U.S. patent application number 11/925843 was filed with the patent office on 2008-06-12 for systems and methods for quantification and classification of fluids in human cavities in ultrasound images.
Invention is credited to Henri Baartmans, Nicolaas Bom, Vikram Chalana, Charles Theodoor Lancee, Gerald McMorrow, Egon J.W. Merks, Jongtae Yuk.
Application Number | 20080139934 11/925843 |
Document ID | / |
Family ID | 9942032 |
Filed Date | 2008-06-12 |
United States Patent
Application |
20080139934 |
Kind Code |
A1 |
McMorrow; Gerald ; et
al. |
June 12, 2008 |
SYSTEMS AND METHODS FOR QUANTIFICATION AND CLASSIFICATION OF FLUIDS
IN HUMAN CAVITIES IN ULTRASOUND IMAGES
Abstract
Ultrasound imaging systems and methods are disclosed. In one
embodiment, an ultrasonography method includes creating a database
that is representative of a tissue, a fluid, or a cavity of a body,
and transmitting ultrasound pulses into a region-of-interest in a
patient. Echoes are received from the region of interest, and based
upon the received echoes, compiling an ultrasonic pattern of the
region-of-interest is compiled. The pattern is processed by
comparing the region-of-interest patterns to the pattern
information stored in the database. A composition within the
region-of-interest of the patient is then determined.
Inventors: |
McMorrow; Gerald; (Kirkland,
WA) ; Chalana; Vikram; (Mill Creek, WA) ; Yuk;
Jongtae; (Redmond, WA) ; Baartmans; Henri; (HT
IJsselstein, NL) ; Bom; Nicolaas; (Berkenwoude,
NL) ; Lancee; Charles Theodoor; (PH Hoogersmilde,
NL) ; Merks; Egon J.W.; (WR Delft, NL) |
Correspondence
Address: |
BLACK LOWE & GRAHAM, PLLC
701 FIFTH AVENUE, SUITE 4800
SEATTLE
WA
98104
US
|
Family ID: |
9942032 |
Appl. No.: |
11/925843 |
Filed: |
October 27, 2007 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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11213284 |
Aug 26, 2005 |
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11925843 |
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10523681 |
Sep 23, 2005 |
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PCT/EP2003/007807 |
Jul 1, 2003 |
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11213284 |
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11010539 |
Dec 13, 2004 |
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10523681 |
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10523681 |
Sep 23, 2005 |
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PCT/EP2003/007807 |
Jul 1, 2003 |
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PCT/EP2003/007807 |
Jul 1, 2003 |
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Current U.S.
Class: |
600/438 |
Current CPC
Class: |
G06T 2207/10132
20130101; A61B 8/0858 20130101; A61B 5/204 20130101; G01S 7/52036
20130101; G06T 7/0012 20130101; A61B 8/4472 20130101; G01S 15/8909
20130101; G01S 15/102 20130101; G06T 2207/30004 20130101; G06T 7/62
20170101; A61B 8/483 20130101; A61B 8/08 20130101 |
Class at
Publication: |
600/438 |
International
Class: |
A61B 8/00 20060101
A61B008/00 |
Foreign Application Data
Date |
Code |
Application Number |
Aug 9, 2002 |
GB |
GB 0218547.8 |
Aug 9, 2002 |
GB |
GB2391625A |
Claims
1. An apparatus for measuring the volume of fluid in a human or
animal body cavity using a non-invasive, ultrasound echo technique,
comprising: a transducer assembly including a plurality of
ultrasound transducers mounted thereon for transmitting and
receiving a plurality of ultrasound signals into the body cavity at
plural angles of incidence and/or from plural spatial locations;
means for activating the transducers to produce transmitted
ultrasound signals; means for detecting body cavity wall echoes
from received ultrasound signals; means for determining, from said
received signals, a body cavity height H and depth D; means for
determining a specific measurement configuration corresponding to
the body cavity filling degree from the ultrasound signals that
intercept the fluid filled body cavity to thereby select an
appropriate predetermined correction factor K corresponding to that
specific measurement configuration, for optimal calculation of the
volume; and means for calculating the fluid volume according to the
formula H.times.D.times.K.
2. The apparatus of claim 1, wherein the body cavity is a bladder
and the volume of fluid measured is a volume of urine.
3. The apparatus of claim 1, wherein the means for activating
includes means for transmitting said plurality of ultrasound
signals in a selected order.
4. The apparatus of claim 1, wherein the means for detecting uses
echo travel time and other beam information from the plurality of
ultrasound signals.
5. The apparatus of claim 1, wherein the means for determining
selects specific ultrasound signals from the plurality of
ultrasound signals corresponding to ultrasound beams that have
intercepted the fluid filled body cavity.
6. The apparatus of claim 1, further including a display means for
instantaneous display of the calculated fluid volume to allow
optimization of transducer positioning by the user.
7. The apparatus of claim 1, wherein the means for deriving
includes a memory storing a plurality of empirically predetermined
correction factors K.
8. The apparatus of claim 1, wherein the array further includes
five transducers.
9. The apparatus of claim 8, wherein the five transducers are
respectively oriented at angles .phi..sub.A, .phi..sub.B,
.phi..sub.C, .phi..sub.D, and .phi..sub.E, to an axis orthogonal to
the plane of the transducer array, the angles being approximately
.phi..sub.A=-25.degree., .phi..sub.B=0.degree.,
.phi..sub.C+25.degree., .phi..sub.D+25.degree.,
.phi..sub.E+40.degree..
10. A method for measuring the volume of fluid in a human or animal
body cavity using a non-invasive, ultrasound echo technique,
comprising the steps of: transmitting a plurality of ultrasonic
beams into the region of the body containing the cavity at plural
angles of incidence and/or from plural spatial locations; receiving
a plurality of ultrasonic signals from the body; determining, from
said received signals, a body cavity height H and depth D;
determining, from the received signals, a specific measurement
configuration corresponding to the body cavity filling degree from
the ultrasound signals that intercept the fluid filled body cavity
to thereby select an appropriate predetermined correction factor K
corresponding to that specific measurement configuration, for
optimal calculation of the volume; and calculating the fluid volume
according to the formula H.times.D.times.K.
11. The method of claim 10, further including the step of
transmitting the plurality of ultrasonic beams into the body from a
transducer array in which a plurality of transducers are arranged
with a predetermined spatial location and mounting angle.
12. An ultrasonography method, comprising: creating a database that
is representative of a tissue, a fluid, or a cavity of a body;
transmitting ultrasound pulses into a region-of-interest in a
patient; receiving echoes from the region of interest, and based
upon the received echoes: compiling an ultrasonic pattern of the
region-of-interest; processing the pattern by comparing the
region-of-interest patterns to the database; and determining a
composition within the region-of-interest of the patient.
13. The method of claim 12, wherein transmitting ultrasound pulses
into the region of interest includes transmitting the pulses to at
least one of a tissue, a fluid, and a cavity.
14. The method of claim 13, wherein transmitting the pulse further
comprises transmitting the pulses to at least one of urine, blood,
amniotic fluid, lung fluids, liver bile, and mixtures thereof.
15. The method of claim 12, wherein processing the pattern includes
calculating at least one of a Goldberg number, a harmonic ratio,
and an attenuation factor.
16. The method of claim 15, wherein processing the pattern further
includes applying a window algorithm to a section of an echo pulse
near a cavity-boundary interface within the region-of-interest.
17. The method of claim 16, wherein applying a window algorithm
further comprises determining the harmonic frequencies associated
with the section of the echo pulse near the cavity-boundary
interface.
18. The method of claim 12, wherein receiving echoes further
comprises receiving at least one of a single dimensional line, a
two-dimensional plane, and a three-dimensional array of
two-dimensional planes.
19. An ultrasonography method, comprising: creating a database that
is representative of a tissue, a fluid, and a cavity of a body;
transmitting ultrasound pulses into a region-of-interest in the
body; receiving echoes from the region of interest, and based on
the echoes: compiling an ultrasonic pattern of the
region-of-interest; processing the pattern by comparing the
region-of-interest patterns to the database; and determining a
volume within the region-of-interest of the body.
20. The method of claim 19, wherein transmitting ultrasound pulses
to the region of interest includes transmitting the pulses into at
least one of a tissue, a fluid, and a cavity.
21. The method of claim 20, wherein transmitting the pulse further
comprises transmitting the pulses into at least one of urine,
blood, amniotic fluid, lung fluids, liver bile, and mixtures
thereof.
22. The method of claim 19, wherein processing the pattern includes
calculating at least one of a Goldberg number, a harmonic ratio,
and an attenuation factor.
23. The method of claim 19, wherein processing the pattern further
includes applying a window algorithm to a section of an echo pulse
near a cavity-boundary interface within the region-of-interest.
24. The method of claim 23, wherein applying a window algorithm
further comprises determining one or more harmonic frequencies
associated with the section of the echo pulse near the
cavity-boundary interface.
25. The method of claim 19, wherein receiving echoes further
comprises receiving at least one of a single dimensional line, a
two-dimensional plane, and a three-dimensional array of
two-dimensional planes
Description
FIELD OF INVENTION
[0001] This invention relates to ultrasound imaging of bodily
tissues, bodily fluids, and fluid-filled cavities.
BACKGROUND OF THE INVENTION
[0002] The following applications are incorporated by reference as
if fully set forth herein: U.S. application Ser. Nos. 11/213,284
filed on Aug. 26, 2005; 11/010,539 filed on Dec. 13, 2004; and
10/523,681 filed on Feb. 3, 2005.
[0003] Ultrasound imaging is accomplished by placing an ultrasound
transducer on a selected location of a body and projecting
ultrasound energy into the body. Acoustic waves reflecting from
internal structures in the body are then received by the transducer
and are processed to form an image of the internal structures. In a
particular ultrasound method, amplitudes of selected harmonics of
the returned signal are processed to form the ultrasound image.
Briefly and in general terms, harmonic generation by the internal
structures in the body is at least partially determined by the
properties of the tissue that reflect the ultrasound energy, so
that the presence of harmonics in the received echo may be used to
generate useful information in the ultrasound image, as discussed
in further detail in A. Bouakaz, E. Merks, C. Lancee, N. Bom,
"Noninvasive Bladder Volume Measurements Based on Nonlinear Wave
Distortions," Ultrasound in Medicine & Biology, 30:4, pp.
469-476, which publication is incorporated by reference herein.
[0004] In selected ultrasound imaging applications, it is often
desirable to distinguish between a bodily fluid and an adjacent
tissue, or between bodily fluids of different types, such as
between blood and various other bodily fluids. For example, when a
selected anatomical portion is imaged using B-mode ultrasound
imaging, a bodily fluid and certain adjacent soft tissues within
the selected portion may be relatively indistinguishable in the
resulting image. Moreover, when blood within the selected portion
and other bodily fluids are imaged using B-mode ultrasound, images
may be generated that similarly fail to properly distinguish the
blood from the other bodily fluids.
[0005] Accordingly, a new imaging system is needed that permits the
diagnostician to easily distinguish or discriminate between
different fluid compositions, or between a bodily fluid and tissue.
It is further desirable to render more easily detectable in an
ultrasound image any boundaries between cavities that contain
fluids of different composition, and between a bodily fluid and a
bodily tissue.**
[0006] It is well known that bladder dysfunction is associated with
a number of clinical conditions requiring treatment. In many of
these cases it is important to accurately determine the volume of
the bladder. Under other conditions such as post-operative
recovery, where there is temporary loss of bladder sensation and/or
loss of the normal voiding mechanism too much distention of the
bladder has to be avoided. Under those conditions voiding by
catheter introduction is carried out. However, serious
disadvantages to unnecessary catheterization range from. the
uncomfortable situation for the patient to serious possibilities of
infection. Thus, a non-invasive quick measurement of bladder
volume, with the patient usually in the supine position, is
indicated. Sometimes the accurate determination of volume is
indicated; sometimes however an indication is sufficient. Questions
that may be asked are for instance: after voiding:" is there still
too much urine left?"; or after surgery" is the bladder filling
above a certain level so that voiding is necessary?"
[0007] Non-invasive procedures for bladder volume estimation are
known, but are either unreliable or expensive or have some other
significant disadvantages. Palpation and auscultatory percussion
are known to be unreliable, while radiography and dye-excretion
techniques are known to be similarly inaccurate. For assessing
bladder volume, catheterization remains the "gold standard".
However, it is invasive, painful and might produce traumas or
infections.
SUBJECT
[0008] The described technique concerns measurement of urine volume
in the human bladder with the use of pulsed ultrasound with a
limited number of ultrasound transducers.
[0009] In a first version a limited number of transducers are
mounted in a transducer assembly. The assembly is positioned
non-invasively at the body skin over the position of the bladder
with the patient in a supine position. For acoustic contact a
coupling gel may be used. Each ultrasound transducer in the
assembly transmits and receives the ultrasound signal in a narrow
beam through the contact plane. During the measurement the
transducers are used in a certain succession. All transducers have
been mounted in the assembly such that in transmission and
reception successively the beams penetrate the area of the bladder
in approximately the sagittal cross sectional plane. The sagittal
plane is here defined as ANTERO-POSTERIOR plane of the body. One
transducer beam direction is dorsal with in addition at least one
transducer beam in the dorsal-caudal and one transducer beam in the
dorsal-cranial direction. The volume is calculated on the basis of
two bladder measurements defined in the sagittal plane as Depth (D)
and Height (H). These measurements are derived on the basis of echo
travel time from echoes originating at the anterior and posterior
bladder wall. Depth is in principle a measurement in dorsal
direction. Height is a measurement approximately in the cranial
direction. The volume is calculated depending on the specific,
filling dependent, measurement configuration following the formula
D.times.H.times.K. Where K is an empirically measured, filling
configuration dependant, correction factor. Beam directions and
examples for D and H are illustrated in FIGS. 1 and 2.
[0010] In a second version of the described technique a single wide
beam ultrasound transducer is positioned non-invasively at the body
skin over the location of the bladder. The wide beam can be created
by the curved surface of the transducer or by a flat acoustically
active surface of for instance a disk shaped transducer supplied
with a curved lens. Ultrasonic signals are transmitted and received
in the wide, cone like, ultrasound beam and propagation is
approximately spherical. Similar to the above described method a
pulsed echo signal is transmitted at fundamental ultrasonic
frequency. In this second version of the described technique echo
data are analyzed as originating from a distance beyond the average
position of the posterior (filled) bladder wall. The received echo
signal will contain information over almost the entire bladder as
encompassed by the wide ultrasound beam. Due to non-linearity,
higher harmonic components will build up during propagation and
thus be reflected in the returning echo.
[0011] Compared to propagation through normal tissue, the presence
of higher harmonics in the signal is greatly stimulated when
propagating through urine. Analyses of presence of higher harmonic
components in relation to the fundamental frequency is used for
indication of presence of urine in the bladder. Neutralizing
patient variation as to obesity etc can also be accomplished by
comparing echo signals received from sequentially transmitted
pulses at low transmit power (linear propagation only) and pulse
transmission at high power (enhancing non-linearity).
STATE OF THE ART
[0012] Non-invasive bladder volume measurement techniques with
ultrasound echography have been described in the art. In principle,
echography measures distance based on echo travel time. Early echo
techniques did use a single ultrasound transducer and echo
presentation was recorded as echo amplitude versus depth. West, K
A: "Sonocystography: A method for measuring residual urine", Scand
J Urol Nephrol 1: pp 68-70, 1967 describes the subsequent use of
some discrete beam directions. He does not have a separate
transducer for each beam direction. His method is only qualitative,
not instantaneous, and based on distance measurement to the dorsal
posterior bladder wall. His method is not adjusted to specific,
filling dependent, measuring configurations. A relation between the
difference in echo travel time between echoes from the posterior an
anterior bladder wall and the independently measured bladder volume
has been reported by Holmes, J H: "Ultrasonic studies of the
bladder", J. Urology, Vol 97, pp. 654-663. His described volume
measurement method is exclusively based on bladder depth
measurement. Since the bladder changes in shape when filling, a
single distance measurement is not precise enough to predict the
entire bladder volume. No filling dependent measurement
configuration is used.
[0013] Diagnostic ultrasound is today well known for real-time
cross-sectional imaging of human organs. For cross-sectional
imaging the sound beam has to be swept electronically or
mechanically through the cross section to be imaged. Echoes are
presented as intensity modulated dot on the display. The
instruments are costly and require a skilled operator. Volume is
sometimes calculated based on bladder contours obtained in two
orthogonal planes with a geometric assumption of bladder shape. For
3-dimensional or volumetric echography the sound beam has to be
swept through the entire organ. This further increases complexity,
acquisition time of the data, and costs of the instrument.
[0014] HAKENBERG ET AL: "THE ESTIMATION OF BLADDER VOLUME BY
SONOCYSTOGRAPHY", J Urol, Vol 130, pp 249-251, have reported a
simple method that is based on measuring the diameters obtained in
a cross sectional image in the midline sagittal bladder plane only.
The bladder volume has been related to bladder Height and Depth as
follows: Volume is Height.times.Depth.times.6.6 ml. This formula
showed a good correlation coefficient (r=0.942) with a relatively
large average error of 30.1%. For this approach a two-dimensional
imaging apparatus was required. The used apparatus is complex and
is different from the method described in this application. It does
not use a single wide beam transducer or a limited number of fixed
transducers in an assembly or a combination of this.
[0015] An ultrasound apparatus for determining the bladder volume
is shown in U.S. Pat. No. 4,926,871 in the name of Dipankar Ganguly
et al. In this text, a number of possibilities are mentioned,
amongst which a scan head embodiment referred to as a sparse linear
array with transducers mounted at predetermined angles with sound
beams pointing towards the same position. The volume is calculated
according to a geometric model. In the claims an apparatus is
described, involving an automatic calculation of bladder volume
from ultrasound measurements in a first and second plane, which are
substantially orthogonal to each other. Sound beams are deflected
by a stepper motor. It requires a skilled operator to manipulate
the scan head in a particular way to obtain the ultrasound
measurements. For the volume calculation method described in this
application no use is made of any geometrical model of the bladder,
whereas only a limited number of sound beams approximately in the
sagittal plane, or a single wide beam is used.
[0016] Volume measurement based on echographic sampling of the
bladder with a hand guided transducer mounted in a panthograph has
been described by Kruczkowski et al:" A non-invasive ultrasonic
system to determine residual bladder volume", IEEE Eng in Medicine
& Biology Soc 10TH Ann Conf, pp 1623-1624. The sampling covers
the entire bladder, follows a given pattern and is not limited to a
single or two cross sections of the bladder. For the calculation he
needs data from many beam directions. The acquisition procedure is
time consuming and thus no instantaneous volume measurement
results. The method described in this application is based on use
of a single, wide beam or the use of a limited number of mutually
fixed sound beams directions with instantaneous volume
indication.
[0017] The hand steered transducer guiding for recording of echo
data from the bladder has subsequently gained in acquisition speed
by introduction of constructions whereby the transducer, and thus
the beam, was mechanically swept. This nevertheless still requires
an acquisition time equivalent to full acquisition procedure and
thus does not yield an instantaneous display of volume. No
instantaneous feedback on optimal positioning is thus available. An
example of such methods is the BLADDERSCAN. In the Bladderscan
Technology (registered trademark of Diagnostic Ultrasound
Corporation) bladder volume is measured by interrogating a
three-dimensional region containing the bladder and then performing
image detection on the ultrasound signals returned from the region
insonated. The three dimensional scan is achieved by performing
twelve planar scans rotated by mechanically sweeping a transducer
through a 97 degree arc in steps of 1.9 degrees. The three
dimensional scanning requirement makes this instrument complex. It
can not be compared with the simple approach described in this
application.
[0018] Yet another ultrasound method "System for estimating bladder
volume" is described by Ganguly et al in U.S. Pat. No. 5,964,710
dated Oct. 12, 1999. This method is based on bladder wall contour
detection with echographically obtained data in a plurality of
planes which subdivide the bladder. In each single plane of the
plurality of planes a number of N transducers are positioned on a
line to produce N ultrasound beams to measure at N positions the
distance from front to back wall in the selected plan. From this
the surface is derived. This procedure is repeated in the other
planes as well. The volume is calculated from the weighted sum of
the plurality of planes. In Ganguly's method the entire border of
the bladder is echographically sampled in 3 dimensions. His method
differs strongly from the method described in this application
whereby only a single wide beam is used or a limited number of
mutually fixed sound directions are used in approximately a
sagittal plane with a filling dependent measurement
configuration.
[0019] U.S. Pat. No. 6,359,190 describes a device for measuring the
volume of a body cavity, such as a bladder or rectum, using
ultrasound. The device is strapped to the body or incorporated into
a garment such as a nappy or trainer pant. The device includes
several transducers each aimed at a different region of the
subject's bladder (a) to ensure that at least one ultrasound beam
crosses the bladder despite variations in the way that the device
has been positioned on the body, and (b) to enable the transducer
with the strongest signal output to be used. An alarm signal may be
output when the bladder reaches a predetermined threshold
volume.
[0020] An important parameter for assessing bladder volume if this
volume has to be derived from a limited number of beams or planes
is the knowledge of bladder shape and position which can
drastically vary with age, gender, filling degree and disease. In
the adult patient the empty bladder has the shape of a triangular
prism and is located behind the pubis. When it is progressively
filled, there is first a distention of the bladder depth followed
by an expansion of the bladder height. The bladder shape is complex
and can not be represented by a single geometrical formula such as
ellipsoid, sphere etc. This explains the large error that several
studies obtained when a single geometric model was used. However
there exists a correlation between the bladder height and the
bladder widening with progressive filling.
[0021] In the first approach of the present invention an instrument
is described which allows assessment of bladder volume by using
only a few ultrasound beams appropriately oriented in approximately
the sagittal plane. The narrow sound beams in principle diverge
relative to each other. This allows covering a wide range of
filling degrees of the bladder, from almost empty, when the bladder
is located behind the pubis, to a full bladder that causes a
substantial bladder height (See FIGS. 1 and 2). From each beam can
be established, by detection of the posterior bladder wall echo, if
this beam does pass a filled bladder. From the knowledge of all
beams that do pass the filled bladder the appropriate filling or
measurement configuration follows. The acoustic beams are
positioned in such a way that the Depth D and Height H of the
bladder can be estimated for the specific measurement
configuration. The volume of urine is then computed from an
empirical formula D.times.H.times.K that does not depend on any
geometric model. K is a known, empirically established correction
factor which is specific for each measurement configuration and has
been established by calibrated bladder measurements on a prior
series of patients. The accuracy of the first approach is thus
based on an a prior known correction factor which is related to a
specific filling degree, which in turn depends on the number of
beams that intercept the filled bladder.
[0022] A second version of the instrument is based on the
measurement of the presence of higher harmonics in the echo signal.
For this approach the echo signal from a depth greater than the
distance from the transducer to the posterior bladder wall must be
analyzed. For a filled bladder in adults in a supine position, this
depth W would be approximately 12 cm.
[0023] It is known that when sound pulses are transmitted at a
fundamental frequency, higher harmonics of this fundamental
frequency may be present in the received echographic signal.
Non-linear distortion increases with distance, insonifying
ultrasound energy and frequency. Attenuation diminishes the
ultrasound amplitude with increasing propagation distance and
reduces the higher harmonic energy. Since attenuation of the
ultrasound signal in urine is low compared to tissue and non-linear
distortion in urine is large compared to tissue it results that
urine is very different from tissue in its ability to generate
higher harmonics. We have measured the presence of higher harmonics
in the echo signal from 12 cm depth when the bladder was filled.
With an empty bladder the echoes obtained from the same depth did
not contain higher harmonics.
[0024] The interest of higher harmonic signals in the ultrasound
technique stems from echo contrast technology. Echo contrast
material contains coated gas containing micro bubbles suspended in
a fluid. These bubbles can create higher harmonic components in the
echo signal due to non-linearity. This is used to indicate presence
of contrast on the diagnostic image. A wide variety of pulse
techniques is used to stimulate echographic visibility of contrast.
These include multi pulse procedures, multi frequency procedures,
power Doppler imaging, pulse coding, pulse inversion and other
imaging methods. A survey is documented in "Ultrasound Contrast
Agents" ISBN 1-85317-858-4 chapter 3 "Contrast-specific imaging
methods" by de Jong et al. With a single transducer with wide sound
beam, such as results with a curved acoustic element or a flat,
disk shaped transducer plus curved lens, the propagating sound beam
would encompass the entire bladder. The transducer must be designed
to optimally transmit the fundamental ultrasound frequency and at
the same time be capable to receive fundamental and higher harmonic
echo signals. Broadband piezo-electric ceramic transducers have
been described as well as combination transducers using ceramic in
transmission and PVDF material in reception. In transmission a
single or multi pulse procedure can be followed. If the returned
echo signal with such a method would, in relation to the
fundamental echo signal, be analyzed for the presence of higher
harmonics, the presence of a certain level of bladder filling or
the volume of urine can be established.
[0025] EP 0271214 describes an ultrasonic device for monitoring the
volume of fluid in the human bladder by using reflected ultrasound
signals to determine not only the position of the bladder back wall
but also energy returned from the bladder back wall. EP '214
proposes that after bladder filling to approximately 60% capacity,
the distance between the back wall and the front wall of the
bladder stops increasing. However, additional reverberation in the
back wall provides an increase in energy in the reflected signal
which can be used to determine further increases in bladder
volume.
SUMMARY
[0026] The present invention comprises ultrasound imaging systems
and methods. In one aspect, an ultrasonography method includes
creating a database that is representative of a tissue, a fluid, or
a cavity of a body, and transmitting ultrasound pulses into a
region-of-interest in a patient. Echoes are received from the
region of interest, and based upon the received echoes, compiling
an ultrasonic pattern of the region-of-interest is compiled. The
pattern is processed by comparing the region-of-interest patterns
to the pattern information stored in the database. A composition
within the region-of-interest of the patient is then
determined.
BRIEF DESCRIPTION OF THE DRAWINGS
[0027] The embodiments of the present invention are described in
detail below with reference to the following drawings.
[0028] FIG. 1 is a side elevational view of a
microprocessor-controlled transceiver according to an embodiment of
the invention;
[0029] FIG. 2 is a representation of an ultrasound scan cone
emanating from the transceiver in a conic shape formed by a
plurality of three-dimensional distributed scan lines;
[0030] FIG. 3A is a representation of an ultrasound scan cone
emanating from the transceiver in a conic shape formed by a
rotational array of two-dimensional scan planes;
[0031] FIG. 3B is a representation of a scan plane of the
rotational array;
[0032] FIG. 4 is a is depiction of the hand-held transceiver in use
for scanning the abdominal area of a patient;
[0033] FIG. 5 is a perspective view of the hand-held transceiver
device sitting in a communication cradle;
[0034] FIG. 5B illustrates a schematic view of an imaging
system;
[0035] FIG. 6 depicts a schematic view of a plurality of
transceivers in connection with a server;
[0036] FIG. 7 is a schematic view of a plurality of bladder wall
measuring systems connected to a server over the Internet or other
network;
[0037] FIG. 8 illustrates a scan plane image with diagrammatic scan
lines overlaid on the image;
[0038] FIG. 9 is a schematic illustration of a scan line passing
through a full bladder and a uterus;
[0039] FIG. 10 is a plot of echo amplitude versus scan line depth
of the FIG. 9 schematic;
[0040] FIG. 11 is a schematic illustration of a scan line passing
through a near empty bladder and the uterus;
[0041] FIG. 12 is a plot of echo amplitude versus scan line depth
of the FIG.
[0042] FIG. 13 is a schematic illustration of a scan line passing
through body cavities and surrounding tissues;
[0043] FIG. 14 is another spectral plot of a window function
processed insonified region of FIG. 13 for a non-pregnant female
subject having homogenous uterine fluid;
[0044] FIG. 15 is another spectral plot of a window function
processed insonified region of FIG. 13 for a non-pregnant female
subject having heterogeneous uterine fluid;
[0045] FIG. 16 is a calibration plot of harmonic power as a
function of the bladder volume;
[0046] FIG. 17 is a calibration plot of harmonic ratio as a
function of blood composition in amniotic fluid;
[0047] FIG. 18 is a method overview flowchart;
[0048] FIG. 19 is an expansion of sub-algorithm 206 of FIG. 18;
[0049] FIG. 20 is a schematic representation of four surface patch
elements; and
[0050] FIG. 21 is a schematic representation of three scan lines
passing through the subserosal and submucosal wall locations of an
organ.
[0051] FIG. 1 Illustrates a sagittal (anteroposterior) cross
sectional plane of a patient in supine position where a transducer
assembly 1 with transducers A, B, C, D and E, is positioned on the
abdominal wall just above the Symphysis Pubis 2 and the ultrasound
beams are indicated to cross the area of the partially filled
bladder 3. From the transducer assembly, the sound beam A
intercepts the bladder area in dorso-caudal direction, soundbeam B
intercepts the bladder in dorsal direction and sound beams C, D,
and E respectively in dorso-cranial direction. In FIG. 1 the
patient's leg is indicated by 4.
[0052] FIG. 2 Illustrates various bladder filling stages from an
almost empty bladder to a strongly filled bladder and the
corresponding measurement configurations. Depth D and Height H have
been defined for each filling situation as indicated and are
calculated from detected bladder wall echoes taking the specific
measurement configuration into account. For each measurement
configuration a specific Depth D and Height H is defined.
[0053] FIG. 3. Illustrates, by way of example for a transducer
assembly with five transducers (here only A and D, necessary for
calculation of H are shown), the calculation of Height H (5) in the
measurement configuration when bladder posterior wall echoes are
detected originating from sound beam A, B, C and D. This is the
"filled bladder" measurement configuration shown in FIG. 2.
Apparently no posterior wall echoes are detected in sound beam E
because the bladder filling is not yet in a strongly filled stage
and thus beam E does not intercept the bladder. Depth D is derived
from beam B (not shown in FIG. 3).
[0054] FIG. 4 Represents a flow chart of the actions of the
principal hardware components. In this block diagram a "useful"
transducer signal occurs when bladder wall echoes are detectable in
its sound beam.
[0055] FIG. 5. Illustrates a top view of five disk shaped
transducers in a possible transducer assembly. The distance between
transducers B, D and C, A, E and their positioning is such that all
sound beams can be assumed to be in approximately a sagittal cross
section through the bladder. Yet another transducer assembly with 4
transducers in a row is also illustrated.
[0056] FIG. 6. Illustrates a cross sectional view showing in the
length direction a possible transducer and related sound beam
orientation when five single transducers are used.
[0057] FIG. 7 Illustrates the sagittal cross sectional plane with a
single wide beam transducer non-invasively positioned on the
abdominal skin surface over the filled bladder 3. Echo signal is
received from a range at depth W.
[0058] FIG. 8. Is a flow chart illustrating the principal steps
taken by the bladder volume measurement instrument based on a
single ultrasound wide beam where detection of presence of higher
harmonics in the received signal from a give depth range is used to
measure volume. Two different transmit levels are used to enhance
the bladder effect and eliminate patient variation.
[0059] FIG. 9. Illustrates the measured received scattered power in
the fundamental frequency to and the higher harmonic frequencies
2f.sub.0 and 3f.sub.0 in a situation with an empty versus a filled
bladder.
[0060] FIG. 10 shows two possible transmit pulse sequences to
enhance the difference between linear and non-linear sound
propagation.
[0061] FIG. 11 Illustrates a possible look-up table based on prior
calibrated patient bladder volume measurements relating presence of
harmonic power in the received echo signal versus volume.
DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS
[0062] The present invention relates to the ultrasound imaging of
tissues and/or fluid-filled cavities having linear or non-linear
acoustic properties. Many specific details of certain embodiments
of the invention are set forth in the following description and in
FIGS. 1 through 21 to provide a thorough understanding of such
embodiments. One skilled in the art, however, will understand that
the present invention may have additional embodiments, or that the
present invention may be practiced without several of the details
described in the following description.
[0063] FIG. 1 is a side elevational view of an ultrasound
transceiver 10 according to an embodiment of the invention. The
transceiver 10 includes a transceiver housing 18 having an
outwardly extending handle 12 that is suitably configured to allow
a user to manually manipulate the transceiver 10. The handle 12
includes a trigger 14 that allows the user to initiate an
ultrasound scan of a selected anatomical portion, and a cavity
selector 16, which will be described in greater detail below. The
transceiver 10 also includes a transceiver dome 20 that contacts a
surface portion of the patient when the selected anatomical portion
is scanned, and a display 24 operable to view processed results
from the ultrasound scan, and to allow operational interaction
between the user and the transceiver 10. Accordingly, the display
24 may be configured to display alphanumeric data that indicates a
proper and/or optimal position of the transceiver 10 relative to
the selected anatomical portion. In other embodiments, two- or
three-dimensional images of the selected anatomical region may be
displayed on the display 24. The display 24 may be a liquid crystal
display (LCD), a light emitting diode (LED) display, a cathode ray
tube (CRT) display, or other suitable display devices operable to
present alphanumeric data and/or graphical images to a user.
[0064] The transceiver 10 further includes a microprocessor (not
shown in FIG. 1) and computational algorithms (also not shown in
FIG. 1) that cooperatively provide enhanced ultrasound harmonic
imaging that permits boundaries between different fluid
compositions to be distinguished. In addition, the transceiver 10
may be suitably configured to distinguish between a bodily fluid
and tissue, between dissimilar tissues, and/or between dissimilar
bodily organs. The computational algorithms will be discussed in
greater detail below. The transceiver 10 may also be coupled to a
computer system (not shown in FIG. 1) that is operable to receive
either digital or analog signals from the transceiver 10 and to
process the signals to generate a desired ultrasound image. In
addition, the computer system may also at least partially control
the operation of the transceiver 10. The computer system may
comprise any microprocessor-based computer or other computer
systems, such as a mainframe that is capable of executing operating
instructions and manipulating data. Accordingly, the computer
system is not limited to a typical desktop or laptop computer
device.
[0065] The operation of the transceiver 10 will now be described.
The transceiver dome 20 of the transceiver 10 is positioned against
a surface portion of a patient that is proximate to the anatomical
portion to be scanned. The user then actuates the transceiver 10 by
depressing the trigger 14. In response, the transceiver 10
transmits ultrasound signals into the body, and receives
corresponding return echo signals that are at least partially
processed by the transceiver 10 to generate an ultrasound image of
the selected anatomical portion. In a particular embodiment, the
transceiver 10 transmits ultrasound signals in a range that extends
from approximately about two megahertz (MHz) to approximately about
ten MHz.
[0066] Still referring to FIG. 1, the cavity selector 16 is
structured to adjustably control the transmission and reception of
ultrasound signals to the anatomy of a patient. In particular, the
cavity selector 16 adapts the transceiver 10 to accommodate various
anatomical details of male and female patients. For example, when
the cavity selector 16 is adjusted to accommodate a male patient,
the transceiver 10 is suitably configured to locate a single
cavity, such as a urinary bladder in the male patient. In contrast,
when the cavity selector 16 is adjusted to accommodate a female
patient, the transceiver 10 is configured to image an anatomical
portion having multiple cavities, such as a bodily region that
includes a bladder and a uterus. Alternate embodiments of the
transceiver 10 may include a cavity selector 16 that is configured
to select a single cavity scanning mode, or a multiple
cavity-scanning mode that may be used with male and/or female
patients. The cavity selector 16 may thus permit a single cavity
region to be imaged, or a multiple cavity region, such as a region
that includes a lung and a heart to be imaged.
[0067] FIG. 2 is a representation of an ultrasound scan cone 30
emanating from the transceiver 10 of FIG. 1 that will be used to
further describe the operation of the transceiver 10. The scan cone
30 has a substantially conic shape formed by a plurality of
three-dimensional distributed scan lines. A scan cone 30 is shown
emanating from the dome 20 of the transceiver 10 in encompassing a
plurality of three-dimensional-distributed scan lines 31A-34E. The
plurality of scan lines 31A-34E represent a line array in
three-dimensional space. The scan lines within the line array are
one-dimensional ultrasound A-lines that emanate from the
transceiver 10 at different coordinate directions that taken as an
aggregate form the scan cone 30. The different coordinate
directions comprise a length r of a given scan line, and a
rotational angle .theta. and a tilt angle .phi.. Thus, one or more
points P along a scan line within the line array 31A-34E are
defined by the distance r and the angular coordinates .phi., and
.theta..
[0068] The plurality of three-dimensional distributed scan lines
31A-34E comprises a plurality of peripheral scan lines 31A-E and a
plurality of internal scan lines 34A-D. The
three-dimensional-distributed A-lines (scan lines) are not
necessarily confined within a scan plane, but instead are directed
to sweep throughout the internal regions and along the periphery of
the scan cone 30. The three-dimensional-distributed scan lines not
only would occupy a given scan plane in a three-dimensional array
of two-dimensional scan planes, but also the inter-scan plane
spaces, from the conic axis to and including the conic periphery.
For example, assume line 34 B is a conical axis line and lines 31C
and 34A are coplanar with line 34B. Lines 34A and 34B are separated
by a tilt angle .phi..sub.1 and lines 31C and 34A are separated by
a tilt angle .phi..sub.2. Similarly, lines 31F and 31C. are
separated by a rotational angle .theta..sub.1 and lines 31D and 34C
are separated by a rotational angle .theta..sub.2.
[0069] The internal scan lines are represented by scan lines 34A-C.
The number and location of the internal scan lines emanating from
the transceiver 10 is variable and may be selected to sufficiently
visualize structures within the scan cone 30. The internal scan
lines are not peripheral scan lines. The peripheral scan lines
31A-F occupy the conic periphery and converge near the apex of the
scan cone 30, thus representing the peripheral limits of the scan
cone 30.
[0070] FIG. 3A is a representation of an ultrasound scan cone
emanating from the transceiver in a conic shape formed by a
rotational array of two-dimensional scan planes. The scan cone 40
emanating from the dome 20 includes a plurality of scan planes 42
assembled as a rotational array. The scan planes within the
rotational array are angularly separated by an angle .theta..
[0071] FIG. 3B is a representation of the scan plane of the
rotational array. A scan plane 42 is formed by a scan line 48 that
rotationally pivots between a first leg 44 and a second leg 46
about a pivot angle .phi.. The depth of the scan plane 42 is
determined by the effective length r of the scan line 48. The area
of the scan plane 42 is determined as a product of the length r of
the scan line 48 and region swept by the scan line 48 as it
migrates about pivot angle .phi.between the first and second legs
44 and 46.
[0072] FIG. 4 is an isometric view of the transceiver 10 positioned
on an external portion of a patient 26 that will be used to
describe a method of data acquisition for identifying fluids in
bodily cavities according to an embodiment of the invention. The
transceiver 10 is positioned against a surface portion of the
patient 26, and a targeting phase is initiated. In a particular
embodiment, the transceiver 10 is then operated in a
two-dimensional continuous acquisition mode, which permits data to
be continuously acquired and presented as a discrete scan plane
image on the display 24 (or another external display device) as the
operator physically moves the transceiver 10 across various
external portions of the patient 26. In this embodiment, the
operator moves the transceiver 10 around an abdominal region and
depresses the trigger 14 of the transceiver 10 to continuously
acquire real-time two-dimensional images that may be continuously
viewed on the display 24, or on another display device. For
example, when an anatomical portion that contains a urinary bladder
is imaged, urine confined within the bladder appears as a dark
region, and a urine fluid area may be calculated. An alphanumeric
indication of the urine fluid area (for example, in cm.sup.2) may
also be calculated and visually presented on the display 24.
Similarly, if the patient 26 is a pregnant female, amniotic fluid
within the uterus may also be imaged and a corresponding amniotic
fluid area may be calculated and displayed on the display 24. After
acquisition of the two dimensional measurements, the volume of
urine and amniotic fluids are measured in the respective bladder
and uterus by acquiring a 3-D scan as a multiple scan plane array
similar to the scan cone 40. Alternatively, if the two-dimensional
measurements are acquired as three-dimensional distributed scan
lines, a three-dimensional scan is accomplished as a
three-dimensional scan cone of three-dimensional distributed scan
lines similar to the scan cone 30. A cavity selector 16 (as shown
in FIG. 1) is engaged to detect and measure the volumes of either
single or multiple cavities in a subject. In a particular
embodiment where the transceiver 10 is positioned over the
symphasis pubis for acquisition of three-dimensional ultrasound
images, a single cavity includes one of the bladder and the uterus,
and a multiple cavity includes the bladder and the uterus.
[0073] FIG. 5 is an isometric view of transceiver 10 according to
another embodiment of the invention. The transceiver 10 is
structured to be received by a support cradle 50. The support
cradle 50 is coupled to a power supply 51 that provides electrical
energy to the cradle 50 that is communicated to the transceiver 10
either conductively or inductively to provide a charging current to
a power supply positioned within the transceiver 10. The support
cradle 50 is also configured to receive ultrasound data from the
transceiver 10 when positioned in the cradle 50, which may be
transferred to an external processor (not shown in FIG. 3) through
a digital communications link 38, such as a link that employs a
universal serial bus (USB), a FIREWIRE bus in conformity with
IEEE-1394, an RS-232 compatible link, or other similar
communications links in conformity with still other protocols. In
other embodiments, the communications link 38 may be a wireless
link, such as wireless local area network (LAN) or a wireless wide
area network (WAN). Alternatively, the cradle 50 may be powered by
the link 38.
[0074] The communications link 38 may also advantageously provide a
means for transferring imaging data from the transceiver 10, and
for transferring software updates or software revisions from the
external processor to the transceiver 10. In a particular
embodiment, the cradle 50 may include a memory device operable to
retain digital data received from the transceiver 10 before the
data is transferred to the external processor through the
communications link 38.
[0075] FIG. 5B illustrates a schematic view of an imaging system
55. The imaging system 55 includes the transceiver 10 positioned in
the supporting cradle 50. The communications link 38 connects the
transceiver 10 housed in the cradle 12 with a computer 62. The
computer 62 may be a desktop, laptop, or other microprocessor-based
portable computing device. Data from the transceiver 10 is routed
through the cradle 50 to the computer 62 via the communications
link 38. The communications link 38 may be a conductive link, as
shown in FIG. 5B, or it may be a wireless radio frequency link or
an optical link, such as a wireless infrared link. Within the
computer 62 are executable programs to implement the algorithms of
the particular embodiments, including the processing of ultrasound
signals, retrieving imaging programs, and instructions to perform
ultrasound enhancement procedures. Various ultrasound images are
developed by processing the ultrasound signal data, including
one-dimensional ultrasound images, two-dimensional images,
three-dimensional renderings, and enhanced images from the
retrieved imaging programs and instructions. The generated images
may be stored within the computer 62.
[0076] FIG. 6 is a partial isometric and diagrammatic view of a
networked imaging system 60 according to another embodiment of the
invention. The imaging system 60 includes one or more transceivers
10 in accordance with one or more of the previously disclosed
embodiments. The one or more transceivers 10 may be positioned in
supporting cradles 50 that are operably coupled to portable
computing devices 62, which in turn, are suitably configured to
receive imaging data from the one or more transceivers 10 through
the communications link 38. The communications link 38 may be a
conductive link, as shown in FIG. 6, or it may be a wireless radio
frequency link or an optical link, such as a wireless infrared
link. The portable computing devices 62 communicate with a server
66 over a communications network 72. Although two transceivers 10
are shown in the networked imaging system 60 shown in FIG. 6, fewer
than two, or more than two transceivers 10 may be present. In
addition, the processing of ultrasound signals may be divided
between the transceiver 10, the portable computing devices 62, and
the server 66. For example, the transceiver 10 may be configured to
process the ultrasound signals and generate an ultrasound image
using algorithms in accordance with other embodiments of the
invention, or alternately, the ultrasound image may be generated by
the portable computing device 62 or even by the server 66 after
receiving ultrasound signals from the transceiver 10. In a
particular embodiment of the networked imaging system 60, the
imaging algorithms that generate the enhanced ultrasound images
reside on the server 66. Each of the portable computing devices 62
accordingly receives signals acquired from the transceivers 10
through the cradles 50 and stores the signals in the portable
computing device 62. The computing device 62 subsequently retrieves
imaging programs and instructions to perform the additional
ultrasound enhancement procedures from the server 66. Thereafter,
each personal computing device 62 generates various ultrasound
images by processing the ultrasound data, including one-dimensional
ultrasound images, two-dimensional images, three-dimensional
renderings, and enhanced images from the retrieved imaging programs
and instructions. The generated images may be stored on the server
66.
[0077] In another particular embodiment, the imaging programs and
the instructions reside exclusively on the server 50 and are
executed on the server 50. Each portable computing device 62
receives the acquired signals from the transceivers 10 through the
cradle 50 and transfers the acquired signals to the portable
computing device 62. The device 62 subsequently communicates the
signals to the server 66 and processes the signals to generate the
desired ultrasound images, including one-dimensional images,
two-dimensional images, three-dimensional renderings, and other
similar images. The ultrasound images may be stored on the server
66, or alternately, the images may be transferred to one or more of
the personal computing devices 62.
[0078] FIG. 7 is a partial isometric and diagrammatic view of a
networked imaging system 80 according to still another embodiment
of the invention. Many of the elements of the present embodiment
have been discussed in detail in connection with other embodiments,
and in the interest of brevity, will not be discussed further. The
networked imaging system 80 includes a public data network 82
interposed between the communications network 72 and the server 66.
The public data network 82 may include a LAN, a WAN, or the
Internet. Accordingly, other computational devices 84 associated
with the public data network 82 may communicate imaging data and/or
ultrasound images with the portable computing devices 62 and the
server 66. Although two transceivers 10 are shown in the networked
imaging system 80 shown in FIG. 7, fewer than two, or more than two
transceivers 10 may be present.
[0079] FIG. 8 is an ultrasound image 90 of a bodily portion of a
patient that will be used to describe other embodiments of the
invention. The image 90 is formed by projecting a plurality of scan
lines 48 downwardly into a selected anatomical portion of the
patient to form the fan-like scan plane 42. The scan plane 42 may
be rotated about an axis that extends through the transceiver 10 to
generate a scan cone 40 (as shown in FIG. 3A) to obtain
three-dimensional imaging information for the selected anatomical
portion. Accordingly, when ultrasound energy is projected into the
selected anatomical portion, various internal structures may
reflect the ultrasound energy, including a bladder, a front bladder
wall and a back bladder wall. The bladder may contain a volume of a
fluid, such as urine, as shown in FIG. 8. The foregoing structures
typically present imaging resolution difficulties. In particular,
an ultrasound image may fail to adequately resolve a fluid-filled
cavity, or a tissue that forms a boundary of the fluid-filled
cavity, or still other structural details present in the imaged
anatomical portion. Moreover, the foregoing structures generally
respond to ultrasound energy in a non-linear manner, so that
reflected ultrasound echoes include one or more harmonics of a
fundamental ultrasound frequency.
[0080] One measure of the non-linear behavior of various fluids and
tissues is the Goldberg number (G). G is a dimensionless quantity
that generally relates ultrasound attenuation to harmonic
distortion due to non-linear effects in a tissue or a fluid when
subjected to ultrasound energy. Accordingly, when G is about one,
non-linear effects are comparable to attenuation effects in the
tissue. When G is much greater than one, such as for water or
urine, the nonlinear processes are dominant. When G is less than
one, as in soft tissue, attenuation effects are more dominant. For
example, it is known that fatty tissue has a Goldberg number of
approximately about 0.27, while blood, liver, and muscle have a G
value of approximately about one. In contrast, fluids such as urine
have a G value of approximately about 104.
[0081] With reference now also to FIG. 9, the scan plane 42 may
include at least one scan line 102 that extends through a bladder
and a uterus of a female patient. In this condition, referred to as
a "case 1" condition in FIG. 9; the bladder includes a relatively
large volume of urine. Typically, the bladder and the uterus appear
as low echogenicity regions when the anatomical region shown in
FIG. 9 is scanned. Known image processing software (incorporated
herein by reference from one or more of the references listed in
the priority claim section) may be used to image the shallowest
region of low echogenicity. Since the low echogenicity region is
generally preferentially selected, so that no imaging ambiguity
exists, and the bladder is therefore readily identifiable.
[0082] Referring now to FIG. 10, a typical echo amplitude response
for the anatomical region of FIG. 9 is shown. The echo amplitude
response as depicted in FIG. 10 may be obtained through application
of one or more of the algorithms incorporated by reference herein.
For example, the computational algorithms disclosed in U.S. Pat.
No. 6,676,605 to Barnard et al, U.S. Pat. No. 5,235,985 to McMorrow
et al, and U.S. Pat. No. 4,926,871 to Ganguly et al, may be
used.
[0083] When both organs are relatively filled with a fluid (as
shown in FIG. 9), the edges of the bladder and uterus are
relatively detectable and are thus generally distinguishable. In
this case, the relatively full bladder presents a relatively
U-shaped valley at a shallower bodily depth. In contrast, a
corresponding U-shaped plateau presented by the uterus is generally
identifiable at a greater bodily depth. Thus, while the embodiments
of the present invention can improve accuracy and diagnosis in the
foregoing situation where both organs are relatively filled with a
bodily fluid, the embodiments are also suited to imaging bodily
regions when one or both of the organs is less than full.
[0084] FIG. 11 is a diagrammatic representation of the anatomical
region of FIG. 9, wherein the at least one scan line 104 extends
through the bladder and the uterus when the bladder contains a
relatively low volume of urine. This condition is referred to as
"case 2" in FIG. 11. Because the bladder volume is greatly reduced
in the Case 2 situation in comparison to the Case 1 condition shown
in FIG. 9, the low echogenicity region is now principally located
in the uterus. In current image processing methods, an average of
the low echogenicity regions is compared to a threshold value to
distinguish between the bladder and the uterus, which may tend to
contribute to imaging errors. FIG. 12 is an echo amplitude response
corresponding to the anatomical region of FIG. 11. The relatively
empty bladder presents a relatively narrow valley. In contrast, the
uterus generates a relatively wider U-shaped valley. As a
consequence, the bladder is less readily distinguishable from the
uterus when the bladder contains a low volume of urine. The
disclosed embodiments better address the foregoing problems by
using algorithms that more accurately detect ultrasound signal
harmonic differences between cavity-residing fluids adjacent to
enclosing tissue interfaces.
[0085] FIG. 13 is a diagrammatic representation that will be used
to illustrate a method for imaging an anatomical region according
to an embodiment of the invention. An A-mode scan line 106 is
projected into a first bodily cavity 110, for example, a bladder
and a uterus as depicted in FIGS. 9 and 11 for the respective case
1 (which corresponds to a relatively full bladder) and case 2
(which corresponds to a nearly empty bladder). The projected
ultrasonic waveform is altered by the tissue along the distance
d.sub.1 of the scan line 106, which corresponds to tissue preceding
the cavity 110, and by the distance d.sub.2 that corresponds to an
interior portion of the cavity 110 along the scan line 106. Other
intervening cavities along the scan line 106, such as a uterus, may
also alter the projected ultrasound waveform.
[0086] The first cavity 110 is hypoechoic and designated as region
R.sub.H. Other differentially echoic regions illustrated in the
scan plane 42 include hypoechoic regions R.sub.3, R.sub.4, and
R.sub.5. Ultrasound energy passing through and within the body
cavities 110 along scan line 106 may be subjected to an image
analysis algorithm to determine respective volumes of each cavity,
namely VI corresponding to the cavity 110. Signals reflected from
the back wall or other boundaries of the first body cavity 110 are
window function processed in a window region designated as WR1. The
WR1 region spans a portion of the tissue adjacent and distal to the
cavity 110 backwall interface and further spans a portion of the
cavity space along the scan line 106 proximate to the cavity
110-backwall interfaces.
[0087] In the WRI region, window function processing determines the
raw data comprising the fundamental frequency f.sub.0 and a
selected higher order harmonic 2f.sub.0 that is generated within
the WRI space along the scan line 106. The magnitude of the higher
order harmonic generated within WRI varies because different
tissues and/or fluids are encountered by the scan line 106 as it
projects into the body. Consequently, a fluid volume and a fluid
composition within the cavity 110 alters the magnitude of the
higher order harmonic 2f.sub.0 near the scan line 106 that is
proximate to the back wall interface of the cavity 110.
[0088] FIG. 14 is a spectral plot of the insonified region of FIG.
13 that corresponds to a non-pregnant female with a uterus and
nearly full bladder. The fundamental frequency f.sub.0 has a peak
value 140 and the higher order harmonic 2f.sub.0 has a peak value
142. The fundamental and harmonic peaks 140 and 142 are a result of
window function processing the corresponding echo amplitude
response. The magnitude of the harmonic peak 142 may be normalized
by dividing the peak 142 by the fundamental frequency peak 140.
Accordingly, it is noted that a high Goldberg number stemming from
urine in the nearly full bladder corresponds to a high magnitude
for the frequency ratio. Different urine volumes and/or the
presence of other organs, such as a uterus may also alter the
magnitude of the frequency ratio. The magnitude of the second
harmonic peak 142 in the first cavity 110 is affected by the
presence of the uterus and the urine volume and urine composition
contained within the bladder. The composition and volume of the
urine may thus be determined.
[0089] FIG. 15 is a spectral plot of the insonified region of FIG.
13 in a non-pregnant female with a uterus and nearly empty bladder
and having greater amounts of blood and tissue containing uterine
fluid. The spectral plot is within the window region WRI and shows
the fundamental peak 150 and the higher order harmonic peak 152
resulting from frequency domain processing of the corresponding
echo response. As shown, the harmonic spectrum within WRI of the
non-pregnant female exhibits a generally lower Goldberg number than
a non-pregnant female with a substantially greater fluid volume.
The magnitude of the frequency ratio (2f.sub.0/f.sub.0) is
correspondingly lower. Urine fluids mixed with blood at variable
compositions may also alter the magnitude of the higher order
harmonic peak 152 shown in FIG. 15. Accordingly, the reflected
spectral components generated within the first cavity 110 may have
a still lower harmonic ratio (2f.sub.0/f.sub.0) as compared to
unmixed uterine fluids depicted in FIG. 14, since the mixed fluid
mixtures generally exhibit a lower Goldberg number.
[0090] With reference still to FIGS. 14 and 15, an embodiment of
the invention will now be described. The present embodiment is
based on non-linear wave propagation and variations in the
attenuation of ultrasound energy in body fluids. The back wall
ultrasound spectrum is processed to determine a reflected harmonic
content, and this harmonic content is compared to the content of
the fundamental ultrasound energy by forming the frequency ratio.
The resulting value may then be adjusted for differences in
attenuation at a selected frequency j between f.sub.0 and 2f.sub.0
in the intervening tissue i, as described below.
[0091] For a selected window, the total attenuation of the
fundamental frequency component may be expressed as follows in
equation E1:
A.sub.f0=2d.sub.1.sigma..sub.11+2d.sub.2.sigma..sub.21.apprxeq.2d.sub.1.-
sigma..sub.11 dB E1 [0092] where: .sigma..sub.ij=attenuation
coefficient of a tissue i at a frequency j and distances d.sub.1
and d.sub.2 are as shown in FIG. 13.
[0093] While the total attenuation of the higher order harmonic
frequency component may be expressed as follows in equation E2:
A.sub.2f0=2d.sub.1.sigma..sub.12+2d.sub.2.sigma..sub.22.apprxeq.2d.sub.1-
.sigma..sub.12 dB E2
[0094] A difference in the attenuation of higher order harmonic
component to the fundamental component is therefore defined by
equation E3:
A.sub.ratio=2d.sub.1(.sigma..sub.21-.sigma..sub.11) dB. E3
[0095] For example, in soft tissue having an attenuation factor of
about 1.1 dB/cm, the attenuation coefficient .sigma..sub.12 is
approximately about 3.7 dB/cm, when f.sub.0=3.7 MHz. Accordingly,
the amplitude ratio becomes in equation E4:
A.sub.ratio=2d.sub.1(3.7 dB/cm) E4
[0096] Based upon the foregoing, harmonic power amplitudes and
frequency ratios may be derived and associated with known fluid
volumes and fluid compositions derived from living subjects and/or
non-living experimental devices, which are then encoded into
readily accessible look-up tables, calibration plots or other
suitable means for encoding data having utility in measuring the
fluid volumes or identifying fluid compositions in an insonified
subject.
[0097] FIG. 16 is an example of a calibration plot of harmonic
power as a function of the bladder volume in a subject. The
harmonic power may be obtained from a look-up table that includes
data corresponding to different bodily tissues and fluids. The
calibration plot thus permits the determination of a urine volume
when the higher order amplitude and the fundamental frequency
amplitudes are expressed as the ratio 2f.sub.0/f.sub.0. Alternate
embodiments may include calibration plots for other fluid
compositions that stem from data in respective look-up tables that
are enhanced by the foregoing window functions. For example, a
mixture of amniotic fluid and blood, or amniotic fluid/blood/urine
mixtures would have a particular calibration plot. The plot shown
in FIG. 16 may alternately be expressed using a suitable curve
fitting procedure, such as, for example, a linear least squares
procedure, a spline procedure, a polynomial procedure or other
known curve fitting procedures.
[0098] In still another particular embodiment, the Goldberg number
may be used to distinguish between a urinary bladder and a uterus
in post-void scans in a female subject. Because the uterus
generally appears as a dark structure in ultrasound images, it may
be erroneously identified as the urinary bladder in post-void
scans. To avoid this, the present embodiment provides that the
harmonic amplitudes are calculated on a post-void scan of a female
bladder. If a selected combination of harmonic amplitudes is higher
than a selected threshold value, the scan likely contains a fluid.
Otherwise, the scan likely contains the uterus of the female
subject.
[0099] In still yet another particular embodiment, the harmonic
amplitudes may be calculated based upon one or more selected
ultrasound lines, or upon all of the ultrasound lines within the
total image. In addition, the harmonics may be calculated within a
region-of-interest or lines-of-interest when guided by features
detected on the B-mode image, such as a region behind a posterior
wall of a fluid-filled cavity.
[0100] In still another particular embodiment, inter-patient
variability in the harmonic amplitudes may be normalized or
otherwise adjusted using a combination of transmission features,
such as, for example, frequency, peak-to-peak voltage, pulse
length, and other transmission features. Other features may be
extracted from B-mode images, such as a depth of an anterior wall
of a fluid-filled cavity.
[0101] FIG. 17 is a calibration plot of a harmonic ratio expressed
as a function of a blood composition in the amniotic fluid. A
dashed calibration line may be used to graphically determine a
blood percentage composition. The Goldberg number may also be used
to differentiate between various bodily fluid compositions, such
as, for example, a transudate and an exudate fluid developed in a
lung of a subject, or in compositions of urine. It may also be used
to detect the presence of blood within a bodily fluid, such as in
amniotic fluid, or in urine. The harmonic ratio differences between
the fluids may thus be used to identify a composition of a fluid.
For example, as shown in FIG. 17, a harmonic ratio of 3 corresponds
to an amniotic fluid having a blood composition of approximately
20%. Similarly, a harmonic ratio near 4.5 corresponds to an
amniotic fluid having approximately 60% blood composition.
[0102] In still yet another particular embodiment, the Goldberg
number may also be used to distinguish between blood and amniotic
fluid in the uterus of a pregnant female. Because blood has a
significantly lower Goldberg number compared to amniotic fluid, the
second harmonic distortion resulting from a region containing blood
is different from a region containing amniotic fluid. The Goldberg
number and the harmonic ratio may thus be utilized to differentiate
between blood and amniotic fluid and to confirm an identification
of the fluid that is isonofied. For example, certain pregnant
subjects having a low amniotic fluid volume. Accordingly, the
uterus becomes more engorged with blood, making identification of
the amniotic fluid regions more difficult. To address this problem,
the embodiments of the present invention utilize the Goldberg
number and harmonic ratio to assist in the identification process.
In a further example, blood appears very dark and similar in
appearance to amniotic fluid in a B-mode image. Thus, while
measuring an amniotic fluid volume in B-mode images, blood may be
detected in the umbilical cord or in the vessels in walls of the
uterus and erroneously identified as amniotic fluid. Using the
second harmonic distortion in conjunction with the Goldberg Number
assists in discriminating between amniotic fluid and blood.
[0103] In still another particular embodiment, the Goldberg number
may also be used to identify, classify, and measure a volume of a
fluid in the lungs for pleural effusion.
[0104] In another embodiment, an ultrasound image may be created
that shows selected combinations of harmonic amplitudes throughout
the ultrasound image that permits detection of various fluid
regions that would typically be indicated by presence of higher
harmonics. The selected combination of the harmonic reflections may
be embodied in a software program, or alternately as an improvement
to an existing software program in conventional harmonic imaging
ultrasound equipment. The harmonic ratio may then be normalized by
the factor, A.sub.ratio. The results may then be compared to an
empirically derived look up table or calibration plot that
describes one of the attenuation or harmonic characteristics for
each kind of tissue (see FIG. 16). Accordingly, tissues and bodily
fluids with low Goldberg numbers and low harmonic ratios can thus
be differentiated from tissue and body fluids with higher Goldberg
numbers and higher harmonic ratios. The use of the foregoing
harmonic ratios and look-up tables (or equivalent calibration
plots) in the manner described provides another basis to further
differentiate and enhance a display of fluid and tissue regions.
For example, a low volume to near empty bladder may be suitably
differentiated from a uterine cavity and adjacent tissue.
[0105] In another embodiment of the invention, an additional
adjustment may be performed to compensate for a "shock formation
distance" that may occur along a given scan line. Shock formation
distances relate to Goldberg numbers as a consequence of an energy
transfer that occurs in the tissues and nearby fluids as the
fundamental ultrasound frequency is transformed to a harmonic
frequency. For example, an excitation frequency of 3.7 MHz results
in significant harmonic generation in human tissue. Thus, the
distance d.sub.2 (FIG. 13) may also be used to adjust the harmonic
amplitude ratio.
[0106] FIG. 18 illustrates a flow chart of a method 200 of
measuring fluid volumes and classifying fluid compositions,
according to an embodiment of the invention. The method 200 begins
by creating a database that includes attenuation and/or harmonic
characteristics of tissues, fluids, and cavities of a body at block
202. The database may be further characterized by sex, age,
morphological, physiological, and pathological states. The method
200 continues by isonifying a selected region of a patient at block
204. Thereafter, ultrasonic patterns of the insonified region of
the patient are compiled at block 206 (see FIG. 19 below). The
method 200 continues by processing the ultrasonic patterns and
comparing the processed patterns to the database information at
block 208. For example, the processed patterns may be compared
using the volume calibration plot of FIG. 16 and the composition
calibration plot of FIG. 17. Thereafter, the method 200 concludes
by determining at least one of a composition of the insonified
region and a volume of the insonified region based upon the
comparison of the patient's ultrasonic patterns to the database's
content at block 210. Ultrasonic measurement data obtained from the
subject at block 210 may then be applied to a volume calibration
plot similar to FIG. 16 to obtain volume measurements. Similarly,
ultrasonic measurement data obtained from the subject at block 210
can be applied to a composition calibration plot similar to FIG. 17
to obtain fluid composition measurements.
[0107] Still referring to FIG. 18, the processed patterns from the
subject may be compared to the look-up table or calibration plot as
depicted in FIG. 16 to obtain volume information. To obtain a
compositional determination or classification of a given detected
fluid, in accord with FIG. 17, the type of bodily fluids contained
within a cavity may be ascertained through a comparative analysis
of the Goldberg numbers, harmonic ratios, and attenuation factors
within the insonified region. The Goldberg numbers, harmonic
ratios, and attenuation factors stored within a database may then
be accessed to determine the fluid composition. A fluid composition
may thus be determined by accessing a calibration plot,
interpolating from a look-up table, or applying regression
analysis, as previously described. The volume and compositional
look up tables or calibration plots illustrated in FIGS. 16 and 17
may be obtained from ultrasound information databases derived from
simulated human models, accumulating clinical measurements obtained
from patients stored in a separate database, or a combination of
simulated models and clinical measurements. Other databases may
include simulated animal models with or without a veterinary-based
database. Yet other databases may include a combination of human
and veterinary sourced databases.
[0108] Similarly, tissue types or combinations thereof within the
insonified region are determined by comparative analysis between
the Goldberg numbers, harmonic ratios, and attenuation factors
presented by the insonified region of the patient to those same
numbers, ratios, and factors stored in the database.
[0109] FIG. 19 is an expansion of sub-algorithm 206 of FIG. 18.
Sub-algorithm 206 permits the determination of volume of an organ
wall, the mass of an organ wall, the internal organ volume defined
by an inner perimeter of an organ wall, and the outer organ volume
defined by the outer perimeter of an organ wall from echogenic
patterns received from an insonified region. The ultrasonic
patterns of the insonified region having at least one organ of the
patient are compiled at process block 206-2. Once the wall
locations are identified, the wall locations, demodulated magnitude
data, and a subset of quadrature amplitude demodulated signals in
the region of the anterior bladder wall are directed to the
microprocessor for further analysis according to the algorithm
illustrated in FIG. 19 for the particular embodiments. First,
ultrasound data is acquired relative to the bladder, uterus, or
other organs as shown in the first block 206-2. In general,
bladder-specific data can be acquired by a user who manipulates the
transceiver 10 while viewing the received data on a display screen
and then positioning the transceiver 10 as necessary so that an
organ or organs, such as a bladder and uterus, are sufficiently
within the field of view of the cone as depicted in FIGS. 2 and
3A.
[0110] Referring again to FIG. 19, and limiting the discussion to a
specific organ, for example a bladder, echogenic data is collected
by the transceiver 10. After obtaining ultrasound bladder data, the
ultrasound data is processed to determine if the bladder contains
approximately 200 to approximately 400 ml, as shown in the second
process block 206-4 represented as a decision diamond. If "No" to
the query "200 ml.ltoreq.volume.ltoreq.400 ml?", then the bladder
is allowed to accumulate approximately 200 to approximately 400 ml,
as shown in the third process block 206-6, or, if "Yes", meaning
the bladder already contains the preferred approximate 200-400 ml
volume, then the locations of the bladder walls, as shown in the
fourth block 206-8, may be undertaken. The determination of organ
wall locations and other such exterior boundaries within an
ultrasound scan are within the capability of ultrasound devices
presently on the market. In general, however, the process
determines the length of a scan line from the transceiver dome to
the bladder wall. The data, including wall locations, is stored in
the memory of the computer 62 and is used to determine whether or
not the bladder volume is within a range of approximately 200 to
approximately 400 ml. If the bladder volume is within that range,
the ultrasound data is used to determine the actual surface area
from the wall locations, as indicated in the fifth block 206-10.
The application of previously described methods using harmonic
ratios, powers, and Goldberg G-numbers may be used to enhance the
accuracy of thickness, area, volume, and mass determinations of
bladders holding fluids within the approximate 200-400 ml range.
The surface area calculation is explained with regard to FIG. 21
below and allows for calculation of an outer bladder wall surface
area defined by subserosal locations 372A and 372B and an inner
bladder wall surface area defined by submucosal locations 374A and
374B. While calculating the surface area in the fifth block 206-10,
reflected ultrasound waves are received from the anterior bladder
wall, as indicated in the sixth block 206-12. Although these tasks
are preferably conducted in parallel, they may alternatively be
processed in series. Thereafter, as shown in the seventh block
206-16, the bladder wall thickness is determined from the coherent
signals that overlap at the wall locations. The determination of
bladder wall thickness is explained in greater detail below.
Finally, as shown in the seventh block 206-16, the bladder wall
distance is computed as a difference between panterior and
posterior submucosal bladder wall locations. Thereafter, at the
eighth process block 206-20, the internal bladder volume is
computed as a function of the internal bladder wall distances and
the area of the internal bladder wall.
[0111] The volume restriction indicated in the previous paragraph
is included as the range of bladder volumes that allow for an
optimal measurement of the bladder wall mass, bladder wall volume,
and internal bladder volumes. The volume and mass calculations may
be performed at a volume not in this range, but will result in a
less accurate measurement that can be corrected by application of
the foregoing described methods using harmonic ratios, powers, and
Goldberg G-numbers. For example, bladders having less than 200 ml
or that are near empty, the foregoing described methods using
harmonic ratios, powers, and Goldberg G-numbers will improve the
accuracy of determining bladder wall thicknesses, volumes and mass,
and internal and outer bladder volumes. For bladders having fluid
volumes substantially greater than 400 ml, for example bladder
volumes of 1000 ml to multi-liters, the invention will utilize scan
lines greater than 20 cm to accommodate the larger bladder sizes.
The invention may be applied to measure the thicknesses, masses,
and volumes of internal organs of human and animals. The length of
the scan lines is adjusted to match the dimension of the internal
organ scanned.
[0112] The surface area measurement of fifth block 206-4 is
performed by integrating the area of interpolating surface patch
functions defined by the wall locations. The mathematical
calculations are provided below in greater detail.
[0113] The surface of the bladder is defined to be S. This surface
corresponds to the actual surface of the bladder determined by
analysis of the wall locations of the bladder. Since this shape is
not known in advance, modeling the bladder as a sphere or an
ellipsoid provides only a crude approximation of the surface.
Instead, the surface S is defined as a construction of a series of
individual surface patches s.sub.i,j, where i and j count through
the latitude and longitude components of the surface, similar to
the division of the Earth's surface into lines of latitude and
longitude. The area of the bladder surface, S, is defined as the
sum of all the individual surface patches, so that
S=.SIGMA.s.sub.i,j.
[0114] FIG. 20 is a schematic representation of four surface patch
elements. As depicted in three dimensions in FIG. 20, by way of
example, five scan planes 320-328 are seen transmitted
substantially longitudinally across a subserosal wall location 332
referenced to a tri-axis plotting grid 340. The five scan planes
include the first scan plane 320, the second scan plane 322, the
third scan plane 324, the fourth scan plane 326, and the fifth scan
plane 328. The scan planes are represented in the preceding
formulas as subscripted variable j. Substantially normal to the
five longitudinal scan planes are five latitudinal integration
lines 360-368 that include a first integration line 360, a second
integration line 362, a third integration line 364, a fourth
integration line 366, and a fifth integration line 368. The
integration lines are represented in the preceding formulas as
subscripted variable i.
[0115] By way of example, four surface patch functions are
highlighted in FIG. 20 as the subserosal wall location 372. The i
and j subscripts mentioned previously correspond to indices for the
lines of latitude and longitude of the bladder surface. For the
purposes of this discussion, i will correspond to lines of
longitude and j will correspond to lines of latitude although it
should be noted the meanings of i and j can be interchanged with a
mathematically equivalent result. Using the scan plane and
integration line definitions provided in FIG. 20, the four surface
patch functions are identified, in the clockwise direction starting
in the upper left, as s.sub.322,362, s.sub.324,362, s.sub.324,364,
and s.sub.322,364.
[0116] The surface patches are defined as functions of the patch
coordinates, s.sub.i,j(u,v). The patch coordinates u and v, are
defined such that 0<u, v<1 where 0 represents the starting
latitude or longitude coordinate (the i and j locations), and 1
represents the next latitude or longitude coordinate (the i+1 and
j+1 locations). The surface function could also be expressed in
Cartesian coordinates where
s.sub.i,j(u,v)=x.sub.i,j(u,v)i+y.sub.i,j(u,v)j+z.sub.i,j(u,v)k
where i, j, k, are unit vectors in the x-, y-, and z-directions
respectively. In vector form, the definition of a surface patch
function is given in Equation 1. k, are unit vectors in the x-, y-,
and z-directions respectively. In vector form, the definition of a
surface patch function is given in equation E5.
s i , j ( u , v ) = [ x i , j ( u , v ) y i , j ( u , v ) z i , j (
u , v ) ] E5 ##EQU00001##
[0117] With the definitions of surface patch functions complete,
attention can turn to the surface area calculation represented in
the fifth block 206-10 of FIG. 20. The surface area of S, A(S), can
be defined as the integration of an area element over the surface
S, as shown in equation E6.
A ( S ) = .intg. s A E6 ##EQU00002##
[0118] Since S is composed of a number of the patch surface
functions, the calculation for the area of the surface S can be
rewritten as the sum of the areas of the individual surface patch
functions as in equation E7.
A ( S ) = i , j A ( s i , j ) . E7 ##EQU00003##
[0119] Similarly to equation E5 for the entire surface, the area of
the surface patch is the integration of an area element over the
surface patch, shown in equation E8.
A ( S i , j ) = .intg. s i , j A i , j E8 ##EQU00004##
[0120] The integration over the surface patch function can be
simplified computationally by transforming the integration over the
surface to a double integration over the patch coordinates u and v.
The transformation between the surface integration and the patch
coordinate integration is shown in equation E9.
.intg. s i , j A i , j = .intg. u = 0 1 .intg. v = 0 1
.differential. s i , j .differential. u .times. .differential. s i
, j .differential. v v u E9 ##EQU00005##
[0121] By substituting Equation 5 into Equation 4, and Equation 4
into Equation 3, the area for the entire surface can be calculated.
The result of these substitutions is shown in equation E10.
A ( S ) = i , j .intg. u .intg. v .differential. s i , j
.differential. u .times. .differential. s i , j .differential. v v
u E10 ##EQU00006##
[0122] The surface patch function may be any function that is
continuous in its first derivatives. In the embodiment shown, a
cubic B-spline interpolating function is used for the interpolating
surface patch function although any surface function may be used.
This interpolating function is applied to each of the Cartesian
coordinate functions shown in equation E5. The interpolating
equation for the x-coordinate of the s.sub.i,j patch function is
given in equation E11.
x.sub.i,j(u,v)=uM.sub.bX.sub.i,jM.sub.b.sup.tv.sup.t E11
u = [ u 3 u 2 u 1 ] , v = [ v 3 v 2 v 1 ] , ##EQU00007##
[0123] where t denotes matrix and vector transpose,
M b = [ - 1 3 - 3 1 3 - 6 3 0 - 3 0 3 0 1 4 1 0 ] , and
##EQU00008## X i , j = [ x i - 1 , j - 1 x i - 1 , j x i - 1 , j +
1 x i - 1 , j + 2 x i , j - 1 x i , j x i , j + 1 x i , j + 2 x i +
1 , j - 1 x i + 1 , j x i + 1 , j + 1 x i + 1 , j + 2 x i + 2 , j -
1 x i + 2 , j x i + 2 , j + 1 x i + 2 , j + 2 ] ##EQU00008.2##
[0124] Similar calculations are performed for the y.sub.i,j and
z.sub.i,j components of the surface patch function.
[0125] Since the interpolating functions for each of the patch
functions is a cubic surface, the integration may be performed
exactly using a quadrature formula. The formula used in this
application is shown in equation E12.
A ( s i , j ) = i , j 1 4 ( .differential. s i , j .differential. u
.times. .differential. s i , j .differential. v u = 3 - 3 6 , v = 3
- 3 6 + .differential. s i , j .differential. u .times.
.differential. s i , j .differential. v u = 3 - 3 6 , v = 3 + 3 6 +
.differential. s i , j .differential. u .times. .differential. s i
, j .differential. v u = 3 + 3 6 , v = 3 - 3 6 + .differential. s i
, j .differential. u .times. .differential. s i , j .differential.
v u = 3 + 3 6 , v = 3 + 3 6 ) E12 ##EQU00009##
[0126] Recalling the fact that s.sub.i,j(u,v) is defined as a
vector function in Cartesian coordinates (Equation 1), the norm of
the cross product of the partial derivatives can be written as
follows in equation E13.
.differential. s i , j .differential. u .times. .differential. s i
, j .differential. v = ( .differential. y i , j .differential. u
.differential. z i , j .differential. v - .differential. z i , j
.differential. u .differential. y i , j .differential. v ) 2 + (
.differential. z i , j .differential. u .differential. x i , j
.differential. v - .differential. z i , j .differential. u
.differential. x i , j .differential. v ) 2 + ( .differential. x i
, j .differential. u .differential. y i , j .differential. v -
.differential. y i , j .differential. u .differential. x i , j
.differential. v ) 2 E13 ##EQU00010##
[0127] When the physical x-, y-, and z-locations are used in the
interpolating function, the surface are will be calculated in the
square of the units of x, y, and z. At this point the calculation
in the fifth block 206-10 of FIG. 20 is complete.
[0128] The second component to the mass calculation is a
measurement of the thickness of the bladder muscle wall. This
thickness is defined to be the normal thickness between the
subserosal and submucosal surfaces of the bladder wall.
[0129] The wall thickness is calculated from the fractal dimension
of the RF signal in the region of the wall thickness. The fractal
dimension increases due to the multiplicity of interface
reflections through the bladder muscle. The increase and decrease
of fractal dimension through the bladder muscle wall can be modeled
as a parabola where the fractal dimension is a function of the
depth in the region of the bladder wall. The thickness of the
bladder is then determined to be the region of the parabola model
that is at least 97% of the maximal value of the fractal dimension.
The calculations are reviewed below in equation E14.
fd r = log ( max ( RF r = r - w / 2 , r + w / 2 ) - min ( RF r = r
- w / 2 , r + w / 2 ) + w w ) log ( n w ) E14 ##EQU00011##
[0130] The wall thickness is calculated from the fractal dimension
of the RF signal in the region of the wall thickness. The fractal
dimension increases due to the multiplicity of interface
reflections through the bladder muscle. The increase and decrease
of fractal dimension through the bladder muscle wall can be modeled
as a parabola where the fractal dimension is a function of the
depth in the region of the bladder wall. The thickness of the
bladder is then determined to be the region of the parabola model
that is at least 97% of the maximal value of the fractal dimension.
The calculations are reviewed below in equation 15.
[0131] The fractal dimension calculation corresponds to the fourth
block 206-12 of FIG. 20. The fractal dimension is calculated for a
window of length w. In the current embodiment the value of w is 5,
the number of sample points along a scan line, although that value
can be varied. The fractal dimension is calculated from the
difference between the maximum RF signal value in the window
centered at a given depth, r, and the minimum of that same window.
The length of the window, w, is added to this difference, and the
result is then normalized with the length of the window. The
logarithm of that result is then divided by the logarithm of the
ratio of the total number of samples in a scan line, n, to the
length of the window. The calculation of the fractal dimension at
each depth along a scan line is shown in Equation 10. This fractal
dimension measure is calculated for the central n-w samples in a
scan line.
[0132] After the measurements of the fractal dimension have been
calculated based on the ultrasound signal, the thickness of the
bladder wall may be calculated. The following calculations
correspond to the seventh block 206-16 of FIG. 19.
[0133] The fractal dimension, fd, of the RF signal in the region of
the bladder muscle wall is then modeled as a parabolic equation as
a function of depth, r. The model of the equation for a single
depth point is given in equation E15. In that equation, there are 3
parameters (a, b, and c) that define the parabola with the depth
along a scan line r, and the addition of a random element
.epsilon.. The subscript i indicates a specific value of r,fd, and
.epsilon..
fd.sub.i=ar.sub.i.sup.2+br.sub.i+C+.epsilon..sub.i E15
[0134] An equation of the form in equation E15 is obtained for each
depth point in the region of the wall. The number of observations
is variable and depends on the thickness of the bladder wall as
observed by the ultrasound signal. Assuming a set of n
observations, the subscript i would count the observations from 1
to n. The set of n equations of the form in equation 15 may be
compressed into a matrix equation given in equation E16.
fd = X .beta. + where fd = [ fd 1 fd 2 fd n ] , X = [ r 1 2 r 1 1 r
2 2 r 2 1 r n 2 r n 1 ] , .beta. = [ a b c ] , and = [ 1 2 n ] E16
##EQU00012##
[0135] Each row of the fd, and .epsilon., and the X matrix
correspond to one of the n observations. The parabola parameters of
equation E16 are collected in the vector P.
[0136] The next step is to estimate the values of the parameters of
the parabola in the set of n equations of the form in equation E15
or in the matrix equation E16 based on the set of observations. A
least-squares estimation of the parameters is used, and the
calculation for these estimates is shown in equation E17. In E17,
the t superscript indicates matrix transpose, and the -1
superscript indicates the matrix inverse. Parameters with hats (A)
indicate that the value is the least-squares estimate of those
parameters.
{circumflex over (.beta.)}=(X.sup.tX).sup.-1X.sup.tfd E17
[0137] The estimates of the parabola parameters ({circumflex over
(.beta.)}=.left brkt-bot.a {circumflex over (b)} c.right
brkt-bot.') can be substituted into the parabola model to calculate
the estimated fractal dimension at each depth r, as shown in
equation E18. The location of the maximum fractal dimension can be
determined by setting the first derivative of the parabola model to
equal 0 (equation E19) and solving for r. The location where the
fractal dimension is maximal is given in equation E20.
f d ^ ( r ) = a ^ r 2 + b ^ r + c ^ E18 f d ^ ( r ) r = 2 a ^ r + b
^ = 0 E19 r fd max = - b ^ 2 a ^ E20 ##EQU00013##
[0138] To determine the maximal fractal dimension as defined by the
parabolic model, simply substitute equation 20 into equation 18 and
solve for fd.sub.max. The resulting value is shown in equation
E21.
f d ^ max = - b ^ 2 + 4 c ^ 4 a ^ . E21 ##EQU00014##
[0139] To determine the locations where the fractal dimension is
97% of the maximum value, multiply equation E21 by 0.97, substitute
the result into equation E18 and solve for r using the quadratic
formula. The locations where the fractal dimension is 97% of the
maximum value, r.sub.97%, are given in equation E22.
r 97 % = - b ^ .+-. b ^ 2 - 4 a ^ ( c ^ + 0.97 b ^ 2 + 4 c ^ 4 a ^
) 2 a ^ E22 ##EQU00015##
[0140] Two values for r.sub.97% will be calculated from Equation
18. The difference between those two values will identify the
thickness of the bladder muscle wall along the given scan line.
Since these scan lines may or may not be perpendicular to the
bladder muscle surface and bladder wall thickness must be measured
along a line perpendicular to the bladder surface, a collection of
these measurements are combined to determine the actual thickness
of the bladder wall.
[0141] These measurements could be made at any surface of the
bladder muscle wall. In FIG. 22, three scan lines are shown to
cross the bladder muscle in two locations: the anterior wall
closest to the transducer, and the posterior wall furthest from the
transducer. The parabolic model described previously can be applied
twice on each to determine the thickness of both the anterior and
posterior wall. The maximum and minimum and mean values of these
thicknesses are used in the mass calculation and historical
tracking of data. In the embodiment shown, this final thickness
determination marks the end of the process identified in the
seventh block 206-16 of FIG. 19.
[0142] FIG. 21 is a schematic representation of three scan lines
passing through the subserosal and submucosal wall locations of an
organ, here schematically illustrated for a bladder. Three scan
lines 362, 364, and 366 penetrate the bladder. The dotted portion
of the lines represents the portion of the scan lines that passes
through the bladder muscle wall at an anterior or front wall
location 370A and a posterior or back wall location 370B. The first
362, the second 364, and the third 366 scan lines are shown
transmitting through the front subserosal wall location 372A and
front submucosal wall location 374A. Similarly, the first 362, the
second 364, and the third 366 scan lines are shown transmitting
across the internal bladder region 375 and through the back
submucosal wall location 374B and back submserosal wall location
372B. The front and back subserosal locations 372A and 372B occupy
an outer bladder wall perimeter and the front and back submucosal
locations 374A and 374B occupy an inner bladder wall perimeter. A
bladder wall thickness value 376 is obtained for the respective
differences along each scan line 362-366 between the subsersosal
wall locations 372A and the submucosal wall locations 374A, or the
subserosal wall locations 372B and the submucosal wall locations
374B. The maximum and minimum and mean values of these thicknesses
are used in the bladder wall mass calculation and historical
tracking of data. In the preferred embodiment, the bladder is
assumed to have a uniform wall thickness, so that a mean wall
thickness value is derived from the scanned data and used for the
determination of the internal wall volume 375. Only three scan
lines are shown in a plane, each separated by 7.5 degrees from each
other. Both the number of scan lines in the plane and the angles
separating each scan line within a plane may be varied.
[0143] Once the bladder wall thickness and the inner and outer
surface area have been measured, the volume of the internal bladder
region 375 may be calculated by the determining the respective
differences between the front and back submucosal wall locations
374A and 374B along each scanline penetrating the bladder region
375. The difference between the front and back submucosal wall
locations 374A and 374B defines an inter-submucosal distance. The
internal volume of the bladder region 375 is then calculated as a
function of the inter-submucosal distances of the penetrating scan
lines and the area of the subserosal boundary or internal bladder
perimeter. The volume of internal region 375 is assumed to be the
surface area times a function of the inter-submucosal distances,
where the assumption is further based on a uniform wall subserosal
boundary at all points around the internal bladder perimeter. In
the embodiment shown, this volume calculation corresponds to the
eighth block 206-20 of FIG. 19.
[0144] The methods to obtain the wall-thickness data, the mass
data, and the volume of internal region 375 via downloaded digital
signals can be configured by the microprocessor system for remote
operation via the Internet web-based system. The Internet web-based
system ("System For Remote Evaluation Of Ultrasound Information
Obtained By A Program Application-Specific Data Collection Device")
is described in patent application Ser. No. 09/620,766, herein
incorporated by reference. The internet web-based system has
multiple programs that collect, analyze, and store organ thickness
and organ mass determinations. The alternate embodiment thus
provides an ability to measure the rate at which internal organs
undergo hypertrophy with time and permits disease tracking, disease
progression, and provides educational instructions to patients.
[0145] In summary, the foregoing method supplements current
algorithms in novel ways that advantageously allow different bodily
fluids and/or tissues to be distinguished by volume and
composition. In particular, different regions having low
echogenicity may be properly distinguished. This feature
advantageously permits shadowed regions of a fetus such as the arms
and the legs of the fetus to be distinguished from a head region.
In a diagnostic method directed to the detection of an aortic
aneurysm, the foregoing embodiments may be used to differentiate
shadowed regions resulting from bowel gas from other low echo
regions. In a gall bladder imaging scan, the foregoing embodiments
may be used to determine whether bile or other bodily fluids are
within the field of view of the ultrasound-scanning device.**
[0146] The first method describes a simple device that allows the
assessment of bladder 23 volume, using only a few beams
appropriately oriented. Under the assumption that there exists a
correlation between the bladder height and width, a simple approach
has been developed. It consists of a limited number of acoustic
beams positioned in such a way that the depth D and the height H of
the bladder could be estimated in approximately a single sagittal
plane. The volume of urine is then computed from an empirical
formula that does not assume any geometric model.
[0147] In operation of the apparatus of the present invention, the
transducer assembly 1 is placed on the abdomen of the patient in
the supine position, just above the symphysis pubis 2. We are
presenting a particular configuration of the assembly 1.
Nevertheless, various configurations can be derived from this model
and several modifications could be achieved (number of transducers,
position, orientation, etc. . . . ) without departing from the
initial ideas. The device proposed as an example is composed of
five disc shaped transducers A, B, C, D and E (focused or
non-focused) positioned in the assembly at predetermined distance
from each other (FIG. 5, top panel) and oriented at predetermined
angles .phi..sub.A, .phi..sub.B, .phi..sub.C, .phi..sub.D, and
.phi..sub.E (FIG. 6). Referring to FIG. 5 (top panel), it appears
that the transducers A, B, C, D and E are oriented in two different
planes. The distance between these two planes is small compared to
the bladder 3 size and thus we can assume that the information
received from each transducer represent the characteristics of
approximately a single sagittal or anteroposterior plane. The
orientation of each beam has been determined from the knowledge of
the bladder 3 position and shape when it is filling up as measured
in a patient series. The first beam of the transducer assembly 1
(soundbeam from transducer A) is oriented in such a way that it
reaches the bottom of the bladder, passing just above the symphysis
pubis 2. The remaining beams are positioned for successively
intercepting the bladder 3 when it expands with increasing filling
degree.
[0148] Computation of the Depth D and Height 5: Depending on the
number of beams that are intercepting the bladder 3 and on the
geometrical configuration of the transducer assembly (1), the
distances H and D are determined by different mathematical
procedures. For most measurement configurations the depth D of the
bladder is determined by the distance between echoes derived from
front and back wall of the bladder estimated from Transducer B.
[0149] The Height H (5) calculation in the specific measurement
configuration (here we selected as an example the "filled bladder"
configuration of FIG. 2) when posterior bladder wall echoes are
detected in signals obtained in beam A, B, C, and D, but not in
beam E is illustrated in FIG. 3. For the other filling geometries
the height is calculated in a corresponding way. The mathematical
procedure is as follows:
cos .0. A = [ AA 2 ] / [ AA 1 ] => [ AA 2 ] = cos .0. A [ AA 1 ]
( 1 ) sin .0. A = [ A 1 A 2 ] / [ AA 1 ] => [ A 1 A 2 ] = sin
.0. A [ AA 1 ] ( 2 ) cos .0. D = [ DD 2 ] / [ DD 1 ] => [ DD 2 ]
= cos .0. D [ DD 1 ] ( 3 ) cos .0. A = [ D 1 D 2 ] / [ DD 1 ] =>
[ D 1 D 2 ] = sin .0. D [ DD 1 ] ( 4 ) cos .0. A = [ AA 2 ] / [ AA
1 ] => [ AA 2 ] = cos .0. A [ AA 1 ] ( 5 ) cos .0. A = [ AA 2 ]
/ [ AA 1 ] => [ AA 2 ] = cos .0. A [ AA 1 ] ( 5 ) ID 1 = [ D 1 D
2 ] + [ A 1 A 2 ] + [ AD ] ( 6 ) => Height = [ A 1 D 1 ] = [ A 1
I ] 2 + [ ID 1 ] 2 ( 7 ) ##EQU00016##
[0150] Volume computation: The volume of urine is correlated to the
bladder diameter (Height 27 and Depth 26) by the empirical
formulae:
Height*Depth*K
[0151] where K is a correction factor. Depending on the number of
beams that allow the determination of the bladder dimensions (from
1 to 5) and others parameters such as the age, the gender, the
correction factor is different. For a given situation (parameters
other than number of beam are fixed), the correction factors KL,
K2, K3, K4 and K5 are optimized using linear regression
analysis.
[0152] The process executed by the hardware is illustrated in the
flow chart of FIG. 4.
[0153] After positioning the transducer assembly correctly over the
bladder area the measurement procedure is started by pressing the
start button which during the (short) measurement procedure remains
depressed. Subsequently the transducers are activated for
transmission of ultrasound pulses and reception of echoes and
possible detection of bladder wall echoes in a specific order.
Thereafter it is established, when a clear posterior bladder wall
echo is detected, which ultrasound beams, this we call here the
beams of "useful" transducers, penetrate the filled bladder. From
this, the filling situation or measurement geometry is established.
As a result the proper correction factor can be selected. After
calculation of the volume the value is stored in memory and
displayed. During the measurement procedure the transducer assembly
is slightly moved and memory data are refreshed if a larger volume
is measured. The highest value will correspond with the correct
bladder volume. This is displayed.
[0154] In a general aspect, therefore, the apparatus may use beam
information comprising at least: angle of incidence (known from the
transducer mounting angle), spatial position (known from the
transducer position in the array) and echo travel time (deduced
from the reflected beam). Other beam parameters or information from
reflected beams may also be used in accordance with known
ultrasound techniques, such as frequency, pulse rate etc.
[0155] For determining body cavity and height, the apparatus may
select only beams corresponding to those that have intercepted the
fluid filled body cavity.
[0156] The arrangements described in connection with FIGS. 1 to 6
illustrate use of five transducers. This configuration was selected
in order to achieve a selected degree of accuracy of measurement
over a complete expected range of total volumes in a human adult.
In the preferred configuration, accuracy of measurement of the
order of 100 ml over a range encompassing a bladder fill level from
0 to approximately 800 ml has been exhibited. It will be understood
that a smaller number of transducers could be used when either the
desired measurement accuracy can be reduced, or when the total fill
range covered can be reduced.
[0157] For example, using just three transducers, it has been shown
to be possible to cover a fill range of 0 to approximately 500 ml
with an accuracy of 100 ml.
[0158] Similarly, four transducers has been shown to cover a range
0 to approximately 700 ml, and two transducers, a range of 0 to
approximately 300 ml.
[0159] Such configurations can be used when it is only necessary to
indicate gross ranges of bladder filling, or to indicate a
clinically important threshold fill level.
[0160] In other embodiments, the apparatus may be provided with an
input device such as a keypad or computer interface so that the
user can enter patient information, such as gender, weight and age.
This information can then be used to ensure correct selection of an
available correction factor, K, from a memory of the apparatus.
[0161] The apparatus may also be provided with means for inputting
calibration data, such as absolute measurements of bladder fill
level separately deduced from conventional measurements. These can
be stored by the apparatus and used to optimise stored K values as
part of an iterative, `self-learning` process. In other words, the
apparatus may incorporate an algorithm for automatically adjusting
predetermined correction factors stored therein based on
calibration data entered into the machine for comparison with
measurement data taken by the apparatus.
[0162] The apparatus may also comprise a means for indicating
correct caudal-cranial positioning of the transducer array on the
body over the bladder. For example, in a normal measurement as
suggested in figure, it is expected that at least transducers A, B
and C will indicate a bladder present condition, whereas
transducers D and E might, or might not indicate bladder present,
according to the bladder fill level. In the event that, for
example, no signal is indicated by A, or by A and B, but signal is
indicated by D or D and E, then it can be deduced that the
transducer assembly is positioned too far in the cranial direction.
This could be indicated on the display of the device.
[0163] In summary, the described first method differs greatly from
known other apparatus:
[0164] 1) The device is composed of a limited number of static
single element transducers;
[0165] 2) The arrangement of the transducer is not similar to the
arrangement of a linear array;
[0166] 3) The transducers are oriented towards the bladder with
specific angles allowing the estimation of the urine volume over a
wide range of volumes;
[0167] 4) The method for automatic volume computation does not
assume any geometrical model for the bladder shape;
[0168] 5) It is valid for any bladder shape since the volume is
computed with an empirical formula for various filling ranges;
[0169] 6) It is not based only on the measurement of distances
between the front and back wall or area in different planes;
[0170] 7) It uses an automatic detection of the bladder height and
depth depending on the number of beams that intercept the
bladder;
[0171] 8) It optimizes the correction factor depending on the
degree of filling (or other factors, such as age, gender, weight,
that may influence the calculations);
[0172] 9) The device includes a closed loop to easily find the
optimal position;
[0173] 10) The optimal position corresponds to the largest volume
computed;
[0174] 11) The device works instantaneously.
DETAILED DESCRIPTION OF THE SECOND METHOD
[0175] The second version of the device is based on a different
principle. The approach consists of using a single acoustic beam
with a very wide width such that it encloses approximately the
entire volume of the bladder when it is filled up. Such a wide beam
width can be obtained using a single element transducer with a
defocusing lens as drawn in FIG. 7 or a curved single element
transducer.
[0176] The schematic principle of transducer positioning is
illustrated in FIG. 7. The sagittal cross section through the
bladder is shown. The cone like shape of the acoustic beam allows
to encompass approximately the full bladder volume, and therefore
any harmonic distortion detected in the echo signal returning from
a region beyond the posterior wall of the bladder around depth W,
would correlate to the amount of fluid contained in the
bladder.
[0177] It has been demonstrated that the propagation of ultrasound
waves is a nonlinear process. The nonlinear effects, which increase
with higher intensities, have been predicted and demonstrated at
frequencies and intensities used in the diagnostic range either in
water or in human body (A Baker et al.: "Distortion and
High-Frequency Generation Due to Non-Linear Propagation of Short
Ultrasonic Pulses From A Plane Circular Piston", J. Acoustic Soc Am
92(3), pp 1699-1705). The distortion is due to slight
non-linearities in sound propagation that gradually deform the
shape of the propagating sound wave, and result in development of
harmonic frequencies which were not present in the transmitted wave
close to the transducer. This manifests itself in the frequency
domain as the appearance of additional harmonic signals at integer
multiples of the original frequency.
[0178] These effects occur most strongly when ultrasound waves
propagate within liquids with relatively low acoustic attenuation
such as water, amniotic fluid or urine. Indeed, acoustic
propagation in fluids gives rise to extreme nonlinear effects at
diagnostic frequencies. Within soft tissues, nonlinear processes
also take place but are modified as a result of the different
acoustic characteristics of these tissues, most notably their high
acoustic absorption. Indeed, water and amniotic fluids (urine) are
significantly different from tissue.
[0179] It is known from literature (A C Baker: "Prediction Of
Non-Linear Propagation In Water Due To Diagnostic Medical
Ultrasound Equipment", Phys Med Biol 1991 VOL 36, NO 11, PP
1457-1464; T Szabo et al.: "Effects of Non-Linearity On The
Estimation Of In-Situ Values Of Acoustic Output Parameters", J
Ultrasound Med 18:33-41,1999; M Hamilton et al.: "Nonlinear
Acoustics", Academic Press) that the non-linearity of a medium is
characterized by the coefficient of non-linearity B. Typical values
for P are 3. 6 for water, 4 for blood and 6.5 for fatty tissue.
[0180] In addition to being nonlinear, all the media have
acoustical loss due to absorption. The acoustical loss is described
by the power law: A=AOFB where ao is constant and b ranges from 1
to 2 depending on the medium. For water, the rate of absorption of
an ultrasound wave propagating through it is quadratically related
to the frequency (b=2). However, the rate of energy loss due to
absorption is considered small and most of the time the
dissipation-less theory is applicable over short ranges. However,
biological media have large rates of energy loss and the frequency
dependence has an exponential value of 1 to 1.5.
[0181] By considering both attenuation due to absorption loss and
non-linearity, the exchange of energy between the two processes is
complicated, because attenuation diminishes the amplitude of the
generated harmonic components with propagation distance while
non-linearity builds up these harmonics. So, harmonic distortion
generally tends to enrich the higher harmonic components at the
expense of the lower ones (energy transfer), while absorption damps
out the higher components more rapidly than the lower ones. It is
therefore difficult to reach a balance in which a given component
loses as much energy by absorption as it gains from nonlinear
distortion. Moreover, since the conditions for stability depend on
the amplitude of the wave, which slowly decreases with propagation
distance, the wave can never be completely stable, only relatively
so.
[0182] The balance between the nonlinear process and the
attenuation process is given by the Goldberg number .GAMMA. (Szabo
et al.), which represents a measure of which process dominates.
When .GAMMA.=1, nonlinear effects are comparable to attenuation
effects. If .GAMMA. is higher than 1, nonlinear processes dominate
and when the Goldberg number is below 1, attenuation effects take
over. As indication, for acoustic pressures of 500 kPa and LMPA, at
a transmit frequency of 3 MHz, the Goldberg number is respectively
86.5 and 43.2 for water. It is only 2.8 and 1.4 for liver-like
tissue respectively at these pressures. For both settings, the
parameter shows that for water, non-linearity is up to thirty times
greater than for tissue.
[0183] The approach used here is based on the
"non-linearity/attenuation" characteristic in differentiating
between fluid media and soft tissue media. As described above, a
single element transducer is placed in front of the bladder. The
transducer generates a wide acoustic beam that is able to enclose
the full bladder volume. Depending on the volume of urine contained
in the bladder (bladder filling) and thus crossed by the acoustic
beam, the amount of harmonic distortion generated in the back of
the bladder will change. A radio frequency (RF) backscattered
signal might be selected from a region of interest located
preferably in the backside of the bladder. The amount of energy
comprised in the second harmonic or higher harmonic components of
the received RF echo signal can be extracted and correlated to the
amount of volume of urine that has been encompassed by the acoustic
beam. Since harmonic generation is different in tissue than in
fluids, only the volume of urine that has been crossed by the
acoustic beam would generate more harmonic energy. When the bladder
is empty or below a certain volume level, no harmonic distortion
occurs, whereas maximal distortion will be obtained for a full
volume.
[0184] FIG. 9 illustrates the principle of the invention. Top panel
shows two situations. The bladder is either empty (Panel A left
side) or filled up with urine (Panel A right side). At a certain
distance beyond the bladder (around 12 cm from the transducer), a
region of interest of 1.5 cm width at depth W (see FIG. 7) is
selected. Power spectra corresponding to echo signal recorded from
the regions of interest are displayed in panel B.
[0185] The spectrum corresponding to the empty bladder (solid line)
shows only a fundamental component. The harmonic distortion is very
weak so that no harmonic frequencies are generated. However, the
echo signal corresponding to the filled bladder situation (dashed
line) demonstrates clear distortion where a second harmonic
component with a significant energy is generated. The third
harmonic component can be also present with lesser energy depending
on the urine volume that has been crossed by the acoustic beam.
[0186] FIG. 9 demonstrates that depending on the volume contained
in the bladder that the acoustic beam has intersected, the amount
of generated second harmonic energy varies. When the acoustic beam
crosses only tissue or when the volume of urine is very small,
harmonic distortion is the lowest with no or very low harmonic
energy. If the bladder is filled up or if the volume of urine is
above a certain level (threshold), harmonics are generated. The
generation of a harmonic component (second and/or higher harmonics)
can be used for volume measurement, or simply as an indicator of
filling of the bladder to a certain volume extent. The criterion
can be such that if a certain amount of second harmonic (or higher
harmonics) is generated in the echo signal, the device would
indicate that the critical volume (or threshold) (say in adult
patients around 450 ml) has been reached.
[0187] To avoid and eliminate any differences due to patient to
patient variations, a normalization procedure needs to be performed
a priori. Such a normalization procedure might consist of recording
a first signal at very low transmit acoustic power from the same
region of interest as described in the previous section. Such power
would allow only linear propagation of the ultrasonic waves and
avoid any harmonic generation. The echo signal would therefore have
undergone only attenuation effects.
[0188] In the following transmit-receive sequence, the transmit
acoustic power is increased with a certain factor (e) and a new
recording is performed from the same region of interest. This
measure with a much higher acoustic pressure is carried out to
allow harmonic distortion to occur in the tissue. The echo signal
in this case will undergo both attenuation and distortion effects.
The first echo signal (linear case) will be re-scaled by the factor
that corresponded to the increase in transmit power (e), and then
used as a reference signal. Consequently, each patient has his own
reference hence eliminating any variations such as obesity,
INHOMOGENEITIES, etc.
[0189] A block diagram of a possible steps describing the second
method is given in the flow chart of FIG. 8. The two transmitted
signals might be transmitted with a very low repetition rate as
indicated in FIG. 10. The first packet of transmit signals with low
acoustic amplitude are used for calibration. The echoes received
from those signals are averaged to reduce the noise level.
[0190] The number of signals can be chosen such that a high
signal-to-noise ratio is obtained. The second packet of signals
with higher amplitudes are used to induce nonlinear propagation and
harmonic distortion. The echoes received from these signals are
averaged and then the harmonic energy is filtered and then compared
to the calibration echo.
[0191] In order to estimate the volume of urine in the bladder, a
look-up table can be created beforehand. Such a table, saved in the
hard disk of the electronic device, will contain the correspondence
between the harmonic energy and the volume of urine. Such a table
can be extracted from a curve similar to the one given in FIG. 11.
Such a curve can be obtained from a "learning" patient set of
measurements. Look-up tables may eventually be produced for
specific patient groups for age; gender and/or weight as an input
parameter.
[0192] The described second method differs greatly from known other
apparatus:
[0193] 12) The device is composed of a single element defocused
ultrasound transducer with a conical beam shape;
[0194] 13) The single acoustic beam entirely encompasses the
volumetric area of a possibly filled bladder.
[0195] 14) The method is based on measurement of non-linear
properties and attenuation behavior of propagating ultrasound waves
as influenced by a urine filled bladder.
[0196] 15) The method incorporates a technique to eliminate patient
variation due to fat or skin properties.
[0197] 16) The method for automatic volume computation does not
assume any geometrical model for the bladder shape;
[0198] 17) It is valid for any bladder shape since the received
signal "integrates" all volume effects in the ultrasound beam.
[0199] 18) All known other methods use bladder wall echoes as a
basis to calculate volume.
[0200] 19) The device works instantaneously. Other embodiments are
intentionally within the scope of the accompanying claims.
[0201] While preferred and alternate embodiments of the invention
have been illustrated and described, as noted above, many changes
can be made without departing from the spirit and scope of the
invention. Accordingly, the scope of the invention is not limited
by the disclosure of these preferred and alternate embodiments.
Instead, the invention should be determined entirely by reference
to the claims that follow.
* * * * *