The duration of the pulses and the number of pulses in the burst are controllable, as described, for example, in U.S. Pat. No. 4,550,067 (Maslak, et al.), assigned to the assignee of the present invention. The low pass filter 104 is designed as described above to substantially eliminate energy at the harmonic. The low pass filter 104 can be an analog filter such as a suitable Bessel or Chebyshev filter.

FIG. 9 shows another transmit beamformer 120 that can be used in the system 10 of FIG. 1. The transmit beamformer 120 includes a pulse generator 122 which supplies pulses to a low pass filter 124. The output of the low pass filter is supplied to a high voltage amplifier 126. Because the low pass filter 124 is applied prior to the amplifier 126, the low pass filter 124 can be designed with smaller components and hence can be higher order and more effective in suppressing harmonic energy.

It should be noted that conventional transducer arrays have a finite bandwidth, such as 75% at the -6 dB levels. Considering the example where the fundamental frequency is 2.5 MHz and the harmonic frequency is 5 MHz, the transducer will typically be centered between the transmit and receive frequencies (as for example at 3.75 MHz), with a relatively high bandwidth. Notice that when this transducer is operated at 2.5 MHz with a symmetrical band shape, the transducer will skew the transmitted spectrum toward higher frequencies. In the design of the transmitted pulse, the spectrum should preferably be modified to take account of the skewing effect of the finite bandwidth effects of the transducer and the fact that the transducer is working away from its center frequency in transmit.

Of course, the techniques described above can be used in systems 10 which provide apodization in the normal manner. If desired, apodization profiles can be modified if experience shows that the contributions from the end transducer elements (which are focused at deep ranges) are too attenuated. In this case, it may be preferable to increase the weighting at the end elements to compensate for this effect to some extent. By operating the end transducer elements at higher power levels than central transducer elements, more nearly uniform power levels are obtained at various ranges in the tissue being imaged.

A further modification is to use fewer transmit lines than receive lines. Since contrast agent tends to be consumed by the insonification process, it is preferable to reduce the number of transmit lines by reducing the frequency of firings and/or the spatial density of lines. Reduced firings and density can be achieved by firing one fourth the number of transmit lines and forming four synthetic receive lines closely spaced around each transmit line, which may be slightly defocused if required. Wright, et al. U.S. patent application Ser. No. 08/418,640, (now U.S. Pat. No. 5,667,373) filed Apr. 7, 1995, assigned to the assignee of the present invention, discloses synthetic line systems that can be adapted for this purpose. Furthermore, frames of low intensity ultrasound may be interleaved with high intensity frames.

Multiple transmit zone firing (compound focus) may also yield a better result by spreading the beam between two selected focal regions. These techniques can be implemented on the transmit beamformer described in U.S. patent application Ser. No. 08/286,652, as described above.

As discussed above, ultrasonic energy responsive to any of the various transmit waveforms is transmitted substantially simultaneously along an ultrasound line with two or more focal regions. For example, transmit waveforms with Gaussian envelopes and associated with different focal regions are summed. As discussed above, the waveforms preferably have the same fundamental center frequency. The summed signal excites the transducer to generate a beam of ultrasonic energy focused at two or more depths. The energy reflects from structures in the target, such as tissue and any contrast agent included in the tissue.

In alternative embodiments, two transmit beams are transmitted substantially simultaneously along two different scan lines. For example, one transmit beam is focused at a depth along one scan line, and another transmit beam is focused at the same or a different depth along a parallel scan line with a different origin, a scan line at a different angle or a scan line at a different angle with a different origin. Transmit beams are focused along different scan lines by changing the delays associated with waveforms at each transducer element. In this embodiment, each transmit beam may be associated with the same fundamental center frequency. Alternatively, different fundamental center frequencies are transmitted substantially simultaneously for each transmit beam. In this embodiment, the fundamental center frequencies and corresponding bandwidths are preferably chosen to minimize transmission for any transmit beam at a harmonic of any of the fundamental frequencies.

When multiple transmit beams are simultaneously transmitted and focused along different scan lines for fundamental imaging, acoustic clutter may remain. When used for harmonic imaging, the clutter may be reduced.

Also as discussed above, the transmit beamformer of U.S. application Ser. No. 08/286,652 is operable to generate and sum waveforms corresponding to two or more focal depths along the beam. A desired time domain signal or excitation waveform is-divided by the carrier to obtain the desired complex envelope using complex values for both the time domain signal and the carrier. The complex envelope is sampled at a relatively low frequency for generating the excitation waveforms. The waveforms are summed to generate an ultrasound beam with two or more focal points (i.e. two or more transmit beams along a same scan line).

By way of example, the parameters provided for waveforms focused at a first depth in Appendices 1, 3 and 4 may also be used for each of the two or more focal depths, except for the number of beams parameter. See U.S. Pat. No. 5,675,554. The number of transmit beams is set to 2 or more and corresponds to the number of focal points. Each transmit beam corresponds to a different focal depth along the same transmit or scan line (i.e. two transmit beams correspond to transmission of a transmit beam focused at two focal depths) and at the same fundamental center frequency.

As discussed above, harmonic energy at the second harmonic is imaged. Using any of various receive beamformers, such as the receive beamformer described in U.S. patent application Ser. No. 08/432,615 (see U.S. Pat. No. 5,685,308), ultrasonic energy from the target is received at a harmonic of the fundamental frequency. Other beamformers may be used. See Appendix 2 for various parameters associated with harmonic energy reception. Other parameters may be used. The number of receive beams is 2 in this example. For two focal depths, two or more receive beams may be used, each using the parameters disclosed in Appendix 2. For example, two receive beams corresponding to a same scan line and the two focal depths of a transmission are used. Alternatively, two or more receive beams for two or more transmit beams along different scans lines are used. In yet another alternative, multiple receive beams corresponding to different scan lines are used for each single transmitted beam (i.e. synthetic lines as discussed above).

Compound focus transmission is performed with simultaneous transmission to two or more focal regions, as discussed above, or with sequential transmission to each of the two or more focal regions. In one embodiment, a same aperture for each focal point is used for simultaneous transmission to two focal points. For a third or fourth focal region, a waveform corresponding to the focal regions is summed with the other waveforms as discussed above. The contribution to the summed signal from one waveform may be negligible for some elements in the aperture, such as at the end of a linear array when one of the beams is transmitted with a small active aperture.

For sequential compound focus, the number of simultaneous transmit beams is preferably 1, but may be 2 or more. The number of simultaneous receive beams may be 1 or more. A first transmit beam corresponding to a first focus depth is transmitted. Harmonic echoes responsive to the first transmit beam are received along a portion of or the entire depth of the scan line. A second transmit beam corresponding to a second focus depth is then transmitted. Harmonic echoes responsive to the second transmit beam are received along a portion of or the entire depth of the scan line. Both sets of transmit and receive beams are along the same beam path or scan line and at the same fundamental center frequency. Any echo information from both transmit beams representing the same range is combined or one value is selected.

For additional focus depths, additional sets of transmit and receive beams are processed sequentially with the sets discussed above. For example, each of four sets corresponds to 20, 40, 60 and 90 mm focal depths. To image an area of the target, a frame of data representing the area is obtained. The frame of data corresponds to a plurality of scan lines. Each scan line corresponds to two or more focus depths.

Preferably, the target is free of contrast agent throughout an entire imaging session as discussed below. Alternatively, contrast agents are injected during the imaging session.

A same aperture for each focal point may be used for sequential transmission to two or more focal points. For two focal points, two waveforms are sequentially generated for each transducer element in the aperture. In this example, the sequential transmissions uses the same aperture width with different delay profiles for each focal point. Alternatively, a different aperture width may be used for each focal point.

In alternative embodiments, the various foci associated with depths and lines are combined in any of the embodiments discussed above. For example, a transmit beam focused at a depth or discrete point and a transmit beam having a line focus are transmitted. The beams are transmitted sequentially or simultaneously along the same or different scan lines. In one embodiment, the beams are transmitted along the same scan line. The single depth may overlap with the line focus or occur before or after the near or outer focal limits, respectively, of the line focus. The depths discussed above correspond to ranges from the transducer. Furthermore and as discussed above, the fundamental center frequencies associated with each transmit beam may be the same or different.

In yet another alternative embodiment, compound focus of two or more beams each having a line focus is used. Each transmit beam may be transmitted sequentially or simultaneously, at the same or different fundamental center frequency, and along the same or different scan lines. For transmit beams along the same scan line, the line focus for one beam may overlap the line focus for another beam. For example, the near and/or outer limit of one line focus occurs between the near and outer limits for another line focus. Alternatively, the lines do not overlap (i.e. the outer limit of a line focus closer to the transducer occurs at a depth before the near limit of another line focus).

Compound focus provides a spreading or distribution of higher intensity through a greater depth of field or over a wider region. A rapid evolution of field intensity, coupled with remaining at a high intensity, may result in more effective harmonic imaging. Compound focus may provide a peak intensity, such as the maximum allowed Mechanical Index, for multiple points in a scan region or along a scan line as compared to one point associated with a single focal point.

Compound focus or other transmission techniques discussed herein may be used with various types of waveforms, such as Gaussian waveforms. It has been discovered that bipolar waveforms derived from switched DC levels can be filtered to provide effective suppression of ultrasonic energy at the harmonic. This can represent a substantial advantage, because bipolar, switched DC pulse sequences can be generated in a relatively inexpensive beamformer.

FIG. 10 shows a block diagram of one channel of a transmit beamformer 200 that utilizes such bipolar switched DC transmit waveforms. As shown in FIG. 10, each channel of the transmit beamformer 200 includes a binary waveform generator 202 that responds to inputs from an input clock, a select line defining the number of waveform cycles to be included in each pulse of energy, and a start input from a focal delay memory 204 that is in turn responsive to an acoustic line select input. Once the appropriate acoustic line has been selected, the focal delay memory 204 provides a start signal to the binary waveform generator 202 at the appropriate time, and the binary waveform generator 202 generates two output waveforms, which are 180.degree. out of phase in this example.

These two waveforms are applied to respective pulse amplifiers 206, 208. Each of the pulse amplifiers 206, 208 responds to a high voltage input which can be switched to one of a plurality of levels such as 10, 20 and 50 volts in this example. The pulse amplifiers 206, 208 operate with positive and negative high voltage rails, respectively. Typically, each pulse amplifier 206 includes a transistor circuit used to switch the high voltage through the pulse amplifier 206, 208 to the output of the amplifier 206, 208. Each pulse amplifier outputs an amplified binary waveform, and the waveforms are summed at a summing node prior to being applied as an input to an analog lowpass filter 210. The signal applied as an input to the analog lowpass filter is a bipolar, switched DC transmit waveform. The output of the analog lowpass filter 210 is a filtered bipolar transmit waveform, which is applied to an ultrasonic transducer element 214 via a protection circuit 212. The protection circuit 212 is included to protect the receive electronics from high voltage transmit pulses. This circuit is typically based on a diode clamping circuit. Returning ultrasonic echoes from the target are converted by the transducer element 214 into electrical signals that are passed via the protection circuit 272 to a receiver (not shown), that typically includes a receive beamformer.

The analog lowpass filter 210 may be of any standard type. The filter 210 may comprise various resistors, inductors, and capacitors, and may include Bessel, Butterworth or Chebyshev filters, for example. Preferably, the filter 210 achieves approximately 20 dB of suppression at the second harmonic (two times the fundamental frequency). Ideally, the filter 210 may achieve a far higher degree of suppression. This may be achieved by using a high order filter, or a high suppression filter such as a Chebyshev filter. If desired, a multipole filter or a band pass filter can be used. In some applications, the components of the analog lowpass filter 210 are made switchable so that the cutoff frequency of the filter 210 can be varied according to the frequency characteristics of the transducer element 214 being used. This switching may be accomplished using relays or semiconductor analog switches such as those supplied by Supertex as HV2 series type switches.

Although the filter 210 has been described above with respect to passive components, it should be understood that the filter 210 may be implemented using active filter circuits, including operational amplifiers for example. Such active filter circuits are well-known in the art, and are described for example in Horowitz and Hill, The Art of Electronics, Chapter 4 (Cambridge University Press, 1984).

Unipolar pulse sequences can also be formed using circuitry similar to that described in Maslak U.S. Pat. No. 4,140,022, which provides pulses having controllable amplitude, period, and pulse count. See also the description of pulse sequence transmitters in IEEE 1980 Ultrasonic Symposium, pp. 757-762 (Karrer, et al. "A Phased Array Acoustic Imaging System for Medical Use").

It is not required in all embodiments that the lowpass filter operate on the high voltage transmit waveform. An alternative beamformer 220 (FIG. 11) sums the two chopped DC output signals of a binary waveform generator 222 at a summer 223, and applies this summed signal as an input to a lowpass filter 224. The output of the lowpass filter 224 is applied to an adjustable gain amplifier 226 which amplifies the filtered signal to a voltage level appropriate for application to the transducer element 230 via the protection circuit 228. This approach allows the waveform to be generated and filtered at a relatively low voltage. This allows components rated at a lower voltage to be used in the filter 224, which may provide an important cost reduction.

As another alternative, the lowpass filter 210, 224 may include two components, one for each of the unipolar components of the transmit waveform prior to the point in the signal path where they are summed. See FIG. 11a, where the two components of the lowpass filter are shown at 232, 234

FIG. 12 provides waveforms that illustrate the operation of the transmit beamformers 200, 220. As shown in FIG. 12, the waveforms 240, 242 represent the output waveforms from the binary waveform generator 202, 222. Note that the positive excursions of the waveform 240 are out of phase by 180.degree. with respect to the negative excursions of the waveform 242. The waveforms 240, 242 are best considered as unipolar components of the bipolar transmit waveform. The summation of the waveforms 240, 242 is shown at waveform 244. This summed waveform is bipolar in that it includes both positive and negative components, and is generated as described above as a sum of switched DC signals. The waveform 244 represents the bipolar transmit waveform that is applied as an input to the analog lowpass filters 210, 224.

FIGS. 13 and 14 provide further information regarding the unfiltered bipolar transmit waveform 244. As shown in FIG. 13, this waveform displays excursions from +1 to -1 in an arbitrary amplitude scale, and it includes four cycles at a frequency of 2.5 MHz. FIG. 14 shows the frequency spectrum 246 of the bipolar transmit waveform 244 of FIG. 13. Note that maximum ultrasonic power is exhibited at the fundamental frequency (2.5 MHz), and that there is substantial power reduction (greater than -30 dB) at harmonics of the fundamental, such as 5 and 7.5 MHz.

FIGS. 15 and 16 are corresponding graphs for the filtered bipolar transmit waveform supplied as an output of the analog lowpass filter 210, 224 of FIGS. 10 and 11 and the summer 223 of FIG. 11 a. This filtered bipolar transmit waveform is identified as reference numeral 248, and includes four cycles with a fundamental frequency of 2.5 MHz. As shown in FIG. 16, the frequency spectrum 250 of the waveform 248 has substantially reduced power (by greater than 30 dB) at all frequencies greater than about 3.6 MHz.

FIG. 17 provides a filter transfer function for a five pole Butterworth filter (F.sub.C =3.75 MHz) that can be used for the lowpass filter 210, 224. This transfer function 252 reduces ultrasonic power at the harmonic 5 MHz with respect to the fundamental 2.5 MHz by more than 10 dB, and by about 13 dB in this example.

Of course, it will be understood by those skilled in the art that the voltage applied to the transducer elements in the beamformers 200, 220, 220' of FIGS. 10-11a described above will typically also be a function of transducer element number. Typically, apodization is employed so that end transducer elements are operated at a lower voltage than are the center transducer elements. This may be achieved by switching the high voltage input to the pulse amplifiers 206, 208 or by adjusting the gain select input for the amplifier 226 to different voltage levels according to the element position of the transducer element in the array. For example, the adjustable amplifier 226 can be a voltage controlled amplifier and the voltage control can be determined by apodization requirements.

FIGS. 18-23 relate to switched filter circuits that may be used in a method or a system for ultrasonic imaging in which ultrasonic energy is transmitted at a fundamental frequency to a target, and reflected ultrasonic energy at a harmonic of the fundamental frequency is received for imaging.

As shown in FIG. 18, the filter circuit 300 defines two signal paths. The first signal path proceeds from the input 302 via the conductor 306 to the output 304. The second signal path proceeds from the input 302 via the conductor 308 to the output 304. In this embodiment the second signal path includes a filter 312 and a isolation circuit 314.

The input 302 is intended to be coupled to the output of a ultrasonic transmitter, which can operate to generate unipolar, bipolar or shaped waveforms having substantial energy at the fundamental frequency. The output 304 of the filter circuit 300 is intended for connection either directly or indirectly to an ultrasonic transducer. For example, the output 304 may be coupled (either directly or indirectly) to a transducer connector or to a transducer cable. Similarly, the output 304 may be coupled to the input of an ultrasonic receiver as shown in FIG. 18, to a receiver protection circuit (not shown) or to a transmit/receive switch (not shown).

A switch 310 is controlled to select either the conductor 306 or the conductor 308. In a first mode of operation the switch 310 is positioned as shown in FIG. 18, and the transmitter is coupled to the transducer via the first signal path including the conductor 306. In this mode of operation the filter 312 is by-passed. In a second mode of operation the switch 310 can be placed by the control in connection with the conductor 308, and transmit signals from the transmitter to the transducer then pass through the filter 312 and the isolation circuit 314.

The filter 312 can be any suitable filter designed to suppress energy at the harmonic of interest. For example, the filter 312 can be a band reject filter centered about the second harmonic (or other desired harmonic such and third harmonic) of the fundamental frequency. Alternately, the filter 312 can be a bandpass filter centered on the fundamental frequency or a low pass filter.

The isolation circuit 314 is not required in all embodiments. However, it provides the advantage of avoiding electrical loading of the transducer elements or of the receiver input stage. Such loading is acceptable, but better performance may be obtained using the isolation circuit 314 in many cases.

The filter circuit 320 of FIG. 19 is similar to that of FIG. 18, except two switches 330, 331 are provided, operated in parallel by the control. When the switches 330, 331 are positioned as shown in FIG. 19, the input 322 is connected via the conductor 326 to the output 324, and both the filter 332 and the isolation circuit 334 are completely isolated from the signal path. When the switches 330, 331 are in the opposite position, in which the selected signal path includes the conductor 328, signals passing from the transmitter to the transducer are passed through the filter 332 on the isolation circuit 334.

In FIG. 19 the receiver input is connected adjacent the input 332, and the switches 330, 331 can automatically be controlled such that they are placed in the state shown in FIG. 19 during receive operations and are automatically switched to the opposite state for transmit operations.

The filter circuit 340 of FIG. 20 is similar to the circuit 300 of FIG. 18, with two exceptions. In this case the switch 350 has been placed adjacent the input 342, and the isolation circuit 354 has been moved adjacent the output 344. Thus, signals pass through the isolation circuit 354 regardless of whether the switch 350 routes transmit pulses via the conductor 346 or via the filter 352 and the conductor 348.

As shown in FIG. 21, a filter circuit 360 can be provided which provides two signal paths including conductors 366 and 368, respectively, between an input 362 and an output 364 for each of 128 channels. The following discussion refers only to channel 1, but the other channels are configured similarly. In this case the filter circuit 360 includes a 128 channel 1:2 multiplexer 370 that switches the transmitter signal applied to the input 362 either to output terminal 1 (connected to conductor 366) or to output terminal 129 (connected to filter 372, isolation circuit 374 and conductor 368). Thus, in a first mode of operation of the multiplexer 370, transmitter output signals on each of 128 channels are routed via the respective inputs 362 and the respective conductors 366 to the respective outputs 364, thereby bypassing the filters 372 and the isolation circuits 374. In a second mode of operation as selected by the control, transmit signals are routed from the inputs 362 via the respective filters 372, isolation circuits 374, and conductors 368 to the respective outputs 364, thereby placing the filters 372 and the isolation circuits 374 in the signal paths.

In any of the embodiments of FIGS. 18-21 the isolation circuits can be eliminated, they can be placed as shown in FIG. 20 in a common portion of the first and second signal paths, or they may be placed as shown in FIGS. 18, 19 and 21 in only one of the signal paths. Also as shown in FIGS. 18-21 the receiver can if desired be connected to the combined signal paths either upstream or downstream of the filter. In the event the receiver is connected upstream of the filter as shown in FIG. 19, the switch or switches should be controlled to prevent the filter from operating in the receive mode, at least in those cases where the receiver is responsive to harmonic ultrasonic energy. In cases such as FIG. 18 where the receiver is connected to the signal path downstream of the filter, the switch 310 can be controlled to bypass the filter 312 selectively, as for example when the transmitter is used in an alternative mode of operation in which both the transmitter and the receiver are centered at the same fundamental frequency. As shown in FIG. 21, the switches for the filters may be part of a pre-existing transducer multiplexer.

FIG. 22 is a schematic diagram of a filter circuit 380 that is similar to the filter circuit 320 of FIG. 19 with two exceptions. First, the filter circuit 380 does not include an isolation circuit. As explained above, such isolation circuits are optional. Second, the filter circuit 380 includes a total of three signal paths 386, 388, 389 between the input 382 and the output 384. The switches 390, 391 route signals from the transmitter to the output via anyone of these three signal paths. The signal path 386 includes no filter while the signal paths 388, 389 include separate respective filters 392, 393. The filters 392, 393 may for example be band pass filters appropriate for two different fundamental frequencies. In some applications, the unfiltered signal path 386 can be eliminated.

The filter circuit 400 of FIG. 23 includes a filter circuit 406 interposed between an input 402 and an output 404. In this case the filter circuit 406 includes component filters 408, 410, and a switch 412 is used to connect the signal path containing the desired one of the component filters 408, 410 to the filter circuit 406. In this way, the filter characteristics of the filter circuit 406 can be adjusted as appropriate for the particular application by controlling the switch 412 to select the appropriate one of the filters 408, 406. As used herein, the term "filter" is intended broadly to encompass both filter circuits such as low pass, band pass, or band blocking filter circuits as well filter components such as capacitors, inductors and other electrical components that affect the filter characteristics of a filter circuit.

Switched filters similar to any of those shown in FIGS. 18-23 may be used in the systems shown in FIGS. 2, 8, 9, 10, 11 and 11a. For example, switched filters (using any of the approaches described above) may be substituted for the unswitched filter 58 of FIG. 2, the filter 104 of FIG. 8, the filter 124 of FIG. 9, the filter 210 of FIG. 10, the filter 224 of FIG. 11, or the filters 232, 234 of FIG. 11a.

As an alternative or in addition to the filtering circuits discussed above, other circuitry may be used to tune each transducer element for transmit and receive processing. This other circuitry tunes the transducer element to pass frequencies in a desired band, such as fundamental frequencies for transmit or harmonic frequencies for receive processing. Other frequencies, such as frequencies outside the pass band, may be suppressed by the tuned transducer frequency response. Preferably, the frequency response of the transducer element is adjusted differently for transmit and receive processing.

Referring to FIG. 24, a tuning circuit is generally shown at 500. The circuit 500 includes a transducer element 502 connected in parallel to a switch 504 and an inductor 506. A transmit waveform input 508, an isolation circuit 510 and an inductor 512 connect in series with the transducer element 502. A diode protection clamp 514, a low voltage pre-amplifier 516 and a receive signal output 518 also connect to the transducer element 502.

Any of various transducers and associated transducer elements 502 may be used. In one embodiment, an Acuson C3 curved linear radiology transducer is used. This transducer has a 3.5 MHz center frequency without tuning (i.e. without the tuning circuitry shown in FIG. 24). The C3 transducer center frequency is appropriate for use in harmonic imaging, since lower transmit frequencies, such as 2-3 MHz, may be more effective for harmonic processing. Furthermore, the impedance characteristics of the C3 transducer may be more easily adjusted by the tuning circuit 500 than other transducers. Other transducers may be used.

The inductors 506 and 512 are used to tune the transducer. The inductors are of any impedance and may be determined experimentally. In one embodiment, the inductor 512 is 7.5 .mu.H, and the inductor 506 is 1.5 .mu.H.

The transmit waveform is input and passes through the inductor 512. For transmission, the switch 504 is open. In one embodiment, the switch 504 comprises an analog switch with digital control, such as the Supertex HV2 series switch (Supertex Inc., Sunnyvale Calif.). For reception, the switch 504 is closed. The receive signal is amplified by the amplifier 516 and is prevented from passing to the transmit beamformer by the isolation circuit 510.

Referring to FIG. 26, the frequency response associated with the C3 transducer without tuning is shown by a dashed line 520. By passing the transmit waveform through the inductor 512, the frequency response during transmit is tuned as shown by a solid line 522. By switching the inductor 506 into a parallel connection with the transducer element 502, the frequency response during reception is tuned as shown by a dash-dot line 524.

The transmit tuning passes lower frequencies, such as associated with transmission at a 2.0 MHz fundamental center frequency, and filters the transmitted waveform at the harmonic frequencies, such as at 4.0 MHz, as compared to the un-tuned frequency response (see dashed line 520). The reception tuning passes higher frequencies, such as associated with the 4.0 MHz second harmonic center frequency, and filters the received waveform at fundamental frequencies, such as 2.0 MHz, as compared to the un-tuned frequency response (see dashed line 520).

To change the relative sensitivity of the transducer at different frequencies, the value of the inductors 506 and 512 are changed to have greater or lesser inductance.

Referring to FIG. 25, an alternative embodiment of a tuning circuit is generally shown at 600. The circuit 600 includes a transducer element 602 connected in parallel through a cable 620 and a switch 604 to an inductor 606. A transmit waveform input 608 and an inductor 612 connect in series with the transducer element 602 through the switch 604 and another switch 622. A receive signal output 618 connects to the inductor 606 through the switch 622. The receive signal output 618 provides signals to additional circuitry, such as a diode clamp and pre-amplifier.

In operation, the transducer is tuned for transmission by connecting the inductor 612 in series with the transducer element 602 through switches 604 and 622. The transducer is tuned for reception by connecting the inductor 606 in parallel with the transducer element 602 with switches 604 and 622.

In alternative embodiments, tuning is provided in only transmit or only receive. For example, if only transmit tuning is provided, the switch 504 and the inductor 506 are removed. The transducer may be selected for a desired receive frequency response and tuned for a desired transmit response. Other tuning components, such as resistors, different inductors, capacitors, transformers or combinations thereof may be used to tune the transducer through parallel, series or combined network connections.

In the systems described above, bi-polar rectangular or uni-polar transmit waveforms may also be shaped. The amplitude of each transmit waveform is shaped to gradually rise to a maximum value and gradually decrease from the maximum value. Each transmit waveform is shaped by modulating a carrier waveform with an envelope waveform. Alternatively and with respect to transmitting a uni-polar waveform, a shaped low pass off-set waveform with a gradually increasing and decreasing amplitude is summed with a bi-polar waveform in real-time or off-line. Shaping the amplitude of bi-polar rectangular or uni-polar transmit waveforms to reduce energies associated with harmonic frequencies is discussed in U.S. application Ser. No. 08/893,150 (now U.S. Pat. No. 5,913,813)(Attorney Docket No. 5050/220) for Ultrasonic Imaging Method And System For Transmit Signal Generation For An Ultrasonic Imaging System Capable Of Harmonic Imaging, assigned to the assignee of the present invention and filed concurrently herewith, the disclosure of which is hereby incorporated by reference.

As yet another alternative embodiment, the transmit waveforms discussed above may be pulse width modulated. As disclosed in U.S. application Ser. No. 08/893,287 (now U.S. Pat. No. 5,933,614) (Attorney Docket No. 5050/218) for Ultrasonic Imaging Method and Apparatus For Generating Pulse Width Modulated Waveforms With Reduced Harmonic Response, assigned to the assignee of the present invention and filed concurrently herewith, the disclosure of which is hereby incorporated by reference, the duration of each pulse within a burst is selected to reduce the energy transmitted at harmonic frequencies. In particular, the duration of one or more pulses is different than other pulses within the burst. Preferably, the widths of the pulses within the burst gradually increase and then decrease, but other duration patterns may be used.

As yet another alternative embodiment, the transmit waveforms discussed above may be shaped as a function of summation of the waveforms in the acoustic domain. As disclosed in U.S. application Ser. No. 08/893,271 abandoned (Attorney Docket No. 5050/219) for Ultrasonic Imaging Method And System For Harmonic Imaging Pulse Shaping In The Acoustic Domain, assigned to the assignee of the present invention and filed concurrently herewith, the disclosure of which is hereby incorporated by reference, the transmit waveform associated with a first transducer element or elements is shaped relative to a second waveform associated with a second transducer element or elements. For example, the first waveform (1) is delayed by a fraction of a cycle or one or more cycles, (2) is adjusted in amplitude, (3) is transmitted for a different number of cycles or (4) any combination of two or all three of (1), (2), and (3) relative to the second waveform. The first and second waveforms are focused at a point and transmitted. The transmitted waveforms sum in the acoustic domain at the point to form the desired waveform for reduction of energies transmitted in the harmonic frequencies. Preferably, the desired waveform corresponds to an amplitude that rises gradually to a maximum value and decreases gradually from the maximum value.

As an example of generating transmitted waveforms as a function of the resulting summed signal in the acoustic domain, the first and second waveforms are rectangular waves. The first waveform is delayed, in addition to any focusing delay, by may be 1/2 of a cycle relative to the second waveform. At the point in the body, the first and second waveform sum together to form a third waveform. The third waveform has three amplitude levels (0,1,2). The greatest amplitude is associated with an overlap of the first and second transmit waveforms and is preferably at the center of the third waveform. The number of cycles and amplitude shape of the first and second waveforms may also be controlled to create the desired third waveform in the acoustic domain.

Any of the various alternatives discussed above, such as pulse width modulation, filtering, generation of waveforms with multiple amplitudes and summation of waveforms in the acoustic domain, may be used in combination. The combination may include more than two of the alternatives discussed above.

All of the harmonic imaging techniques described above can be used in both tissue and contrast agent harmonic imaging modes. In the tissue harmonic imaging mode, no additional non-linear contrast agent is added to the target, and only the non-linear characteristics of the tissue are relied on to create the ultrasonic image. Medical ultrasound imaging is typically conducted in a discrete imaging session for a given subject at a given time. For example, an imaging session can be limited to an ultrasound patient examination of a given tissue of interest over a period of 1/4 to 1 hour, though other durations are possible. In this case no additional non-linear contrast agent is introduced into the tissue at any time during the imaging session. In the contrast agent harmonic imaging mode, any one of a number of well known non-linear contrast agents such as those described above can be added to the target in order to enhance the non-linear harmonic response of the tissue. For this reason, it should be understood that the introduction of an added non-linear contrast agent into the tissue being imaged is not implied in any of the following claims unless such added non-linear contrast agent is expressly recited.

From the foregoing, it should be apparent that improved systems and methods for imaging contrast agent have been disclosed. These systems can use the transmit beamformers described in U.S. Pat. No. 5,608,690, which focuses different frequency components at different ranges. Also, as described above, these techniques can be used with other beamformers which utilize other transmit waveforms. Of course, the various aspects of this invention relating to harmonic suppression and line focus (or other spreading of the region of maximum intensity) can be used separately from one another, rather than in combination as described above. It is intended that the foregoing detailed description be regarded as illustrative rather than limiting. It is the following claims, including all equivalents, which are intended to define the scope of this invention.

APPENDIX 1__________________________________________________________________________Parameters for Transmit Beamformer(all references are to FIG. 3 of U.S. Pat. No. 5,675,554)Parameter Corresponds to Value__________________________________________________________________________h2 FIR filter h2 of T312 (14641)h3 FIR filter h3 of T324 (10-201)Ku2 Upsampler of T326 2Ku1 Upsampler of T312 2 (determined by: KUI = Nb*4/Ku2)h4 FIR filter h4 of T326 (232)Nb Number of transmit beams 1cw.sub.-- on 0 -> pulsed modeEnvelope Type, `Envelope` of FIG. 4 0 -> realbeam 0Ns B-mode: 17 F-mode: 23h1 Envelope in T304 B-mode: (0.0039, 0.0156, 0.0430, 0.1133, 0.2461, 0.4531, 0.7031, 0.9141, 0.9961, 0.9141, 0.7031, 0.4531, 0.2461, 0.1133, 0.0430, 0.0156, 0.0039) (33% Gaussian, sampled at Fe = 10 MS/s) F-mode: (0.0039, 0.0078, 0.0195, 0.0430, 0.0898, 0.1719, 0.2930, 0.4531, 0.6406, 0.8203, 0.9492, 0.9961, 0.9492, 0.8203, 0.6406, 0.4531, 0.2930, 0.1719, 0.0898, 0.0430, 0.0195, 0.0078, 0.0039) (25% Gaussian, sampled at Fe = 10 MS/s)phi Phase to be applied at T310 See definition of phi belowFs Sampling Freq at O/P of T328 40 MS/sFe Sampling Freq of envelope in T304 10 MS/s (determined from above based on: Fe = Fs/Ku1/Ku2Fpa select v.sub.-- phi = Fc/F0 0 -> use modulation frequeency for focusingFm/F0 v = Fc/F0 0.5 (Fm = 2.5 MHz; F0 = 5.0 MHz)FO 5.0 MHzTogether these terms, plus the fact that the envelope is real, are usedto calcuiate the phaseapplied in complex multiplier T310, as described in Application Ser. No.08/432,056. Inparticular, this phase is broken into:envelope phase: zero because envelope is realfine focusing: calculated from the difference between quantized and idealdelay, using thephase alignment frequency vphi = vmodulation: complex multiplier T310 includes a component whichcorresponds tomodulation to a frequency (Fm/F0 - 1)*F0. This modulation, in combinationwith latermodulation in complex multiplier T318, results in an overall effectivemodulation frequency Fm.phi = phi.sub.-- D + phi.sub.-- E = phi.sub.-- Rphi.sub.-- D (Phase portion of delay (fine focusing)) = -2.(pi)v.sub.-- phi.tau.sub .-- phi tau.sub.-- phi = low order portion of the delay word representing fractional units of T0 (1/F0) as in Pat. No. 5,675,554. This is the portion of the specified focusing delay which is applied via phasing rather than true time delay.phi.sub.-- E = 0 (Waveform sample phase is zero because envelope isreal)phi.sub.-- R = 2.(pi).kul.(v - 1).n/4 (n is the successive samplenumber)This is a phase rotation of 2.(pi).(Fm - F0).t where t = kul.t/(4.F0)__________________________________________________________________________ APPENDIX 2__________________________________________________________________________Parameters for Receive Beamformer(all references are to FIG. 3 of U.S. Pat. Application Ser. No.08/432,615)Parameter Corresponds to Value__________________________________________________________________________Fs Sampling rate Fs at ADC 40 MS/sNb Number of receive beams 2Kd1 Downsampler of R162 2h1 FIR filter hl of R160 (232) (selected based on Kd1)h1 bypass disabledh2 FIR filter h2 of R164 (10-1)h2 bypass disabledh3 FIR filter h3 of R167 (14j-8-10j84j-1)h3 bypass disabledKd2 Downsampler of R169 4 (determined by: Kd2-Nb*4/Kd1)Fb 5 MS/s (determined from above, based on Fb = Fs/Kd1/Kd2)Fp select 0 -> Fp = FstartFp select 1 -> Fp = Fstart - Fdownslope*RFstart/FO Fstart B-mode: Fstart/FO = 1.1172 (5.5859 MHz) F-mode: Fstart/F0 = 1.0781 (5.3906 MHz)Fdownslope (DELTA F sub downshift) B-mode: Fdownslope = 1.793E-4 FO/TO <-> 5821 Hz/mm F-mode: Fdownslope = 1.144E-4 FO/TO <-> 3716 Hz/mmTbreak Tbreak >= 1792 TO (276 mm)Fupslope 0Base Band Filter= = =Ntaps 16type realL/M ratio 1/1 (based on L = 1, M = 1)hbbf Base band filter coefficients (based on L = 1, M = 1) B-mode: hbbf = (0, 0, 0, 0, 0, 0.0195, 0.0996, 0.2285, 0.2988, 0.2285, 0.0996, 0.0195, 0, 0, 0, 0) F-mode: hbbf = (0, 0, 0, 0.0020, 0.0137, 0.0469, 0.1172, 0.1992, 0.2383, 0.1992, 0.1172, 0.0469, 0.0137, 0.0020, 0, 0) Baseband filter is an FIR operating on the output of the receive beamformer__________________________________________________________________________ APPENDIX 3__________________________________________________________________________Revised Parameters for Transmit BeamformerAll parameters are as defined in Appendix 1 and are setequal to the values of Appendix 1 except as noted.H3 = [10 - 1]Ns = 17 (B-mode) 17 (F-mode)B-mode Envelope = [.01171875 .03125 .078125 .16796875 .32033125 .52734375 .75 .9296875 .99609375 .9296875 .75 .52734375 .3203125 .16796875 .078125 .03125 .01171875] (29% Gaussian, sampled at Fe = 10 MS/s)F-mode Envelope = [.015625 .0390625 .08984375 .19140625 .34375 .546875 .765625 .93359375 .99609375 .93359375 .765625 .546875 .34375 .19140625 .08984375 .0390625 .015625] (30% Gaussian, sampled at Fe = 10 MS/s)__________________________________________________________________________ APPENDIX 4______________________________________Second Revised Parameters for Transmit Beamformer(Adapted for use with Acuson 3V2C Probe)All parameters are as defined in Appendix 1 and are setequal to the values of Appendix 1 except as noted.H3 = [10 - 1]Ku2 = 4Ku1 = 1Ns = 27Envelope = [0.015625 0.03125 0.0546875 0.08984375 0.140625 0.21484375 0.30859375 0.41796875 0.546875 0.6796875 0.80078125 0.90625 0.97265625 0.99609375 0.97265625 0.90625 0.80078125 0.6796875 0.546875 0.41796875 0.30859375 0.21484375 0.140625 0.08984375 0.0546875 0.03125 0.01562500] (33% Gaussian, sampled at Fe = 14 MS/s)Fs = 56 MS/sFe = 14 MS/sF0 = 3.5 MHz______________________________________

Claims

1. A method for ultrasonically receiving data for imaging a target with distributed energy transmissions, said method comprising the following steps:

(a) transmitting ultrasonic energy centered at a fundamental frequency and focused at a first depth along a scan line;

(b) transmitting substantially simultaneously with step (a) ultrasonic energy centered at the fundamental frequency and focused at a second depth along the scan line; and

(c) receiving ultrasonic energy from the target at a harmonic of the fundamental frequency, the ultrasonic energy associated with at least the first and second depths.

2. The method of claim 1 wherein said target is free of contrast agent throughout an entire imaging session.

3. The method of claim 1 wherein said target comprises tissue and contrast agent.

4. The method of claim 1 wherein the transmission of step (a) is generated with a set of transducer elements and the transmission of step (b) is generated with the substantially same set of transducer elements.

5. The method of claim 1 further comprising (d) transmitting substantially simultaneously with steps (a) and (b) ultrasonic energy centered at the fundamental frequency and focused at a third depth along the scan line.

6. The method of claim 1 further comprising (d) tuning a transducer for steps (a) and (b); and wherein steps (a) and (b) comprise transmitting Gaussian waveforms.

7. The method of claim 1 further comprising (d) tuning a transducer for step (c).

8. The method of claim 7 further comprising (e) tuning the transducer for steps (a) and (b), wherein step (e) is operative to enhance a first frequency spectrum oriented near the fundamental frequency and step (d) is operative to enhance a second frequency spectrum oriented near the harmonic.

9. A method for ultrasonically receiving data for imaging a target with distributing energy transmissions, said method comprising the following steps:

(a) transmitting ultrasonic energy centered at a fundamental frequency and focused at a first depth along a scan line;

(b) receiving ultrasonic energy associated with at least the first depth in the target at a harmonic of the fundamental frequency;

(c) transmitting ultrasonic energy centered at the fundamental frequency and focused at a second depth along the scan line;

(d) receiving ultrasonic energy associated with at least the second depth in the target at a harmonic of the fundamental frequency; and

(e) performing steps (c) and (d) after steps (a) and (b);

wherein the target is free of contrast agent throughout an entire imaging session.

10. The method of claim 9 wherein the transmission of step (a) is generated with a set of transducer elements and the transmission of step (c) is generated with the substantially same set of transducer elements.

11. The method of claim 9 further comprising:

(f) transmitting ultrasonic energy centered at the fundamental frequency and focused at a third depth along the scan line; and

(g) receiving ultrasonic energy associated with at least the third depth in the target at a harmonic of the fundamental frequency.

12. The method of claim 9 further comprising (f) repeating steps (a)-(e) for a plurality of scan lines.

13. The method of claim 9 further comprising (f) tuning a transducer for steps (a) and (c); and wherein steps (a) and (c) comprise transmitting Gaussian waveforms.

14. The method of claim 9 further comprising (f) tuning a transducer for steps (b) and (d).

15. The method of claim 14 further comprising (g) tuning the transducer for steps (a) and (c), wherein step (g) is operative to enhance first frequency spectrum oriented near the fundamental frequency and step (f) is operative to enhance a second frequency spectrum oriented near the harmonic.

16. A method for ultrasonically generating data for imaging as a function of a frame of data representing spatial areas in a target, said method comprising the following steps:

(a) transmitting a first ultrasonic energy beam centered at a fundamental frequency and focused at a first depth along a scan line;

(b) receiving ultrasonic energy associated with at least the first depth in the target at a harmonic of the fundamental frequency;

(c) transmitting a second ultrasonic energy beam centered at the fundamental frequency and focused at a second depth along the scan line;

(d) receiving ultrasonic energy associated with at least the second depth in the target at a harmonic of the fundamental frequency;

(e) performing steps (c) and (d) after steps (a) and (b);

(f) repeating steps (a)-(e) for a plurality of scan lines; and

(g) generating the frame from data associated with the plurality of scan lines.

17. The method of claim 16 wherein said target is free of contrast agent throughout an entire imaging session.

18. The method of claim 16 wherein said target comprises tissue and contrast agent.

19. The method of claim 16 wherein the transmission of step (a) is generated with a set of transducer elements and the transmission of step (c) is generated with the substantially same set of transducer elements.

20. The method of claim 16 further comprising (h) transmitting ultrasonic energy centered at the fundamental frequency and focused at a third depth along each scan line.

21. The method of claim 16 further comprising (h) tuning a transducer for steps (b) and (d); and wherein steps (a) and (c) comprise transmitting Gaussian waveforms.

22. The method of claim 21 further comprising (i) tuning the transducer for steps (a) and (c), wherein step (i) is operative to enhance a first frequency spectrum oriented near the fundamental frequency and step (h) is operative to enhance a second frequency spectrum oriented near the harmonic.

23. A method for ultrasonically receiving data for imaging a target with distributed energy transmissions, said method comprising the following steps:

(a) transmitting ultrasonic energy centered at a first fundamental frequency in a first transmit beam focused at a first region along a first scan line;

(b) transmitting substantially simultaneously with step (a) ultrasonic energy centered at a second fundamental frequency in a second transmit beam focused at a second region along a second scan line, the second scan line different than the first scan line; and

(c) receiving ultrasonic energy along at least first and second receive lines from the target at harmonics of the first and second fundamental frequencies, respectively, the first and second receive lines associated with at least the first and second regions, respectively.

24. The method of claim 23 wherein said target is free of contrast agent throughout an entire imaging session.

25. The method of claim 23 wherein said target comprises tissue and contrast agent.

26. The method of claim 23 further comprising (d) tuning a transducer for steps (a) and (b); and

wherein steps (a) and (b) comprise transmitting Gaussian waveforms.

27. The method of claim 23 wherein step (c) comprises forming at least two receive lines associated with each of the first and second scan lines.

28. The method of claim 23 wherein the second fundamental frequency comprises a same frequency as the first fundamental frequency.

29. A method for ultrasonically receiving data for imaging a target with distributed energy transmissions, said method comprising the following steps:

(a) transmitting first ultrasonic energy centered at a first fundamental frequency in a first transmit beam focused at a first depth;

(b) transmitting second ultrasonic energy centered at a second fundamental frequency in a second transmit beam having a line focus; and

(c) receiving ultrasonic energy from the target at harmonics of the first and second fundamental frequencies, the ultrasonic energy associated with at least the first depth and with depths associated with the line focus.

30. The method of claim 29 wherein:

step (a) comprises transmitting the first ultrasonic energy focused at the first depth along a scan line; and

step (b) comprises transmitting the second ultrasonic energy having the line focus along the scan line.

31. The method of claim 30 wherein:

the line focus comprises an outer focal limit and a near focal limit; and

the first depth is at a different depth than depths along the line focus.

32. The method of claim 31 wherein the first depth is nearer a transducer than the near focal limit.

33. The method of claim 29 wherein the first and second ultrasound energies are transmitted substantially simultaneously.

34. The method of claim 29 wherein the first and second ultrasound energies are transmitted sequentially.

35. The method of claim 29 further comprising step (d) of tuning a transducer; and

wherein steps (a) and (b) comprise transmitting Gaussian waveforms.

36. The method of claim 29 wherein the target is free of contrast agent throughout an entire imaging session.

37. The method of claim 29 wherein the second fundamental frequency comprises a same frequency as the first fundamental frequency.

38. A method for ultrasonically receiving data for imaging a target with distributed energy transmissions, said method comprising the following steps:

(a) transmitting first ultrasonic energy centered at a fundamental frequency in a first transmit beam having a first line focus;

(b) transmitting second ultrasonic energy centered at the fundamental frequency in a second transmit beam having a second line focus; and

(c) receiving ultrasonic energy from the target at a harmonic of the fundamental frequency, the ultrasonic energy associated with at least the first and second line foci.

39. The method of claim 38 wherein:

step (a) comprises transmitting the first ultrasonic energy focused along a scan line; and

step (b) comprises transmitting the second ultrasonic energy along the scan line.

40. The method of claim 39 wherein:

the first and second line foci comprises first and second outer focal limits and first and second near focal limits, respectively; and

the first outer focal limit is at a depth closer to a transducer than the second near focal limit.

41. The method of claim 39 wherein:

the first and second line foci comprises first and second outer focal limits and first and second near focal limits, respectively; and

the first outer focal limit is at a depth between the second near and outer focal limits.

42. The method of claim 38 wherein the first and second ultrasound energies are transmitted substantially simultaneously.

43. The method of claim 38 wherein the first and second ultrasound energies are transmitted sequentially.

44. The method of claim 38 further comprising step (d) of tuning a transducer; and

wherein steps (a) and (b) comprise transmitting Gaussian waveforms.

45. The method of claim 38 wherein the target is free of contrast agent throughout an entire imaging session.

46. A system for ultrasonically receiving data for imaging a target with distributed energy transmissions, the system comprising:

means for transmitting ultrasonic energy centered at a first fundamental frequency in a first transmit beam focused at a first region along a first scan line;

means for transmitting substantially simultaneously with step (a) ultrasonic energy centered at a second fundamental frequency in a second transmit beam focused at a second region along a second scan line, the second scan line different than the first scan line; and

means for receiving ultrasonic energy along at least first and second receive lines from the target at harmonics of the first and second fundamental frequencies, respectively, the first and second receive lines associated with at least the first and second regions, respectively.

47. The system of claim 46 wherein the first fundamental frequency comprises a same frequency as the second fundamental frequency.

48. A system for ultrasonically receiving data for imaging a target with distributed energy transmissions, said system comprising:

means for transmitting first ultrasonic energy centered at a first fundamental frequency in a first transmit beam focused at a first depth;

means for transmitting second ultrasonic energy centered at a second fundamental frequency in a second transmit beam having a line focus; and

means for receiving ultrasonic energy from the target at harmonics of the first and second fundamental frequencies, the ultrasonic energy associated with at least the first depth and with depths associated with the line focus.

49. The system of claim 48 wherein the first fundamental frequency comprises a same frequency as the second fundamental frequency.

50. A system for ultrasonically receiving data for imaging a target with distributed energy transmissions, said system comprising:

means for transmitting first ultrasonic energy centered at a fundamental frequency in a first transmit beam having a first line focus;

means for transmitting second ultrasonic energy centered at the fundamental frequency in a second transmit beam having a second line focus; and

means for receiving ultrasonic energy from the target at a harmonic of the fundamental frequency, the ultrasonic energy associated with at least the first and second line foci.

51. A method for ultrasonically receiving data for imaging a target with distributed energy transmissions, said method comprising the following steps:

(a) transmitting ultrasonic energy centered at a fundamental frequency in a first transmit beam focused at a first region along a first scan line;

(b) transmitting substantially simultaneously with step (a) ultrasonic energy centered at the fundamental frequency in a second transmit beam focused at a second region along a second scan line, the second scan line different than the first scan line; and

(c) receiving ultrasonic energy along at least first and second receive lines from the target, the ultrasonic energy centered at a harmonic of the fundamental frequency, the first and second receive lines associated with at least the first and second regions, respectively.

52. The method of claim 51 wherein said target is free of contrast agent throughout an entire imaging session.

53. The method of claim 51 wherein said target comprises tissue and contrast agent.

54. The method of claim 51 further comprising (d) tuning a transducer for steps (a) and (b); and

wherein steps (a) and (b) comprise transmitting Gaussian waveforms.

55. The method of claim 51 wherein step (c) comprises forming at least two receive lines associated with each of the first and second scan lines.

56. The method of claim 51 wherein the first and second regions correspond to a same depth.