US 20080139934 A1 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.
Claims(25) 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×D×K.2. The apparatus of 3. The apparatus of 4. The apparatus of 5. The apparatus of 6. The apparatus of 7. The apparatus of 8. The apparatus of 9. The apparatus of _{A}, φ_{B}, φ_{C}, φ_{D}, and φ_{E}, to an axis orthogonal to the plane of the transducer array, the angles being approximately φ_{A}=−25°, φ_{B}=0°, φ_{C}+25°, φ_{D}+25°, φ_{E}+40°.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×D×K.11. The method of 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 14. The method of 15. The method of 16. The method of 17. The method of 18. The method of 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 21. The method of 22. The method of 23. The method of 24. The method of 25. The method of Description This invention relates to ultrasound imaging of bodily tissues, bodily fluids, and fluid-filled cavities. 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. 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, “ 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. 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.** 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?” 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. 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. 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×H×K. Where K is an empirically measured, filling configuration dependant, correction factor. Beam directions and examples for D and H are illustrated in 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. 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). 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. 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. 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×Depth×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. 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. 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. 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. 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. 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. 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. 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 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. 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. 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. 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. 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. The embodiments of the present invention are described in detail below with reference to the following drawings. 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 The transceiver The operation of the transceiver Still referring to The plurality of three-dimensional distributed scan lines The internal scan lines are represented by scan lines The communications link In another particular embodiment, the imaging programs and the instructions reside exclusively on the server 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. With reference now also to Referring now to When both organs are relatively filled with a fluid (as shown in The first cavity In the WRI region, window function processing determines the raw data comprising the fundamental frequency f With reference still to For a selected window, the total attenuation of the fundamental frequency component may be expressed as follows in equation E1: -
- where: σ
_{ij}=attenuation coefficient of a tissue i at a frequency j and distances d_{1 }and d_{2 }are as shown inFIG. 13 .
- where: σ
While the total attenuation of the higher order harmonic frequency component may be expressed as follows in equation E2: A difference in the attenuation of higher order harmonic component to the fundamental component is therefore defined by equation E3: For example, in soft tissue having an attenuation factor of about 1.1 dB/cm, the attenuation coefficient σ 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. 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. 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. 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. 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. 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. 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 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 Still referring to 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. Referring again to 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. The surface area measurement of fifth block 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 By way of example, four surface patch functions are highlighted in The surface patches are defined as functions of the patch coordinates, s
With the definitions of surface patch functions complete, attention can turn to the surface area calculation represented in the fifth block
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.
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.
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.
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.
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
where t denotes matrix and vector transpose,
Similar calculations are performed for the y 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.
Recalling the fact that s
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 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. 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.
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. The fractal dimension calculation corresponds to the fourth block 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 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 ε. The subscript i indicates a specific value of r,fd, and ε. 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.
Each row of the fd, and ε, and the X matrix correspond to one of the n observations. The parabola parameters of equation E16 are collected in the vector P. 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. The estimates of the parabola parameters ({circumflex over (β)}=└â {circumflex over (b)} ĉ┘′) 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.
To determine the maximal fractal dimension as defined by the parabolic model, simply substitute equation 20 into equation 18 and solve for fd
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
Two values for r These measurements could be made at any surface of the bladder muscle wall. In Once the bladder wall thickness and the inner and outer surface area have been measured, the volume of the internal bladder region The methods to obtain the wall-thickness data, the mass data, and the volume of internal region 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.** The first method describes a simple device that allows the assessment of bladder In operation of the apparatus of the present invention, the transducer assembly Computation of the Depth D and Height The Height H (
Volume computation: The volume of urine is correlated to the bladder diameter (Height 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. The process executed by the hardware is illustrated in the flow chart of 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. 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. For determining body cavity and height, the apparatus may select only beams corresponding to those that have intercepted the fluid filled body cavity. The arrangements described in connection with 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. 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. 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. 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. 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. 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. In summary, the described first method differs greatly from known other apparatus: 1) The device is composed of a limited number of static single element transducers; 2) The arrangement of the transducer is not similar to the arrangement of a linear array; 3) The transducers are oriented towards the bladder with specific angles allowing the estimation of the urine volume over a wide range of volumes; 4) The method for automatic volume computation does not assume any geometrical model for the bladder shape; 5) It is valid for any bladder shape since the volume is computed with an empirical formula for various filling ranges; 6) It is not based only on the measurement of distances between the front and back wall or area in different planes; 7) It uses an automatic detection of the bladder height and depth depending on the number of beams that intercept the bladder; 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); 9) The device includes a closed loop to easily find the optimal position; 10) The optimal position corresponds to the largest volume computed; 11) The device works instantaneously. 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 The schematic principle of transducer positioning is illustrated in 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. 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. 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. 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. 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. The balance between the nonlinear process and the attenuation process is given by the Goldberg number Γ (Szabo et al.), which represents a measure of which process dominates. When Γ=1, nonlinear effects are comparable to attenuation effects. If Γ 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. 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. 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. 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. 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. A block diagram of a possible steps describing the second method is given in the flow chart of 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. 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 The described second method differs greatly from known other apparatus: 12) The device is composed of a single element defocused ultrasound transducer with a conical beam shape; 13) The single acoustic beam entirely encompasses the volumetric area of a possibly filled bladder. 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. 15) The method incorporates a technique to eliminate patient variation due to fat or skin properties. 16) The method for automatic volume computation does not assume any geometrical model for the bladder shape; 17) It is valid for any bladder shape since the received signal “integrates” all volume effects in the ultrasound beam. 18) All known other methods use bladder wall echoes as a basis to calculate volume. 19) The device works instantaneously. Other embodiments are intentionally within the scope of the accompanying claims. 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. Referenced by
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