ULTRASOUND 3D IMAGING SYSTEM
The present invention relates to an ultrasound imaging system in which the scan head includes a beamformer circuit that performs far field subarray beamforming or includes a sparse array selecting circuit that actuates selected elements. When using a hierarchical two-stage or three-stage beamforming system, three dimensional ultrasound images can be generated in real-time. The invention further relates to flexible printed circuit boards in the probe head. The invention furthermore relates to the use of coded or spread spectrum signalling in ultrasound imaging systems. Matched filters based on pulse compression using Golay code pairs improve the signal-to-noise ratio thus enabling third harmonic imaging with suppressed sidelobes. The system is suitable for 3D full volume cardiac imaging.
CROSS REFERENCE TO RELATED APPLICATION
This application is a continuation-in-part of U.S. application Ser. No. 12/570,856 filed Sep. 30, 2009, which is a continuation-in-part of International Application No. PCT/US09/56995 filed on Sep. 15, 2009, which is a continuation-in-part of U.S. application Ser. No. 12/286,555 filed on Sep. 30, 2008 and claims priority to U.S. Application No. 61/192,063 filed on Sep. 15, 2008 by Chiang et al., entitled: “Ultrasound 3D Imaging System.” This application also claims priority to U.S. application Ser. No. 11/474,098 filed Jun. 23, 2006 and of International Application No. PCT/US2007/014526 filed Jun. 22, 2007. The entire contents of the above applications are incorporated herein by reference.
BACKGROUND OF THE INVENTION
Medical ultrasound imaging has become an industry standard for many medical imaging applications. Techniques have been developed to provide three dimensional (3D) images of internal organs and processes using a two dimensional (2D) transducer array. These systems require thousands of beamforming channels. The power required to operate such systems has resulted in the use of an analog phase shift technique with a digital delay beamformer that results in a compromise of image quality.
There is a continuing need for further improvements in ultrasound imaging technologies enabling improved real-time three dimensional imaging capability. In addition, this improved capability should support continuous real-time display for a fourth dimensional 4D function.
SUMMARY OF THE INVENTION
The present invention relates to a system for ultrasound medical imaging that provides three dimensional (3D) imaging using a two dimensional (2D) array of transducer elements in a probe housing. Embodiments of the invention provide systems and methods for medical imaging having high resolution and numerous imaging modalities.
In a preferred embodiment, the probe housing contains a first beamforming circuit that transmits beamformed data to a second housing having a second beamforming circuit. The first beamforming circuit provides a far-field subarray beamforming operation. The resulting beamformed data is transmitted from the scan head to a second housing having the second beamforming circuit that provides near-field beamsteering and beamfocusing.
A preferred embodiment provides a scan head that can be connected to a conventional ultrasound system in which the scan head provides the inputs to the conventional beamforming processing function. The scan head beamformer can utilize a low power charge domain processor having at least 32 beamforming channels.
A preferred embodiment of the invention employs a sparse array where only a fraction of the transducer elements need to be activated. By selecting the four corner elements of the array to provide proper mean lobe bandwidth, minimizing average sidelobe energy and clutter, eliminating periodicity and maximizing peak to side lobe ratio, quality images are produced. To steer the beams across the volume or region of interest, different transducer elements must be actuated in proper sequence to maintain the peak to sidelobe ratio. The system processor can be programmed to provide the desired sequence for transducer actuation to direct the beam at different angles. Alternatively, a discrete controller can be used to control sparse array actuation. A preferred embodiment provides a scan head with integrated switching circuits for sequentially selecting sparse array actuation elements for sequential multiple beamforming. The scan head can be connected to a conventional ultrasound system in which the scan head provides the inputs to the conventional beamforming processing functions. In another embodiment, the transmit array elements and receive array elements can be operated independently with the transmit elements comprising a sparse array and the receive elements being a near fully populated array. In a preferred embodiment, the multiplexer and beamformer circuits can be integrated into an interface system, or alternatively, into a host processing system, leaving a 2D transducer array mounted in the probe housing.
The present invention utilizes nondestructive sensing at each stage of the delay elements in the beamformer. So with a 65 stage delay line, for example, there are 64 usable outputs with one at each stage. The time resolution can be in the range of ⅛λ to 1/16λ.
Using high voltage multiplexers in the probe and the nondestructive sensing allows for time multiplexed sequential beamforming. It is now possible to sequentially change tap selection of each delay line to form multiple beams.
In addition to the three dimensional (3D) display capability, a fourth dimension or time resolved image display can be used to record and display a sequence of images recorded at 10 frames per second or higher, for example. This enables viewing of rapidly changing features such as blood or fluid flow; heart wall movement etc. at video frames rates of 30 frames per second.
Another preferred embodiment of the invention utilizes a three stage beamformer system in which a first stage performs a first beamforming operation on data received from a transducer array, which generates first beamformed data that is followed by a second stage that performs a second beamforming operation to provide second stage beamformed data that is then delivered to a third beamforming stage that performs a third beamforming operation.
The stages can be performed using charge domain processors. Data can also be converted from analog to digital form before the first stage, or the second stage, at the third stage or thereafter. One stage can utilize parallel beamforming operations and a second stage provides serial beamforming.
A preferred embodiment of the invention performs real time imaging of large volumes such as the human heart without having to take gated images of different portions of the heart in sequence and then stitch the images together. This can be done using beamforming architectures in which multiple beams can be transmitted in a single pulse. This provides for the collection of adult heart images within a single cardiac cycle or heartbeat. This can be done with a narrowband phase shifting beamforming system and/or with a time domain beamforming system. By using parallel and serial beamforming components distributed in the transducer probe housing and the system main processor housing lightweight portable and cart mounted systems can be used for real time full volume cardiac imaging.
In medical ultrasound imaging, there is a need for harmonic imaging where the transmitted waveform is of one fundamental frequency Fo, and the received signal of interest is a higher harmonic, generally the 2nd harmonic (2 Fo), or the third harmonic (3 Fo). The harmonic signal of interest is generated by the image targets in the body, and harmonics in the transmitted waveform are not of interest. Therefore it is important to suppress harmonic components from the transmitted waveform to obtain a clearer response.
The transmit circuit can generate either square waves or sinusoidal excitations. Square wave pulse generators are generally less costly than sinusoidal pulse generators, and the square wave pulsers are widely used in ultrasound imaging equipment. However a typical square wave itself contains high levels of the third harmonic. The present invention uses a modified square waveform which produces significantly reduced amount of the third harmonic, thus providing substantially improved harmonic imaging using the less expensive wave pulsers.
BRIEF DESCRIPTION OF THE DRAWINGS
FIGS. 3F(1)-3F(4) illustrate preferred embodiments of multiplexers used for switching in ultrasound transducer systems.
FIGS. 4A-4D illustrate a gated acquisition sequence for cardiac imaging.
DETAILED DESCRIPTION OF THE INVENTION
The objective of the beamforming system is to focus signals received from an image point onto a transducer array. By inserting proper delays in a beamformer to wavefronts that are propagating in a particular direction, signals arriving from the direction of interest are added coherently, while those from other directions do not add coherently or cancel. For real-time three-dimensional applications, separate electronic circuitry is necessary for each transducer element. Using conventional implementations, the resulting electronics rapidly become both bulky and costly as the number of elements increases. Traditionally, the cost, size, complexity and power requirements of a high-resolution beamformer have been avoided by “work-around” system approaches. For real-time three-dimensional high-resolution ultrasound imaging applications, an electronically steerable two-dimensional beamforming processor based on a delay-and-sum computing algorithm is chosen.
The concept of an electronically-adjustable acoustic conformal lens is to divide the surface of a 2D transducer array into plane “tiles” of relatively small subarrays. As described in U.S. Pat. No. 6,292,433 the entire contents of which incorporated herein by reference, and illustrated in
To scan an image plane using a steerable beamformer system a process such as that shown in
A detailed diagram of an electronically-controlled beamforming system in accordance with the invention is shown in
Use of the scan head with a second stage digital beamformer is shown in
Without using this hierarchical subarray far-field and then near-field beamforming approach, for an 80×80 element 2D array, a cable consisting of six thousand and four hundred wires is needed to connect the transducer array to a conventional beamforming system. As shown in
The beamformer processing system is a time domain processor that can simultaneously process the returns of a large 2D array providing a low-power, highly integrated beamformer system capable of real time processing of the entire array in a portable system. While a system with 192 parallel received channels supports a matrix 2D array probe for a real-time 3D/4D imaging application, the hierarchical multi-stage beamforming can be used with a low-power compact ultrasound system.
In a preferred embodiment, the hierarchical beamforming can also be applied to a one dimensional (1D) array for the real-time 2D imaging application. For example, for a 128-element 1D array, a group of 8 adjacent elements can be grouped together as a subarray. Within each subarray, each of the 8 elements has its own short programmable delay and then the outputs of the eight delays are summed together and then applied to a common long delay. It is important to note that two different methods that can be used for this two-stage implementation. In the first implementation, all the beamforming circuits including both the short and long delays are placed in the back-end processor, so for a 128 element 1D array, 128 connection cables are used as I/O cables between the transducer probe and the backend processor. An alternative implementation is to integrate all the subarray processors within the transducer probe, i.e., for a 128-element array, all 16 subarray processors each with 8 programmable delays are integrated within the transducer probe, so only 16 cable elements are needed to connect the front-end integrated probe with the back-end processors. Within the back-end, only 16 long delay beamforming circuits are needed to complete the beamforming function. Similarly, for a 64-element array with integrated eight 8-element subarray processors in the probe, the back-end processor can be simplified to only 8 beamforming circuits, only 8 cable elements are needed to connect the front-end integrated probe with the back-end processor. Furthermore, low-power transmit circuitry and A/D converters can be integrated into the front-end probe, so a wireless communication link can be used to connect the front-end probe and the back-end processor. A wireless USB connection or a wireless FireWire connection can be used.
The construction of a 64×48 element 2D transducer probe 485 with integrated 4×4 sub-array processors is illustrated in
As can be seen in
In this chip, there are 16 tapped delay lines, each receiving returns from its corresponding receive element. During the receive mode after a transmission pulse, 4 sequential beams which are summed outputs from the 16 tapped delay lines are formed at every sampling clock. The tap output of each delay line is controlled by a 4-beam time-multiplex buffer memory. With each new digital update, a corresponding non-destructively sensed delayed sample is clocked out of the tapped delay line. With the four digital updates sequentially applied to the buffer memory, four delayed samples for each of the four beams are then sequentially clocked out.
The initial tap positions of each delay line are pre-loaded in memory before the scanning starts. During receive mode, at every sampling clock, returned echoes are sampled and clocked into its corresponding delay line. A multiplier is incorporated at the output of each tapped delay line to provide the beamshaping, apodization function. For example, if the center frequency of the transducer is 2 Mhz, the tapped delay line samples the returned echo at a 8 Mhz rate. The tap outputs are sequentially non-destructively sensed at a 32 Mhz rate to generate the 4 beams. That is, after a returned echo loaded into a delay line, 32 ns later the tap output of the delay sample of this delay line for the 1st beam is clocked out and applied to the multiplier, another 32 ns later, the delayed tap output for the 2nd beam is clocked out to the multiplier, the procedure follows, until 128 ns later the tap out for the 4th beam is clocked out. The 16 multiplier outputs are summed together to form a single beam at 32 Mhz rate. It is important to note that for dynamic focusing, each beam needs two digital update bits; one for tap update and one for interpolation. In this chip, each channel has an analog input and a digital input; the two update bits are sequentially loaded into the chip. To support the 4-beam sequential outputs, the two-digital update bits of each beam are dynamically loaded into the chip at a 64 Mhz rate, thereby allowing continuous subarray beamforming function at a 8 Mhz analog input sampling rate. If the range depth is 15 cm, for a 2 Mhz probe oversampled by 4, the total received beamforming includes 2000 points. In this embodiment, the memory size on the circuit board shown in
Typical ultrasound transducers use the same element for transmit and receive. The high voltage transmit pulse is sent to a particular element, and the echo from the same element travels back to the system via the same cable wire.
In some applications, it is desirable, or necessary to use separate elements for transmit and receive. One such application is the use of different transducer materials for transmit and receive, so that the transmitter and receiver elements can be made with different frequency responses that is a first frequency response and a second frequency responses different from the first frequency response. This is particularly useful for harmonic imaging where the receiver center frequency is double or triple that of the transmitter center frequency. The transmit multiplexer (TR_MUX) integrated circuit chip allows one single cable wire to connect to the transmit element and the receive element by providing a fast high voltage switch that connects the cable wire to the transmit element during the transmit period, then to the receive element in the receive period as shown in FIG. 3F(1). Off-the-shelf high voltage multiplexer chips are not suitable for this application, as their intended use is to multiplex among elements for aperture selection, a slow process that happens at ultrasound scan line boundaries. In order to support switching between transmit period and receive period, the switching time needs to be on the order of less than a wave period to a few of wave periods, otherwise there will be a large dead time near the surface of the transducer that cannot be imaged. For example, a switch turn on/off time of 1 micro-second can switch from transmit to receive in one wave period at a1 MHz transmit frequency, or two wave periods when the transmit frequency goes up to 2 MHz. It can be advantageous to amplify the received signals as shown in FIG. 3F(2).
Another application that requires a fast in-the-probe transmit/receive switch is a 2D array probe where the receive elements are first formed into sub-arrays to reduce the number of cable wires for receiving as shown in FIG. 3F(3). The sub-array beamforming circuit is usually a low voltage device, so it is necessary to isolate the sub-array from the cable wire during the high voltage transmit period.
Yet another application uses two levels of the TR_MUX chips to allow low voltage amplification circuits be used on a shared transmit/receive element as shown in FIG. 3F(4). The same element is used for transmit and receive. The receive signal is amplified prior to being sent to the system circuit via the shared cable wire. In this case, the low voltage receive amplification circuitry is detached from the cable and the element during each transmit period, when high voltage pulses are used. During each receive period, the element is attached to the low voltage amplification circuit and then to the cable wire.
The chip shown in
In operation as shown in
An implementation of 64 element 1D array with integrated first stage subarray processor can also be implemented using the design of
A preferred embodiment of a 64 element (or more, e.g. 128 or 256 elements) 1D array 496 with integrated subarray processor is shown in
Alternatively, the beamforming processor can be mounted on a printed circuit board has to provide the subarray beamforming function for a total of 64 receive elements which are divided into 8 subarrays with each subarray consisting of 8 adjacent elements. 8 subarray processors each can perform the beamforming function, i.e., time-delay-and-sum function, for 8 adjacent receive elements are incorporated on the circuit board. A photomicrograph of a 16 channel subarray beamformer chip shown in
As pointed before for dynamic focusing, each beam needs two digital update bits; one for tap update and one for interpolation. To support the 4-beam sequential outputs, the two-digital update bits of each beam are dynamically loaded into the chip at an eight times input-sampling rate, so to allow continuous subarray beamforming function at the analog input sampling rate. If the range depth is 15 cm, for a 2 Mhz probe oversampled by 4, the total received beamforming are 2000 points. In the embodiment, the memory size on the board shown in
Existing are medical ultrasound systems with matrix-array transducers can provide real-time 3D (RT-3D) echocardiography along with state-of-the-art 2D imaging. The major advantages of RT-3D acquisition compared with 2D image include shorter acquisition times, reduced operator dependence, and the ability to manipulate images offline to extract any number of desired views for data analysis. Furthermore, quantitative data regarding Left Ventricle volumes and ejection fraction are more precisely obtained using the 3D technique. Although the term “real-time” is applied to all of the currently available 3D echocardiographic technology, it is important to recognize that in the current scanners, “live 3D” refers to true real-time images that are acquired without electrocardiographic gating. However, this type of real-time 3D imaging has a narrow sector with only a partial volume and is not suitable for imaging the left ventricle.
To obtain full-volume 3D images in current scanners, electrocardiography is used to gate the image acquisition. Four to 7 subvolumes are acquired over 4 to 7 cardiac cycles and then merged to obtain a complete data set, shown in
As indicated, about 128 by 96 beams are required to provide the complete coverage of the left ventricle. In a conventional implementation, during the first cardiac cycle with electrocardiographic gating, 32 by 96 beams are used to acquire part of the 3D image (
A preferred embodiment of the present invention generates 16 scanning beams for each transmit pulse, as a result, it generates a true “live 3D” image with 128×96 scanning beams operating at least at a six 3D volumetric images per second rate. The speed of sound in tissue is about 1500 cm/sec, the round-trip propagation time for a sound wave penetrating a 15-cm depth is about 200 microseconds. For 3D imaging, such as of the heart including both left and right ventricles, as shown in
Systems used to generate at least 16 beams for each transmit pulse are shown in the embodiments of
An ultrasound system using a probe such as that shown in
A preferred embodiment of the invention for 2D array beamforming, each minimizing noise and cable loss with improved S/N performance, are described in
By systematically varying beamformer delays and shading along a viewing angle of a 2D transducer array, returned echoes along the line of sight representing the 3D radiation sources can be used to create the scanned image at the scanned angle. The system can provide continuous real-time large area scanned images throughout a large field of view at 20 frames/s or more. At this frame rate, the system can be used to display continuous 3D images vs. time, thus providing 4D information of the scanned object. As shown in
For example, if a computing circuit with 16-serial addressable outputs is built in with the CDP subarray and the 2nd stage time delay lines, for each transmit pulse, 16 beams or scan lines each along a different scan angle can be created. For 256-pulses with a down-range depth of 15 cm, the system can generate a 4096-beams with a 64×64 pixel resolution at a frame rate of 20 frames/s. The system is fully programmable; the beamforming electronics can be adjusted to zoom-in to a smaller field-of-view for high-resolution or higher frame rate images. For example, using 192-transmit pulses with the same down-range depth of 15 cm, the system can generate a 3072-beams with a 64×48 pixel resolution at a 30 frame/s frame rate.
The array described addresses ultrasound imaging applications using a two-dimensional 2 cm×2 cm array at a frequency of 3 MHZ. The need for resolution on the order of less than half the wavelength dictates as large an aperture as possible that can be housed within a compact package. To interrogate a 90 degree scanning volume and also minimize the impact of grating lobes, an element pitch or separation of less than 0.25 mm is desirable, leading to a 80×80 element array. Using the subarray processing technique described above, a scan head with integrated subarray beamforming circuits followed by a 2nd stage near-field beamsteering/beamfocusing system provides a practical implementation. However, the implementation still requires at least 32 subarray chips to be integrated on a scan head. An alternative pseudo random array design approach can be used to achieve this resolution with a much less amount of processing components in the scanned head.
To make a sparse array practical, the combination of low insertion loss and wide bandwidth performance is important for realizing acceptable imaging performance with low illumination levels. Quarter-wave matching layers with low acoustic impedance, but physically solid backing results in a robust array that loses only 3-4 dB in the conversion of received signal energy to electrical energy. Array band-widths of 75% or more are typical of this design and construction process. Also, the transducer array employs element positioning and an interconnect system suitable for the beamformer circuitry. The electronics are mounted on printed-circuit boards that are attached to the transducer elements via flexible cables. In practice, a majority of the array elements are connected to outputs using the flexible cables. However, only a small fraction of the total number of elements are wired to the circuit boards. Nevertheless, the large number of array element connections are sufficient to insure a unique pattern of active-element locations in the final array.
As an example of a sparse array, assuming a 2×2 cm array with 256 active elements, the resulting filling factor is 4%. The output signal to noise ratio of the array is proportional to the number of active elements, so this filling factor corresponds to a loss in sensitivity of −13 dB when compared to a filled array of the same dimensions. To compensate for this loss, a transmitted signal of wider bandwidth is chosen to increase array sensitivity. In the approach presented here, the sensitivity is increased on the order of 10 dB. Further details regarding sparse array devices can be found in U.S. Pat. No. 6,721,235, the contents of which is incorporated herein by reference.
Positioning the elements of the array follows the approach in which care must be taken to eliminate any periodicity that would produce grating lobes that compete with the main lobe. Pseudorandom or random arrays can be used (
The primary goal of the optimization method is to minimize the average side lobe energy. Specifically, this is done by interactively evaluating the optimization criterion:
where the weighting function, W(ux, uy), applies more weight to regions in the array response that require side lobe reduction. The optimization method begins with no weighting (i.e., W(ux, uy)=1) and proceeds by choosing successively better weighting functions that satisfy the optimization criterion. Since the side lobes that require the greatest reduction are related to the previously computed beampattern, B(ux,uy), the weighting is chosen such that W(ux,uy)=B(ux,uy). This is done in an interactive manner until convergence.
Basically, a random array is capable of producing an imaging point spread function that has a main lobe to average side lobe ratio of N, where N is the total number of active elements in the array. For the 256-element sparse array example, the resulting ratio is −13 dB. Using a wide bandwidth approach improves this ratio by 10 dB. Based on the preceding optimization criterion, a pseudorandom placement of the array elements was generated (
A commercially available window-based 3D visualization software can be used to visualizing, manipulating, and analyzing the 3D multiple-beams volume image data generated by the electronically-adjustable acoustic conformal lens system. Traditionally, a clinician with 2D ultrasound images for diagnosis would look at the 2D scanned images slice by slice and mentally reconstruct the information into a 3D representation to judge the anatomy of the patient. This procedure requires the clinician to have well-founded experience as well as a highly sophisticated understanding of human anatomy. To create a “complete” image to the 3D structures, the clinician has to take all available slices into account. Looking at hundreds of slices is too time-consuming, even for a single patient. 3D visualization based on 3D volume data can help overcome this problem by providing the clinician with a 3D representation of the patient's anatomy reconstructed from the set of multiple-scanned beamforming data.
A commercially available software tool such as KB-Vol3D of KB-VIS technologies, Chennai, India, provides display or viewing 3D features such as:
- Fast Volume-Rendering
- Shaded Surface Display
Shaded-Surface module allows easy visualization of surfaces in the volume. Surfaces may be created by intensity-based thresholding. Alternatively, the Seeding option allows selection of specific connected structures of interest.
- MIP (Maximum Intensity Projection) with Radials
- MPR (Multiple-Plane-Reformating) with Oblique & Double-Oblique and 3D correlation
- MRP Slabs & Multi-Cuts
- Curved MPR
- Color & Opacity Presets with Editor
- Region-Growing and Volume Measurements
- Cutaway Viewing with Slab-Volume and Interactive Real-time VOI
Volume-interiors are easily visualized using the “Cutaway-Viewing” tool. A Cut-Plane is used to slice through the volume, revealing the interior regions. The cut-plane is easily positioned and oriented using the mouse.
The VOI (Volume-of-Interest) tool allows interactive, real-time Volume-of-Interest display. The user can isolate and view sub-volumes of interest very easily and in real-time, using easy click-and-drag mouse operation.
- Image Save in Multiple Formats
Images displayed by KB-Vol3D can be captured to various image formats (including DICOM, JPEG, and BMP etc.)
- Movie Capture in AVI Format
Visualization operations can also be captured to an AVI movie .le and played on Windows Media Player, QuickTime, and Real Player etc.
The invention can be implemented using a scan head 12 connected to a portable computer 14 as shown in
A 2D array configuration using sparse-array for transmission and non-overlapped fully-populated array is used for receiving. For an N×M element array, only m-elements with optimized sparse array placement are used for transmition and then the remaining NM-m elements are used as the receiving array. For example, for a 40×60-element 2D array, 256-elements are used as transmit element, the placement of the transmit elements are optimized based on selection criteria, the remaining 2144 element are used as received elements. This embodiment simplifies the multiplexer requirement needed for a 2D array, in which case the multiplexer can be mounted in the interface housing.
An example of the element locations for the near fully-populated 40 by 60 receive array 50 is shown in
The azimuth and elevation cross-sections of the beampattern of the above mentioned receive array are shown in
An example of the final element locations for the 256 transmit sparse array 60 are shown in
A low-power ultrasound system capable of electronically scanning a two-dimensional, 2D, matrix array to generate real-time three-dimensional, 3D, volumetric images with 64 by 64, 4096, scanning beams at a greater than 20 3D images per second is described. For each transmit pulse, the system is capable of generating 16 received beams. In addition, the design is able to drive a one and one-half dimensional array and also support wide-bandwidth encoded transmit waveform for pulse compressing to improve the system sensitivity. Wide bandwidth enables the use of chirped or coded waveforms (PN sequence) that can extend the length of the low power transmit burst without a loss of axial resolution. The combination of these features results in an imaging array with electronic systems that will fit within a portable hand-carried device.
The beamformer processing system is a time domain processor that will simultaneously process the returns of a large 2D array, the low-power; highly integrated beamformer that provide a real time processing of the entire array and will thus provide a low cost unit that can be hand carried.
There is a strong need for a real-time 3D ultrasound imaging using a 2D matrix array. In this section, the minimal number of receive beamforming channels required in an ultrasound system to support a real-time 3D imaging is analyzed. It is shown that a minimum of 192 parallel received beamforming channels is required to support a reasonable sized such as 48×64-element array.
An example of a system having an electronically-adjustable acoustic conformal lens is to divide the surface of a 2D transducer array into plane “tiles” of relatively small subarrays can be formed in U.S. Pat. No. 6,292,433, the contents of which is incorporated herein by reference; beamforming of the entire array can be separated into two stages, first a small-aperture subarray beamforming followed by a second stage large-aperture coherent summing of the outputs from each of the subarrays. As depicted in the tiles/subarrays can be made small enough so that when an object is placed within the field-of-view of the imaging system, the incident radiation from the object toward each “tile” can be treated using a far-field approximation. However, near-field beamforming capability has been incorporated in the actual implementation of the subarray beamforming system to allow a broader application. Additional delay elements are incorporated as second-stage processing to allow all subarrays to be coherently summed. The delay-and-sum beamformer allows each subarray to “look” for signals radiating from a particular direction. By adjusting the delays associated with each element of the array, the array's look direction can be electronically steered toward the source of radiation. The delay line requirement for each element in the sub-array can be less than a hundred stages. Only long delays for global summing are needed for the final near field focusing. A detailed diagram of an electronically-controlled beamforming system in accordance with the invention is shown in FIG. 14A of U.S. Pat. No. 6,292,433. This system consists of a bank of parallel time-delay beamforming processors. Each processor consists of two components: a 2D sub-array beamformer for small-aperture beamsteering/focusing and an additional time delay processor to allow hierarchical near-field beamforming of outputs from each corresponding subarray. As can be seen in
It is easy to understand that, without using this hierarchical subarray small aperture and then large aperture beamforming approach, for an 80×80 element 2D array, a cable consisting of six thousand and four hundred wires is needed to connect the transducer array to a conventional beamforming system. As shown in FIG. 14A of U.S. Pat. No. 6,292,433 referenced above, the number of inputs to each subarray processor equals the total number of delay elements in the subarray, each sub-array only has a single output. That is to say, the number of inputs to a subarray equals the number of transducer elements associated with that subarray. The number of subarray outputs equals the total transducer array element number divided by the number of subarrays. For example, if one selects to use a 5×5 subarray to implement this hierarchical beamforming system, after the first stage subarray beamforming, the total number of wires needed to connect to the 2nd stage near-field beamforming is reduced by a factor of 25. More specifically, as mentioned above, without using this 2D subarray beamforming, 6400 wires are needed to connect an 80×80 2D transducer array to a conventional back-end beamforming. Using a 5×5 subarray processing bank first, the number of wires required to connect to the backend beamforming system is reduced to 256. Based on this example of the invention, a bank of 256 5×5 element subarrays beamformer can be integrated with a 80×80 element 2D array in the scan head, so a cable consisting of 256 wires is adequate to connect the integrated scan head with the back-end near-field beamforming system.
It is important to note that 5×5 subarray small-aperture beamforming processors can be easily integrated in a small size silicon integrated circuit, eight of such 5×5 subarray beamforming can be integrated on one integrated circuit. Note that subarrays have generally between 9 and 64 transducer elements corresponding to a 3×3 subarray up to an 8×8 subarray. The preferred range is at or between 4×4 and a 6×6 array for a square array geometry. Rectangular subarrays can also be used preferably either 3×4, 4×5, or 4×6. Note that a ¼λ error minimum criteria is uses. Only 32 integrated circuit devices need be incorporated into the scanhead, it can reduce the cable size from 6,400 wires down to 256 wires. Similarly, for a 64×48 element 2D array, using a 4×4 subarray processing bank in the transducer housing first, the number of back-end beamforming channels is reduced to 192.
In the present invention, preferred embodiments for a 2D array beamforming, each minimizing noise and cable loss with improved signal to noise ratio performance, are described in
In a preferred embodiment of the system a large-aperture beamforming system is incorporated into the main processor housing of the ultrasound imaging system as shown in connection with
The speed of sound in tissue is about 1500 cm/sec so that the round-trip propagation time for a sound wave penetrating a 15-cm depth is about 20 microseconds. For a real-time 3D imaging, at least 64×64 scanning beams at a frame rate greater than 20 3D volumetric images per second are needed to provide diagnostic quality images. For each transmit beam, the real-time 3D imaging system has to be able to form at least 16 beams for each transmit pulse to support the preferred 3D frame rate requirement. In this section, both a serial time-multiplexed beamforming and a parallel simultaneous time-domain beamforming implementation are addressed.
To achieve a 16 beam scanning requirement, a combination of serial and parallel architecture can be used, i.e., the system can use front-end time-multiplexed serial beamforming elements technique to form two beams, then followed by 8 parallel beamforms at the back-end processor, or the system can form 4 serial beams, for each serial output beam, the back-end processor then forms 4 parallel beams, and so forth.
By systematically varying beamformer delays and shading along a viewing angle of a 2D transducer array, returned echoes along the line of sight representing the 3D radiation sources can be used to create the scanned image at the scanned angle. The system can provide continuous real-time large area scanned images throughout a large field of view at 20 frames/s or more. As shown in
For example, if a computing circuit with 16-serial addressable outputs is incorporated with the processor subarray and the 2nd stage time delay lines, for each transmit pulse, 16 beams or scan lines each along a different scan angle can be created. For 256-pulses with a down-range depth of 15 cm, the system can generate a 4096-beams with a 64×64 pixel resolution at a frame rate of 20 frames/s. The system is fully programmable; the beamforming electronics can be adjusted to zoom-in to a smaller field-of-view for high-resolution or higher frame rate images. For example, using 192-transmit pulses with the same down-range depth of 15 cm, the system can generate a 3072-beams with a 64×48 pixel resolution at a 30 frame/s frame rate.
The objective of a beamforming system is to focus signals received from an image point onto a transducer array. By inserting proper delays in a beamformer to align wavefronts that are propagating in a particular direction, signals arriving from the direction of interest are added coherently, while those from other directions do not add coherently or cancel. The time-of-flight from the radiation source to the focal point can be calculated and stored in memory for every channel from multiple directions of arrival in parallel. In a conventional implementation, separate electronic circuitry is necessary for each beam; for a multi-beam system, the resulting electronics rapidly become both bulky and costly as the number of beams increases. For example, beamforming for a linear 192 element array requires 192 parallel delay lines each with a programmable delay length of greater than 128λ. To form four parallel beams, for example, a total of 768 programmable long delay lines are required. To simplify the required electronics for multiple beams, a hierarchical two stage beamforming system is described.
The concept of hierarchical beamforming is to separate the time-of-flight calculation into two parts: the first part is a short delay for coarse-resolution, small aperture beamforming, followed by a long delay for fine resolution, large aperture beamforming. Shown in
To exemplify operation of a two-stage delays, a differential delay profile must be generated for all elements in the 1D or 2D array. To do this, the differential delay equation is calculated and all of the differential delays as a function of angles Theta and Phi, at a given range, are tabulated. For example, as shown in
In a 2 stage delay system, the tabulated data from the preceding step are broken into a coarse delay and a fine delay. To determine how to partition the coarse delay and the fine delay, the maximum differential delay error is constrained (typically set to have a maximum differential delay error less than or equal to 1 sample). The tabulated delays (from the preceding step) are also used to determine when a receive element is enabled. For example,
A block diagram of an hierarchical two-stage parallel beamforming system 958 is shown in
Furthermore, within each small-aperture, short-delay line, a time-of-flight control circuit is used to select the tap position output from a charge-domain processing circuit that non-destructively senses the tapped-delay line output. Each receiver has a multiplier for beam shading/apodization. Within each processor, all the multipliers share a common output. The summed charge is then applied to a matched filter to decode and to compress the returned echoes to produce an imaging pulse with a reduced signal-to-noise ration. An analog to digital (A/D) or a converter on-chip charge-domain A/D converter can be used so that hierarchical summing can be carried out digitally.
In a preferred embodiment, it is important to employ high speed digital communication connection between the beamformer output and the backend processor. As described previously, the analog returned echoes received by each transducer element is converted to a digital signal by an analog to digital converter (A/D) during signal processing. As shown in the beamformer 974 of
The use of coded or spread spectrum signaling has gained tremendous favor in the communications community. It is now routinely used in satellite, cellular, and wire-line digital communications systems. Shown in
In ultrasound, the concept of using spread spectrum/coded excitation transmit waveform comprises modulating a base sequence of transmit pulses of length P with a code sequence with a code length N. A code pulse sequence of N bursts is often referred to as an N-chip code. An example of a gated 3 MHz sinusoid with a 5-Chip Barker coding [111-11] is shown in
A preferred method of forming a transmission signal is shown in
A 192 channel receive beamforming system capable of forming four parallel, compressed beams for each transmitted, spreaded coded excited waveform is shown in the beamformer system 985 of
A matched filter implementation is shown in
gn=fn-1W1+fn-2W2+fn-3W3+ . . . +fn-K-2WK-2+fn-K-1Wk-1+fn-KWK
Using the example shown in
Consider an ultrasound pulser with 3 cycles of square wave, where an example is shown in
Thus a preferred embodiment of the present invention uses a modified square wave by reducing the pulse high time and pulse low time to two third of the regular square wave. This waveform has a much lower third harmonic component as the regular square wave. Shown in
In current ultrasound system, 2nd harmonic imaging mode has been widely accepted to image tissue and showed considerable improvements in image quality that reduces artifacts. As a result, it allows physicians to make better diagnoses in several applications compared to the fundamental excitation mode. The improvements were attributed to the effects of wave distortion due to nonlinear propagation in tissue. Current system have used the 2nd harmonic imaging. Since the energy in the second harmonic frequency band is much lower than that in the fundamental frequency band, to increase the 2nd harmonic sensitivity the spectral overlap between the fundamental and the 2nd harmonic has to be minimized, however, in doing so, the 2nd harmonic imaging resolution is reduced as a result. The higher harmonics, in particular the 3rd harmonic, not only represent an additional important information for tissue imaging and characterization, but also has the advantage that it is easier to filter out the contribution from the fundamental frequency. Unfortunately, with the current settings of existing systems (MI, frequency), the amount of third harmonic energy returning from tissue is much less than that reflected at the fundamental frequency. It requires the ultrasound system to have excellent sensitivity and dynamic range to display the 3rd harmonic image. Preferred embodiments of the present invention, use 3rd harmonic imaging with coded excitation where additional gain from matched-filter processing is used to overcome the problem with weak 3rd harmonic returns.
Two pertinent requirements for a good 3rd harmonic coded excitation imaging are, first by, minimum 3rd harmonic component in the transmit coded waveform, and secondly, a code selection with a 3rd harmonic receive template that has a minimal sidelobe energy after the 3rd harmonic matched filtering. A transmit coded waveform with minimum 3rd harmonic component. In this section, two 3rd harmonic templates for a Golay coded with minimum sidelobe are presented.
Golay complementary pairs (GCP)coded binarysequences are used as in this method, a complementary pairs of matched filter outputs are summed to generate impulse return with minimum sidelode
where a(z) and b(z) are complementary pairs. As a result, no image artifacts due to match-filter sidelobes. In the following, the 3rd harmonic template of a coded excitation transmit waveform is derived.
Let the Fourier transform of transmit waveform be
- the Fourier transform of 3rd harmonic excited by this transmit waveform should be
- assuming that the 3rd harmonic is caused by the 3rd power term of the Taylor series of tissue non-linearity. Then the 3rd harmonic waveform is the Inverse Fourier transform. If this is used as a match filter template, there are never significant side lobes since the power spectrum has been changed. For this reason, the following Fourier transform is used to generate match filter templates.
Based on the above derivation, the fundamental and 3rd harmonic template of a 10-bit Colay code pair is shown in
Another technique in selecting the 3rd harmonic template is by inserting two zeros after each code word of the fundamental template. The third harmonic matched filter outputs, i.e. the autocorrelation of this alternative 3rd harmonic template is shown in
The claims should not be read as limited to the recited order or elements unless stated to that effect. All embodiments that come within the scope and spirit of the following claims and equivalents thereto are claimed as the invention.
1. A medical ultrasound imaging system comprising:
- a two dimensional transducer array in a probe housing and a first plurality of subarray beamformers in the probe housing such that the transducer array generates a plurality of beams for cardiac imaging.
2. The system of claim 1 further comprising a plurality of multiplexor circuits in the probe housing.
3. The system of claim 1 wherein the array of transducer elements operates as a sparse array.
4. The system of claim 1 further comprising a second plurality of beamformers in a second housing, the second plurality of beamformers being in communication with the probe housing, the second beamformers receiving first image data from the first plurality of subarray beamformers having a plurality of second delay lines that receive first image data, plurality subarray beamformers operating in parallel to provide three dimensional image data.
5. The system of claim 1 wherein the second beamformer comprises a second stage beamformers.
6. The system of claim 5 further comprising a third stage beamformer.
7. The system of claim 1 wherein the system weighs less than 15 lbs.
8. The system of claim 7 wherein the system comprises a transducer array in a probe housing that is connected to a processor housing.
9. The system of claim 1 further comprising a sparse array transmission system.
10. The system of claim 1 further comprising a sparse array receiver system.
11. The system of claim 1 further comprising a matched filter.
12. The system of claim 1 further comprising a program performing a transmission waveform that is oversampled.
13. The system of claim 1 wherein the probe housing further comprises a plurality of flexible cables, each cable connecting a transducer subarray to a circuit board.
14. The system of claim 1 wherein the probe housing encloses a plurality of circuit boards, each circuit board having at least one of the first plurality of subarray beamformers, a memory that stores beamformer control data and a multiplexor circuit.
15. The system of claim 1 further comprising a flexible circuit.
16. The system of claim 15 wherein the flexible circuit comprises a flexible cable.
17. The system of claim 15 wherein the flexible circuit comprises a flexible printed circuit.
18. The system of claim 11 wherein the matched filter comprises a plurality of weights associated with stages of a delay line.
19. The system of claim 1 wherein each beam in the plurality of subarray beamformers is compressed.
20. The system of claim 1 wherein the second plurality of beamformers comprise a plurality of digital beamformers.
21. The system of claim 20 wherein the first plurality of beamformers comprise charge domain processors.
22. The system of claim 1 further comprising a program for storing a spread spectrum excitation waveform.
23. The system of claim 1 further comprising a system processor for performing scan conversion.
24. The system of claim 1 further comprising a system processor for performing Doppler processing.
25. The system of claim 1 wherein the probe housing further comprises a catheter or probe body with the transducer array mounted on a distal region.
26. The system of claim 1 wherein the system comprises a cardiac imaging system that simultaneously images the left and right ventricles.
27. The system of claim 1 wherein the second housing comprises a processor housing, a display and a control panel weighing less than 15 lbs.
28. The system of claim 1 wherein the array has an aperture sized to detect a heart volume of at least 200 ml/m2 in less than one cardiac cycle.
29. The system of claim 1 wherein the system simultaneously generates a plurality of beams for full volume cardiac imaging in less than one cardiac cycle.
30. A medical ultrasound imaging system comprising:
- a transducer array in a probe housing and a beamformer device for processing ultrasound data, the array being connected to a transmit circuit that actuates transmission of an ultrasound pulse from the array without a third harmonic of a transmission pulse frequency; and
- a control circuit that controls the transmission pulse.
31. The system of claim 30 further comprising a match filter to filter detected image data to detect a third harmonic with suppressed sidelobes.
32. The system of claim 1 wherein the array of transducer elements operates as a sparse array.
33. The system of claim 1 further comprising first plurality of beamforming in the probe housing and a second plurality of beamformers in a second housing, the second plurality of beamformers being in communication with the probe housing, the second beamformers receiving first image data from the first plurality of subarray beamformers having a plurality of second delay lines that receive first image data, plurality subarray beamformers operating in parallel to provide three dimensional image data.
34. The system of claim 33 wherein the second beamformers comprises a second stage beamformers.
35. The system of claim 30 wherein the control circuit actuates transmission of a modified square wave with a suppressed third harmonic.
36. A method for full volume cardiac ultrasound imaging comprising:
- transmitting ultrasound signals with a two dimensional array of transducer elements in a probe housing, the probe housing including circuitry to transmit a plurality of beams and form full volume cardiac images at a rate of at least 4 full volume cardiac images per second.
37. The method of claim 36 further comprising using a second beamformer device in a second housing, the second beamformer device being in communication with the probe housing, the second beamformer device receiving first beamformed image data from first subarrays, the second beamformer device having a plurality of second beamformers, the first subarrays operating in parallel to provide image data.
38. The method of claim 36 wherein the array of transducer elements comprises subaperture arrays.
39. The system of claim 36 further comprising using an array of transducer elements that is connected with a flexible circuit having the first beamformer device mounted thereon.
40. The method of claim 36 further comprising using the second beamformer that comprises a second stage beamformer.
41. The method of claim 40 further comprising using a third stage beamformer.
42. The method of claim 36 further comprising using a system that comprises a system processor, a display and a control panel that weighs less than 15 lbs.
43. The method of claim 36 further comprising using a system that comprises a probe housing that is connected to a processor housing.
44. The method of claim 36 further comprising using a flexible circuit board in the probe housing.
45. The method of claim 36 further comprising using a flexible cable in the probe housing that connects the transducer array to a circuit board assembly.
46. The method of claim 36 further comprising using parallel and serial beamforming.
47. The method of claim 36 further comprising performing a transmission waveform that is oversampled.
48. The method of claim 42 further comprising performing scan conversion and Doppler processing.
49. The method of claim 36 further comprising using a beamformer device to process image data from a volume of at least 200 ml/m2 during a single cardiac cycle.
50. The method of claim 36 further comprising emitting a transmission pulse having a transmission frequency, the pulse being emitted without a third harmonic of the transmission frequency.
51. The method of claim 50 further comprising using a match filter to process image data to detect a third harmonic of the transmission frequency.
52. The method of claim 51 further comprising using a Golay code match filter.
53. The method of claim 50 further comprising using an autocorrelation of a match filter template.
54. The method of claim 36 further comprising determining a sum of a first autocorrelation of a match filter template and a second autocorrelation of a match filter template to detect a third harmonic with sidelobe cancellation.
55. The method of claim 50 wherein the transmission pulse is a modified square wave pulse.
56. The method of claim 36 further comprising using a transducer probe housing having a controller connected to a plurality of subarray beamformer circuits.
57. The method of claim 56 wherein the probe housing further comprises a plurality of transmit circuits that controls a corresponding plurality of transducer subarrays.
58. The method of claim 56 wherein the probe housing contains a plurality of memory circuits, each memory circuit connected to the plurality of subarray beamformer circuits.
59. The method of claim 57 further comprising using a plurality of multiplexer circuits connected to the transmit circuits and the beamformer circuits in the probe housing.
60. The method of claim 56 further comprising simultaneously beamforming image data using a plurality of at least 16 transmission beams in parallel to form a 3D image.
International Classification: A61B 8/14 (20060101);