METHOD AND APPARATUS FOR COHERENCE
FILTERING OF ULTRASOUND IMAGES
FIELD OF THE INVENTION
This invention generally relates to digital ultrasound
imaging systems and, in particular, to methods for improving
medical ultrasound images by means of data-dependent filtering.
BACKGROUND OF THE INVENTION
A conventional ultrasound imaging system comprises
an array of ultrasonic transducer elements which are used to
transmit an ultrasound beam and then receive the reflected beam
from the object being studied. Such scanning comprises a series
of measurements in which the steered ultrasonic wave is
transmitted, the system switches to receive mode after a short time
interval, and the reflected ultrasonic wave is received and stored.
Typically, transmission and reception are steered in the same
direction during each measurement to acquire data from a series
of points along an acoustic beam or scan line. The receiver is
dynamically focused at a succession of ranges along the scan line
as the reflected ultrasonic waves are received.
For ultrasound imaging, the array typically has a
multiplicity of transducer elements arranged in one or more rows
and driven with separate voltages. By selecting the time delay (or
phase) and amplitude of the applied voltages, the individual
transducer elements in a given row can be controlled to produce
ultrasonic waves which combine to form a net ultrasonic wave that
travels along a preferred vector direction and is focused at a
selected point along the beam. Multiple firings may be used to
acquire data representing the same anatomical information. The
beamforming parameters of each of the firings may be varied to
provide a change in maximum focus or otherwise change the
content of the received data for each firing, e.g., by transmitting
successive beams along the same scan line with the focal point of
each beam being shifted relative to the focal point of the previous
beam. By changing the time delay and amplitude of the applied
voltages, the beam with its focal point can be moved in a plane to
scan the object.
The same principles apply when the transducer probe
is employed to receive the reflected sound in a receive mode. The
voltages produced at the receiving transducer elements are
summed so that the net signal is indicative of the ultrasound
reflected from a single focal point in the object. As with the
transmission mode, this focused reception of the ultrasonic energy
is achieved by imparting separate time delays (and/or phase
shifts) and gains to the signal from each receiving transducer
element. The output signals of the beamformer channels are then
coherently summed to form a respective pixel intensity value for
each point of focus, corresponding to a sample volume in the
object region or volume of interest. These pixel intensity values
are log-compressed, scan-converted and then displayed as an
image of the anatomy being scanned.
Tissue types and anatomical features are most easily
differentiated in an ultrasound image when they differ in image
brightness. Image brightness on conventional medical ultrasound
imaging systems is a function of the receive beamformed signal
amplitude, i.e., after coherent summation of the delayed receive
signals on each transducer element. More precisely, the logarithm
of the beamformed signal amplitude is displayed, with user-
adjustable gain and contrast, and, if desired, a choice of a handful
of grayscale mapping tables.
A human kidney usually appears in an ultrasound
image as a darkish, ellipsoidal region (corresponding to the renal
cortex) with a bright, irregularly shaped interior (the medulla). One
criterion used by sonographers to evaluate ultrasound image
quality is the contrast (i.e., the displayed difference in brightness)
between the renal cortex and the medulla. This can be artificially
increased by adjusting the grayscale maps manually after the fact,
but this approach is of little practical value. Much more desirable
would be the identification of another tissue contrast mechanism
which could be used in addition to the receive amplitude to
distinguish tissue types.
SUMMARY OF THE INVENTION
The present invention is a method and apparatus for
improving medical ultrasound images by utilizing data-dependent
filtering. The filter increases contrast between tissue types by
distinguishing them on the basis of the degree of coherence of the
receive ultrasound signals. The method also provides some
suppression of speckle noise without significantly degrading
resolution. The method can be implemented in real-time with only
a modest change to the hardware of an existing ultrasound
imaging system. The invention can be incorporated in the
beamforming system of a digital ultrasound imaging system having
either a baseband beamformer or a pure time-delay beamformer
(also known as an RF beamformer).
In accordance with the method of the invention, a
quantity, called the coherence factor, is calculated for each pixel in
the image. The coherence factor is defined to be the ratio of two
quantities: the amplitude of the receive signals summed coherently
and the amplitude of the receive signals summed incoherently.
The coherence data is stored in buffer memory and is optionally
spatially filtered and mapped. The amplitude data is concurrently
acquired and stored in buffer memory.
The system of the invention can be selectively
operated to display the coherence information alone, the
amplitude information alone, or a combination of the coherence
and amplitude information. In accordance with the preferred
embodiment, this combination consists of multiplying, sample by
sample, the receive beamformed amplitude by the coherence
factor, and then displaying the modified amplitude conventionally,
i.e., by log-compressing and scan-converting.
BRIEF DESCRIPTION OF THE ACCOMPAYINGRAWINGS
FIG. 1 is a block diagram of an ultrasonic imaging
system which incorporates the present invention.
FIG. 2 is a block diagram showing of a receiver which
forms part of the system of FIG. 1.
FIG. 3 is a block diagram showing the receiver of FIG.
2 in more detail.
FIG. 4 is a block diagram showing the detection
processor of FIG. 2 in more detail.
FIG. 5 is a graph showing mappings for the coherence
factor C in accordance with first and second preferred
embodiments of the invention. The solid line is the default (no
mapping) and the dashed lines show two linear mappings with
thresholds.
DETAILED DESCRIPTION OF THE PREFERRED
EMBODIMENTS
As shown in FIG. 1, the ultrasonic imaging system
incorporating the invention includes a transducer array 10
comprised of a plurality of separately driven transducers 2, each of
which produces a burst of ultrasonic energy when energized by a
pulsed waveform produced by a transmitter 12. The ultrasonic
energy reflected back to transducer array 10 from the object under
study is converted to an electrical signal by each receiving
transducer 2 and applied separately to a receiver 14 through a set
of transmit/receive (T/R) switches 16. Transmitter 12, receiver 14
and switches 16 are operated under control of a digital controller
18 responsive to commands by a human operator. A complete
scan is performed by acquiring a series of echoes in which
switches 16 are set to their transmit positions, transmitter 12 is
gated ON momentarily to energize each transducer 2, switches 16
are then set to their receive positions, and the subsequent echo
signals produced by each transducer 2 are applied to receiver 14.
The separate echo signals from each transducer 2 are combined
in receiver 14 into a single echo signal which is used to produce a
line in an image on a display system 20.
Transmitter 12 drives transducer array 10 such that
the ultrasonic energy produced is directed, or steered, in a beam.
To accomplish this, transmitter 12 imparts a time delay Ti to the
respective pulsed waveforms 24 that are applied to successive
transducers 2. By adjusting the time delays Ti appropriately in a
conventional manner, the ultrasonic beam can be directed away
from axis 25 by an angle 6 and focused at a fixed range ft. A
sector scan is performed by progressively changing the time
delays Ti in successive excitations. The angle 6 is thus changed
in increments to steer the transmitted beam in a succession of
directions.
The echo signals produced by each burst of ultra-
sonic energy reflect from objects located at successive ranges
along the ultrasonic beam. The echo signals are sensed
separately by each transducer 2 and a sample of the echo signal
magnitude at a particular point in time represents the amount of
reflection occurring at a specific range. Due to the differences in
the propagation paths between a reflecting point P and each
transducer 2, however, these echo signals are not detected
simultaneously. Receiver 14 amplifies the separate echo signals,
imparts the proper time delay to each, and sums them to provide a
single echo signal which accurately indicates the total ultrasonic
energy reflected from point P located at range R along the
ultrasonic beam oriented at the angle 6.
To simultaneously sum the electrical signals pro-
duced by the echoes impinging on each transducer 2, time delays
are introduced into each separate channel 34 (see FIG. 2) of
receiver 14. The beam time delays for reception are the same
delays Ti as the transmission delays described above. However,
the time delay of each receive channel continuously changes
during reception of the echo to provide dynamic focusing of the
received beam at the range ft from which the echo signal
emanates.
Under direction of digital controller 18, receiver 14
provides delays during the scan such that steering of receiver 14
tracks the direction 9 of the beam steered by transmitter 12 and
samples the echo signals at a succession of ranges R so as to
provide the proper delays and phase shifts to dynamically focus at
points P along the beam. Thus, each emission of an ultrasonic
pulse waveform results in acquisition of a series of data points
which represent the amount of reflected sound from a
corresponding series of points P located along the ultrasonic
beam.
A scan converter 19 receives the series of data points
produced by receiver 14 and converts the data into the desired
image. More particularly, the scan converter converts the acoustic
image data from polar coordinate (R-?) sector format or Cartesian
coordinate linear array to appropriately scaled Cartesian coordi-
nate display pixel data at the video rate. This scan-converted
acoustic data is then supplied to a display monitor (not shown) of
a display system 20, which images the time-varying amplitude of
the signal envelope as a grey scale.
As shown in FIG. 2, receiver 14 comprises three
sections: a time-gain control section 26, a receive beamforming
section 28 and a mid-processor 30. Time-gain control (TGC)
section 26 includes a respective amplifier 32 for each of the
receive channels 34, and a time-gain control circuit 36 is provided
for controlling gain of amplifiers 32. The input of each amplifier 32
is coupled to a respective one of transducers 2 to amplify the echo
signal which it receives. The amount of amplification provided by
amplifiers 32 is controlled through a control line 38 driven by TGC
circuit 36, the latter being set by hand operation of potentiometers
40.
The receive beamforming section of receiver 14
includes a multiplicity of receive channels 34, each receive
channel 34 receiving the analog echo signal from a respective
amplifier 32 at a respective input 42. The analog signals are
digitized and produced as a stream of signed digitized samples.
These samples are respectively delayed in the receive channels
such that when they are summed with samples from each of the
other receive channels, the amplitude of the summed signals is a
measure of the strength of the echo signal reflected from a point P
located at range R on the steered beam ?.
As shown in FIG. 3, each receive channel 34 supplies,
in addition to the delayed signed samples, the amplitude, or
absolute value, of the delayed signed samples. As shown in FIG.
3, the delayed signed samples are provided to a coherent
summation bus 44, while the amplitudes of the delayed, signed
samples are provided to an incoherent summation bus 46.
Coherent summation bus 44 sums the delayed signed samples
from each receive channel 34 using pipeline summers 48 to
produce coherent sum A. Incoherent summation bus 46 sums the
amplitudes of the delayed signed samples from each receive
channel 34 using pipeline summers 50 to produce incoherent sum
B.
Receiver midprocessor section 30, as shown in FIG. 2,
receives the coherently summed beam samples from summers 48
via output A and receives the incoherently summed beam samples
from summers 50 via output B. Midprocessor section 30
comprises a detection processor 52, which is shown in more detail
in FIG. 4.
Detection processor 52 calculates and applies a
coherence factor C in accordance with the present invention. The
coherence factor is calculated for each pixel in the image and is
defined to be the ratio of two quantities: the amplitude of the sum
of the receive signals and the sum of the amplitudes of the receive
signals, or

where Si is the delayed signal for the i-th transducer element. This
ratio is calculated in detection processor 52, shown in FIG. 4, by
calculating the absolute value of the coherent sum A in a summer
54 and then calculating the ratio of the absolute value of the
coherent sum A to the incoherent sum B in a divider 56, i.e., C =
|A| IB.
For the case of a pure time-delay beamformer, the
signal from each channel is a real, signed quantity and the
coherent sum is the arithmetic sum of these signals. The
incoherent sum is the arithmetic sum of the absolute value of each
signal, i.e., a sum of non-negative numbers.
For the case of a baseband beamformer, the channel
signals are complex numbers / + iQ, with real part / and imaginary
part Q. The coherent sum is the sum of these complex numbers
and is also complex. The absolute value of this coherent sum is a
real, non-negative number, i.e., (i2 + Q2). This is the usual
signal which is log-compressed, scan-converted and displayed.
The incoherent sum for the baseband beamformer is the sum of
the absolute values of each (complex) channel signal, i.e., a real,
non-negative number.
Thus the coherence factor C is a real, non-negative
quantity. The minimum value of C is zero, since it is the ratio of
two non-negative numbers. The denominator in Eq. (1) can
vanish only if all the si's are zero. In this case, the numerator also
vanishes, so C is defined as zero in this case. The maximum
value of C is unity. This follows from Bessel's inequality:

where A and B are any two vectors. C equals unity only when s, is
a constant independent of /', which is when the receive signals are
perfectly coherent, i.e., identical, across the transducer array.
Spatially filtering the coherence factor can be
advantageous because — like the normal amplitude image — the
coherence factor suffers from speckle noise. The coherence
information can be spatially filtered to reduce this speckle noise
without significantly degrading the apparent resolution of the final
image in those cases (transparent overlay and modified grayscale,
described below) in which the coherence data is not displayed
independently. For example, the coherence factor can be filtered
with a simple 5x5 filter which substitutes the average of the 25
values for the center value in the 5 x 5 filter kernel. The use of
spatial filtering increases the contrast between the bright and dark
areas of the kidney, for example, and within fat and muscle layers.
In accordance with a further optional aspect of the
invention, the coherence factor can be mapped before it is
displayed or applied to the amplitude image, in order to optimize
the coherence data for particular imaging applications. For
example, the alternate mapping M1 shown in Fig. 5 will zero out
the data (C=0) when the coherence factor C falls below a
predetermined threshold. Similarly alternate mapping M2 zeros
out the data at another threshold. This can be useful in cases
where the primary diagnostic concern is identifying blood vessels
in an image.
The coherence factor C provides independent in-
formation about the tissue and can be displayed as a separate
image or as a transparent color map overlaid on the B-mode
image. Alternatively, the coherence information can be combined
with the amplitude information and displayed as a single grayscale
image. In the simplest case, this combination consists of
multiplying, sample by sample, the receive beamformed amplitude
by the coherence factor, and then displaying the modified
amplitude conventionally (by log-compressing and scan-
converting).
FIG. 4 depicts a system which can be selectively
operated to display the coherence information alone, the
amplitude information alone, or a combination of the coherence
and amplitude information. In accordance with the preferred
embodiment of the invention, the amplitude of the coherent sum,
i.e., \A\, is placed into an R-? memory buffer 58 which holds the
samples for each range R and for each scan line direction 8. The
coherence factor C, calculated as described above, is placed into
a separate R-? memory buffer 60. As mentioned above, the
coherence information may optionally be filtered and scaled. The
filtering and scaling operations are performed in buffer 60 by
applying a two-dimensional filter 62 and a coherence map 64.
The filtered and scaled coherence factor data is indicated by
output C in FIG. 4
The output signal |A| of memory 58 is supplied to the
input of a three-position switch 66. When switch 66 is set to
position 1, the input of switch 66 is coupled to a first input of a
multiplier 70. When switch 66 is set to position 2, the input of
switch 66 is not used. When switch 66 is set to position 3, the
input of switch 66 is coupled to a memory 72 which stores log-
compression look-up tables.
Similarly, the output C of memory 60 is coupled to the
input of a three-position switch 68. When switch 68 is set to
position 1, the input of switch 68 is coupled to a second input of
multiplier 70. When switch 68 is set to position 2, the input of
switch 68 is coupled to scan converter 19. When switch 68 is set
to position 3, the input of switch 68 is not used.
In a first operating mode, only the coherence data is
displayed. This is accomplished by setting both of switches 66
and 68 to position 2 so that the output signal C is supplied directly
to scan converter 19 and the scan-converted coherence data is
displayed on a linear scale by display system 20, shown in FIG. 1.
In a second operating mode, only the amplitude data is
displayed. This is accomplished by setting both of switches 66
and 68 to position 3 so that the output signal \A\ is supplied
directly to log-compression memory 72. The amplitude data is log-
compressed in memory 72 and then scan-converted by scan
converter 19 in a conventional manner. The log-compressed,
scan-converted amplitude data is then displayed by the display
system.
In a third operating mode, the product of the
coherence and amplitude data is displayed. This is accomplished
by setting both of switches 66 and 68 to position 1 so that the
output signals \A\ and C axe sent to respective inputs of multiplier
70. Multiplier 70 multiplies, sample by sample, the amplitude data
by the respective coherence factors. The modified amplitude data
is then log-compressed, scan-converted and displayed in
conventional manner.
While only certain preferred features of the invention have
been illustrated and described, many modifications and changes will
occur to those skilled in the art. It is, therefore, to be understood
that the appended claims are intended to cover all such
modifications and changes as fall within the true spirit of the
invention.
The method for imaging ultrasonic scatterers of the present invention is used for
obtaining ultrasound images of any object which produces an ultrasound echo.
The inventive method does not produce a medicinal, surgical, curative,
diagnostic, therapeutic or other treatment of human beings or animals.
1. A system for imaging -ultrasound scatterers,
comprising:
an ultrasound transducer array for transmitting
ultrasound beams and detecting ultrasound eerrbes reflected by
said ultrasound scatterers, said transducer array comprising a
multiplicity of transducer elements;
transmitter means coupled to said transducer array for
forming a transmit beam for each one of a multiplicity of sample
volumes;
receiver means comprising a multiplicity of receive
channels for receiving respective amplitude signals from said
multiplicity of transducer elements;
means for forming an incoherent sum of the received
amplitude signals derived from ultrasound echoes reflected by a
single sample volume, a respective incoherent sum being formed
for each of said multiplicity of sample volumes; and
means for displaying an image comprised of pixels
wherein the intensity of each pixel is a function of the incoherent
sum formed for the corresponding one of said multiplicity of
sample volumes.
2. The system of claim 1, further comprising:
means for forming a coherent sum of the received
amplitude signals derived from ultrasound echoes reflected by a
single sample volume, a respective coherent sum being formed for
each of said multiplicity of sample volumes; and
meas for forming a ratio for each of said multiplicity of
sample volumes, wherein said ratio equals the absolute value of
the coherent sum for a respective sample volume divided by the
incoherent sum for said respective sample volume.
3. The system of claim 2, wherein the intensity of
each pixel in said displayed image is linearly proportional to the
ratio derived for the corresponding one of said multiplicity of
sample volumes.
4. The system of claim 2,) further comprising means
for forming, for each pixel of said image, a product equal to the
absolute value of the coherent sum for a respective sample
volume multiplied by the ratio for said respective sample volume.
5. The system of claim 4, wherein intensity of each
pixel in said displayed image is logarithmically proportional to the
product derived for the corresponding one of said multiplicity of
sample volumes.
6. The system of claim 2, further comprising a two-
dimensional filter for filtering the ratios for said multiplicity of
sample volumes prior to display.
7. The system of claim 2, further comprising mapping
means for mapping the ratios for said multiplicity of sample
volumes prior to display.
8. A method for imaging ultrasound scatterers,
comprising the steps of:
transmitting ultrasound beams focused at respective
sample volumes of a multiplicity of sample volumes, at least a
plurality of said sample volumes containing ultrasound scatterers;
detecting, at a multiplicity of detection locations for
each sample volume, ultrasound echoes reflected from said
multiplicity of sample volumes;
producing a respective amplitude signal in response to
detection of an ultrasound echo from each of said multiplicity of
detection locations;
forming an incoherent sum of the amplitude signals
derived from ultrasound echoes reflected by a single sample
volume, a respective incoherent sum being formed for each of said
multiplicity of sample volumes; and
displaying an image comprised of pixels wherein the
intensity of each pixel is a function of the incoherent sum formed
for the corresponding one of said multiplicity of sample volumes.
9. The method of claim 8, further comprising the steps
of:
forming a coherent sum of the amplitude signals
derived from ultrasound echoes reflected by a single sample
volume, a respective coherent sum being formed for each of said
multiplicity of sample volumes; and
forming a ratio for each of said multiplicity of sample
volumes, wherein said ratio equals the absolute value of the
coherent sum for a respective sample volume divided by the
incoherent sum for said respective sample volume.
10. The method of claim 9, wherein the intensity of
each pixel in said displayed image is linearly proportional to the
ratio derived for the corresponding one of said multiplicity of
sample volumes.
11. The method of claim 9, further comprising the step
of forming, for each pixel of said image, a product equal to the
absolute value of the coherent sum for a respective sample
volume multiplied by the ratio for said respective sample volume.
12. The method of claim 11, wherein the intensity of
each pixel in said displayed image is logarithmically proportional
to the product derived for the corresponding one of said
multiplicity of sample volumes.
13. The method of claim 9, further comprising the step
of spatially filtering the ratios for said multiplicity of sample
volumes prior to display.
14. The method of claim 9, further comprising the step
of mapping the ratios for said multiplicity of sample volumes prior
to display.
15. A system for imaging ultrasound scatterers,
comprising:
an ultrasound transducer array for transmitting
ultrasound beams and detecting ultrasound echoes reflected by
said ultrasound scatterers, said transducer array comprising a
multiplicity of transducer elements;
transmitter means coupled to said transducer array for
forming a transmit beam for each one of a multiplicity of sample
volumes;
receiver means comprising a multiplicity of receive
channels for receiving respective amplitude signals from said
multiplicity of transducer elements;
means for forming an incoherent sum of the received
amplitude signals derived from ultrasound echoes reflected by a
single sample volume, a respective incoherent sum being formed
for each of said multiplicity of sample volumes;
means for forming a coherent sum of the received
amplitude signal derived from ultrasound echoes reflected by a
single sample volume, a respective coherent sum being formed for
each of said multiplicity of sample volumes;
means for forming, for each of said multiplicity of
sample volumes, a ratio equal to the absolute value of the
coherent sum for a respective sample volume divided by the
incoherent sum for said respective sample volume;
first memory means for storing the absolute value of
the coherent sum for each of said multiplicity of sample volumes;
second memory means for storing the ratio for each of
said multiplicity of sample volumes;
means for forming a product for each of said multi-
plicity of sample volumes, said product being equal to the absolute
value of the coherent sum for a respective sample volume
multiplied by the ratio for said respective sample volume;
means for log-compressing data coupled to the output
of the produc forming means;
switching means for coupling said first and second
memory means to said product forming means in a first switching
state and coupling said first memory means to said log-
compressing means in a second switching state; and
means for displaying an image comprised of pixels,
coupled to said log-compressing means.
16. The system of claim 15, wherein the displaying
means comprises means for scan-converting data coupled to the
output of said log-compressing means, and wherein said switching
means is adapted to couple said second memory means to said
scan-converting means in a third switching state.
17. The system of claim 15, further comprising a two-
dimensional filter for filtering the ratios for said multiplicity of
sample volumes prior to display.
18. The system of claim 15, further comprising
mapping means for mapping the ratios for said multiplicity of
sample volumes prior to display.

A method and apparatus for improving medical
ultrasound images employs data-dependent filtering. A quantity,
called the coherence factor, is calculated for each pixel in the
image. The coherence factor is defined to be the ratio of two
quantities: the amplitude of the receive signals summed coherently
and the amplitude of the receive signals summed incoherently.
The coherence data is stored in buffer memory and is optionally
spatially filtered and mapped. The amplitude data is concurrently
acquired and stored in buffer memory. The system can be
selectively operated to display the coherence information alone,
the amplitude information alone, or a combination of the
coherence and amplitude information.