U.S. patent application number 10/114183 was filed with the patent office on 2003-02-06 for ultrasound scan conversion with spatial dithering.
This patent application is currently assigned to TeraTech Corporation. Invention is credited to Broadstone, Steven R., Chiang, Alice M., Gilbert, Jeffrey M..
Application Number | 20030028113 10/114183 |
Document ID | / |
Family ID | 27504383 |
Filed Date | 2003-02-06 |
United States Patent
Application |
20030028113 |
Kind Code |
A1 |
Gilbert, Jeffrey M. ; et
al. |
February 6, 2003 |
Ultrasound scan conversion with spatial dithering
Abstract
An ultrasound imaging system includes a scan conversion process
for converting ultrasound data into a standard display format
conversion and can be performed on a personal computer by
programming the computer to convert data from polar coordinates to
cartesian coordinates suitable for display on a computer monitor.
The data is provided from scan head enclosure that houses an array
of ultrasonic transducers and the circuitry associated therewith,
including pulse synchronizer circuitry used in the transmit mode
for transmission of ultrasonic pulses and beam forming circuitry
used in the receive mode to dynamically focus reflected ultrasonic
signals returning from the region of interest being imaged.
Inventors: |
Gilbert, Jeffrey M.;
(ElCerrito, CA) ; Chiang, Alice M.; (Weston,
MA) ; Broadstone, Steven R.; (Woburn, MA) |
Correspondence
Address: |
THOMAS O. HOOVER, ESQ.
BOWDITCH & DEWEY, LLP
161 Worcester Road
P.O. Box 9320
Framingham
MA
01701-9320
US
|
Assignee: |
TeraTech Corporation
|
Family ID: |
27504383 |
Appl. No.: |
10/114183 |
Filed: |
April 2, 2002 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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10114183 |
Apr 2, 2002 |
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09447144 |
Nov 23, 1999 |
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6379304 |
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09447144 |
Nov 23, 1999 |
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09203877 |
Dec 2, 1998 |
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6248073 |
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09203877 |
Dec 2, 1998 |
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PCT/US97/24291 |
Dec 23, 1997 |
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PCT/US97/24291 |
Dec 23, 1997 |
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08773647 |
Dec 24, 1996 |
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5904652 |
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08773647 |
Dec 24, 1996 |
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PCT/US96/11166 |
Jun 28, 1996 |
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PCT/US96/11166 |
Jun 28, 1996 |
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08599816 |
Feb 12, 1996 |
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5690114 |
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08599816 |
Feb 12, 1996 |
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08496804 |
Jun 29, 1995 |
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5590658 |
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08599816 |
Feb 12, 1996 |
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08496805 |
Jun 29, 1995 |
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5839442 |
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Current U.S.
Class: |
600/447 |
Current CPC
Class: |
A61B 8/4488 20130101;
A61B 8/465 20130101; A61B 2560/045 20130101; G01S 15/8979 20130101;
G01S 15/899 20130101; G01S 7/52046 20130101; A61B 7/04 20130101;
A61B 8/12 20130101; A61B 8/463 20130101; G01S 7/52023 20130101;
A61B 8/4455 20130101; G01S 7/003 20130101; A61B 8/00 20130101; G01S
7/52053 20130101; G01S 7/52044 20130101; A61B 8/56 20130101; G01S
7/52079 20130101; G01S 15/8915 20130101; G01S 15/8934 20130101;
G01S 7/52026 20130101; G01S 7/5208 20130101; G01S 7/52025 20130101;
G01S 7/52063 20130101; G01S 7/52034 20130101; G01S 7/52074
20130101; G01S 7/52073 20130101 |
Class at
Publication: |
600/447 |
International
Class: |
A61B 008/00 |
Claims
What is claimed:
1. A method of processing image data with an ultrasound imaging
device comprising of: providing a portable ultrasound imaging
system including a transducer array within a handheld probe, an
interface unit connected to the handheld probe with a first
interface, the interface unit having a beamforming device and being
connected to a data processing system with a second interface;
transmitting data from the handheld probe to the interface unit
with the first interface; performing a beamforming operation with
the beamforming device in the interface unit; and transmitting data
from the interface unit to the data processing system with the
second interface such that the data processing system receives a
beamformed electronic representation of the region of interest.
2. The method of claim 1 wherein the data processing system further
comprises a portable computer having a flat panel display.
3. The method of claim 1 further comprising generating a colored
image of the object.
4. The method of claim 1 further comprising providing a beamforming
circuit including a programmable delay device.
5. The method of claim 1 further comprising a hand-held probe
having a circuit board on which circuit elements are mounted, the
circuit elements including a charged coupled device integrated
circuit connected to an analog to digital converter.
6. The method of claim 1 wherein the step of providing a data
processing system further comprises providing a battery powered
portable computer having a graphical user interface.
7. The method of claim 1 further comprising displaying the image in
one of a plurality of windows on a display connected to the data
processing system.
8. The method of claim 1 further comprising performing a scan
conversion with a remap array.
9. The method of claim 1 wherein the first interface comprises an
electrical cable.
10. The method of claim 1 wherein the second interface comprises an
electrical cable.
11. The method of claim 1 wherein the second interface comprises a
wireless connection.
12. A portable ultrasound system for imaging a region of interest
comprising: a handheld probe housing in which a transducer array is
mounted; an interface unit connected to the handheld probe housing
with a first interface, the interface unit including a beamforming
device; and a data processing system connected to the interface
unit with a second interface such that the data processing system
receives a beamformed representation of the region of interest.
13. The system of claim 12 further comprising a flat panel display
connected to the data processing system that displays an image of
the region of interest.
14. The system of claim 12 wherein the probe housing further
comprises a beamforming circuit board having a programmable delay
device.
15. The system of claim 12 further comprising a circuit board
within the probe housing, the circuit board having a beamforming
integrated circuit mounted thereon.
16. The system of claim 12 further comprising a display and a
battery in the data processing system such that the battery
provides power to the probe housing.
17. The system of claim 12 further comprising a digital signal
processor in the interface unit.
18. The system of claim 12 wherein the data processing system
comprises a personal computer having a graphical user
interface.
19. The system of claim 12 further comprising a data transmitter
that forwards isochronous data from the interface unit to the data
processing unit.
20. The system of claim 12 wherein the second interface provides
asynchronous data transfer between the interface unit and the data
processing system.
21. The system of claim 12 wherein asynchronous signals are
transmitted from the data processing unit to the interface
unit.
22. The system of claim 12 wherein the transducer array comprises a
phased array device.
23. The system of claim 12 wherein the beamforming device comprises
32 channels or more.
24. The system of claim 12 wherein the interface unit further
comprises a transmit/receiver circuit and a preamplifier
circuit.
25. The system of claim 12 wherein the interface unit further
comprises a memory and a system controller.
26. The system of claim 12 wherein the beamforming device comprises
a CDP beamformer.
27. The system of claim 12 wherein the data processor includes a
scan conversion system and a high standard high speed
communications port.
28. The system of claim 12 wherein the second interface comprises a
wireless connection.
29. The system of claim 12 wherein the data processing system is
programmed to perform scan conversion with a remap array.
30. A portable ultrasound system for imaging a region of interest
comprising: an ultrasound probe system including a transducer array
and a beamforming device; and a data processing system connected to
the ultrasound probe system with an isochronous transfer connector
from the beamforming device to the data processing system to
provide a high-speed transmission link such that the data
processing system receives a beamformed representation of the
region of interest.
31. The system of claim 30 further comprising a flat panel display
connected to the data processing system that displays an image of
the region of interest.
32. The system of claim 30 wherein the probe system further
comprises a beamforming circuit having a programmable delay
device.
33. The system of claim 30 further comprising an interface unit
connected to a probe housing and a circuit board within the
interface unit, the circuit board having a beamforming integrated
circuit mounted thereon.
34. The system of claim 33 further comprising a display and a
battery in the data processing system such that the battery
provides power to the probe housing.
35. The system of claim 33 further comprising a digital signal
processor in the interface unit.
36. The system of claim 30 wherein the data processing system
comprises a personal computer having a graphical user
interface.
37. The system of claim 30 further comprising a data transmitter
that forwards isochronous data from the probe system to the data
processing unit.
38. The system claim of 33 wherein the transmission link provides a
connection between the interface unit and the data processing
unit.
39. The system of claim 30 wherein asynchronous signals are
transmitted from the data processing system to the probe.
40. The system of claim 33 wherein the interface unit comprises a
wireless interface.
41. The system of claim 30 wherein the transducer array comprises a
phased array device.
42. The system of claim 30 wherein the probe system comprises a
handheld housing having a transducer array and an interface
unit.
43. The system of claim 30 wherein the probe comprises a beamformer
device.
44. The system of claim 42 wherein the interface unit comprises a
beamforming device.
45. The system of claim 30 wherein the transmission link comprises
a wireless interface.
46. The system of claim 30 further comprising a first cable
connecting the probe to an interface unit and a second cable for
the transmission link between the interface unit and the data
processor.
Description
CROSS REFERENCES TO RELATED APPLICATIONS
[0001] This is a Continuation application of U.S. Ser. No.
09/447,144, filed on Nov. 23, 1999 which is a Continuation
application of U.S. Ser. No. 09/203,877, filed on Dec. 2, 1998
which is a Continuation application of International Application
No. PCT/US97/24291 filed on Dec. 23, 1997 which is a
Continuation-in-part application of U.S. Ser. No. 08/773,647 filed
on Dec. 24, 1996 which is a Continuation-in-part of International
Application No. PCT/US96/11166, filed on Jun. 28, 1996, which is a
Continuation-in-Part application of U.S. Ser. No. 08/599,816, filed
on Feb. 12, 1996, which is a Continuation-in-Part of U.S. Ser. Nos.
08/496,804 and 08/496,805 both filed on Jun. 29, 1995, the entire
contents of the above applications are being incorporated herein by
reference.
BACKGROUND OF THE INVENTION
[0002] Conventional ultrasound imaging systems typically include a
hand-held scan head coupled by a cable to a large rack-mounted
console processing and display unit. The scan head typically
includes an array of ultrasonic transducers which transmit
ultrasonic energy into a region being imaged and receive reflected
ultrasonic energy returning from the region. The transducers
convert the received ultrasonic energy into low-level electrical
signals which are transferred over the cable to the processing
unit. The processing unit applies appropriate beam forming
techniques such as dynamic focusing to combine the signals from the
transducers to generate an image of the region of interest.
[0003] Typical conventional ultrasound systems include transducer
arrays having a plurality, for example 128, of ultrasonic
transducers. Each transducer is associated with its own processing
circuitry located in the console processing unit. The processing
circuitry typically includes driver circuits which, in the transmit
mode, send precisely timed drive pulses to the transducer to
initiate transmission of the ultrasonic signal. These transmit
timing pulses are forwarded from the console processing unit along
the cable to the scan head. In the receive mode, beam forming
circuits of the processing circuitry introduce the appropriate
delay into each low-level electrical signal from the transducers to
dynamically focus the signals such that an accurate image can
subsequently be generated.
[0004] For phased array or curved linear scan heads, the ultrasound
signal is received and digitized in its natural polar (r,.theta.)
form. For display, this representation is inconvenient, so it is
converted into a rectangular (x,y) representation for further
processing. The rectangular representation is digitally corrected
for the dynamic range and brightness of various displays and
hard-copy devices. The data can also be stored and retrieved for
redisplay. In making the conversion between polar and rectangular
coordinates, the (x,y) values must be computed from the (r,.theta.)
values because the points on the (r,.theta.) array and the
rectangular (x,y) grid are not coincident.
[0005] In prior scan conversion systems, each point on the (x,y)
grid is visited and its value is computed from the values of the
two nearest 6 values by linear interpolation or the four nearest
neighbors on the (r,.theta.) array by bi-linear interpolation. This
is accomplished by use of a finite state machine to generate the
(x,y) traversal pattern, a bi-directional shift register to hold
the (r,.theta.) data samples in a large number of digital logic and
memory units to control the process and ensure that the correct
asynchronously received samples of (r,.theta.) data arrive for
interpolation at the right time for each (x,y) point. This prior
implementation can be both inflexible and unnecessarily complex.
Despite the extensive control hardware, only a single path through
the (x,y) array is possible.
SUMMARY OF THE INVENTION
[0006] In a preferred embodiment of the invention, scan data is
directed into a computer after beamforming and scan conversion is
performed to convert the scan data into a display format. In a
preferred embodiment, scan conversion can be performed entirely
using a software module on a personal computer. Alternatively a
board with additional hardware can be inserted to provide selected
scan conversion functions or to perform the entire scan conversion
process. For many applications, the software system is preferred as
additional hardware is minimized so the personal computer can be a
small portable platform, such as a laptop or palmtop computer.
[0007] Scan conversion is preferably performed using a spatial
dithering process described in greater detail below. Spatial
dithering simplifies the computational requirements for scan
conversion while retaining image resolution and quality. Thus, scan
conversion can be performed on a personal computer without the need
for more complex interpolation techniques and still provide
conversion at frame rates suitable for real time ultrasound
imaging.
[0008] Preferably, the scan conversion procedure includes an input
array, a remap array, and an output array. The remap array is an
array of indices or pointers, which is the size of the output image
used to determine where to get each pixel from the input array. The
numbers in each position in the remap array indicate where in the
input data to take each pixel will go into the output array in the
same position. Thus, the remap array and output array can be
thought of as having the same geometry while the input array and
output array have the same type of data, i.e., actual image
data.
[0009] The input array has new data for each ultrasound frame,
which means that it processes the data and puts the data in the
output array on every frame. In accordance with a preferred
embodiment of the invention, there is a new ultrasound frame
approximately every 1130 second. Consequently, the remap array data
can be generated relatively slowly (but still well under about one
second) as long as the routine operation of computing a new output
image from a new input data set is performed at the frame rate of
approximately 30 frames per second. This allows a general purpose
personal computer to perform the task of generating the data for
the remap array without compromising performance, but also without
having to dedicate additional hardware to the task. In a computing
system having a digital signal processor (DSP), the DSP can perform
the computations of the remap array.
[0010] Alternatively, certain scan conversion functions can be
performed by hardware inserted into the personal computer on a
circuit board. This board or a card can be inserted and used as an
interface to deliver data in the proper form to the PC bus
controller.
BRIEF DESCRIPTION OF THE DRAWINGS
[0011] The foregoing and other objects, features and advantages of
the invention will be apparent from the following more particular
description of preferred embodiments of the invention, as
illustrated in the accompanying drawings in which like reference
characters refer to the same parts throughout the different views.
The drawings are not necessarily to scale, emphasis instead being
placed upon illustrating the principles of the invention.
[0012] FIG. 1 is a block diagram of a conventional imaging array as
used in an ultrasound imaging system.
[0013] FIG. 2A is a schematic illustration of the relationship
between a linear ultrasound transducer array and a rectangular scan
region in accordance with the present invention.
[0014] FIG. 2B is a schematic illustration of the relationship
between a curved linear ultrasound transducer array and a curved
scan region in accordance with the present invention.
[0015] FIG. 2C is a schematic illustration of the relationship
between a linear ultrasound transducer array and a trapezoidal scan
region in accordance with the present invention.
[0016] FIG. 2D is a schematic illustration of a phased array scan
region.
[0017] FIG. 3 is a schematic pictorial view of a preferred
embodiment of the ultrasound imaging system of the present
invention.
[0018] FIG. 4A is a schematic functional block diagram of a
preferred embodiment of the ultrasound imaging system of the
invention.
[0019] FIG. 4B is a schematic functional block diagram of an
alternative preferred embodiment of the ultrasound imaging system
of the invention.
[0020] FIG. 5A is a schematic diagram of a beamforming and
filtering circuit in accordance with the invention.
[0021] FIG. 5B is a schematic diagram of another preferred
embodiment of a beamforming and filtering circuit in accordance
with the invention.
[0022] FIG. 5C is a schematic diagram of another preferred
embodiment of a beamforming and filtering circuit in accordance
with the invention.
[0023] FIG. 5D is a schematic diagram of a low pass filter in
accordance with the invention.
[0024] FIG. 5E is an example of an interface circuit board in
accordance with the invention.
[0025] FIG. 5F is a preferred embodiment of an integrated
beamforming circuit in accordance with the inventions.
[0026] FIG. 6 is a graphical illustration of the passband of a
filter in accordance with the invention.
[0027] FIG. 7A is a schematic diagram of input points overlayed on
a display.
[0028] FIG. 7B is a schematic diagram of a display of FIG. 6 having
input data converted to pixels.
[0029] FIG. 8 is a schematic diagram of a preferred embodiment of a
general purpose image remapping architecture.
[0030] FIGS. 9A-9B are a flow chart illustrating a remap array
computation technique in accordance with the invention.
[0031] FIG. 10 is a flow chart of an output frame computation
engine.
[0032] FIGS. 11A-11B are schematic pictorial views of two
user-selectable display presentation formats used in the ultrasound
imaging system of the invention.
[0033] FIGS. 12A-12B are functional block diagrams of a preferred
graphical user interface.
[0034] FIG. 13 illustrates a dialog box for ultrasound image
control.
[0035] FIGS. 14A-14D illustrate display boxes for entering system
information.
[0036] FIGS. 15A-15C illustrates additional dialog boxes for
entering probe or FOV data.
[0037] FIGS. 15D-15J illustrate additional display and dialog boxes
for a preferred embodiment of the invention.
[0038] FIG. 16 illustrates imaging and display operations of a
preferred embodiment of the invention.
[0039] FIGS. 17A-17C illustrate preferred embodiments of integrated
probe systems in accordance with the invention.
[0040] FIG. 18 illustrates a 64 channel integrated controller of a
transmit/receive circuit for an ultrasound system.
[0041] FIG. 19 illustrates another preferred embodiment of a
transmit and receive circuit.
[0042] FIG. 20 illustrates a Doppler Sonogram system in accordance
with the invention.
[0043] FIG. 21 illustrates a color flow map based on a fast fourier
transform pulsed Doppler processing system in accordance with the
invention.
[0044] FIG. 22 illustrates a processing system a waveform
generation in accordance with the invention.
[0045] FIG. 23 is a system for generating a color flow map in
accordance with the invention.
[0046] FIG. 24 is a process flow sequence for computing a color
flow map in accordance with the invention.
[0047] FIG. 25 is a process flow sequence for generating a color
flow map using cross correlation method.
DETAILED DESCRIPTION OF THE INVENTION
[0048] A schematic block diagram of an imaging array 18 of N
piezoelectric ultrasonic transducers 18(1)-18(N) as used in an
ultrasound imaging system is shown in FIG. 1. The array of
piezoelectric transducer elements 18(1)-18(N) generate acoustic
pulses which propagate into the image target (typically a region of
human tissue) or transmitting media with a narrow beam 180. The
pulses propagate as a spherical wave 185 with a roughly constant
velocity. Acoustic echoes in the form of returning signals from
image points I.sub.p or reflectors are detected by the same array
18 of transducer elements, or another receiving array and can be
displayed in a fashion to indicate the location of the reflecting
structure.
[0049] The acoustic echo from the image point I.sub.p in the
transmitting media reaches each transducer element 18(1)-18(N) of
the receiving array after various propagation times. The
propagation time for each transducer element is different and
depends on the distance between each transducer element and the
image point I.sub.p. This holds true for typical ultrasound
transmitting media, i.e. soft bodily tissue, where the velocity of
sound is at least relatively constant. Thereafter, the received
information is displayed in a manner to indicate the location of
the reflecting structure.
[0050] In two-dimensional B-mode scanning, the pulses can be
transmitted along a number of lines-of-sight as shown in FIG. 1. If
the echoes are sampled and their amplitudes are coded as
brightness, a grey scale image can be displayed on a cathode ray
tube (CRT) or monitor. An image typically contains 128 such scanned
lines at 0.75.degree. angular spacing, forming a 90.degree. sector
image. Because the velocity of sound in water is
1.54.times.10.sup.5 cm/sec, the round-trip time to a depth of 16 cm
will be 208 .mu.s. Thus, the total time required to acquire data
along 128 lines of sight (for one image) is 26.6 ms. If other
signal processors in the system are fast enough to keep up with
this data acquisition rate, two-dimensional images can be produced
at rates corresponding to standard television video. For example,
if the ultrasound imager is used to view reflected or back
scattered sound waves through the chest wall between a pair of
ribs, the heart pumping can be imaged in real time.
[0051] The ultrasonic transmitter is typically a linear array of
piezoelectric transducers 18(1)-18(N) (typically spaced
half-wavelength apart) for steered arrays whose elevation pattern
is fixed and whose azimuth pattern is controlled primarily by delay
steering. The radiating (azimuth) beam pattern of a conventional
array is controlled primarily by applying delayed transmitting
pulses to each transducer element 18(1)-18(N) in such a manner that
the energy from all the transmitters summed together at the image
point I.sub.p produces a desired beam shape. Therefore, a time
delay circuit is needed in association with each transducer element
18(1)-18(N) for producing the desired transmitted radiation pattern
along the predetermined direction.
[0052] As previously described, the same array 18 of transducer
elements 18(1)-18(N) can be used for receiving the return signals.
The reflected or echoed beam energy waveform originating at the
image point reaches each transducer element after a time delay
equal to the distance from the image point to the transducer
element divided by the assumed constant speed of the propagation of
waves in the media. Similar to the transmitting mode, this time
delay is different for each transducer element. At each receiving
transducer element, these differences in path length should be
compensated for by focusing the reflected energy at each receiver
from the particular image point for any given depth. The delay at
each receiving element is a function of the distance measured from
the element to the center of the array and the viewing angular
direction measured normal to the array.
[0053] The beam forming and focusing operations involve forming a
sum of the scattered waveforms as observed by all the transducers,
but in this sum, the waveforms must be differentially delayed so
they will all arrive in phase and properly weighted in the
summation. Hence, a beam forming circuit is required which can
apply a different delay on each channel, and vary that delay with
time. Along a given direction, as echoes return from deeper tissue,
the receiving array varies its focus continually with depth. This
process is known as dynamic focusing.
[0054] After the received beam is formed, it is digitized in a
conventional manner. The digital representation of each received
pulse is a time sequence corresponding to a back-scattering cross
section of ultrasonic energy returning from a field point as a
function of range at the azimuth formed by the beam. Successive
pulses are pointed in different directions, covering a field of
view from -45.degree. to +45.degree.. In some systems, time
averaging of data from successive observations of the same point
(referred to as persistence weighting) is used to improve image
quality.
[0055] FIGS. 2A-2D are schematic diagrams illustrating the
relationship between the various transducer array configurations
used in the present invention and their corresponding scan image
regions. FIG. 2A shows a linear array 18A which produces a
rectangular scanning image region 180A. Such an array typically
includes 128 transducers.
[0056] FIG. 2B is a schematic diagram showing the relationship
between a curved linear transducer array 18B and the resulting
sectional curved image scan region 180B. Once again, the array 18B
typically includes 128 adjacent transducers.
[0057] FIG. 2C shows the relationship between a linear transducer
array 18C and a trapezoidal image region 180C. In this embodiment,
the array 18C is typically formed from 192 adjacent transducers,
instead of 128. The linear array is used to produce the trapezoidal
scan region 180C by combining linear scanning as shown in FIG. 2A
with phased array scanning. In one embodiment, the 64 transducers
on opposite ends of the array 18C are used in a phased array
configuration to achieve the curved angular portions of the region
180 C at its ends. The middle 64 transducers are used in the linear
scanning mode to complete the rectangular portion of the region
180C. Thus, the trapezoidal region 180C is achieved using a
sub-aperture scanning approach in which only 64 transducers are
active at any one time. In one embodiment, adjacent groups of 64
transducers are activated alternately. That is, first, transducers
1-64 become active. Next, transducers 64-128 become active. In the
next step, transducers 2-65 are activated, and then transducers
65-129 are activated. This pattern continues until transducers
128-192 are activated. Next, the scanning process begins over again
at transducers 1-64.
[0058] FIG. 2D shows a short linear array of transducers 18D used
to perform phased array imaging in accordance with the invention.
The linear array 18D is used via phased array beam steering
processing to produce an angular slice region 180D.
[0059] FIG. 3 is a schematic pictorial view of an ultrasound
imaging system 10 of the present invention. The system includes a
hand-held scan head 12 coupled to a portable data processing and
display unit 14 which can be a laptop computer. Alternatively, the
data processing and display unit 14 can include a personal computer
or other computer interfaced to a CRT for providing display of
ultrasound images. The data processor display unit 14 can also be a
small, lightweight, single-piece unit small enough to be hand-held
or worn or carried by the user. Although FIG. 3 shows an external
scan head, the scan head of the invention can also be an internal
scan head adapted to be inserted through a lumen into the body for
internal imaging. For example, the head can be a transesophogeal
probe used for cardiac imaging.
[0060] The scan head 12 is connected to the data processor 14 by a
cable 16. In an alternative embodiment, the system 10 includes an
interface unit 13 (shown in phantom) coupled between the scan head
12 and the data processing and display unit 14. The interface unit
13 preferably contains controller and processing circuitry
including a digital signal processor (DSP). The interface unit 13
can perform required signal processing tasks and can provide signal
outputs to the data processing unit 14 and/or scan head 12. For
user with a palmtop computer, the interface unit 13 is preferably
an internal card or chip set. When used with a desktop or laptop
computer, the interface unit 13 can instead be an external
device.
[0061] The hand-held housing 12 includes a transducer section 15A
and a handle section 15B. The transducer section 15A is maintained
at a temperature below 41.degree. C. so that the portion of the
housing that is in contact with the skin of the patient does not
exceed this temperature. The handle section 15B does not exceed a
second higher temperature preferably 50.degree. C.
[0062] FIG. 4A is a schematic functional block diagram of one
embodiment of the ultrasound imaging system 10 of the invention. As
shown, the scan head 12 includes an ultrasonic transducer array 18
which transmits ultrasonic signals into a region of interest or
image target 11, such as a region of human tissue, and receives
reflected ultrasonic signals returning from the image target. The
scan head 12 also includes transducer driver circuitry 20 and pulse
synchronization circuitry 22. The pulse synchronizer 22 forwards a
series of precisely timed and delayed pulses to high voltage driver
circuits in the drivers 20. As each pulse is received by the
drivers 20, the high-voltage driver circuits are activated to
forward a high-voltage drive signal to each transducer in the
transducer array 18 to activate the transducer to transmit an
ultrasonic signal into the image target 11.
[0063] Ultrasonic echoes reflected by the image target 11 are
detected by the ultrasonic transducers in the array 18. Each
transducer converts the received ultrasonic signal into a
representative electrical signal which is forwarded to
preamplification circuits 24 and time-varying gain control (TGC)
circuitry 25. The preamp circuitry 24 sets the level of the
electrical signals from the transducer array 18 at a level suitable
for subsequent processing, and the TGC circuitry 25 is used to
compensate for attenuation of the sound pulse as it penetrates
through human tissue and also drives the beam forming circuits 26
(described below) to produce a line image. The conditioned
electrical signals are forwarded to the beam forming circuitry 26
which introduces appropriate differential delay into each of the
received signals to dynamically focus the signals such that an
accurate image can be created. Further details of the beam forming
circuitry 26 and the delay circuits used to introduce differential
delay into received signals and the pulses generated by the pulse
synchronizer 22 are described in the incorporated International
Application PCT/US96/11166.
[0064] In one preferred embodiment, the dynamically focused and
summed signal is forwarded to an A/D converter 27 which digitizes
the summed signal. Digital signal data is then forwarded from the
A/D 27 over the cable 16 to a color doppler processing circuit 36.
It should be noted that the A/D converter 27 is not used in an
alternative embodiment in which the analog summed signal is sent
directly over the system cable 16. The digital signal is also
demodulated in a demodulation circuit 28 and forwarded to a scan
conversion circuit 37 in the data processor and display unit
14.
[0065] As also shown a scan head memory 29 stores data from a
controller 21 and the data processing and display unit 14. The scan
head memory 29 provides stored data to the pulse synchronize 22,
the TGC 25 and the beam former 26.
[0066] The scan conversion circuitry 37 converts the digitized
signal data from the beam forming circuitry 26 from polar
coordinates (r,.theta.) to rectangular coordinates (x,y). After the
conversion, the rectangular coordinate data can be forwarded to an
optional post signal processing stage 30 where it is formatted for
display on the display 32 or for compression in a video compression
circuit 34. The post processing 30 can also be performed using the
scan conversion software described hereinafter.
[0067] Digital signal data from the A/D connector 27 is received by
a pulsed or continuous Doppler processor 36 in the data processor
unit 14. The pulsed or continuous Doppler processor 36 generates
data used to image moving target tissue 11 such as flowing blood.
In a preferred embodiment, with pulsed Doppler processing, a color
flow map is generated. The pulsed Doppler processor 36 forwards its
processed data to the scan conversion circuitry 28 where the polar
coordinates of the data are translated to rectangular coordinates
suitable for display or video compression.
[0068] A control circuit, preferably in the form of a
microprocessor 38 inside of a personal computer (e.g., desktop,
laptop, palmtop), controls the high-level operation of the
ultrasound imaging system 10. The microprocessor 38 or a DSP
initializes delay and scan conversion memory. The control circuit
38 controls the differential delays introduced in both the pulsed
synchronizer 22 and the beam forming circuitry 26 via the scan head
memory 27.
[0069] The microprocessor 38 also controls a memory 40 which stores
data used by the scan conversion circuitry 28. It will be
understood that the memory 40 can be a single memory or can be
multiple memory circuits. The microprocessor 38 also interfaces
with the post signal processing circuitry 30 and the video
compression circuitry 34 to control their individual functions. The
video compression circuitry 34 compresses data to permit
transmission of the image data to remote stations for display and
analysis via a transmission channel. The transmission channel can
be a modem or wireless cellular communication channel or other
known communication method.
[0070] The portable ultrasound imaging system 10 of the invention
can preferably be powered by a battery 44. The raw battery voltage
out of the battery 44 drives a regulated power supply 46 which
provides regulated power to all of the subsystems in the imaging
system 10 including those subsystems located in the scan head 12.
Thus, power to the scan head can be provided from the data
processing and display unit 14 over the cable 16.
[0071] FIG. 4B is a schematic functional block diagram of an
alternative preferred embodiment of the ultrasound imaging system
of the invention. In a modified scan head 12', demodulation
circuitry is replaced by software executed by the microprocessor 38
in a modified data processing and display unit 14'. In particular,
the digital data stream from the A/D converter 27 is buffered by a
FIFO memory 37. The microprocessor executes software instruction to
demodulate, perform scan conversion, color doppler processing, post
signal processing and video compression. Thus many hardware
functions of FIG. 4A are replaced by software stored in memory 40
in FIG. 4B, reducing hardware size and weight requirements for the
system 10'.
[0072] Additional preferred embodiments for beam forming circuitry
of ultrasound systems are depicted in FIGS. 5A, 5B, and 5C. Each of
these implementations requires that sampled-analog data be
down-converted, or mixed, to a baseband frequency from an
intermediate frequency (IF). The down-conversion or mixing is
accomplished by first multiplying the sampled data by a complex
value (represented by the complex-valued exponential input to the
multiplier stage), and then filtering the data to reject images
that have been mixed to nearby frequencies. The outputs of this
processing are available at a minimum output sample rate and are
available for subsequent display or Doppler processing.
[0073] In FIG. 5A, a set of sampling circuits 56 is used to capture
a data 54 represented by a packets of charge in a CCD-based
processing circuit fabricated on an integrated circuit 50. Data are
placed in one or more delay lines and output, at appropriate times
using memory and control circuitry 62, programmable delay circuits
58, to an optional interpolation filter 60. The interpolation
filter can be used to provide refined estimates of the round-trip
time of a sound wave and thereby provide better focus of the
returned signals from an array of sensors. In FIG. 5A, two
processing channels 52, of an array of processors, are depicted.
The outputs from the interpolation filters are combined, at an
analog summing junction 66, to provide a datum of beamformed output
from the array.
[0074] Data obtained using an ultrasound transducer resembles the
output of a modest-bandwidth signal modulated by the center
frequency of the transducer. The center frequency, or
characteristic frequency, of the transducer is equivalent to the
IF. In a sample-analog system (e.g., using CCDs),
.OMEGA.=2.pi.f.sub.I/f.sub.s, where fi is the intermediate
frequency and f.sub.I is the sampling frequency. The value n
corresponds to the sample-sequence number (i.e., n=0,1,2,3,4, . . .
). The outputs of the multiplier 68 are termed, in-phase (I) or
quadrature (Q) samples. In general, both I and Q values will be non
zero. When the IF is chosen to equal the f.sub.s/4, however, the
multiplier output will only produce either I or Q values in a
repeating sequence, I, Q, -I, -Q, I, Q, -I . . . In fact, the input
data are only scaled by 1 and -1. Thus, if the input data, a, are
sequentially sampled at times, a[0], a[1], a[2], a[3], a[4], . . .
, a[n], the output data are a[0], j*a[1], -a[2]. -j*a[3], a[4], . .
. , a[n], the output data are a[0], j*a[1], -a[2], -j*[3], a [4], .
. .
[0075] The I and Q outputs 74, 76 are each low-pass filtered 70, 72
to reject signal images that are mixed into the baseband. The
coefficients of the low-pass filters can be designed using a
least-mean square (LMS or L2-norm) or Chebyshev (L-infinity norm)
criteria. In practice, it is desirable to reduce the number of
coefficients necessary to obtain a desired filter characteristic as
much as possible.
[0076] An example of a CCD implementation of a low-pass filter is
illustrated in FIG. 5D. The device 90 consists of a 13-state tapped
delay line with five fixed-weight multipliers 94 to implement the
filter coefficients. As can be seen in the illustration of FIG. 6,
the ripple in the passband is under 0.5 dB and the stopband
attenuation is less than -30 dB of full scale.
[0077] The output of the low-pass filters are then decimated 78 by
at least a factor of 2. Decimation greater than 2 may be warranted
if the bandwidth of the ultrasound signal is bandlimited to
significantly less than half the sampling frequency. For most
ultrasound signals, a decimation factor greater than 2 is often
used because the signal bandwidth is relatively narrow relative to
the sampling frequency.
[0078] The order of the decimation and the low-pass filters may be
interchanged to reduce the clocking frequency of the low-pass
filters. By using a filter bank, the coefficients for the I and Q
low-pass filters can be chosen such that each filter only accepts
every other datum at its input. This "alternating clock" scheme
permits the layout constraints to be relaxed when a decimation rate
of 2 is chosen. These constraints can be further relaxed if the
decimation factor is greater than 2 (i.e., when the signal
bandwidth .ltoreq..ltoreq.f/2).
[0079] The down-converted output data are passed on for further
processing that may include signal-envelope detection or Doppler
processing. For display, the signal envelope (also referred to as
the signal magnitude) is computed as the square root of the sum of
the squares of the I and Q outputs. For the case when IF=f.sub.s/4,
that is either I=0 or Q=0, envelope detection becomes trivial. The
I and Q data are often the inputs to Doppler processing which also
uses the signal envelope to extract information in the positive-
and/or negative-frequency sidebands of the signal. In FIG. 5A, only
one down-conversion stage is required following the ultrasound
beamforming.
[0080] In FIG. 5B, a down-conversion stage has been placed in each
processing channel 52 following the sampling circuits 56. Here the
production of I and Q data 86, 88 is performed exactly as before,
however, much sooner in the system. The primary advantage of this
approach is that the data rate in each processing channel can be
reduced to a minimum, based on the ultrasound signal bandwidth and
hence the selection of the low-pass filter and decimation factor.
In this implementation, all processing channels 52 will use the
same complex-value multipliers and identical coefficients and
decimation factors in the filter stage. As in the preceding
implementation, complex-valued data are delayed and interpolated to
provide beamformed output.
[0081] The ultrasound front end depicted in FIG. 5C is nearly
identical to that in FIG. 5B. The difference is that the
interpolation stage 85, 87 has been removed and replaced by
choosing unique values in the complex-valued multipliers to provide
a more-precise estimate of the processing-channel delay. This
approach has the disadvantage that the output of the multiplier
will always exhibit I and Q values that are non zero. This is a
consequence of the varying sampling rate around the unit circle, in
a complex-plane diagram, of the multiplier input. Thus, this
approach can provide a more precise estimate of the sample delay in
each channel, but at the expense of producing fully complex-valued
data at the output of each processing channel. This modification
may require more post-processing for envelope and Doppler detection
than that presented in the previous implementations.
[0082] A preferred embodiment of a system used to interface between
the output of the beamforning or filtering circuit and the computer
is to provide a plug in board or card (PCMCIA) for the
computer.
[0083] The board 700 of FIG. 5E illustrates an embodiment in which
16 bits of digital beamformed data are received over the cable from
the scanhead by differential receivers 702. A clock signal is also
received at registers 704 along with converted differential data.
The first gate array 708 converts the 16 bits to 32 bits at half
the data rate. The 32 bit data is clocked into the FIFO 712 which
outputs add-on data 716. The second gate array 710 has access to
all control signals and outputs 714 to the PCI bus controller. This
particular example utilizes 16 bits of data, however, this design
can also be adapted for 32 bits or more.
[0084] Alternatively, a card suitable for insertion in a slot or
port of a personal computer, laptop or palmtop computer can also be
used. In this embodiment the differential receivers input to
registers, which deliver data to the FIFO and then to a bus
controller that is located on the card. The output from the
controller is connected directly to the PCI bus of the computer. An
alternative to the use of differential drivers and receivers to
interconnect the scan head to the interface board or card is to
utilize the IEEE 1394 standard cable also known as "firewire."
[0085] An example of a preferred embodiment of an integrated
beamforming circuit 740 is illustrated in FIG. 5F. The circuit 740
includes a timing circuit 742, and 5 delay circuits 760 attached to
each side of summing circuit 754. Each circuit 760 includes a
sampling circuit 746, a CCD delay line 752, a control and memory
circuit 750, a decoder 748, and a clocking driver circuit 744. The
circuitry is surrounded by contact pads 756 to provide access to
the chip circuitry. The integrated circuit is preferably less than
20 square millimeters in area and can be mounted on a single board
in the scan head as described in the various embodiments set forth
in the above referenced incorporated application. A sixteen, thirty
two, or sixty four delay line integrated circuit can also be
implemented utilizing a similar structure.
[0086] FIG. 7A is a schematic diagram of input points overlayed on
a display. As illustrated, input points I.sub.p received from the
ultrasound beam 180 do not exactly align with the rectangular
arranged pixel points P of a conventional display 32. Because the
display 32 can only display pixelized data, the input points
I.sub.p must be converted to the rectangular format.
[0087] FIG. 7B is a schematic diagram of a display of FIG. 6 having
input data converted to pixels. As illustrated, each image point
I.sub.p is assigned to a respective pixel point P on the display 32
to form an image.
[0088] One purpose of scan conversion is to perform the coordinate
space transformation required for use with scan heads that are not
flat linear, such as phased array, trapezoidal or curved linear
heads. To do this, data must be read in one order and output data
must be written in another order. Many existing systems must
generate the transformation sequences on the fly, which reduces the
flexibility and makes trapezoidal scan patterns more difficult.
[0089] Because scan conversion is reordering the data, it can also
be used to rotate, pan and zoom the data. Rotation is useful for
viewing the image with the scan head depicted at the top, left,
right, or bottom of the image, or an arbitrary angle. Zooming and
panning are commonly used to allow various parts of the image to be
examined more closely.
[0090] In addition to zooming into one area of the object, it is
useful to be able to see multiple areas simultaneously in different
regions of the screen. Often the entire image is shown on the
screen but certain regions are replaced with zoomed-in-views. This
feature is usually referred to as "window-in-a-window." Current
high-end systems provides this capability for one window, but it is
preferred that an imaging system allow any number of zoomed
regions, each of which having an arbitrary size and shape.
[0091] The use of irregular scan patterns can ease system design
and allow greater scan head utilization. In particular, this allows
reduction or hiding of dead time associated with imaging deep
zones. In the case of deep zone imaging, the beam is transmitted
but received at some later time after the wave has had time to
travel to the maximum depth and return. More efficient use of the
system, and thus a higher frame rate or greater lateral sampling,
can be obtained if other zones are illuminated and reconstructed
during this dead time. This can cause the scan pattern to become
irregular (although fixed and explicitly computed). The flexible
scan conversion described below corrects for this
automatically.
[0092] FIG. 8 is a schematic diagram of a preferred embodiment of a
general purpose image remapping architecture. In accordance with a
preferred embodiment of the invention, data is preferably brought
directly into the PC after beamforming and the remainder of the
manipulation is performed in software. As such, additional hardware
is minimized so the personal computer can be a small portable
platform, such as a laptop or palmtop computer.
[0093] Preferably, there is an input array 142, a remap array 144
and an output array 146. The remap array 144 is an array of indices
or pointers, which is the size of the output image used to
determine where to get each pixel from the input array 142. The
numbers in each position in the remap array 144 indicate where in
the input data to take each pixel which will go into the output
array 146 in the same position. Thus, the remap array 144 and
output array 196 can be thought of as having the same geometry
while the input array 142 and output array 146 have the same type
of data, i.e., actual image data.
[0094] The input array 142 has new data for each ultrasound frame,
which means that it processes the data and puts the data in the
output array 146 on every frame. In accordance with the invention,
there is a new ultrasound frame at a rate of at least 20 frames per
second and preferably approximately every {fraction (1/30)} second.
However, the remap array 144 is only updated when the head type or
viewing parameters (i.e., zoom and pan) are updated. Consequently,
the remap array 144 data can be generated relatively slowly (but
still well under about one second or else it can become cumbersome)
as long as the routine operation of computing a new output image
from a new input data set is performed at the frame rate of
approximately 30 frames per second. This allows a general purpose
personal computer to perform the task of generating the data for
the remap array 144 without compromising performance, but also
without having to dedicate additional hardware to the task. In- a
computing system having a digital signal processor (DSP), the DSP
can perform the computations of the remap array 144.
[0095] In a preferred embodiment of the invention, input memory for
the input array 142 can be either two banks of Static Random Access
Memory (SRAM) or one bank of Video Random Access Memory (VRAM),
where the input is serial access and the output is random access.
The VRAM bank, however, may be too slow and refresh too costly. The
remap memory for the remap array 144 is preferably sequential
access memory embodied in VRAM, or Dynamic Random Access Memory
(DRAM), although random access SRAM will also work. The output
memory for the output array 146 can be either a frame buffer or a
First-In First-Out (FIFO) buffer. Basically, the scan conversion is
done on demand, on the fly. Scan conversion is preferably performed
by software in the PC. If scan conversion is done in hardware,
however, the PC is merely storing data, thus reducing system
complexity. Thus, an architecture in accordance with the invention
is preferably just two random access input buffers, a sequential
access remap buffer and small (if any) FIFO or bit of pipelining
for the output buffer. This implies the output frame buffer is in
PC memory.
[0096] In accordance with a preferred embodiment of the invention,
a spatial dithering technique employing error diffusion is used in
ultrasound scan conversion. Typical dithering is done in the pixel
intensity domain. In accordance with the invention, however,
dithering is used in ultrasound scan conversion to approximate
pixels in the spatial domain and not in the pixel intensity domain.
Spatial dithering is used to approximate values that fall between
two input data points. This happens because only discrete radii are
sampled but pixels on the display screen can fall between two radii
and need to be filtered. Spatial dithering must be used to
interpolate between longitudinal sample points.
[0097] Recall that the remap array 144 stores the mapping of each
output point to an input point. The input data points are typically
in polar coordinates while the output points are in rectilinear
coordinates. Although the remap array 144 merely contains indices
into the input array 142, they can be considered to contain radius
(r) and angle (.theta.) values. Ideally, these values have
arbitrary precision and do not have to correspond to actual sampled
points. Now consider that these arbitrary precision numbers must be
converted into integer values. The integer radius values correspond
to discrete samples that were taken and are limited by the radial
sampling density of the system. The integer angle values correspond
to discrete radial lines that were scanned and are thus limited by
the number of scan angles. If spatial dithering is applied, these
floating point values can be mapped into fixed integer values
without having the artifacts that appear with discrete rounding
without error diffusion.
[0098] FIGS. 9A-9B are a flow chart illustrating a remap array
computation technique in accordance with the invention. At step
205, the scan heads are checked to see if there has been any
change. If the scan heads have been changed, processing continues
to step 210 where the new head type is configured. After step 210,
or if there has been no change in the scan heads (step 205)
processing continues to step 215. At step 215, the display window
is checked to see if there is any zooming, panning or new
window-in-window feature. If so, processing continues to step 220
where the user inputs the new viewing parameters. After step 220,
or if there is no window change at step 215, processing continues
to step 225 where the remap array is cleared to indicate a new
relationship between the input and output arrays.
[0099] At step 230, the program chooses a window W to process. At
step 235, all line error values LE and all sample error values SE
are initialized to zero. At step 240, a point counter P is
initialized to point to the top left pixel of the window W.
[0100] At step 245, the application computes a floating point line
number L.sub.FP and sample offset S.sub.FP for each point in a view
V. For a phased array, this would be a radius r and an angle
.theta.. At step 250, any previously propagated error terms
L.sub.E, S.sub.E (discussed below) are added to the floating point
values L.sub.FP, S.sub.FP for the point P. At step 255, floating
point terms are rounded to the nearest integer L.sub.R, S.sub.R,
which correspond to actual sampled points. At step 260, the
application computes rounding errors as:
L.sub.RE=L.sub.FP-L.sub.R;
S.sub.RE=S.sub.FP-S.sub.R.
[0101] At step 265, the errors are propagated to the pixel points
to the right, below left, below, and below right relative to the
current point P.
1 PROPAGATE ERRORS L.sub.E (right) = L.sub.E (right) + L.sub.RE *
7/16 L.sub.E (below left) = L.sub.E (below left) + L.sub.RE * 3/16
L.sub.E (below) = L.sub.E (below) + L.sub.RE * 5/16 L.sub.E (below
right) = L.sub.E (below right) + L.sub.RE * 1/16 S.sub.E (right) =
S.sub.E (right) + S.sub.RE * 7/16 S.sub.E (below left) = S.sub.E
(below left) + S.sub.RE 3/16 S.sub.E (below) = S.sub.E (below) +
S.sub.RE * 5/16 S.sub.E (below right) = S.sub.E (below right) +
S.sub.RE * 1/16
[0102] At step 270, the application computes a data index based on
a scan data ordered index:
REMAP(P)=Index(L.sub.R, S.sub.R).
[0103] At step 275, a check is made to see if there are more points
in the window. If there are more points to be processed, the
pointer P is incremented to the next point at step 280. Processing
then returns to step 245. Once all points in the window have been
processed, processing continues to step 285.
[0104] At step 285, a check is made to see if there are more
windows to be processed. If so, processing returns to step 230.
Otherwise, processing is done.
[0105] Because the dithering maps one source to each output pixel,
the same remapping architecture can be used to make real-time scan
conversion possible in software, even on portable computers. Thus,
the complicated dithering operation is only performed during
initialization or when viewing parameters are changed. However, the
benefits of the dithering are present in all the images.
[0106] FIG. 10 is a flow chart of an output frame computation
engine. At step 305, beamforming, demodulated input data is read
into memory. At step 310, the output pixel index P is initialized.
At step 315, the output array is set equal to the remapped input
array according to the following:
OUTPUT(P)=INPUT(REMAP(P)).
[0107] At step 320, the output pixel index P is incremented. At
step 325, a check is done on the pixel index P to see if the image
has been formed. If not, processing returns to step 315. Once all
the pixels in the image have been computed, processing continues to
step 330 where the output image is optionally smoothed. Finally, at
step 335, the output image is displayed.
[0108] Although dithering does remove the mach-banding and moire
pattern artifacts which occur with simple rounding, dithering can
introduce high-frequency noise. It is this high-frequency noise
whose average value allow for the smooth transition effects. To the
untrained eye, these artifacts are far less objectionable than
those obtained with the simple rounding or nearest-point case, but
may be objectionable to ultrasound technicians.
[0109] These artifacts can be greatly reduced or potentially
eliminated by employing a low-pass spatial filter to smooth the
image after the remapping process. The filter can be a box filter
or non-symmetrical filters can be matched to a desired input
resolution characteristic. Filters can be applied in the
rectilinear domain that match the orientation or angle of point
coordinates at the particular location.
[0110] Basically, it is desirable to have a matched filter whose
extent is similar to or proportional to distances between points
being dithered. Thus, a high magnification is preferably
accompanied by a large filter with much smoothing, whereas in
places with a spacing of the sampled radius r or angle (.theta.) is
small (on the order of one pixel), no filtering may be
required.
[0111] Because the remapping operation is basically two loads and a
store, it can be performed using a standard personal computer. The
remapping algorithm when encoded in assembly language has been
shown to work on a 166 MHZ Pentium-based PC to obtain very-near
real-time operation. In addition, the demodulation has been
performed on the PC when written in assembly language while still
achieving near real-time operation. Text and graphics labels are
preferably effected by storing fixed values or colors in the
beginning of the input buffer and then mapping to those places
where those colors are to be used. If effect, shapes or text are
drawn in the remap array, which will open and automatically be
overlayed on all of the images at no computational cost.
[0112] FIGS. 11A-11B are schematic pictorial views of display
formats which can be presented on the display 32 of the invention.
Rather than displaying a single window of data as is done in prior
ultrasound imaging systems, the system of the present invention has
multiple window display formats which can be selected by the user.
FIG. 11A shows a selectable multi-window display in which three
information windows are presented simultaneously on the display.
Window A shows the standard B-scan image, while window B shows an
M-scan image of a Doppler two-dimensional color flow map. Window C
is a user information window which communicates command selections
to the user and facilitates the user's manual selections. FIG. 11B
is a single-window optional display in which the entire display is
used to present only a B-scan image. Optionally, the display can
show both the B-mode and color doppler scans simultaneously by
overlaying the two displays or by showing them side-by-side using a
split screen feature.
[0113] FIG. 12 is functional block diagram of a preferred graphical
user interface. A virtual control 400 includes an ultrasound image
control display 410, a probe model properties display 420, and a
probe specific properties display 500. The virtual control display
400 is preferably coded as dialog boxes in a Windows
environment.
[0114] FIG. 13 illustrates a dialog box for the ultrasound image
control 410. Through the ultrasound image control display 410, the
user can select a probe head type 412, a zone display 414, a
demodulation filter 416, and an algorithm option 418. The user also
can initiate the ultrasound scan through this dialog box.
[0115] The probe model properties display 420 includes model type
425, safety information 430, image Integrated Pulse Amplitude (IPA)
data 435, doppler IPA data 440, color IPA data 445, probe geometry
450, image zones data 455, doppler zones data 460, color zones data
465, image apodization 470, doppler apodization 475, and color
apodization 480. These are preferably encoded as dialog boxes.
Through the model-properties dialog box 425, a user can enter
general settings for the probe model.
[0116] FIG. 14A illustrates a dialog box for entering a viewing
probe model properties. Entered parameters are downloaded to the
ultrasound probe.
[0117] FIG. 14B illustrates a dialog box for entering and viewing
safety information 430. As illustrated, a user can enter general
settings 432 and beam width table data 434 per governing
standards.
[0118] FIG. 14C illustrates a dialog box for entering and viewing
image IPA data 435. The dialog box displays beamformed output
values, listed in volts as a function of image display zones for
various drive voltages. Similar dialog boxes are used to enter the
doppler and color IPA data 440, 445.
[0119] FIG. 14D illustrates a dialog box for effecting the image
apodization function 470. As illustrated, the operator can enter
and view general settings 472 and vector information 474. The user
can select active elements for array windowing (or
apodization).
[0120] The probe specific property display 500 includes dialog
boxes for entering probe specifics 510, image Field Of-View (FOV)
data 520, doppler FOV data 530, and color FOV data 540. Through the
probe specifics dialog box 510, the user can enter general settings
512, imaging static information 514, doppler static information
516, and FOV settings 518.
[0121] FIG. 15A illustrates a dialog box for entering, and viewing
probe specific information. Any number of probes can be
supported.
[0122] FIG. 15B-15C illustrate dialog boxes for entering image FOV
data 520. As illustrated, a user can enter general settings 522,
breakpoint PGC data 524, zone boundaries 526, and zone duration 528
data. Dialog boxes for the doppler and color FOV data displays
530,540 are similar and are all the entry of general settings 532,
542, breakpoint TGC data 534, 544, and PRF data 536, 546.
[0123] FIGS. 15D-15J illustrate additional windows and control
panels for controlling an ultrasound imaging system in accordance
with the invention. FIG. 15D shows a viewing window for the region
of interest and a control panel situated side by side with the scan
image. FIG. 15E shows controls for the doppler field of view and
other selectable settings. FIG. 15F shows the color field of view
controls. FIG. 15G shows properties of the probe. FIG. 15h shows
the color IPA data for a probe. FIG. 151 shows the probe geometry
settings for a linear array. FIG. 15J shows settings for doppler
apodization.
[0124] FIG. 16 illustrates the zoom feature of a preferred
embodiment of the imaging system in accordance with the invention.
In this particular illustration detailed features of a phantom, or
internal anatomical features 600 of a patient that are shown on
screen 32, can be selected and enlarged within or over a display
window. In this particular example, a region 602 is selected by the
user and is enlarged at window 604. A plurality of such regions can
be simultaneously enlarged and shown on screen 32 in separate or
overlying windows. If two scan heads are in use, different views
can be shown at the same time, or previously recorded images can be
recalled from memory and displayed beside an image presented in
real time.
[0125] The architecture of the integrated front-end probe approach
was designed to provide small size, low power consumption and
maximal flexibility in scanning, including: 1) multi-zone focus on
transmission; 2) ability to drive a variety of probes, such as
linear/curved linear, linear/trapezoidal, and sector scan; 3)
ability to provide M-mode, B-mode, Color Flow Map and Doppler
Sonogram displays; 4) multiple, selectable pulse shapes and
frequencies; and 5) different firing sequences. Different
embodiments for the integrated front-end system 700 are shown in
FIGS. 17A, 17B and 17C. Modules unique to this invention are the
blocks corresponding to: beamforming chip 702, transmit/receive
chip 704, preamplifier/TGC chip 706.
[0126] The block labeled "front-end probe" (front-end controller)
directly controls the routine operation of the ultrasound scan head
by generating clock and control signals provided to modules 702,
704, 706 and to the memory unit 708. These signals are used to
assure continuous data output and to indicate the module for which
the data appearing at the memory-unit output are intended. Higher
level control of the scan head 710, as well as initialization, data
processing and display functions, are provided by a general purpose
host computer 720, such as a desktop PC, laptop or palmtop. Thus,
the front-end controller also interfaces with the host computer,
e.g. via PCI bus or Fire Wire 714 to allow the host to write
control data into the scanhead memory unit and receive data back.
This is performed at initialization and whenever a change in
parameters (such as number and/or position of zones or type of scan
head) is required when the user selects a different scanning
pattern. The front-end controller also provides buffering and
flow-control functions, as data from the beamformer must be sent to
the host via a bandwidth-constrained link, to prevent data
loss.
[0127] The system described permits two different implementations
of the Color Flow Map (CFM) and Doppler Sonogram (DS) functions.
FIG. 17A shows a hardware-based 722 implementation, in which a
dedicated Doppler-processing chip is mounted on a back-end card 724
and used as a co-processor to the host computer 720 to accomplish
the CFM and DS computations. FIG. 17B shows a software
implementation in which the CFM and DS computations are performed
by the host computer.
[0128] FIG. 17C shows yet another system integration, in which the
transducer array and the front-end processing units are not
integrated into a single housing but are connected by coaxial
cables. The front-end units include the front-end controller, the
memory and the three modules 704 (transmit/receive chip), 706
(preamp/TGC chip) and 702 (the beamforming chip) as shown in the
Figure.
[0129] "FireWire" refers to IEEE standard 1394, which provides
high-speed data transmission over a serial link. This allows use of
high-volume, low cost commercial parts for the interface. The
standard supports an asynchronous data transfer mode that can be
used to send commands and configuration data to the probe head
memory. It can also be used to query the status of the head and
obtain additional information, such as the activation of any
buttons or other input devices on the head. Additionally, the
asynchronous data transfer mode can be used to detect the type of
probe head attached. An isochronous transfer mode can be used to
transfer data back from the beamformer to the host. These data may
come directly from the A/D or from the demodulator or some
combination. If Doppler processing is placed in the probe head, the
Doppler processed data can be sent via FireWire. Alternatively the
data can be Doppler processed via software or hardware in the host.
There also exists a wireless version of the FireWire standard,
allowing communication via an optical link for untethered
operation. This can be used to provide greater freedom when the
probe head is attached to the host using wireless FireWire.
[0130] The preamp/TGC chip as implemented consists of integrated 32
parallel, low-noise, low-power, amplifier/TGC units. Each unit has
60-dB programmable gain, a noise voltage less than 1.5nV1{square
root}{square root over (Hz)} and dissipates less than 11 mW per
receiver channel.
[0131] As shown in FIG. 18, the multi-channel transmit/receive chip
consists of a global counter, a global memory and a bank of
parallel dual-channel transmit/receiver controllers. Within each
controller 740, there are local memory 745, delay comparator,
frequency counter & comparator, pulse counter & Comparator,
phase selector, transmit/receive select/demux switch (T/R switch),
and level shifter units.
[0132] The global counter 742 broadcasts a master clock and bit
values to each channel processor 740. The global memory 744
controls transmit frequency, pulse number, pulse sequence and
transmit/receive select. The local delay comparator 746 provides
delay selection for each channel. For example, with a 60 MHZ clock,
and a 10-bit global counter, a delay of up to 17 .mu.s can be
provided for each channel. The local frequency counter 748 provides
programmable transmit frequency. A 4-bit counter with a comparator
provides up to sixteen different frequency selections. For example,
using a 60-MHZ master clock, a 4-bit counter can be programmed to
provide different transmit frequencies such as 60/2=30 MHz, 60/3=20
MHz, 60/4=15 MHz, 60/5=12 MHz, 60/6=10 MHz and so on. The local
pulse counter 750 provides different pulse sequences. For example,
a 6-bit counter with a comparator can provide programmable
transmitted pulse lengths from one pulse up to 64 pulses. The
locally programmable phase selector which provides sub-clock delay
resolution.
[0133] While typically the period of the transmit-chip determines
the delay resolution, a technique called programmable subclock
delay resolution allows the delay resolution to be more precise
than the clock period. With programmable subclock delay resolution,
the output of the frequency counter is gated with a phase of the
clock that is programmable on a per-channel basis. In the simplest
form, a two-phase clock is used and the output of the frequency
counter is either gated with the asserted or deasserted clock.
Alternatively, multiple skewed clocks can be used. One per channel
can be selected and used to gate the coarse timing signal from the
frequency counter. For example, for a 60-MHz master clock, a
two-to-one phase selector provides 8-ns delay resolution and a
four-to-one phase selector provides 4-ns delay resolution.
[0134] Also shown are the integrated transmit/receiver select
switch 754, T/R switch and the integrated high-voltage level
shifter 750 for the transmit pulses. A single-chip transmit/receive
chip capable of handling 64 channel drivers and 32-channel
receivers can be used, each channel having a controller as shown in
FIG. 18.
[0135] In another implementation, shown in FIG. 19, the T/R
select/mux switch and the high-voltage level shifter are separated
from the other components 760 on a separate chip 762 to allow use
of different high-voltage semiconductor technologies, such as
high-breakdown silicon CMOS/JFET or GaAs technology for production
of these components.
[0136] The basic method for pulsed-Doppler ultrasound imaging is
illustrated in FIG. 20. The waveform consists of a burst of N
pulses 770. After each pulse as many range (depth) samples as
needed are collected. The time evolution of the velocity
distribution of material within the range gate is displayed as a
sonogram 772, a two-dimensional display in which the horizontal
axis represents time and the vertical axis velocity (as assessed by
Doppler shift). Different regions can be interrogated by moving the
range gate and varying its size. A Doppler sonogram can be
generated using single-range-gate Doppler processing, as shown in
FIG. 20. The operation of this method is as follows. A sequence of
N ultrasonic pulses is transmitted at a pulse repetition frequency
f.sub.prf along a given viewing angle. The return echoes are range
gated and only returns 774 from a single range bin are used,
meaning that only the returned signals corresponding to a region at
a selected distance (e.g. from depth d to d+.delta.d) from the
transducer array along the selected viewing angle are processed to
extract Doppler information. The velocity profiles of scatterers in
the selected region can be obtained by computing the Doppler shifts
of the echoes received from the scatterers. That is, Fourier
transformation 776 of the received time-domain signal provides
frequency information, including the desired Doppler, f.sub.d. The
velocity distribution of the scatterers in the region of interest
can be obtained from the relationship: 1 f d = 2 v c f c
[0137] where c is the speed of sound in the transmitting medium and
f.sub.c, is the center frequency of the transducer. As an example,
if N=16 and f.sub.prf=1 KHz, the above equation can be used to
generate a sonogram 772 displaying 16 ms of Doppler data. If the
procedure is repeated every N/f.sub.prf seconds, a continuous
Doppler sonogram plot can be produced.
[0138] Another embodiment involves a pulse-Doppler process for
color flow map applications. It is clinically desirable to be able
to display flow rates and patterns over a large region in real
time. One method for approaching this task using ultrasound is
called color flow mapping (CFM). Color flow mapping techniques are
an extension of the single-gated system described above. In CFM,
velocities are estimated not only along a single direction or line
segment, but over a number of directions (multiple scan lines)
spanning a region of interest. The velocity information is
typically color-coded (e.g. red indicates flow toward the
transducer, blue away) and superimposed over a B-mode image that
displays the underlaying anatomy.
[0139] A color-flow map 780 based on pulsed-Doppler processing is
shown in FIG. 21. The basic single-range bin system of FIG. 20 can
be extended to measure a number of range gates by sampling at
different depths and retaining the samples in storage for
additional processing. Note that this does not increase the
acquisition time, as data are collected from the same RF line.
Sweeping the beam over an area then makes it possible to assemble
an image of the velocities in a 20 region of interest. In
operation, the data from J range bins 782 along a single direction
are processed in parallel. After N pulse returns are processed, the
outputs represent a J.times.N range-vs-Doppler distribution, which
in turn can be used to generate a J.times.N velocity distribution
profile. The mean velocity at each depth d.sub.kk=1,2 . . . J, is
used to generate a single point or cell on the color-flow map; in
each cell, the standard deviation is used to assess turbulence. If
the procedure is repeated every N/f.sub.prf seconds for every J
range bins (e.g. spaced J/2 range bins apart) and for every scan
line in the region of interest, a 2D color-flow map plot can be
produced.
[0140] It is important to note that instead of an FFT-based
computation, a cross correlation technique, as described in the
publication of Jorgen A. Jensen, "Estimation of Blood Velocities
Using Ultrasound," University Press 1996, the contents of which is
incorporated herein by reference, can also be used to produce a
similar color flow map.
[0141] The range gate size and position can be determined by the
user. This choice determines both the emitted pulse length and
pulse repetition frequency. The size of the range gate is
determined by the length of the pulse. The pulse duration is
T.sub.p=21.sub.g/C=M.sub.fo
[0142] if the gate length is l.sub.g, and M is the number of sine
periods. The depth of the gate determines how quickly pulse echo
lines can be acquired. The maximum rate is
f.sub.prf=c/2d.sub.o
[0143] where d.sub.o is the distance to the gate.
[0144] The generic waveform for the pulse-Doppler ultrasound
imaging is shown in FIG. 22 where the waveform consists of a burst
of N pulses 800. As many as range depth samples as needed are
collected following each pulse in the burst. FIG. 22 also shows a
block diagram 810 of a conventional signal processor for this
imaging technique, where the returned echoes received by each
transducer are sampled and coherently summed prior to in-phase and
quadrature demodulation. The down converted/basebanded returns are
converted to a digital representation, and then stored in a buffer
memory until all the pulse returns comprising a coherent interval
are received. The N pulse returns collected for each depth are then
read from memory, a weighting sequence, v(n), is applied to control
Doppler sidelobes, and an N-point FFT is computed. During the time
the depth samples from one coherent interval are being processed
through the Doppler filter, returns from the next coherent interval
are being processed through the Doppler filter, returns from the
next coherent interval are arriving and are stored in a second
input buffer. The FFT 818 output is passed on to a display unit or
by time averaging Doppler samples for subsequent display.
[0145] The CDP device described here performs all of the functions
indicated in the dotted box of FIG. 22, except for A/D conversion,
which is not necessary because the CDP device provides the analog
sampled data function. This CDP Pulsed-Doppler Processor (PDP)
device has the capability to compute a matrix-matrix product, and
therefore has a much broader range of capabilities than needed to
implement the functions shown within the dotted lines.
[0146] The PDP device computes the product of two real-valued
matrices by summing the outer products formed by pairing columns of
the first matrix with corresponding rows of the second matrix.
[0147] In order to describe the application of the PDP to the
Doppler filtering problem, we first cast the Doppler filtering
equation into a sum of real-valued matrix operations. The Doppler
filtering is accomplished by computing a Discrete Fourier Transform
(DFT) of the weighted pulse returns for each depth of interest. If
we denote the depth-Doppler samples g(kj), where k is the Doppler
index, O.ltoreq.k.ltoreq.N-1, and j is the depth index, then 2 g (
k , j ) = n = 0 n - 1 v ( n ) f ( n , j ) exp ( - j2 kn / N )
[0148] The weighting function can be combined with the DFT kernel
to obtain a matrix of Doppler filter transform coefficients with
elements given by
W(k,n)=W.sub.k,n=v(n)exp(-j2nkn/N)
[0149] The real and imaginary components of the Doppler filtered
signal can now be written as 3 g r , kj = n = 0 N = 1 ( W r , kn f
r , nj - W i , kn f i , nj ) g r , kj = n = 0 N = 1 ( W r , kn f r
, nj + W i , kn f i , nj )
[0150] In the above equations, the double-indexed variables may all
be viewed as matrix indices. Therefore, in matrix representation,
the Doppler filtering can be expressed as matrix product operation.
It can be seen that the PDP device can be used to perform each of
the four matrix multiplications, thereby implementing the Doppler
filtering operation.
[0151] A block diagram of the PDP device described in this
invention is shown in FIG. 22. The device includes a J-stage CCD
tapped delay line, J CCD multiplying D/A converters (MDACs)
J.times.K accumulators, a J.times.K Doppler sample buffer, and a
parallel-in-serial out (PISO) output shift register. The MDACs
share a common 8-bit digital input on which elements from the
coefficient matrix are supplied. The tapped delay line performs the
function of a sample-and hold, converting the continuous-time
analog input signal to a sampled analog signal.
[0152] A two-PDP implementation 840 for color flow mapping in a
ultrasound imaging system is shown in FIG. 23. In this device,
during one pulse return interval, the top PDP component computes
all the terms of the form W.sub.k,f.sub.r and W.sub.if.sub.r as
shown in the above, while the bottom component computes the terms
of the form -W.sub.if.sub.i, and W.sub.kf.sub.i. The outputs of
each component are then summed to alternately obtain g.sub.r and
g.sub.i.
[0153] Doppler and color flow map processing involves a significant
amount of computation. This processing may be accomplished in
software using a general-purpose microprocessor. The presence of
instructions optimized for matrix-matrix operations, such as the
Intel MMX feature set, can substantially improve performance. A
software flow chart for color-flow map computation based on the FFT
computation algorithm is shown in FIG. 24. After initialization
900, the downconverted data is obtained 902 and the pointer P is at
the beginning of the scan line 904, the data is averaged and stored
906, a weighting function is applied 908, the FFT is computed 910,
the magnitude z(k) is computed for each frequency 912 followed by
the computation of first and second moments 914 and display thereof
in color 916. The painter is incremented 918 and each scan line is
processed as needed.
[0154] A software flow chart for color-flow map computation based
on the cross-correlation computation is showing in FIG. 25.
[0155] After initiation 940, the scan line data is obtained 942,
followed by the range bin data 944. The cross correlation is
computed 946 and averaged 948, and the velocity distribution 950,
first and second moments 952 are obtained and displayed 954. The
range bin data is increased 956 and the process repeated.
[0156] While this invention has been particularly shown and
described with references to preferred embodiments thereof, it will
be understood by those skilled in the art that various changes in
form and details may be made therein without departing from the
spirit and scope of the invention as defined by the appended
claims.
* * * * *