U.S. patent application number 13/498043 was filed with the patent office on 2012-07-12 for ultrasound 3d imaging system.
Invention is credited to Steven Broadstone, Alice Chiang, William Wong.
Application Number | 20120179044 13/498043 |
Document ID | / |
Family ID | 46455804 |
Filed Date | 2012-07-12 |
United States Patent
Application |
20120179044 |
Kind Code |
A1 |
Chiang; Alice ; et
al. |
July 12, 2012 |
ULTRASOUND 3D IMAGING SYSTEM
Abstract
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.
Inventors: |
Chiang; Alice; (Weston,
MA) ; Wong; William; (Milton, MA) ;
Broadstone; Steven; (Woburn, MA) |
Family ID: |
46455804 |
Appl. No.: |
13/498043 |
Filed: |
September 30, 2010 |
PCT Filed: |
September 30, 2010 |
PCT NO: |
PCT/US10/50959 |
371 Date: |
March 23, 2012 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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12570856 |
Sep 30, 2009 |
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13498043 |
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Current U.S.
Class: |
600/447 ;
600/459; 600/467 |
Current CPC
Class: |
A61B 8/145 20130101;
G01S 15/8913 20130101; G01S 15/8927 20130101; G01S 15/8959
20130101; G01S 7/5208 20130101; G01S 7/52095 20130101; G01S 15/8925
20130101; G01S 15/8961 20130101; A61B 8/565 20130101; A61B 8/4488
20130101; A61B 8/52 20130101; A61B 8/08 20130101; G01S 7/52079
20130101; A61B 8/483 20130101; A61B 8/54 20130101; A61B 8/4472
20130101; A61B 8/0883 20130101; G01S 15/8993 20130101; G10K 11/346
20130101; A61B 8/4483 20130101; A61B 8/488 20130101; A61B 8/466
20130101; A61B 8/4411 20130101; A61B 8/543 20130101; G01S 7/52082
20130101; A61B 8/4444 20130101; G01S 7/52085 20130101; A61B 8/14
20130101 |
Class at
Publication: |
600/447 ;
600/459; 600/467 |
International
Class: |
A61B 8/14 20060101
A61B008/14 |
Claims
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/m.sup.2 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/m.sup.2 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.
Description
CROSS REFERENCE TO RELATED APPLICATION
[0001] 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
[0002] 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.
[0003] 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
[0004] 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.
[0005] 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.
[0006] 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.
[0007] 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.
[0008] 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
1/8.lamda. to 1/16.lamda..
[0009] 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.
[0010] 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.
[0011] 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.
[0012] 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.
[0013] 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.
[0014] In medical ultrasound imaging, there is a need for harmonic
imaging where the transmitted waveform is of one fundamental
frequency F.sub.o, and the received signal of interest is a higher
harmonic, generally the 2.sup.nd harmonic (2 F.sub.o), or the third
harmonic (3 F.sub.o). 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.
[0015] 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
[0016] FIG. 1 illustrates the use of a two dimensional tiled array
for ultrasound imaging in accordance with the invention.
[0017] FIG. 2 illustrates a steerable two dimensional array in
accordance with the invention.
[0018] FIG. 3A illustrates the use of a first beamformer device for
far field beamsteering and focusing and a second time delay
beamformer for near field beamforming.
[0019] FIG. 3B illustrates a first analog subarray beamformer
forwarding data to a digital beamformer.
[0020] FIG. 3C illustrates a scanhead for a two dimensional
transducer probe.
[0021] FIG. 3D illustrates a preferred embodiment utilizing a
flexible circuit board and cable assembly.
[0022] FIG. 3E is a photomicrograph of a preferred embodiment with
an integrated circuit beamformer device having 16 channel subarray
beamformers that can form 4 sequential beams.
[0023] FIGS. 3F(1)-3F(4) illustrate preferred embodiments of
multiplexers used for switching in ultrasound transducer
systems.
[0024] FIG. 3G illustrates a switch timing diagram in accordance
with a preferred embodiment of the invention.
[0025] FIG. 3H is a photomicrograph of a 16 channel high voltage
multiplexer integrated circuit chip in accordance with a preferred
embodiment of the invention.
[0026] FIG. 3I is a schematic diagram of an 8 channel multiplexer
chip in accordance with a preferred embodiment of the
invention.
[0027] FIGS. 4A-4D illustrate a gated acquisition sequence for
cardiac imaging.
[0028] FIG. 4E illustrates a full cardiac imaging ultrasound scan
using a plurality of beams with a single transcript pulse with at
least six 3D volumetric images per second.
[0029] FIG. 4F illustrates an ultrasound system using a probe such
as that shown in FIG. 3C.
[0030] FIG. 4G illustrates an ultrasound system has a subarray
beamformer with serial beam output.
[0031] FIG. 4H illustrates an ultrasound system with a second stage
subaperture beamformer (509) that generates an output for a third
stage beamformer (510).
[0032] FIG. 4I illustrates an ultrasound system with a controller
integrated into the transducer probe housing.
[0033] FIG. 4J illustrates an ultrasound system with a controller
and transmission circuit integrated in the transducer probe
housing.
[0034] FIG. 4K illustrates an ultrasound system with parallel time
delay processors (519P) producing parallel output data.
[0035] FIG. 5A illustrates a preferred embodiment of a three
dimensional imaging system in accordance with the integrated
Subarray scan head invention.
[0036] FIG. 5B illustrates a preferred embodiment of the integrated
Subarray scan head invention using a charge domain processor for
the 2.sup.nd time delay beamforming.
[0037] FIG. 6A illustrates the use of the integrated subarray scan
head probe of the present invention with a second stage beamforming
ultrasound processor.
[0038] FIG. 6B illustrates use of the integrated Subarray scan head
with a digital beamforming processor.
[0039] FIG. 7 illustrates an ultrasound system in accordance with
the invention.
[0040] FIG. 8A illustrates a sparse array used in accordance with
the invention.
[0041] FIG. 8B graphically illustrates the sparse array
performance.
[0042] FIG. 9A illustrates the use of the integrated sparse array
scan head probe of the present invention connected to a host system
with charge-domain beamforming processing.
[0043] FIG. 9B illustrates the use of the integrated sparse array
scan head probe of the present invention connected to a
conventional digital ultrasound system with m-parallel beamforming
components.
[0044] FIG. 10 illustrates a scan head connected to a portable
computer in accordance with a preferred embodiment of the
invention.
[0045] FIG. 11 illustrates a near fully populated receive array in
which the receiving elements are independent of, and do not
overlap, the transmit array.
[0046] FIG. 12 graphically illustrates the azimuth and elevation
cross-sections of the receive array beampattern.
[0047] FIG. 13 is a magnified portion of the azimuthal beampattern
of FIG. 12 showing the mainlobe and sidelobe structure.
[0048] FIG. 14 illustrates a near fully populated receive array
beampattern.
[0049] FIG. 15 shows selected transmit locations for a sparse array
in accordance with the invention.
[0050] FIG. 16 illustrates a cross-sectional view of the transmit
sparse array beampattern of the embodiment in FIG. 15.
[0051] FIG. 17 illustrates a sparse transmit array beampattern.
[0052] FIG. 18 illustrates that it is possible to limit average
sidelobe energy to less than -35 dB relative to the central peak of
the beampattern.
[0053] FIG. 19 illustrates the 2D differential delay equation.
[0054] FIG. 20 illustrates a differential display profile.
[0055] FIG. 21 illustrates differential delay errors.
[0056] FIGS. 22A-22C illustrate embodiments of system processors in
a four parallel beamforming system.
[0057] FIGS. 23A and 23B illustrated non-coded and coded transmit
waveforms for spread spectrum ultrasound transmission.
[0058] FIGS. 24A-24C illustrate a process for forming a
transmission signal.
[0059] FIGS. 25A-25D illustrate preferred embodiments including a
matched filter.
[0060] FIGS. 26A-26D illustrate a square wave signal and modified
square wave transmit signals for third harmonic imaging.
[0061] FIGS. 27A and 27B illustrate match filter template for first
and second codes of Golay.
[0062] FIGS. 27C and 27D illustrate the autocorrelation of the
fundamental and third harmonics for the filter tremplate of the
first and second Golay codes.
[0063] FIG. 28 shows the sum of the two autocorrelation shown in
FIGS. 27C and 27D.
[0064] FIGS. 29A and 29B illustrate examples of the third harmonic
template of first and second 10 bit Golay codes.
[0065] FIGS. 30A-30C illustrate sidelobe cancellation with the
summed output.
DETAILED DESCRIPTION OF THE INVENTION
[0066] 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.
[0067] 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 FIG. 1 the tiles/subarrays
120 are made small enough so that when an object is placed within
the field-of-view of the imaging system, the incident radiation 122
from the object toward each "tile" can be treated using a far-field
approximation. Additional delay elements are incorporated as
second-stage processing to allow all subarrays to be coherently
summed (i.e., global near-field beamforming can be achieved by
simply delaying and then summing the outputs from all subarrays.)
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. Thus instead of looking in one direction as seen at
124a, the direction of tiles 120 can be steered in different
direction 124b. 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.
[0068] To scan an image plane using a steerable beamformer system a
process such as that shown in FIG. 2 can be used. A raster scan 260
can be used to scan an image plane 262 using a 2D steerable
transducer array 264.
[0069] A detailed diagram of an electronically-controlled
beamforming system in accordance with the invention is shown in
FIG. 3A. This system consists of a bank of parallel time-delay
beamforming processors 330, -330N. Each processor 330 consists of
two components: a 2D sub-array beamformer 332 for far-field
beamsteering/focusing and an additional time delay processor 334 to
allow hierarchical near-field beamforming of outputs from each
corresponding subarray. The sub-arrays 332 include m-programmable
delay lines 340 with tap selectors 342, multiplexers 344 and summed
346 output. As can be seen in FIG. 3A, for a system with
n-sub-arrays, n-parallel programmable 2.sup.nd-stage near field
time delays are needed for individual delay adjustment which are
converted with A/D converter 352 to allow all n-parallel outputs be
summed 354 coherently, in turn, this summed output is filtered 338
and provides the 3D images of the targeted object. A processor 336
controls sub-array operation.
[0070] Use of the scan head with a second stage digital beamformer
is shown in FIG. 3B. In this embodiment, a plurality of N sub-array
beamformers 400, which can be near-field beamformers in a preferred
embodiment, each receive signals from m transducer elements that
have separate delay lines whose outputs are summed and provided to
beamformers 420 so that this beamformer can be a conventional
system with conventional processor 480. A separate sub-array
processor 460 controls beamformers 400.
[0071] Without using this hierarchical subarray far-field and then
near-field beamforming approach, for an 80.times.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. 3A, 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. The
number of inputs to the subarray bank equals the number of 2D array
elements, and the number of outputs from the subarray bank equals
to the total transducer array element number divided by the
subarray element number, i.e., the number of outputs from the
subarray bank reference to the number of inputs is reduced by a
factor equal to the size of the subarray. For example, if one
selects to use a 5.times.5 subarray to implement this hierarchical
beamforming concept, after the first stage subarray beamforming,
the total number of wires needed to connect to the 2.sup.nd 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.times.80
element 2D transducer array to a conventional back-end beamforming
system. Using a 5.times.5 subarray processing bank first, the
number of wires required to connect to the backend beamforming
system is reduced to 256. Based on the current invention, a bank of
256 5.times.5 element subarrays Beamformer can be integrated with a
80.times.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.times.5 subarray far-field beamforming
processors can be easily integrated in a small size silicon (Si)
integrated circuit, eight of such 5.times.5 subarray beamformers
can be integrated on one chip. In this embodiment, only 32 chips or
less are integrated into the scan head, thereby reducing the cable
size from 6,400 wires down to 256 wires.
[0072] 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.
[0073] FIG. 3B demonstrates the hierarchical beamforming
architecture in which the beamforming of a group of neighboring
receive elements is implemented in two stages, i.e., instead of a
single long delay for each of the receive elements, a common long
delay is shared by all elements within the group, but each has its
own shorter, programmable delay in front of the long delay. Within
each group, the outputs from each of the short delays are summed
together then applied to the common long delay. A small group of
neighboring receiving elements having this common long delay
characteristic is defined as a "subarray" of the transducer array.
For example, for application using a 2D matrix array for real-time
3D imaging, the subarray can be a small array with 4.times.4, or
5.times.5 adjacent elements. The first stage programmable delays of
each element within this subarray are integrated inside the
transducer probe; the summed output from each subarray is then
connected to the backend processor. So, for a 64.times.48 element
2D array (3072 or more transducer elements), if a 4.times.4
subarray is used for the first stage beamforming, only 192 I/O
cable elements are needed to connect the front-probe to the backend
processors.
[0074] 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.
[0075] The construction of a 64.times.48 element 2D transducer
probe 485 with integrated 4.times.4 sub-array processors is
illustrated in FIG. 3C. The 64.times.48 element 2D array 487 can
include stacking 48 rows of 1D arrays each with 64 elements.
Connections to the elements of each 64 element 1D transducer array
are through a flex cable, so the transducer head assembly can
include the 2D transducer array and 48 flex cables 486. As shown in
FIG. 3C, each sub-array can include 4.times.4 elements (or other
rectangular or 2D geometry preferably having at least 16 elements
and up to 256 elements per subarray), the 48 flex cables are
grouped into 12 groups, with each adjacent four flex cables
connected to a printed circuit board, i.e., flex cables
corresponding to the row 1 and row 4 1D transducer array are
connected to the first printed circuit board 488, the row 5 to row
8 flex cables are connected to the second printed circuit board 487
and so forth, until the row 45 to row 48 flex cables are connected
to the twelfth printed circuit board. One end of each flex cable is
connected to the transducer elements and the other end of the flex
cable is connected to a 64-element flex connector mounted on the
printed circuit board. Within each printed circuit board, there are
a 16 4.times.4 element subarray processors 489 and a high voltage
multiplexor chip 491. The subarray processor consists of 16
parallel programmable delay lines each with a low-noise
pre-amplifier at its input and a separated programmable multiplier
as apodizer and the outputs of the 16 multipliers are summed
together to form a single outputs. Within each board, there is also
a high-voltage multiplex chip, so to allow the 4.times.64 element
transducer either operated in transmit or receive mode. A memory
chip 490 is also mounted on each printed circuit board to store the
programming delay for each of the delay lines. There are also power
supply cables and digital inputs 492 connected through interface
495 to each printed circuit boards.
[0076] As can be seen in FIG. 3C, the beamforming processor on the
printer circuit board has to provide the subarray beamforming
function for a total of 64.times.4 receive elements which are
divided into 16 subarrays with each subarray consisting of
4.times.4 elements. 16 subarray processors each can perform the
beamforming function, i.e., time-delay-and-sum function, for 16
receives are incorporated on the printer circuit board. A
photomicrograph of a 16 channel subarray beamformer chip is shown
in FIG. 3E, for each transmit pulse, the chip is capable of forming
4 sequential beams for 16 receivers. The chip is based on a 0.35
.mu.m double-poly, four-metal process. The size of the illustrated
chip is 1.2 mm.times.0.6 mm. Thus, 8 such 4.times.4 subarray
processors can be integrated on one chip with a chip size about 1.2
mm by 5 mm, so the circuit board only requires two such subarray
beamforming chips, each including 8 of the 4.times.4 subarray
processors.
[0077] 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.
[0078] 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
1.sup.st beam is clocked out and applied to the multiplier, another
32 ns later, the delayed tap output for the 2.sup.nd beam is
clocked out to the multiplier, the procedure follows, until 128 ns
later the tap out for the 4.sup.th 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 FIG. 3C is
64.times.4.times.4.times.2.times.2000=4 Mbits or more. Furthermore,
a compression method can be used to reduce the memory size.
[0079] 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.
[0080] 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).
[0081] 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.
[0082] 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.
[0083] The chip shown in FIG. 3G is a multiplexer fabricated using
high voltage CMOS process (>80V). The intended use is for
embedding in the acoustic module handle of a medical diagnostic
ultrasound probe where either two sets of transducer elements,
transmit and receiver, or the subarray beamformed outputs and the
transmit element share the same common transducer cable (COM).
Multiplexing is achieved with two sets of high voltage switches,
capable of handling high voltage bi-polar signals and frequencies
up to 20 MHz. The turning on and off of these switches is
controlled by two configuration signals (CONFIG[1:0]) and two
timing signals, TX_TIME and RX_TIME, which indicates when the
system is performing transmit or receive. The CONFIG[1:0] can be
used to configure the functionality of the port pins, for example,
swapping the transmit and receive elements.
[0084] In operation as shown in FIG. 3G, the switches between the
transmit elements and the common cable are turned on during the
transmit period (TX_TIME=1), and off during the receive period
(RX_TIME=1). The switches between the receive elements and the
common cable are turned off during transmit and turned on during
receive. The turn on/off time of the switches is less than 1
microsecond with time skew less than 100 picoseconds. A
photonmicrograph of a 16 channel high voltage multiplexer
integrated circuit chip is shown in FIG. 3H that was manufactured
using a 1 micron, two-poly, two-metal process and has a size of 14
mm.times.8 mm.
[0085] An implementation of 64 element 1D array with integrated
first stage subarray processor can also be implemented using the
design of FIG. 3C except that a single 64-element transducer array
491 is connected to a flex cable 497 with one end of the flex cable
is connected to each of the transducer element, the other end of
the cable is connected to a 64-pin flex connector. The connector is
mounted on a printed circuit board. Within the printed circuit,
there are eight 8-element subarray processors. Each subarray
processor consists of eight programmable delay lines, each delay
line has its separated low-noise pre-amplifier and at the output of
the delay line, there is an apodizer, i.e., a multiplier for the
beamshaping function. The outputs of the eight multipliers are
summed together to form a single analog outputs. A high-voltage
multiplexer circuit chip is also included on the printed circuit
board, to allow the 64 element transducer either operated in
transmit or receive mode. A memory chip is also mounted on the
printed circuit board to store the programming delay for each of
the delay lines. There are also power supply cables and digital
inputs connected to each of the printed circuit boards.
[0086] 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 FIG. 3D. In this implementation, instead of using a
printed circuit board, the subarray processing chip 499a, the high
voltage multiplexer circuit chip 449b and the memory chip 449c are
mounted on the flexible printed circuit or cable directly (497,
498).
[0087] 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 FIG. 3E, for each
transmit pulse, the chip is capable of forming 4 sequential beams
for 16 receivers. The chip also uses the 0.35 .mu.m double-poly,
four-metal process. The size of this chip is 1.2 mm.times.0.6 mm.
As in the prior embodiment, 4 such 16 subarray processor can easily
be integrated on one chip with a chip size about 1.2 mm by 2.4 mm,
so the board only requires one such subarray beamforming chip.
[0088] 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 FIG. 3E is
64.times.4.times.2.times.2000=1 Mbits where compression method can
be used to reduce the memory size. The systems described herein can
be used with a catheter or probe for insertion within body cavities
for cardiac imaging (4D) or other internal organs. The probe or
catheter assembly can include a circuit housing with the first
plurality of beamformers as described herein.
[0089] 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.
[0090] 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 FIGS. 4A-4D. Note
that the healthy adult human heart generally has a volume in a
range of 200-500 ml/m.sup.2. See, for example, Chikos et al.
"Visual assessment of Total Heart volume and Specific Chamber Size
from Standard Chest Radiographs," Am. J. Roentgenol 128:375-380,
March 1977, the entire contents of which is incorporated herein by
reference. Thus it is desirable to image a volume larger than 200
ml/m.sup.2 in order to capture a 3D image of the adult human heart
based on a singe transmission pulse sequence. Thus, the transducer
array has an aperture sufficient to transmit and receive signals
within this volume range substantially simultaneously, i.e. in less
than one cardiac cycle.
[0091] 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 (FIG. 4A). During the second cardiac cycle,
the second set of 32 by 96 beams is used to cover the adjacent part
of the cardiac image, FIG. 4B. The procedure follows (FIGS. 4C-4D),
until the 4.sup.th cardiac cycle with electrocardiographic gating;
the last portion of the cardiac image is then acquired. The four
images are then merged together to provide the complete 3D image.
This technique offers a wide sector angle and generates a
full-volume image after recording a sequence of multiple cardiac
cycles. However, stitching artifacts can occur with movement and in
patients with arrhythmias and respiratory difficulties, resulting
in nondiagnositc images. Therefore, the current RT-3D technique
used for echocardiography is near real time, but not truly real
time.
[0092] 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.times.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 FIG. 4E, 128.times.96
scanning beams are needed to provide a wide sector view angle. The
total round-trip time required for the 128.times.96 beams is
128.times.96.times.200 microseconds=2.45 seconds. The present
invention forms 16 received beams for each transmit pulse, it
follows then that only 2.45/16=0.15 s is needed to generate a
single 3D volumetric image with 128 by 96 beams or six 3D
volumetric images per second. At a six 3D volumetric images per
second rate, the scanner provides real-time 3D diagnostic quality
images for coronary artery disease detection. Thus, preferred
embodiments generate at least four full volume cardiac images per
second.
[0093] Systems used to generate at least 16 beams for each transmit
pulse are shown in the embodiments of FIGS. 4F-4K. systems
employing a phase shifting approach can also be used to achieve
full volume imaging.
[0094] An ultrasound system using a probe such as that shown in
FIG. 3C is illustrated in the diagram of FIG. 4F in which a 2D
transducer with integrated MUX, memory and subarray beamformer. The
subarray beamformer has a low-noise input amplifier and memory
504(1), 504(m) in the scanhead 502.
[0095] FIG. 4G illustrates a system with 2D transducer with
integrated MUX, memory and subarray beamformer in a probe 505 with
serial beam output 506. In this mode, the back-end time delay
processor has to run at a rate that is q.times.transducer input
sampling rate, i.e., if the center frequency of the transducer is 2
Mhz, the probe is oversampled at 8 Mhz rate. If four serial outputs
are generated by the subarray processor, ie., q=4, the back end
processor has to run at 32 MHz rate. The subarray beamformer has a
low-noise input amplifier
[0096] FIG. 4H illustrates a system with 2D transducer with
integrated MUX, memory and subarray beamformer with serial beam
output. In this mode, the back-end time delay processor has to run
at a rate that is q.times.transducer input sampling rate, i.e., if
the center frequency of the transducer is 2 Mhz, the probe is
oversampled at 8 Mhz rate. If four serial outputs are generated by
the subarray processor, ie., q=4, the back end processor has to run
at 32 MHz rate. In this approach, the back-end processor uses
subaperature beamforming system 508, i.e., the n adjacent receive
channels are grouped together, formed first stage beamforming
device 509 and then share a common long delay line 510. So, the
inputs to the back end are m channels. However, the back-end output
to the summing circuits are reduced to m/n outputs. For example, if
the transducer array are 64.times.48 elements, a 4.times.4 subarray
is used for subarray beamforming, the total outputs from the
integrated probe is then 64.times.48/16=192 cables, i.e., m=192.
However, in the current implementation, a subaperature size of 8 is
used, i.e., n=8, it follows then the total number of long time
delay processors are 192/8=24. The subarray beamformer has a
low-noise input amplifier.
[0097] FIG. 4I includes the elements of the system shown in FIG.
4G, however, a controller 514 is incorporated in the front-end
probe 512. The controller can be formed on a third circuit board
along circuitry for a wireless or cable connection 515 to the
external housing 513 which can be an interface to the main system
processor or can be integrated into a cart system or a portable
system as described herein.
[0098] FIG. 4J is a similar architecture to what described in FIG.
4H, however, a controller 514 is integrated into the front-end.
i.e., the interconnection cable are m=192 channels, however
sub-aperture beamformers are incorporated into the backend process,
n adjacent receive channels are beamformed first and then applied
to the back-end time delay process, if n=8, the outputs to the
summer are only 24 channels. In addition, an optional approach of
integrating transmit chips 517(m) into the front-end integrated
probe is used. However, the transmit channel can be located in the
back-end processor as indicated in FIG. 4I.
[0099] FIG. 4K is similar to what used in FIG. 4J, however, in the
back end processors, p-parallel beams are formed with each of the
subaperture beamformers 519 and parallel delay processor 519P. So,
if p=4, that is for each output beam from the front-end integrated
probe 518, four parallel beams are formed, each of the 4 outputs
are summed together at the summer 519S. This is how 16 beams are
formed for each transmission, i.e., in the probe, q=4, four serial
beams are formed, in the back-end p=4, four parallel beams are
formed. With p times q=16. Please note in FIG. 4K, a transmit
circuit has also been included in the front-end integrated probe.
However, note that the transmit chip can be located at the back-end
processor. In addition, in FIGS. 4F-4K a 2D transducer array is
used, the architecture can also be used with a 1D transducer array
with subaperture beamforming. These systems provide full volume
cardiac imaging that can produce video imaging of the heart
including left and right ventricles at video rates of at least 4
full volume images per second and preferably 6 full volume images
per second or more.
[0100] A preferred embodiment of the invention for 2D array
beamforming, each minimizing noise and cable loss with improved S/N
performance, are described in FIGS. 5A, 5B, 6A and 6B. In there
implementations, the bank of m parallel subarray beamforming
processors 520 and multiplexers 528 are integrated with the 2D
transducer array 525 to create a compact, low-noise, scan head 500.
FIG. 5A depicts a system that the compact scan head is connected to
a dedicated processing module, in which the m-parallel preamp/TGCs
522 transmit/received chips 524 and the 2.sup.nd stage time delay
processing units 526 are housed. This dedicated processing module
communicates with a host computer 540 via FireWire IEEE 1394 or USB
or PCI bus 542. Control and synchronization is performed by the
system controller 544 located in the processing module or housing
546. FIG. 5B depicts the same architecture as stated in FIG. 5A,
except that inside the dedicated processing module, the 2.sup.nd
stage time delay processing units are specifically implemented by
using charge-domain programmable (CDP) time-delay lines 600 in
housing 620 that is connected to handheld probe 660 and computer
housing 648. FIG. 6B depicts a system that the compact sparse array
scan head 700 is connected to a conventional, commercially
available time-domain digital ultrasound imaging system 700 with
n-parallel beamforming channels 760. It is easy to see that in FIG.
6A, the time-delay processor 720 can also be implemented by using
CDP time-delay lines 740. In these embodiments the near-field
beamforming is housed 720, 780 in the same housing with other image
processing functions. These systems are described in International
Application No. PCT/US2007/014526 filed Jun. 22, 2007, designating
the U.S. and U.S. application Ser. No. 11/474,098 filed Jun. 23,
2006, both applications being incorporated herein by reference in
their entirety.
[0101] 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 FIG. 7 a CDP beamforming chip 810, a time multiplexed
computing structure can be used to generate multiple beams, i.e.,
for each transmit pulse, the bank of 2D subarray beamformers 818
and its corresponding 2.sup.nd stage near-field time-delay line are
capable of providing multiple beams sequentially. The computing
circuits sequentially generate the delays required for forming K
beams. The device operates as follows. Once a set of sampled
returned-echoes are loaded in the delay lines with sampling
circuits 814, at time t.sub.1, the delays required for forming beam
1 are computed 812 within each module 822 and applied in parallel
to all delay lines. The sampled return-echoes with proper delays
are coherently summed 802 and filtered 804 to form the first beam.
At time t.sub.2, the delays required for forming beam 2 are
computed within each module and applied in parallel to all delay
lines. The sampled return-echoes with proper delays are coherently
summed to form the second beam. The procedure repeats until the Kth
beam is coherently formed.
[0102] For example, if a computing circuit with 16-serial
addressable outputs is built in with the CDP subarray and the
2.sup.nd 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.times.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.times.48
pixel resolution at a 30 frame/s frame rate.
[0103] The array described addresses ultrasound imaging
applications using a two-dimensional 2 cm.times.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.times.80 element array. Using the
subarray processing technique described above, a scan head with
integrated subarray beamforming circuits followed by a 2.sup.nd
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.
[0104] 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.
[0105] As an example of a sparse array, assuming a 2.times.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.
[0106] 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 (FIG. 8A). The geometry of activated
element placement has been developed to maximize the efficiency of
the beamformers while minimizing grating and side lobe clutter.
Switching between a plurality of different array patterns is used
to provide the most efficient beam pattern at different beam angles
relative to the region or volume of interest being scanned. Thus, a
first pattern can utilize that illustrated in FIG. 8A, which is
than switched to a second pattern for a different scan angle. This
can involve selecting a transducer element within a neighborhood
880 surrounding a given element to scan at a second angle.
[0107] 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:
J = 1 2 u mnx 2 .intg. .intg. s W ( u x , u y ) B ( u x , u y ) u x
u y , ( 1 ) ##EQU00001##
where the weighting function, W(u.sub.x, u.sub.y), applies more
weight to regions in the array response that require side lobe
reduction. The optimization method begins with no weighting (i.e.,
W(u.sub.x, u.sub.y)=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(u.sub.x,u.sub.y), the
weighting is chosen such that
W(u.sub.x,u.sub.y)=B(u.sub.x,u.sub.y). This is done in an
interactive manner until convergence.
[0108] 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 (FIG.
8A).
[0109] FIG. 8B is a plot of the array performance, sensitivity
versus cross range, for a 256-element sparsely-sampled array at 3
MHZ. The peak to maximum side lobe level is approximately 30 dB. To
improve this performance, the system is configured to achieve the
maximum main lobe to clutter level ratio possible, which has been
independently verified.
[0110] FIG. 9B depicts a system that the sparse array scan head 900
is connected to a conventional, commercially available time-domain
digital ultrasound imaging system 940 with m-parallel beamforming
channels. It is easy to see that in FIG. 9A, the time-delay
processor can also be implemented by using CDP time-delay lines 920
in housing 925 that is connected to a separate computer 927. An
array of m multiplexers 906 is used to switch between a sequence of
scan patterns executed using a software program and system
controller 940 or processor 950. The sequence of sparse array
patterns is thus selected to scan at different scan angles of an
object being imaged to provide 3D ultrasound imaging thereof.
[0111] 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.
[0112] A commercially available software tool such as KB-Vol3D of
KB-VIS technologies, Chennai, India, provides display or viewing 3D
features such as: [0113] Fast Volume-Rendering [0114] Shaded
Surface Display
[0115] 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. [0116] MIP (Maximum
Intensity Projection) with Radials [0117] MPR
(Multiple-Plane-Reformating) with Oblique & Double-Oblique and
3D correlation [0118] MRP Slabs & Multi-Cuts [0119] Curved MPR
[0120] Color & Opacity Presets with Editor [0121]
Region-Growing and Volume Measurements [0122] Cutaway Viewing with
Slab-Volume and Interactive Real-time VOI
[0123] 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.
[0124] 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. [0125] Image Save in Multiple
Formats
[0126] Images displayed by KB-Vol3D can be captured to various
image formats (including DICOM, JPEG, and BMP etc.) [0127] Movie
Capture in AVI Format
[0128] Visualization operations can also be captured to an AVI
movie .le and played on Windows Media Player, QuickTime, and Real
Player etc.
[0129] The invention can be implemented using a scan head 12
connected to a portable computer 14 as shown in FIG. 10. the
ultrasound system 10 can also include a cable 16 to connect the
probe 12 to the processor housing 14. Certain embodiments can
employ an interface unit 13 which can include a beamformer device.
Scan head 12 can include a transducer array 15A (2D) and a circuit
housing 15B which can house multiplexer and/or beamforming
components as described in detail in U.S. Pat. Nos. 6,106,472 and
6,869,401, the entire contents of these patents being incorporated
herein by reference.
[0130] A 2D array configuration using sparse-array for transmission
and non-overlapped fully-populated array is used for receiving. For
an N.times.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.times.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.
[0131] An example of the element locations for the near
fully-populated 40 by 60 receive array 50 is shown in FIG. 11. The
2400-element array is depopulated by the 256 sparse array transmit
elements to yield 2144 receive element locations. These elements
are independent and do not overlap the sparse-array transmit
elements. In a preferred embodiment the transmit elements
constitute less then 25% of the total number of array elements, and
preferably less then 15%.
[0132] The azimuth and elevation cross-sections of the beampattern
of the above mentioned receive array are shown in FIG. 12. The
first sidelobe is approximately -13 dB relative to the central
peak. The grating lobes are less than -30 dB relative to the peak.
Given that the 2D array is wider than tall, the azimuthal beamwidth
(plotted in blue (solid)) is slightly narrower than the elevation
beamwidth (plotted in green (dotted)).
[0133] In FIG. 13, a magnifying view of the above mentioned
azimuthal beam pattern demonstrates the detailed mainlobe and
sidelobe structure. For this case, the beamwidth is approximately
1.5 degrees. The beam pattern is nearly identical to the fully
populated 60.times.40 element beam pattern. The receive array beam
pattern is shown in FIG. 14. As stated above, the received sparse
array is comprised of a 2144 elements. There are no sidelobes due
to depopulating the center of the array by 256 (transmit)
elements.
[0134] An example of the final element locations for the 256
transmit sparse array 60 are shown in FIG. 15. The 256 element
locations are confined to the central 32.times.32 elements of the
fully populated array. These elements are independent and do not
overlap the receive array elements. A cross sectional view of the
transmit sparse array beampattern is shown in FIG. 16. The first
sidelobe is approximately -17 dB relative to the central peak. The
grating lobes are less than -27 dB relative to the peak. The sparse
array optimization algorithm minimizes the sidelobe energy +/-45
degrees in elevation and +/-45 degrees in elevation.
[0135] FIG. 17 demonstrates the beam pattern of the sparse transmit
array shown in FIG. 15. The transmit beampattern is designed to
uniformly cover a 4.times.4 beam data pyramid. The transmit sparse
array is comprised of a 256-element subset of the fully populated
2400-element array (approximately 10% fill). The placement of the
transmit/receive array design algorithm required over 750
iterations to minimize the transmit and receive sidelobe energy
within the +/-45 degree azimuth, +/-45 degree elevation region. As
shown in FIG. 18, after 750 iterations, the final sparse
transmit-array element locations limit the average sidelobe energy
to less than -35 dB relative to the central peak of the
beampattern.
[0136] 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.
[0137] 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.
[0138] 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.times.64-element
array.
[0139] 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 FIG. 14A referenced above
for a system with m-subarrays, m-parallel programmable
2.sup.nd-stage near field time delays are needed for individual
delay adjustment to allow all m-parallel outputs be summed
coherently, in turn, this summed output provides the 3D images of
the targeted object.
[0140] It is easy to understand that, without using this
hierarchical subarray small aperture and then large aperture
beamforming approach, for an 80.times.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.times.5 subarray to
implement this hierarchical beamforming system, after the first
stage subarray beamforming, the total number of wires needed to
connect to the 2.sup.nd 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.times.80 2D transducer array to a conventional
back-end beamforming. Using a 5.times.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.times.5 element subarrays beamformer can
be integrated with a 80.times.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.
[0141] It is important to note that 5.times.5 subarray
small-aperture beamforming processors can be easily integrated in a
small size silicon integrated circuit, eight of such 5.times.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.times.3 subarray up to an 8.times.8
subarray. The preferred range is at or between 4.times.4 and a
6.times.6 array for a square array geometry. Rectangular subarrays
can also be used preferably either 3.times.4, 4.times.5, or
4.times.6. Note that a 1/4.lamda. 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.times.48 element 2D array, using a
4.times.4 subarray processing bank in the transducer housing first,
the number of back-end beamforming channels is reduced to 192.
[0142] 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 FIGS.
4-6B. In these embodiments, the bank of m parallel subarray
beamforming processors are integrated with the 2D transducer array
to create a compact, low-noise, scan head. FIG. 4 depicts a system
that the compact scan head is connected to a dedicated processing
module, in which the m-parallel preamp/TGCs, transmit/received
chips and the 2.sup.nd stage time delay processing units are
housed. This dedicated processing module communicates with a host
computer 540 via FireWire, USB or PCI bus. Control and
synchronization is preformed by the system controller located in
the processing module. FIG. 5 depicts the same architecture as
stated in FIG. 4, except, inside the dedicated processing module,
the 2.sup.nd stage time delay processing units are specifically
implemented by using Charge-Domain programmable time-delay lines.
FIGS. 6A and 6B depicts a system that the compact scan head is
connected to a conventional, commercially available time-domain
digital ultrasound imaging system with m-parallel beamforming
channels. It is easy to see that in FIGS. 6A and 6B the time-delay
processor can also be implemented by using CDP time-delay
lines.
[0143] 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 FIGS.
19-24B.
[0144] 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.times.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.
[0145] 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.
[0146] 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 FIG. 7, in a CDP beamforming chip, a time multiplexed computing
structure can be used to generate multiple beams, i.e., for each
transmit pulse, the bank of 2D subarrays and its corresponding
2.sup.nd stage near-field time-delay line are capable of providing
multiple beams sequentially. The computing circuits sequentially
generate the delays required for forming K beams. The device
operates using the following sequence: once a set of sampled
returned-echoes are loaded in the delay lines, at time t.sub.1, the
delays required for forming beam 1 are computed within each module
and applied in parallel to all delay lines. The sampled
return-echoes with proper delays are coherently summed to form the
first beam. At time t.sub.2, the delays required for forming beam 2
are computed within each module and applied in parallel to all
delay lines. The sampled return-echoes with proper delays are
coherently summed to form the second beam. The procedure repeats
until the Kth beam is coherently formed.
[0147] For example, if a computing circuit with 16-serial
addressable outputs is incorporated with the processor subarray and
the 2.sup.nd 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.times.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.times.48 pixel resolution at a 30 frame/s
frame rate.
[0148] 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 128.lamda..
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.
[0149] 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 FIG. 19 is the 3D differential delay equation
for a 2D array. This equation represents differential delay at
array element (x.sub.m, y.sub.m) as a function of range and angles
Theta and Phi (relative to the center of the 2D array). The
equation can be reduced to a 1D array by setting all of the y.sub.m
(y coordinates of the element locations) to 0. The differential
delay can be constrained to a single plane (instead of a volume) by
setting angle Phi=0.
[0150] 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
FIG. 20, the differential delay profile is plotted for an element
near the center of a 2D array.
[0151] 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,
FIG. 21 depicts the differential delay error as a function of range
for a few elements. The worst case differential delay (the data
plotted in blue) is for an element in the corner of a 2D array
(Theta=+45 degs, Phi=+45 degs) attempting to receive image data
from the direction Theta=-45 degs, Phi=-45 degs. For this case, the
maximum differential delay is greater than the constraint (>1
sample error); therefore, the element would not be enabled to
receive until a range greater than approximately 100 samples.
[0152] A block diagram of an hierarchical two-stage parallel
beamforming system 958 is shown in FIG. 22A. A two dimensional
transducer array 960 of a handheld probe, such as 12 in FIG. 10
weighing less than 15 lbs, can be coupled to amplifier 962 before
being connected to the input of the beamforming system 964. The
beamforming system can comprise a plurality of short delay lines,
which are coherently summed at summing circuit 968 where the output
is delivered the longer delay lines 970, which are also summed at
summing circuit 972. First stage coarse beamforming includes
coherently summing returned echoes from a small aperture, for
example, with 8 neighboring receivers in this particular
embodiment. Because of the small size of the aperture, the delay
length of each short delay is only about 8.lamda.. So, for a 192
element inputs, 24 such small aperture, coarse beams are formed.
Each of those 24 beams is then applied to its corresponding long
programmable long delay line for a large aperture, fine-resolution
beamforming. To form four parallel beams, four such beamforming
structures are required. As can be seen in FIG. 22A, this
hierarchical implementation to form 24 coarse beams, only 192 short
delays are required, and then followed by 24 long delays, each one
with a programmable delay length shorter than 128.lamda.. For four
parallel beams, only 192 short delays plus 96 long delays are
needed, it offers a tremendous saving in terms of electronic
components and power.
[0153] 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.
[0154] 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
FIG. 22B, A/D converters 976 can be used at the input of each short
delay line and the time delays performed digitally. Or
alternatively, as shown in the embodiment 980 of FIG. 22C, the A/D
converters 982 can be used at the output of each coarse beam and
the long delay can be performed digitally. The A/D conversion can
be performed using available discrete components, or in a preferred
embodiment, a charge domain A/D converter can be formed on the same
integrated circuit with the charge domain beamformer with
hierarchical summing carried out digitally.
[0155] 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 FIG. 23A is an example of a 5
cycle 3 MHz sinusoid without spread spectrum coding. A coded or
spread spectrum system transmits a broadband, temporally elongated
excitation signal with a finite time-bandwidth product. The
received signal is decoded to produce an imaging pulse with
improved signal to noise ratio. The benefit of using coded signals
in ultrasound imaging systems offers the use of high-resolution
imaging while significantly lowering the peak acoustic power. These
signals also provide signal processing gain that improves the
overall system receiving sensitivity. Direct sequence modulation is
the modulation of a carrier by a code sequence. In practice, this
signal can be AM (pulse), FM, amplitude, phase or angle modulation.
It can also be a pseudorandom or PN sequence that can comprise a
sequence of binary values that repeat after a specified period of
time.
[0156] 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 FIG. 23B. Each "chip"
corresponds to 1 cycle of the gated transmit waveform. Thus, FIG.
23B appears nearly identical to that of FIG. 23A except that the
4.sup.th cycle is inverted. In both FIG. 23A and FIG. 23B, the
continuous line represents the continuously sampled sinusoidal
waveform, whereas the cross hatched points is a sampled signal,
where 10 samples are taken per cycle. The coded pulse sequence,
which has a length N.times.P, can effectively reduce the peak power
in the transmitting media by spreading the power spectrum over a
longer time duration. Upon reception of the spread spectrum/coded
returned echoes, a pulse compression matched filter can be used to
decode the received signals to produce an imaging pulse that has
improved signal to noise ration (SNR). The SNR improvement of a
N.times.P coded pulse sequence is 10 log(NP). So, for a Barker code
of length 7 and a two-cycle burst transmit waveform, a SNR of 11.4
dB improvement can be achieved. However, in the present system, the
transmit and received waveforms are oversampled with an oversampled
rate of S. Typically an oversampled rate of S=4 has been used. It
follows then, at the receiver end, a matched filter with tap length
of N.times.P.times.S can be used to decode and to compress the
returned echoes to produce an imaging pulse with a SNR improvement
of 10 log(NPS). In the above example, for N=7, P=2, S=4, a SNR of
17.5 dB can be achieved.
[0157] A preferred method of forming a transmission signal is shown
in FIGS. 24A-24C. The base sequence is a single pulse, as can be
seen in FIG. 24A. Using the 5-chip Barker code [111-11], FIG. 24B
represents the convolution of the base sequence with the Barker
code. Finally, the system transmits an oversampled version of the
continuous waveform, as shown in FIG. 24C, a 6-times oversampled
waveform is used as transmitted wave from
[0158] 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
FIG. 25A. In this implementation, a two-stage hierarchical
beamforming architecture is used, first a small aperture short
delay beamformer 986 output signals that are coherently summed
return echoes from 8 adjacent transducers, a pulse compression
matched filter 987 is then followed to decode the received signal,
this compressed signal 988 is then applied to a long delay line to
complete the beamforming requirement. In system 990 of FIG. 25B, an
A/D converter 992 is incorporated at each of the matched filter
outputs. It follows then the long delay will be carried out
digitally, using a second stage digital delay line
implementation.
[0159] A matched filter implementation is shown in FIG. 25C. The
filter 994 consists of a K-stage tapped delay line receiving
signals from sampling circuit 995 and K programmable multipliers.
The spreaded, coarse beamformed signal, f.sub.n, is continuously
applied to the input of the delay line. At each stage of the delay,
the signals can be non-destructively sensed and multiplied by a tap
weight 996, W.sub.k, where k=1, 2, 3, . . . , K-2, K-1, K. The
weighted signals are summed together with summing circuit 997 to
create a compressed output g.sub.m 998. It can be seen that, at
time t=n,
g.sub.n=f.sub.n-1W.sub.1+f.sub.n-2W.sub.2+f.sub.n-3W.sub.3+ . . .
+f.sub.n-K-2W.sub.K-2+f.sub.n-K-1W.sub.k-1+f.sub.n-KW.sub.K
[0160] Using the example shown in FIGS. 24A-24C, if the system
transmit a 6-time oversampled, 5-chip Barker code, and the weights
of the matched filter are selected as a time reversal of the
transmitted 5-chip Barker code excitation waveform, the matched
filter produces a cross correlation output 999 that is the
compressed, decoded pulsed signal (see FIG. 25D), with a filter
gain of 10 log(5.times.6)=15 dB.
[0161] Consider an ultrasound pulser with 3 cycles of square wave,
where an example is shown in FIG. 26A. The frequency spectrum of
such a waveform has a third harmonic component at about 4 dB below
the fundamental frequency. Shown in FIG. 26B shows a modified
square wave that has been combined with a sinusoidal signal.
[0162] 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 FIG. 26C is a modified 3 cycle square wave with pulse
widths reduced to 2/3 of regular square wave. Using Modified Square
Wave pulsers, a low harmonic transmitted waveform approximates a
sinusoidal waveform, compared with regular square wave. This allows
an ultrasound imaging system to use less expensive pulser design
for harmonic imaging.
[0163] In current ultrasound system, 2.sup.nd 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
2.sup.nd harmonic imaging. Since the energy in the second harmonic
frequency band is much lower than that in the fundamental frequency
band, to increase the 2.sup.nd harmonic sensitivity the spectral
overlap between the fundamental and the 2.sup.nd harmonic has to be
minimized, however, in doing so, the 2.sup.nd harmonic imaging
resolution is reduced as a result. The higher harmonics, in
particular the 3.sup.rd 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 3.sup.rd harmonic image. Preferred embodiments of
the present invention, use 3.sup.rd harmonic imaging with coded
excitation where additional gain from matched-filter processing is
used to overcome the problem with weak 3.sup.rd harmonic
returns.
[0164] Two pertinent requirements for a good 3.sup.rd harmonic
coded excitation imaging are, first by, minimum 3.sup.rd harmonic
component in the transmit coded waveform, and secondly, a code
selection with a 3.sup.rd harmonic receive template that has a
minimal sidelobe energy after the 3.sup.rd harmonic matched
filtering. A transmit coded waveform with minimum 3.sup.rd harmonic
component. In this section, two 3.sup.rd harmonic templates for a
Golay coded with minimum sidelobe are presented.
[0165] 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
a(z)a(z-1)+b(z)b(z-1)=2N
where a(z) and b(z) are complementary pairs. As a result, no image
artifacts due to match-filter sidelobes. In the following, the
3.sup.rd harmonic template of a coded excitation transmit waveform
is derived.
[0166] Let the Fourier transform of transmit waveform be
H(j.omega.)=.parallel.H(j.omega.).parallel.e.sup.j.alpha.(.omega.)
[0167] the Fourier transform of 3rd harmonic excited by this
transmit waveform should be
[0167] H 3 rdH ( j.omega. ) = H 3 ( j .omega. 3 ) = H ( j .omega. 3
) 3 j3.alpha. ( .omega. 3 ) ##EQU00002## [0168] 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.
[0168] H ( j .omega. 3 ) j3.alpha. ( .omega. 3 ) ##EQU00003##
[0169] Based on the above derivation, the fundamental and 3.sup.rd
harmonic template of a 10-bit Colay code pair is shown in FIGS. 27A
and 27B, where GCP1=[-1 1 1 -1 1 -1 1 1 1 -1], GCP2=[-1 1 1 1 1 1 1
-1 -1 1] The fundamental and 3.sup.rd harmonic matched filter
outputs, i.e. the autocorrelation of fundamental and 3.sup.rd
harmonic templates of the selected Golay pair is shown in FIGS. 27C
and 27D, respectively. Finally, the same fundamental and third
harmonic outputs of the two complementary codes are shown in FIG.
28. The 3.sup.rd harmonic template of the first 10 bit Colay code
is shown in FIG. 29A. A 12 times oversampling for fundamental and 4
times oversampling are used in deriving the code. The 3.sup.rd
harmonic template of the second 10 bit Colay code is shown in FIG.
29B. Again, a 12 times oversampling for fundamental and 4 times
oversampling are used in deriving this code.
[0170] Another technique in selecting the 3.sup.rd 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 3.sup.rd harmonic
template is shown in FIGS. 30A and 30B. It is demonstrated in FIG.
30C that there is no sidelobe of this matched filter output.
[0171] 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.
* * * * *