U.S. patent application number 11/712633 was filed with the patent office on 2007-06-28 for ultrasonic diagnostic imaging with blended tissue harmonic signals.
This patent application is currently assigned to Philips Ultrasound, Inc.. Invention is credited to Michalakis Averkiou, Jeffry E. Powers, David N. Roundhill.
Application Number | 20070149879 11/712633 |
Document ID | / |
Family ID | 26708866 |
Filed Date | 2007-06-28 |
United States Patent
Application |
20070149879 |
Kind Code |
A1 |
Roundhill; David N. ; et
al. |
June 28, 2007 |
Ultrasonic diagnostic imaging with blended tissue harmonic
signals
Abstract
An ultrasonic diagnostic imaging system and method are described
which produce tissue harmonic images containing both fundamental
and harmonic frequency components. Such a blended image takes
advantage of the performance possible with the two types of
ultrasonic echo information and can advantageously reduce near
field clutter while improving signal to noise performance in the
far field of the image.
Inventors: |
Roundhill; David N.;
(Woodinville, WA) ; Averkiou; Michalakis;
(Nicosia, CY) ; Powers; Jeffry E.; (Bainbridge
Island, WA) |
Correspondence
Address: |
DORSEY & WHITNEY LLP;INTELLECTUAL PROPERTY DEPARTMENT
SUITE 3400
1420 FIFTH AVENUE
SEATTLE
WA
98101
US
|
Assignee: |
Philips Ultrasound, Inc.
|
Family ID: |
26708866 |
Appl. No.: |
11/712633 |
Filed: |
February 28, 2007 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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09617318 |
Jul 17, 2000 |
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11712633 |
Feb 28, 2007 |
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09247343 |
Feb 8, 1999 |
6283919 |
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09617318 |
Jul 17, 2000 |
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08943546 |
Oct 3, 1997 |
5879303 |
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09247343 |
Feb 8, 1999 |
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08723483 |
Sep 27, 1996 |
5833613 |
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08943546 |
Oct 3, 1997 |
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Current U.S.
Class: |
600/447 ;
600/458 |
Current CPC
Class: |
G01S 7/52026 20130101;
A61B 8/488 20130101; G01S 7/52074 20130101; G01S 15/8993 20130101;
G01S 7/5206 20130101; A61B 8/463 20130101; G01S 7/52033 20130101;
G01S 7/52036 20130101; G01S 15/895 20130101; G01S 7/52046 20130101;
A61B 8/14 20130101; A61B 8/08 20130101; G01S 7/52077 20130101; G01S
7/52038 20130101; G01S 15/8979 20130101 |
Class at
Publication: |
600/447 ;
600/458 |
International
Class: |
A61B 8/00 20060101
A61B008/00 |
Claims
1-27. (canceled)
28. A method of imaging a biological sample, comprising the steps
of: generating an initial ultrasonic signal; directing the
ultrasonic signal into and along a propagation path in the sample,
wherein the sample causes finite, non-linear amplitude distortion
of the ultrasonic signal along the propagation path and thereby
produces a distorted ultrasonic signal comprised of a first order
component signal and higher order harmonic component signals at a
first and higher order harmonic frequencies respectively, and
further wherein the sample also reflects the distorted ultrasonic
signal including the first order and the higher order harmonic
components; receiving the higher order harmonic components of the
reflected distorted ultrasonic signal produced by the distortion of
the initial ultrasonic signal along the propagation path and caused
by said sample; forming an image principally from one of said
received higher order harmonic components of the reflected
distorted ultrasonic signal; and displaying said formed image.
29. A method according to claim 28, wherein the removing step
includes the step of high-pass filtering the received, reflected
distorted signal to remove therefrom the first order component
thereof.
30. A method according to claim 28 wherein: the generating signal
includes the steps of generating first and second ultrasonic
signals; the directing step includes the steps of directing the
first and second ultrasonic signals into the sample; the receiving
step includes the step of receiving any first and second signals
reflected and distorted by said sample; the forming step includes
the steps of i) subtracting the received second distorted signal
from the received first distorted signal to produce a resultant
signal, and ii) forming the image from said resultant signal.
31. A method according to claim 28 wherein: the higher order
harmonic component signals include a second order harmonic
component and further, higher order components; and the forming
step includes the step of forming the image principally from the
second order component of the received reflected distorted
ultrasonic signal.
32. A method according to claim 28, further including the step of
maintaining the sample substantially free of any contrast agent
while directing the initial ultrasonic signal into and along the
propagation path in the sample.
33. A method according to claim 28, wherein: the generating step
includes the step of generating a series of ultrasonic pulse
signals; and the directing step includes the step of directing the
series of ultrasonic pulse signals into and along the propagation
path in the sample.
34. A method according to claim 28, wherein the sample linearly
reflects the distorted ultrasonic signal produced by the distortion
of the initial ultrasonic signal along the propagation path and
caused by the sample.
35. A system for imaging a biological sample, comprising: means for
generating an initial ultrasonic signal; means for directing the
initial ultrasonic signal into and along a propagation path in the
sample, wherein the sample causes finite, non-linear amplitude
distortion of the fundamental signal along the propagation path,
and said distortion produces a distorted ultrasonic signal
comprised of a first order component and higher order harmonic
components at a first and higher order harmonic frequencies
respectively, and wherein the sample also reflects the distorted
ultrasonic signal including the first order and the higher order
harmonic components thereof; means for receiving the higher order
harmonic components of the reflected distorted ultrasonic signal
produced by the distortion of the initial ultrasonic signal along
the propagation path and caused by said sample; means for forming
an image principally from one of said received higher order
harmonic component signals of the reflected distorted ultrasonic
signal; and means for displaying said formed image.
36. A system according to claim 35, wherein the means for removing
the first order component from the received distorted signal
includes a high-pass filter to filter the received, reflected
distorted signal to remove therefrom the first order component
thereof.
37. A system according to claim 35, wherein: the means for
generating the ultrasonic signal includes means for generating
first and second ultrasonic signals; the means for directing the
ultrasonic signal into the sample includes means for directing the
first and second ultrasonic signals into the sample; the receiving
means includes means for receiving any first and second signals
reflected and distorted by said sample; the means for forming the
image includes i) means for subtracting the received second
distorted signal from the received first distorted signal to
produce a resultant signal, and ii) means for forming the image
from said resultant signal.
38. A system according to claim 35, wherein: the higher order
harmonic components include a second order harmonic component and
further, higher order harmonic components; and the forming means
includes means for forming the image principally from the second
order harmonic component of the received reflected distorted
ultrasonic signal.
39. A system according to claim 35, for use with a sample that is
substantially free of contrast agent while the initial ultrasonic
signal is directed into and along the propagation path.
40. A system according to claim 35, wherein: the generating means
includes means to generate a series of ultrasonic pulse signals;
and the directing means includes means to direct the series of
ultrasonic pulse signals into and along the propagation path in the
sample.
41. A system according to claim 35, wherein the sample linearly
reflects the distorted ultrasonic signal produced by the distortion
of the initial ultrasonic signal along the propagation path and
caused by the sample.
42. A method of imaging a biological sample, comprising the steps
of: generating a transmit ultrasonic signal, said transmit signal
being at a fundamental frequency and having negligible energy in
the second harmonic bandwidth of the fundamental frequency;
directing the transmit ultrasonic signal into and along a
propagation path in the sample, wherein the sample causes finite,
non-linear amplitude distortion of the transmit signal along the
propagation path and thereby produces a distorted ultrasonic signal
comprised of a first order component signal and higher order
harmonic component signals at a first and higher order harmonic
frequencies respectively, and further wherein the sample also
reflects the distorted ultrasonic signal including the first order
and the higher order harmonic component signals; receiving the
higher order harmonic components of the reflected distorted
ultrasonic signal produced by the distortion of the input
ultrasonic signal along the propagation path and caused by said
sample; forming an image principally from one of said received
higher order harmonic component signals of the reflected distorted
ultrasonic signal; and displaying said formed image.
43. A method according to claim 42, wherein: the generating step
includes the step of using a phased array transducer-receiver unit
to generate the transmit signal; and the directing step includes
the steps of i) using the transducer-receiver unit to focus the
transmit signal on a focal point in the sample, and ii) using
electrical circuitry in the transducer-receiver unit to move the
focal point around the sample.
44. A method according to claim 42, wherein said received higher
order harmonic components include a second order harmonic component
signal, and the forming step includes the step of forming the image
principally from the received second order harmonic component
signal.
45. A system for imaging a biological sample, comprising: means for
generating a transmit ultrasonic signal, said transmit signal being
at a fundamental frequency and having negligible energy in the
second harmonic bandwidth of the fundamental frequency; means for
directing the transmit signal into and along a propagation path in
the sample, wherein the sample causes finite, non-linear amplitude
distortion of the transmit ultrasonic signal along the propagation
path, and said distortion thereby produces a distorted ultrasonic
signal comprised of a first order component signal and higher order
harmonic component signals at a first and higher order harmonic
frequencies respectively, and wherein the sample also reflects the
distorted ultrasonic signal including the first order and the
higher order harmonic component signals; means for receiving the
higher order harmonic components of the reflected distorted
ultrasonic signal produced by the distortion of the initial
ultrasonic signal along the propagation path and caused by said
sample; means for forming an image principally from one of said
received higher order harmonic component signals; and means for
displaying said formed image.
46. A system according to claim 45, wherein the generating means
includes a phased array transducer-receiver unit to generate the
transmit signal.
47. A system according to claim 46, wherein: the
transducer-receiver unit focuses the transmit signal on a focal
point in the sample; and the directing means includes electrical
circuitry in the transducer-receiver unit to move the focal point
around the sample.
48. A system according to claim 45, wherein said received higher
order harmonic components include a second order harmonic component
signal, and the forming means includes means for forming the image
principally from the received second order harmonic component
signal.
Description
[0001] This is a divisional application of U.S. patent application
Ser. No. 08/943,546, filed Oct. 3, 1997 and entitled "ULTRASONIC
DIAGNOSTIC IMAGING OF RESPONSE FREQUENCY DIFFERING FROM TRANSMIT
FREQUENCY" which claims the benefit of U.S. Provisional Application
No. 60/032,771 filed Nov. 26, 1996.
[0002] This invention relates to ultrasonic diagnosis and imaging
of the body and, in particular, to new methods and apparatus for
ultrasonically imaging with a response frequency which differs from
the transmitted frequency.
[0003] Ultrasonic diagnostic imaging systems have been used to
image the body with the enhancement of ultrasonic contrast agents.
Contrast agents are substances which are biocompatible and exhibit
uniquely chosen acoustic properties which return readily
identifiable echo signals in response to insonification. Contrast
agents can have several properties which enables them to enhance an
ultrasonic image. One is the nonlinear characteristics of many
contrast agents. Agents have been produced which, when insonified
by an ultrasonic wave at one frequency, will exhibit resonance
modes which return energy at other frequencies, in particular,
harmonic frequencies. A harmonic contrast agent, when insonified at
a fundamental frequency, will return echoes at the second, third,
fourth, and higher harmonics of that frequency.
[0004] It has been known for some time that tissue and fluids also
have inherent nonlinear properties. Tissue and fluids will, even in
the absence of a contrast agent, develop and return their own
non-fundamental frequency echo response signals, including signals
at harmonics of the fundamental. Muir and Carstensen explored these
properties of water beginning in 1980, and Starritt et al. looked
at these properties in human calf muscle and excised bovine
liver.
[0005] While these non-fundamental frequency echo components of
tissue and fluids are generally not as great in amplitude as the
harmonic components returned by harmonic contrast agents, they do
exhibit a number of characteristics which may be advantageously
used in ultrasonic imaging. One of us (M. Averkiou) has done
extensive research into these properties in work described in his
doctoral dissertation. In this exposition and other research, the
present inventors have seen that the main lobe of a harmonic beam
is narrower than that of its fundamental, which they have found has
implications for clutter reduction when imaging through narrow
orifices such as the ribs. They have seen that the sidelobe levels
of a harmonic beam are lower than the corresponding sidelobe levels
of the fundamental beam, which they have found has implications for
off-axis clutter reduction. They have also seen that harmonic
returns from the near field are also relatively less than returning
energy at the fundamental frequency, which they have found has
implications for near field clutter rejection. As will be seen,
these properties may be exploited in the methods and constructed
embodiments of the present invention.
[0006] In accordance with the principles of the present invention,
an ultrasonic imaging system and method are provided for imaging
tissue and fluids from response frequencies which differ from the
transmitted frequency, in particular echoes returned from the
tissue or fluids at a harmonic of a transmitted fundamental
frequency. The imaging system comprises a means for transmitting an
ultrasonic wave at a fundamental frequency, means for receiving
echoes at a harmonic frequency, and an image processor for
producing an ultrasonic image from the harmonic frequency
echoes.
[0007] In a preferred embodiment of the present invention the
transmitting and receiving means comprise a single ultrasonic
probe. In accordance with a further aspect of the present
invention, the probe utilizes a broadband ultrasonic transducer for
both transmission and reception.
[0008] In accordance with yet another aspect of the present
invention, partially decorrelated components of received harmonic
echoes are produced and utilized to remove artifacts from the
harmonic image, providing clearly defined images of tissue
boundaries such as that of the endocardium. In a preferred
embodiment the partially decorrelated components are produced by
processing the harmonic echoes through different passbands.
[0009] The methods of the present invention include the use of
harmonic echoes to reduce near-field or multipath clutter in an
ultrasonic image, such as that produced when imaging through a
narrow acoustic window such as the ribs. In accordance with yet a
further aspect of the present invention, harmonic and fundamental
echoes are blended in a common image to reduce clutter, image at
appreciable depths, and overcome the effects of depth-dependent
attenuation.
[0010] In the drawings:
[0011] FIG. 1 illustrates in block diagram form an ultrasonic
diagnostic imaging system constructed in accordance with the
principles of the present invention;
[0012] FIGS. 2, 3, 4, and 5 illustrate certain properties of
harmonic echoes which may be advantageously applied to ultrasonic
imaging applications; and
[0013] FIGS. 6 and 7 illustrate passband characteristics used to
explain the performance of the embodiment of FIG. 1;
[0014] FIG. 8 illustrates typical fundamental and harmonic
frequency passbands of an embodiment of the present invention;
[0015] FIG. 9 illustrates an FIR filter structure suitable for use
in the embodiment of FIG. 1;
[0016] FIG. 10 illustrates in block diagram form a portion of a
preferred embodiment of the present invention;
[0017] FIG. 11 illustrates the operation of the normalization
stages of the embodiment of FIG. 10;
[0018] FIG. 12 is a block diagram of one of the multiplier
accumulators used in the filters of the embodiment of FIG. 10;
[0019] FIG. 13 illustrates typical fundamental and harmonic
frequency passbands of the embodiment of FIG. 10;
[0020] FIG. 14 illustrates the blending of fundamental and harmonic
signal components into one ultrasonic image; and
[0021] FIG. 15 illustrates the passbands of a time varying filter
used in the formation of blended images.
[0022] Referring first to FIG. 1, an ultrasonic diagnostic imaging
system constructed in accordance with the principles of the present
invention is shown in block diagram form. A central controller 120
commands a transmit frequency control 117 to transmit a desired
transmit frequency band. The parameters of the transmit frequency
band, f.sub.tr, are coupled to the transmit frequency control 117,
which causes the transducer 112 of ultrasonic probe 110 to transmit
ultrasonic waves in the fundamental frequency band. In a
constructed embodiment a band of frequencies located about a
central frequency of 1.67 MHz is transmitted. This is lower than
conventional transmitted imaging frequencies, which generally range
from 2.5 MHz and above. However, use of a typical transmit
frequency of 3 or 5 MHz will produce harmonics at 6 and 10 MHz.
Since higher frequencies are more greatly attenuated by passage
through the body than lower frequencies, these higher frequency
harmonics will experience significant attenuation as they return to
the probe. This reduces the depth of penetration and image quality
at greater imaging depths, although the harmonic signals, created
as they are during the propagation of the transmitted wave through
tissue, do not experience the attenuation of a full round trip from
the transducer as the fundamental signals do. To overcome this
problem, the central transmit frequency in the illustrated
embodiment is below 5 MHz, and preferably below 2.5 MHz, thereby
producing lower frequency harmonics that are less susceptible to
depth dependent attenuation and enabling harmonic imaging at
greater depths. A transmitted fundamental frequency of 1.67 MHz
will produce second harmonic return signals at 3.34 MHz in the
illustrated embodiment. It will be understood, of course, that any
ultrasonic frequency may be used, with due consideration of the
desired depth of penetration and the sensitivity of the transducer
and ultrasound system.
[0023] The array transducer 112 of the probe 110 transmits
ultrasonic energy and receives echoes returned in response to this
transmission. The response characteristic of the transducer can
exhibit two passbands, one around the fundamental transmit
frequency and another about a harmonic frequency in the received
passband. For harmonic imaging, a broadband transducer having a
passband encompassing both the transmitted fundamental and received
harmonic passbands is preferred. The transducer may be manufactured
and tuned to exhibit a response characteristic as shown in FIG. 6,
in which the lower hump 60 of the response characteristic is
centered about the transmitted fundamental frequency f.sub.t, and
the upper hump 62 is centered about the received harmonic frequency
f.sub.r of the response passband. The transducer response
characteristic of FIG. 7 is preferred, however, as the single
dominant characteristic 64 allows the probe to be suitable for both
harmonic imaging and conventional broadband imaging. The
characteristic 64 encompasses the transmitted fundamental frequency
f.sub.t, and also the harmonic receive passband bounded between
frequencies f.sub.L and f.sub.c, and centered about frequency
f.sub.r. As discussed above, a low fundamental transmit frequency
of 1.67 MHz will result in harmonic returning echo signals at a
frequency of 3.34 MHz. A response characteristic 64 of
approximately 2 MHz would be suitable for these fundamental and
harmonic frequencies.
[0024] Tissue and cells in the body alter the transmitted
fundamental frequency signals during propagation and the returned
echoes contain harmonic components of the originally transmitted
fundamental frequency. In FIG. 1 these echoes are received by the
transducer array 112, coupled through the T/R switch 114 and
digitized by analog to digital converters 115. The sampling
frequency f.sub.s of the A/D converters 115 is controlled by the
central controller. The desired sampling rate dictated by sampling
theory is at least twice the highest frequency f.sub.c of the
received passband and, for the preceding exemplary frequencies,
might be on the order of at least 8 MHz. Sampling rates higher than
the minimum requirement are also desirable.
[0025] The echo signal samples from the individual transducer
elements are delayed and summed by a beamformer 116 to form
coherent echo signals. The digital coherent echo signals are then
filtered by a digital filter 118. In this embodiment, the transmit
frequency f.sub.t is not tied to the receiver, and hence the
receiver is free to receive a band of frequencies which is
different from the transmitted band. The digital filter 118
bandpass filters the signals in the passband bounded by frequencies
f.sub.L and f.sub.c in FIG. 7, and can also shift the frequency
band to a lower or baseband frequency range. The digital filter
could be a filter with a 1 MHz passband and a center frequency of
3.34 MHz in the above example. A preferred digital filter is a
series of multipliers 70-73 and accumulators 80-83 as shown in FIG.
9. This arrangement is controlled by the central controller 120,
which provides multiplier weights and decimation control which
control the characteristics of the digital filter. Preferably the
arrangement is controlled to operate as a finite impulse response
(FIR) filter, and performs both filtering and decimation. For
example, only the first stage output 1 could be controlled to
operate as a four tap FIR filter with a 4:1 decimation rate.
Temporally discrete echo samples S are applied to the multiplier 70
of the first stage. As the samples S are applied, they are
multiplied by weights provided by the central controller 120. Each
of these products is stored in the accumulator 80 until four such
products have been accumulated (added). An output signal is then
produced at the first stage output 1. The output signal has been
filtered by a four tap FIR filter since the accumulated total
comprises four weighted samples. Since the time of four samples is
required to accumulate the output signal, a 4:1 decimation rate is
achieved. One output signal is produced for every four input
samples. The accumulator is cleared and the process repeats. It is
seen that the higher the decimation rate (the longer the interval
between output signals), the greater can be the effective tap
number of the filter.
[0026] Alternatively, temporally separate samples are delayed by
delay elements .tau. and applied to the four multipliers 70-73,
multiplied, and accumulated in the accumulators 80-83. After each
accumulator has accumulated two products, the four output signals
are combined as a single output signal. This means that the filter
is operating as an eight tap filter with a 2:1 decimation rate.
With no decimation, the arrangement can be operated as a four tap
FIR filter. The filter can also be operated by applying echo
signals to all multipliers simultaneously and selectively time
sequencing the weighting coefficients. A whole range of filter
characteristics are possible through programming of the weighting
and decimation rates of the filter, under control of the central
controller. The use of a digital filter provides the advantage of
being quickly and easily changed to provide a different filter
characteristic. A digital filter can be programmed to pass received
fundamental frequencies at one moment, and harmonic frequencies at
the next. The digital filter can thus be operated to alternately
produce images or lines of fundamental and harmonic digital
signals, or lines of different alternating harmonics in a
time-interleaved sequence simply by changing the filter
coefficients during signal processing.
[0027] Returning to FIG. 1, to image just a non-fundamental
frequency, the digital filter 118 is controlled by the central
controller 120 to pass echo signals at a harmonic frequency for
processing, to the exclusion of the fundamental frequency. The
harmonic echo signals from the tissue are detected and processed by
either a B mode processor 37 or a contrast signal detector 128 for
display as a two dimensional ultrasonic image on the display
50.
[0028] The filtered echo signals from the digital filter 118 are
also coupled to a Doppler processor 130 for conventional Doppler
processing to produce velocity and power Doppler signals. The
outputs of these processors are coupled to a 3D image rendering
processor 162 for the rendering of three dimensional images, which
are stored in a 3D image memory 164. Three dimensional rendering
may be performed as described in U.S. Pat. No. 5,720,291, and in
U.S. Pats. 5,474,073 and 5,485,842, the latter two patents
illustrating three dimensional power Doppler ultrasonic imaging
techniques. The signals from the contrast signal detector 128, the
processors 37 and 130, and the three dimensional image signals are
coupled to a video processor 140 where they may be selected for two
or three dimensional display on an image display 50 as dictated by
user selection.
[0029] It has been found that harmonic imaging of tissue and blood
can reduce near field clutter in the ultrasonic image. It is
believed that the harmonic response effect in tissue is dependent
upon the energy level of the transmitted waves. Near to an array
transducer which is focused at a greater depth, transmitted wave
components are unfocused and of insufficient energy to stimulate a
detectable harmonic response in the near field tissue. But as the
transmitted wave continues to penetrate the body, the higher
intensity energy will give rise to the harmonic effect as the wave
components begin to focus. While both near and far field regions
will return a fundamental frequency response, clutter from these
signals is eliminated by the passband of the digital filter 118,
which is set to the harmonic frequency band. The harmonic response
from the tissue is then detected and displayed, while the clutter
from the near field fundamental response is eliminated from the
displayed image.
[0030] FIGS. 2, 3, 4, and 5 illustrate some of the properties of
harmonic return signals which can be utilized to advantage in
ultrasonic imaging. It should be appreciated that several of these
properties and their interactions are not yet fully and commonly
understood among the scientific community, and are still the
subject of research and discussion. FIG. 2 illustrates the spatial
response, and specifically the main lobe and sidelobes, of
fundamental and harmonic signals received by a transducer array
112. In this illustration the array is directed to image an area of
the body behind the ribs, such as the heart, and the main lobe is
seen to extend between ribs 10 and 10'. Overlying the ribs is a
tissue interface 12, as from a layer of fat between the skin and
ribs. The FIGURE shows a main lobe of the fundamental signals FL1,
and on either side of the main lobe are sidelobes FL2 and FL3. The
FIGURE also shows the main lobe HL1 of a harmonic of the
fundamental frequency, and sidelobes HL2 and HL3 of the harmonic
main lobe.
[0031] In this example it is seen that the main lobe of the
fundamental echoes is wide enough to encompass portions of the ribs
10,10'. Accordingly, acoustic energy at the fundamental can be
reflected back toward the transducer 112 as indicated by the arrow
9. While some of the energy of this reflection may travel back to
and be received directly by the transducer, in this example some of
the reflected energy is reflected a second time by the tissue
interface 12, as indicated by arrow 9'. This second reflection of
energy reaches the other rib 10', where it is reflected a second
time as shown by arrow 9'' and travels back to and is received by
the transducer 112.
[0032] Since the intent of this imaging procedure is to image the
heart behind the ribs, these echoes reflected by the ribs are
unwanted artifacts which contaminate the ultrasonic image. Unwanted
echoes which are reflected a number of times before reaching the
transducer, such as those following the paths of arrows 9,9',9'',
are referred to as multipath artifacts. Together, these artifacts
are referred to as image "clutter", which clouds the near field and
in some cases all of the image. This near field haze or clutter can
obscure structure which may be of interest near the transducer.
Moreover, the multipath artifacts can be reproduced in the image at
greater depths due to the lengthy multiple paths traveled by these
artifacts, and can clutter and obscure regions of interest at
greater depths of field.
[0033] But when only the harmonic return signals are used to
produce the ultrasonic image, this clutter from the fundamental
frequencies is filtered out and eliminated. The main lobe HL1 of
the received harmonic echoes is narrower than that of the
fundamental, and in this example passes between the ribs 10,10'
without intersecting them. There are no harmonic returns from the
ribs, and no multipath artifacts from the ribs. Thus, the harmonic
image will be distinctly less cluttered and hazy than the
fundamental image, particularly in the near field in this
example.
[0034] FIG. 3 shows a second example in which the main lobes of
both the fundamental and harmonic returns do not intersect the
ribs, and the problem discussed in FIG. 2 does not arise. But in
this example the ribs 10, 10' are closer to the skin surface and
the transducer 112. While the main lobes do not intersect the ribs,
the sidelobes FL2 of the fundamental do reach the ribs, allowing
sidelobe energy to be reflected back to the transducer as shown by
reflection path 9. Again, this will produce clutter in the
fundamental image. But the smaller and narrower sidelobes HL2 of
the received harmonic energy do not reach the ribs. Again, the
harmonic image will exhibit reduced clutter as compared to the
fundamental image.
[0035] FIG. 4 illustrates the fundamental and harmonic beam
patterns in a perspective which is across the lobes of FIGS. 2 and
3, that is, across the axis of the transducer. This drawing
illustrates the relative amplitude responses of the fundamental and
second harmonic beam patterns. Illustrated are the dynamic response
DRF between the main (FL1) and first sidelobe (FL2) of the
fundamental component of the sound beam, and the dynamic response
DRH between the main (HL1) and first sidelobe (HL2) of the second
harmonic component. If responses due to the main lobes are
considered desired signal responses, and responses due to the
sidelobes are considered to be clutter or noise, the signal to
noise ratio of the harmonic is greater than that of the
fundamental. That is, there is relatively less sidelobe clutter in
a harmonic image than in the corresponding fundamental image of the
same transmission, or DRH>DRF.
[0036] FIG. 5 illustrates another comparison of the properties of
fundamental and harmonic signals, which is the relative amount of
energy (in units of acoustic pressure P) emanating from increasing
depths Z in the body at the fundamental and second harmonic
frequencies. The curve denoted Fund. shows the buildup of
propagated acoustic energy at the fundamental frequency. While the
curve is seen to peak at the focus of the array transducer, it is
seen that there is nonetheless an appreciable amount of fundamental
energy at the shallower depths before the focal region. In
comparison, there is comparatively much less energy, and a lesser
buildup of energy, at the harmonic frequency propagated at these
lesser depths of field. Hence, with less energy available for
multipath reverberation and other aberrations, there is less near
field clutter with harmonic imaging than with imaging the
fundamental echo returns from the same transmission.
[0037] FIG. 8 illustrates the bands of received signals and the
digital filter of a typical FIG. 1 embodiment of the present
invention for a transmitted signal of four cycles of a 1.67 MHz
acoustic wave. Transmitting multiple cycles narrows the bandwidth
of the transmitted signal; the greater the number of cycles, the
narrower the bandwidth. In response to this transmission, the
transducer 112 receives a fundamental signal in a bandwidth 90,
which is seen to peak at the transmitted frequency of 1.67 MHz. As
the fundamental frequency band rolls off, the harmonic band 92
comes up, and is seen to exhibit a peak return at the harmonic
frequency of 3.34 MHz. The received signals are applied to a
digital filter with a passband characteristic 94, which is seen to
be centered around the harmonic frequency of 3.34 MHz. As FIG. 8
shows, this passband will substantially suppress signals at the
fundamental frequency while passing the harmonic signals on to
further processing and image formation. When imaging the heart in
this manner, it has been found that the harmonic response of the
endocardial tissue of the heart is quite substantial, and harmonic
tissue images of the heart show a clearly defined endocardial
border.
[0038] Other signal processing techniques besides filtering may be
used to separate out harmonic signals from received echo
information such as cancellation of the fundamental frequencies in
a broadband signal, leaving only the harmonic frequencies. For
example, U.S. Pat. No. 5,706,819 discloses a two pulse technique,
whereby each scanline is insonified by consecutive fundamental
frequency pulses of opposite phase in rapid succession. When the
resultant echoes are received from the two pulses and combined on a
spatial basis, the fundamental frequencies will cancel and the
nonlinear or harmonic frequencies will remain. Thus, the harmonic
frequencies are separated from the broadband echo signals without
the need for a filter circuit.
[0039] FIG. 10 shows a portion of a preferred embodiment of the
present invention in block diagram form, from the beamformer output
through to the image display. This embodiment not only produces
harmonic images of tissue and blood flow, but also overcomes signal
dropout deficiencies of conventional imaging systems which arise
when imaging patients with difficult to image pathology.
Additionally, this embodiment reduces an artifact of coherent
ultrasound images known as speckle. In FIG. 10, the signal and data
lines connecting the blocks of the block diagram all represent
multi-conductor digital data paths, as the processor of the
illustrated embodiment is entirely digital. Scanline echo data from
the beamformer 116 is applied in parallel to the two channels
30a,30b of the processor illustrated in FIG. 10, one of which is a
high frequency channel and the other of which is a low frequency
channel. Each channel of the processor has a normalization stage
32,132 which multiplies the scanline data by a scale factor on a
sample by sample basis to produce gain or attenuation that can vary
with the depth of the body from which each sample returned. The
scale factor for each channel is provided by normalization
coefficients stored in or generated by coefficient circuits 32,132,
which in a preferred embodiment are digital memories. As the
multiplying coefficients are changed along the sequence of scanline
echoes, depth dependent gain or attenuation is produced.
[0040] The function of the normalization stages is two-fold. One is
to compensate for a transducer aperture which expands with depth of
scan. As signals from an increasing number of transducers are used
with increasing depth, the magnitude of the summed beamformed
signals will increase. This increase is offset by reduced gain
(increased attenuation) in the normalization stage, in proportion
to the rate at which channels are added to the beamforming process,
so that the resultant echo sequence will be unaffected by the
changing aperture.
[0041] The second function of the normalization stages is to
equalize the nominal signal amplitudes of the two channels 30a,30b.
The nominal signal amplitudes of the passbands of the two channels
are desirably the same, so that the original relative signal levels
will be preserved after the passbands are summed to create the full
harmonic passband. But ultrasound signals are subject to depth
dependent attenuation which varies with frequency, higher
frequencies being more greatly attenuated with depth than lower
frequencies. To account for this depth dependent attenuation the
coefficients for the normalization stages provide signal gain which
increases with depth. Since the two channels employ different
frequency passbands, the depth dependent gain of the two channels
differs from one channel to the other. In particular, the rate of
gain increase for the higher frequency passband channel is greater
than that of the lower frequency passband channel. This is
illustrated in FIG. 11, which, for purposes of illustration, shows
the normalization gain characteristic of the higher frequency
passband channel separated into two components. The depth dependent
characteristic 200 offsets the effect of an increasing aperture in
the channel, and the depth dependent characteristic 202 compensates
for depth dependent signal attenuation. The low frequency passband
channel may also have a depth dependent gain characteristic but
with a different characteristic 202 for the different rate of
attenuation of the lower frequencies. The high frequency passband
channel has a similar but more rapidly increasing depth dependent
gain characteristic to account for the more rapid rate of
attenuation of the higher frequencies. Each depth dependent gain
characteristic 202 is chosen to offset the effect of depth
dependent gain for the particular frequency passband employed by
that channel.
[0042] In a preferred embodiment the coefficients of the
coefficient circuits apply a gain or attenuation characteristic
which is a combination of the two characteristics 200,202.
Preferably, the coefficient memories 32,132 store multiple combined
gain curves which are changed with memory addressing to match
scanhead characteristics or the type of signals being processed (2D
or Doppler). The rate of gain change may be controlled by the rate
at which the coefficients are changed for the multiplier of each
normalization stage 30,130.
[0043] The normalized echo signals in each channel are coupled to
quadrature bandpass filters (QBPs) in each channel. The quadrature
bandpass filters provide three functions: band limiting the RF
scanline data, producing in-phase and quadrature pairs of scanline
data, and decimating the digital sample rate. Each QBP comprises
two separate filters, one producing in-phase samples (I) and the
other producing quadrature samples (Q), with each filter being
formed by a plurality of multiplier-accumulators (MACs)
implementing an FIR filter. One such MAC is shown in FIG. 12. As an
echo sample of the scanline data is applied to one input of a
digital multiplier 210 a coefficient is applied to the other
multiplier input. The product of the echo sample and the weighting
coefficient is stored in an accumulator 212 where it may be
accumulated with previous products. Other MACs receive the echo
samples at different phases and likewise accumulate weighted echo
samples. The accumulated outputs of several MACs can be combined,
and the final accumulated product comprises filtered echo data. The
rate at which accumulated outputs are taken sets the decimation
rate of the filter. The length of the filter is a product of the
decimation rate and the number of MACs used to form the filter,
which determine the number of incoming echo samples used to produce
the accumulated output signal. The filter characteristic is
determined by the values of the multiplying coefficients. Different
sets of coefficients for different filter functions are stored in
coefficient memories 38,138, which are coupled to apply selected
coefficients to the multipliers of the MACs. The MACs effectively
convolve the received echo signals with sine and cosine
representative coefficients, producing output samples which are in
a quadrature relationship.
[0044] The coefficients for the MACs which form the I filter
implement a sine function, while the coefficients for the Q filter
implement a cosine function. For bandpass filtering, the
coefficients of the active QBPs additionally implement a low pass
filter function that is frequency shifted to form, in combination
with the sine (for I) and cosine (for Q) functions, a bandpass
filter for the quadrature samples. In the instant example,
QBP.sub.1 in channel 30a is producing I and Q samples of the
scanline data in a first, low frequency passband, and QBP.sub.2 in
channel 30b is producing I and Q samples of the scanline data in a
second, higher frequency passband. Thus, the spectrum of the
original broadband echo signals is divided into a high frequency
band and a low frequency band. To complete the dropout and speckle
reduction process, the echo data in the passband produced by
QBP.sub.1 of channel 30a is detected by a detector 40.sub.1 and the
detected signals are coupled to one input of a summer 48. In a
preferred embodiment detection is performed digitally by
implementing the algorithm (I.sup.2+Q.sup.2).sup.1/2. The echo data
in the complementary passband produced by QBP.sub.2 of channel 30b
is detected by a detector 40.sub.2 and these detected signals are
coupled to a second input of the summer 48. When the signals of the
two passbands are combined by the summer 48, the decorrelated
signal dropout and speckle effects of the two passbands will at
least partially cancel, reducing the signal dropout and speckle
artifacts in the 2D image created from the signals.
[0045] Following the detector in each subchannel is a gain stage
formed by multipliers 44.sub.1,44.sub.2 which receive weighting
coefficients from coefficient memories 42.sub.1,42.sub.2. The
purpose of this gain stage is to partition the balance of analog
and digital gains in the ultrasound system for optimal system
performance. Some of the gains in the echo signal path may be
automatically implemented by the ultrasound system, while others,
such as manual gain control and TGC gain, may be controlled by the
user. The system partitions these gains so that the analog gains
preceding the ADCs (analog to digital converters) of the beamformer
are adjusted optimally for the dynamic input range of the ADCs. The
digital gain is adjusted to optimize the brightness of the image.
The two gains together implement gain control changes effected by
the user.
[0046] In the preferred embodiment the gain imparted to the
scanline signals by the multipliers 44.sub.1,44.sub.2 is selected
in concert with the gain of the preceding normalization stage
34,134 in the channel. The gain of each normalization stage is
chosen to prevent the attainment of saturation levels in the QBPs,
as may occur when strong signals from contrast agents or harmonic
imaging are being received. To prevent saturation levels the
maximum gain of the normalization stage is controlled, and any
reduction imposed by reason of this control is restored by the gain
of the succeeding multiplier 44.sub.1,44.sub.2.
[0047] The gain function provided by these multipliers could be
performed anywhere along the digital signal processing path. It
could be implemented by changing the slope of the compression
curves discussed below. It could also, for instance, be performed
in conjunction with the gains applied by the normalization stages.
This latter implementation, however, would eliminate the ability to
effect the saturation control discussed above. The present
inventors have found implementation of this gain function to be
eased when provided after detection, and in the preferred
embodiment by use of a multiplier after detection.
[0048] The signals produced by the gain stages 44.sub.1,44.sub.2
generally exhibit a greater dynamic range than may be accommodated
by the display 50. Consequently, the scanline signals of the
multipliers are compressed to a suitable dynamic range by lookup
tables. Generally the compression is logarithmic, as indicated by
log compression processors 46.sub.1,46.sub.2. The output of each
lookup table is proportional to the log of the signal input value.
These lookup table are programmable so as to provide the ability to
vary the compression curves, and the brightness and dynamic range
of the scanline signals sent on for display.
[0049] The present inventors have found that the use of log
compression to scale the echo signals can affect low level signals
near the baseline (black) level of the signal dynamic range by
exacerbating the degree and the number of echoes with components at
the black level, a manifestation of the destructive interference
arising from the speckle effect of the coherent ultrasonic energy.
When the echo signals are displayed, many of them will be at the
black level, and appear in the image to have been undetected or
dropped out. The embodiment of FIG. 10 reduces this problem by
producing separate, partially decorrelated versions of the echo
signals in the two channels 30a,30b. This embodiment partially
decorrelates the echo signal versions by separating the echo signal
components into two different passbands as shown in FIG. 13. The
two passbands can be completely separated or, as shown in this
example, overlapping. In this example, the lower passband 300a is
centered about a frequency of 3.1 MHz, and the higher passband 300b
is centered about a frequency of 3.3 MHz, a center frequency
separation of only 200 kHz. Even this small degree of separation
has been found sufficient to decorrelate the signal components of
the two passbands sufficiently such that black level signal dropout
in one passband will frequently not align in frequency with its
corresponding component in the other passband. Consequently, when
these decorrelated replicas of the same echo signal are combined by
the summer 48, the signal dropout and speckle artifacts will be
markedly reduced. This is especially significant when trying to
image fine structures at deep depths in the body, such as the
endocardium. A harmonic image of the endocardium is significantly
improved by the artifact elimination effects of the embodiment of
FIG. 10.
[0050] As discussed previously the signal gain of the two passbands
300a,300b of FIG. 13 can be matched to preserve the original signal
levels after summation. However, in a preferred embodiment, the
lower frequency passband is processed with less dynamic range than
the higher frequency passband as shown in FIG. 13. This has the
effect of suppressing the fundamental frequency contributions of
the lower frequency passband (which contains more fundamental
frequency components than the higher frequency band.) This is
accomplished as a component of different compression
characteristics in the log compression processors
46.sub.1,46.sub.2, or elsewhere in the channels 30a,30b subsequent
to the separation of the broadband signal into separate
passbands.
[0051] The processed echo signals at the output of the summer 48
are coupled to a lowpass filter 52. This lowpass filter, like the
QBPs, is formed by combinations of multiplier-accumulators with
variable coefficients, arranged to implement an FIR filter, to
control the filter characteristic. The lowpass filter provides two
functions. One is to eliminate sampling frequency and other
unwanted high frequency components from the processed echo signals.
A second function is to match the scanline data rate to the
vertical line density of the display 50, so as to prevent aliasing
in the displayed image. The FIR filter performs this function by
selectively decimating or interpolating the scanline data. The
filtered echo signals are then stored in an image memory 54. If the
scanlines have not yet been scan converted, that is, they have
r,.theta. coordinates, the scanlines are scan converted to
rectilinear coordinates by a scan converter and greyscale mapping
processor 56. If scan conversion has been performed earlier in the
process, or is not needed for the image data, the processor 56 may
simply convert the echo data to the desired greyscale map by a
lookup table process. The image data may then be stored in a final
image memory or sent to a video display driver (not shown) for
conversion to display signals suitable for driving the display
50.
[0052] It will be appreciated that, due to the advantage of the
quick programmability of a digital filter, the processing described
above can be performed in an embodiment which utilizes a single one
of the channels 30a, 30b to process the echo data from a scanline
twice to alternately produce a line of signals for each of the two
passbands in a time-interleaved fashion. However, the use of two
parallel channels affords twice the processing speed, enabling
harmonic images to be produced in real time and at twice the frame
rate of a time multiplexed embodiment.
[0053] Harmonic images produced from high frequency signals can
suffer from depth dependent attenuation as the echo signals return
from increasing depths in the body. Lower frequency fundamental
signals may experience less attenuation, and hence in some cases
may exhibit better signal to noise ratios at greater depths. The
embodiment of FIG. 14 takes advantage of this characteristic by
blending fundamental and harmonic image data in one image. It is
possible, for instance, to create a normal tissue image of the
heart from fundamental frequencies, and overlay the fundamental
frequency tissue image with a harmonic tissue image of the heart to
better define the endocardial border in the composite image. The
two images, one from fundamental frequency components and another
from harmonic frequency components, may be formed by alternately
switching the digital filter 118 between fundamental and harmonic
frequencies to separately assemble fundamental and harmonic images,
or by employing the two parallel filters of FIG. 10 with two
passbands, one set to pass fundamental frequencies and the other
set to pass harmonic frequencies. In FIG. 14, the filter of channel
30a is set to pass fundamental signal frequencies, and echo signals
passed by this channel are stored in a fundamental image memory
182. Correspondingly, harmonic signal frequencies are passed by
channel 30b and stored in a harmonic image memory. The fundamental
and harmonic images are then blended together by a proportionate
combiner 190, under control of a blend control 192. The blend
control 192 may automatically implement a pre-programmed blending
algorithm, or one directed by the user. For example, the
proportionate combiner 190 may create a blended image which uses
only echo data from the harmonic image at shallow depths, then
combines echo data from both images at intermediate depths, and
finally only uses echo data of the fundamental image at deep
depths. This combines the reduced clutter benefit of harmonic echo
data at shallow depths and the greater penetration and signal to
noise ratio of fundamental echoes received from deeper depths,
while affording a smooth transition from one type of data to the
other at intermediate depths. Other combining algorithms are also
possible, such as simply switching from one type of data to another
at a predetermined depth, or outlining a region of the image to be
displayed with one type of data while the remainder of the image is
displayed using the other type of data.
[0054] It is also possible to employ the two parallel filters and
blend the components together before image formation, thereby
adding a controllable component of the harmonic echo signals to the
fundamental frequency signals to enhance the resultant image. Such
an embodiment could eliminate the need for separate fundamental and
harmonic image memories and would process the signal components
directly to a blended image memory.
[0055] A third technique for producing blended images is to receive
each scanline of the image through a depth-dependent, time varying
filter. Such filters are well known for improving the signal to
noise ratio of received echo signals in the presence of depth
dependent attenuation as shown, for instance, in U.S. Pat. No.
4,016,750. For the production of blended fundamental and harmonic
images, the passband 210 of a time varying filter is initially set
to pass harmonic frequencies f.sub.h, as shown in FIG. 15, as echo
signals begin to be received from shallow depths. When it becomes
desirable to begin supplementing the image with fundamental signal
components at deeper depths, the passband 210 undergoes a
transition to lower frequencies, eventually moving to the
fundamental frequencies f.sub.f as shown by passband 212 in FIG.
15. In the case of a digital filter such as that shown in FIG. 9,
the change in passband frequencies is effected by changing the
filter coefficients with time. As the filter undergoes this
transition, the passband passes fewer harmonic frequencies and
greater fundamental frequencies until eventually, if desired, the
passband is passing only fundamental frequencies at the maximum
image depth. By receiving each scanline through such a time varying
filter, each line in the resultant image can comprise harmonic
frequencies in the near field (shallow depths), fundamental
frequencies in the far field (deepest depths), and a blend of the
two in between.
[0056] Harmonic tissue images of moving tissue can also be formed
by processing the received harmonic tissue echo signals with the
processor described in U.S. Pat. No. 5,718,229, entitled MEDICAL
ULTRASONIC POWER MOTION IMAGING.
[0057] Thus, the present invention encompasses an ultrasonic
imaging system for imaging the nonlinear response of tissue and
fluids of the body to ultrasound by transmitting a fundamental
frequency signal, receiving an echo signal from the tissue at a
non-fundamental, preferably harmonic, frequency, detecting the
non-fundamental frequency echo signals, and forming an image of the
tissue and fluids from the non-fundamental frequency echo signals.
As used herein the term harmonic also refers to harmonic
frequencies of higher order than the second harmonic and to
subharmonics, as the principles described herein are equally
applicable to higher order and subharmonic frequencies.
* * * * *