Abstract:
A method and system for quantifying the extent of vaso vasorum with contrast enhanced ultrasound and correlating that quantitative value to the risk for vascular disease is provided. An ultrasound contrast agent is administered to a subject and images are obtained using an imaging method that identifies the uptake of the contrast agent by tissues. The amount of uptake is measured and the corresponding signal intensity change correlated with the amount of vaso vasorum present. Additionally, deformations of the vasculature are measured from the series of ultrasound images and this information is coupled with the quantification of the vaso vasorum to determine a risk index indicative of a subject&#39;s risk to vascular disease.

Description:
CROSS-REFERENCE TO RELATED APPLICATIONS 
     This application claims the benefit of U.S. Provisional Patent Application Ser. No. 61/185,802, filed on Jun. 10, 2009, and entitled “Method for Assessing Vascular Disease by Quantitatively Measuring Vaso Vasorum.” 
    
    
     BACKGROUND OF THE INVENTION 
     The field of the invention is ultrasound imaging methods and systems. More particularly, the invention relates to employing ultrasound to assess cardiac disease by quantitatively measuring vaso vasorum. 
     There are a number of modes in which ultrasound can be used to produce images of objects. The ultrasound transmitter may be placed on one side of the object and the sound transmitted through the object to the ultrasound receiver placed on the other side (“transmission” mode). With transmission mode methods, an image may be produced in which the brightness of each pixel is a function of the amplitude of the ultrasound that reaches the receiver (“attenuation” mode), or the brightness of each pixel is a function of the time required for the sound to reach the receiver (“time-of-flight”, or “speed of sound” mode). In the alternative, the receiver may be positioned on the same side of the object as the transmitter and an image may be produced in which the brightness of each pixel is a function of the amplitude or time-of-flight of the ultrasound reflected from the object back to the receiver (“refraction”, “backscatter”, or “echo” mode). 
     There are a number of well known backscatter methods for acquiring ultrasound data. In the so-called “A-scan” method, an ultrasound pulse is directed into the object by the transducer and the amplitude of the reflected sound is recorded over a period of time. The amplitude of the echo signal is proportional to the scattering strength of the refractors in the object and the time delay is proportional to the range of the refractors from the transducer. In the so-called “B-scan” method, the transducer transmits a series of ultrasonic pulses as it is scanned across the object along a single axis of motion. The resulting echo signals are recorded as with the A-scan method and their amplitude is used to modulate the brightness of pixels on a display. The location of the transducer and the time delay of the received echo signals locates the pixels to be illuminated. With the B-scan method, enough data are acquired from which a two-dimensional image of the refractors can be reconstructed. Rather than physically moving the transducer over the subject to perform a scan it is more common to employ an array of transducer elements and electronically move an ultrasonic beam over a region in the subject. 
     Ultrasonic transducers for medical applications are constructed from one or more piezoelectric elements sandwiched between a pair of electrodes. Such piezoelectric elements are typically constructed of lead zirconate titanate (“PZT”), polyvinylidene diflouride (“PVDF”), or PZT ceramic/polymer composite. The electrodes are connected to a voltage source, and when a voltage is applied, the piezoelectric elements change in size at a frequency corresponding to that of the applied voltage. When a voltage pulse is applied, the piezoelectric element emits an ultrasonic wave into the media to which it is coupled at the frequencies contained in the excitation pulse. Conversely, when an ultrasonic wave strikes the piezoelectric element, the element produces a corresponding voltage across its electrodes. Typically, the front of the element is covered with an acoustic matching layer that improves the coupling with the media in which the ultrasonic waves propagate. In addition, a backing material is disposed to the rear of the piezoelectric element to absorb ultrasonic waves that emerge from the back side of the element so that they do not interfere. 
     When used for ultrasound imaging, the transducer typically has a number of piezoelectric elements arranged in an array and driven with separate voltages (“apodizing”). By controlling the time delay (or phase) and amplitude of the applied voltages, the ultrasonic waves produced by the piezoelectric elements (“transmission mode”) combine to produce a net ultrasonic wave focused at a selected point. By controlling the time delay and amplitude of the applied voltages, this focal point can be moved in a plane to scan the subject. 
     The same principles apply when the transducer is employed to receive the reflected sound (“receiver mode”). That is, the voltages produced at the transducer elements in the array are summed together such that the net signal is indicative of the sound reflected from a single focal point in the subject. As with the transmission mode, this focused reception of the ultrasonic energy is achieved by imparting separate time delay (and/or phase shifts) and gains to the echo signal received by each transducer array element. 
     Doppler systems employ an ultrasonic beam to measure the velocity of moving reflectors, such as flowing blood cells. Blood velocity is detected by measuring the Doppler shifts in frequency imparted to ultrasound by reflection from moving red blood cells. Accuracy in detecting the Doppler shift at a particular point in the bloodstream depends on defining a small sample volume at the required location and then processing the echoes to extract the Doppler shifted frequencies. 
     A Doppler system is incorporated in a real time scanning imaging system. The system provides electronic steering and focusing of a single acoustic beam and enables small volumes to be illuminated anywhere in the field of view of the instrument, whose locations can be visually identified on a two-dimensional B-scan image. A Fourier transform processor faithfully computes the Doppler spectrum backscattered from the sampled volumes, and by averaging the spectral components the mean frequency shift can be obtained. Typically the calculated blood velocity is used to color code pixels in the B-scan image. 
     In areas of injured endothelial lining, tiny blood vessels referred to as vaso vasorum are formed to supply these areas. These inflamed areas are vulnerable to form plaque. It would therefore be desirable to have a method for not only visualizing the presence of vaso vasorum, but to quantify their presence and effect. 
     SUMMARY OF THE INVENTION 
     The present invention is directed to a method for measuring the risk a tissue of interest has for developing vascular disease. More particularly, the present invention is a method for quantifying the extent of vaso vasorum with contrast enhanced ultrasound and correlating that quantitative value to the risk for vascular disease. An ultrasound contrast agent is administered to a subject and images are obtained using an imaging method that identifies the uptake of the contrast agent by tissues. The amount of uptake is measured and the corresponding signal intensity change correlated with the amount of vaso vasorum present. Additionally, deformations of the vasculature are measured from the series of ultrasound images and this information is coupled with the quantification of the vaso vasorum to determine a risk index indicative of a subject&#39;s risk to vascular disease. 
     The foregoing and other aspects and advantages of the invention will appear from the following description. In the description, reference is made to the accompanying drawings which form a part hereof, and in which there is shown by way of illustration a preferred embodiment of the invention. Such embodiment does not necessarily represent the full scope of the invention, however, and reference is made therefore to the claims and herein for interpreting the scope of the invention. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
         FIG. 1  is a block diagram of an ultrasonic imaging system that employs the present invention; 
         FIG. 2  is a block diagram of a transmitter which forms part of the ultrasonic imaging system of  FIG. 1 ; 
         FIG. 3  is a block diagram of a receiver which forms part of the ultrasonic imaging system of  FIG. 1 ; and 
         FIG. 4  is a flowchart setting forth the steps of an embodiment of the present invention. 
     
    
    
     DETAILED DESCRIPTION OF THE INVENTION 
     Referring particularly to  FIG. 1 , an ultrasonic imaging system includes a transducer array  100  comprised of a plurality of separately driven elements  102  which each produce a burst of ultrasonic energy when energized by a pulse produced by a transmitter  104 . The ultrasonic energy reflected back to the transducer array  100  from the subject under study is converted to an electrical signal by each transducer element  102  and applied separately to a receiver  106  through a set of switches  108 . The transmitter  104 , receiver  106 , and the switches  108  are operated under the control of a digital controller  110  responsive to the commands input by the human operator. A complete scan is performed by acquiring a series of echoes in which the switches  108  are set to their transmit position, the transmitter  104  is gated on momentarily to energize each transducer element  102 , the switches  108  are then set to their receive position, and the subsequent echo signals produced by each transducer element  102  are applied to the receiver  106 . The separate echo signals from each transducer element  102  are combined in the receiver  106  to produce a single echo signal which is employed to produce a line in an image on a display system  112 . 
     The transmitter  104  drives the transducer array  100  such that the ultrasonic energy produced is directed, or steered, in a beam. A B-scan can therefore be performed by moving this beam through a set of angles from point-to-point rather than physically moving the transducer array  100 . To accomplish this the transmitter  104  imparts a time delay, T, to the respective pulses  116  that are applied to successive transducer elements  102 . If the time delay is zero T i =0, all the transducer elements  102  are energized simultaneously and the resulting ultrasonic beam is directed along an axis  118  normal to the transducer face and originating from the center of the transducer array  100 . As the time delay, T i , is increased, the ultrasonic beam is directed downward from the central axis  118  by an angle, θ. The relationship between the time delay increment, T i , added successively to each i th  signal from one end of the transducer array (i=1) to the other end (i=n) is given by the following relationship: 
     
       
         
           
             
               
                 
                   
                     
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     where S is an equal spacing between centers of adjacent transducer elements  102 , c is the velocity of sound in the object under study, R is a range at which the transmit beam is to be focused, and T 0  is a delay offset that insures that all calculated values, T i , are positive values. 
     The first term in this expression steers the beam in the desired angle, θ, and the second is employed when the transmitted beam is to be focused at a fixed range. A sector scan is performed by progressively changing the time delays, T i , in successive excitations. The angle, θ, is thus changed in increments to steer the transmitted beam in a succession of directions. When the direction of the beam is above the central axis  118 , the timing of the pulses  116  is reversed, but the above formula still applies. 
     Referring still to  FIG. 1 , the echo signals produced by each burst of ultrasonic energy emanate from reflecting objects located at successive positions, R, along the ultrasonic beam. These are sensed separately by each segment  102  of the transducer array  100  and a sample of the magnitude of the echo signal at a particular point in time represents the amount of reflection occurring at a specific range, R. Due to the differences in the propagation paths between a focal point, P, and each transducer element  102 , however, these echo signals will not occur simultaneously and their amplitudes will not be equal. The function of the receiver  106  is to amplify and demodulate these separate echo signals, impart the proper time delay to each and sum them together to provide a single echo signal which accurately indicates the total ultrasonic energy reflected from each focal point, P, located at successive ranges, R, along the ultrasonic beam oriented at the angle, θ. 
     Under the direction of the digital controller  110 , the receiver  106  provides delays during the scan such that the steering of the receiver  106  tracks with the direction of the beam steered by the transmitter  104  and it samples the echo signals at a succession of ranges and provides the proper delays to dynamically focus at points, P, along the beam. Thus, each emission of an ultrasonic pulse results in the acquisition of a series of data points which represent the amount of reflected sound from a corresponding series of points, P, located along the ultrasonic beam. 
     Referring particularly to  FIG. 2 , the transmitter  104  includes a set of channel pulse code memories which are indicated collectively at  200 . Each pulse code memory  200  stores a bit pattern  202  that determines the frequency of the ultrasonic pulse  204  that is to be produced. This bit pattern is read out of each pulse code memory  200  by a master clock and applied to a driver  206  which amplifies the signal to a power level suitable for driving the transducer  100 . In the example shown in  FIG. 2 , the bit pattern is a sequence of four “1” bits alternated with four “0” bits to produce a 5 megahertz (“MHz”) ultrasonic pulse  204 . The transducer elements  102  to which these ultrasonic pulses  204  are applied respond by producing ultrasonic energy. 
     As indicated above, to steer the transmitted beam of the ultrasonic energy in the desired manner, the pulses  204  for each of the N channels must be produced and delayed by the proper amount. These delays are provided by a transmit control  208  which receives control signals from the digital controller  110 . When the control signal is received, the transmit control  208  gates a clock signal through to the first transmit channel  200 . At each successive delay time interval thereafter, the clock signal is gated through to the next channel pulse code memory  200  until all the channels to be energized are producing their ultrasonic pulses  204 . Each transmit channel  200  is reset after its entire bit pattern  202  has been transmitted and the transmitter  104  then waits for the next control signal from the digital controller  110 . By operating the transmitter  104  in this manner, ultrasonic energy can be focused on a focal point, P, when practicing the herein described method. This focal point can be steered electronically with the appropriate changes to the timing delays provided by the transmit control  208 . The term “focal point,” as referred to herein, includes not only a single point object in the usual sense, but also a general region-of-interest to which ultrasound energy is delivered in a substantially focused manner. 
     Referring particularly to  FIG. 3 , the receiver  106  is comprised of three sections: a time-gain control (“TGC”) section  300 , a beam forming section  302 , and a mid processor  304 . The time-gain control section  300  includes an amplifier  306  for each of the N receiver channels and a time-gain control circuit  308 . The input of each amplifier  306  is connected to a respective one of the transducer elements  102  to receive and amplify the echo signal which it receives. The amount of amplification provided by the amplifiers  306  is controlled through a control line  310  that is driven by the time-gain control circuit  308 . As is well known in the art, as the range of the echo signal increases, its amplitude is diminished. As a result, unless the echo signal emanating from more distant reflectors is amplified more than the echo signal from nearby reflectors, the brightness of the image diminishes rapidly as a function of range, R. This amplification is controlled by the operator who manually sets TGC linear potentiometers  312  to values which provide a relatively uniform brightness over the entire range of the scan. The time interval over which the echo signal is acquired determines the range from which it emanates, and this time interval is divided into segments by the TGC control circuit  308 . The settings of the potentiometers are employed to set the gain of the amplifiers  306  during each of the respective time intervals so that the echo signal is amplified in ever increasing amounts over the acquisition time interval. 
     The beam forming section  302  of the receiver  106  includes N separate receiver channels  314 . Each receiver channel  314  receives the analog echo signal from one of the TGC amplifiers  306  at an input  316 , and it produces a stream of digitized output values on an I bus  318  and a Q bus  320 . Each of these I and Q values represents a sample of the echo signal envelope at a specific range, R. These samples have been delayed in the manner described above such that when they are summed at summing points  322  and  324  with the I and Q samples from each of the other receiver channels  314 , they indicate the magnitude and phase of the echo signal reflected from a point, P, located at range, R, on the ultrasonic beam. 
     Referring still to  FIG. 3 , the mid processor section  304  receives the beam samples from the summing points  322  and  324 . The I and Q values of each beam sample is a digital number which represents the in-phase and quadrature components of the magnitude of the reflected sound from a point, P. The mid processor  304  can perform a variety of calculations on these beam samples, where choice is determined by the type of image to be reconstructed. For example, if a conventional magnitude image is to be produced, a detection processor indicated at  326  is implemented in which a digital magnitude, M, is calculated from each beam sample according to:
 
 M =√{square root over (I 2   +Q   2 )}  Eqn. (2);
 
     and output at  120  ( FIGS. 1 and 3 ). 
     The detection processor  326  may also implement correction methods that, for example, examine the received beam samples and calculate corrective values that can be used in subsequent measurements by the transmitter  104  and receiver  106  to improve beam focusing and steering. Such corrections are necessary, for example, to account for the non-homogeneity of the media through which the sound from each transducer element travels during a scan. 
     The mid processor may also include a Doppler processor  328 . Such Doppler processors  328  often employ the phase information, φ, contained in each beam sample to determine the velocity of reflecting objects along the direction of the beam (i.e., direction from the transducer  100 ), where: 
     
       
         
           
             
               
                 
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     The mid processor  304  may also include a correlation flow processor  330  that, for example, measures the motion of reflectors by following the shift in their position between successive ultrasonic pulse measurements. 
     Referring particularly now to  FIG. 4 , a method for quantitatively measuring vaso vasorum, and thereby assessing vascular disease, in accordance with the present invention begins with the administration of an ultrasound contrast agent to a subject, as indicated at step  400 . Exemplary ultrasound contrast agents include those with the trade names SonoVue® (Bracco Diagnostics, Princeton, N.J.), Definity® (Lantheus Medical Imaging, North Billerica, Mass.), Optison (GE Healthcare, Waukesha, Wis.), and Imagent® (IMCOR Pharmaceutical Co., San Diego, Calif.). After the contrast agent has been administered to the subject, a series of image frames are acquired, as indicated at step  402 . The images acquired discriminate between the contrast agent and the background tissues. For example, a contrast pulse sequencing method is employed in which background tissue is separable from the contrast agent by way of simultaneously processing received signals from a plurality of transmitted pulses. The phase and amplitude modulation of each pulse is varied so that the interaction of the pulses with the contrast agent results in a response that is separable from background tissues. An exemplary imaging method of this kind is available under the trade name Cadence™ contrast pulse sequencing (Siemens Medical Solutions USA, Inc., Mountain View, Calif.). 
     From the acquired series of image frames, a perfusion rate of the contrast agent into the surrounding vasculature is determined at step  404 . The rate of perfusion of the contrast agent into the surrounding tissues provides a quantitative measure of the presence of vaso vasorum in the vessel. Where an increase in the perfusion of the contrast agent into the vascular wall occurs, an increase in signal intensity is present in the resulting images. The degree of perfusion of the contrast agent into the vascular wall is representative of the presence of vaso vasorum. To calculate the perfusion rate, the change in image intensity over the series of acquired images is analyzed. The signal intensity change in a selected region of interest is fit on a voxel-by-voxel basis to the following signal model:
 
A+B(1−e −kt )  Eqn. (4);
 
     where A is constant indicative of the peak image intensity of contrast agent uptake, B is a constant indicative of the perfusion rate, k is a constant, and t is the time at which a given image frame was acquired. The constant B is calculated from the logarithm of the measured signal intensity change. It is contemplated that values of the constant, B, greater than 0.50 indicate the presence of vaso vasorum in the blood vessel of interest. It is also contemplated that the peak image intensity value, A, can be utilized to determine the presence of vaso vasorum, in as much as larger peak values are likely representative of the presence of more vaso vasorum in the vessel wall, which in turn provide a larger uptake of the contrast agent. 
     The acquired series of image frames are then also analyzed using a tracking technique that measures deformations in the vessel wall, as indicated at step  406 . An exemplary method of this kind is available under the trade name Velocity Vector Imaging™ (Siemens Medical Solutions USA, Inc., Mountain View, Calif.). Using a motion tracking method, such as the one provided by Velocity Vector Imaging™, radial deformations and rotations in a vessel wall are determined. Additionally, longitudinal and cross-sectional blood flow velocities through the blood vessel of interest can be calculated and utilized to assess the risk for vascular disease. This information, along with the perfusion rate calculated previously, is utilized to produce an index value, as indicated at step  408 . The index value indicates those tissues of interest that are at risk for a particular vascular disease. 
     After the index value has been produced, it is reported to the system operator, ultrasound technologist, clinician, or other healthcare professional, as indicated at step  410 . For example, an index map is produced, in which voxel values in the index map correspond to the index value calculated for the corresponding voxel location in the acquired series of image frames. An exemplary index map includes a discontinuous color coding scheme that indicates those regions where vaso vasorum are present and the degree of vulnerability for those regions to develop vascular disease. For example, an index value in the 75-100 percentile range is coded as red, 50-75 percentile range is coded as orange, 25-50 percentile range is coded as yellow, and 0-25 percentile range is coded as blue. By way of this example, those regions coded as red indicate areas at very high risk for vascular disease, while those coded as orange are at high risk, those coded as yellow are areas at moderate risk, and those coded as blue are at low risk. Alternatively, when the index values include values in the range identified as “risk”, a report can be produced indicating that the subject is at risk for particular a vascular disease. 
     Additionally, the quantified presence of vaso vasorum provided by the calculated perfusion rate in the lumen of the blood vessel can be utilized alone to assess the risk of the patient to developing vascular disease. For example, different threshold values of perfusion rate can be used to identify different risk groups. By way of example, the following ranges of values for the constant, B, can be used: 0-0.50, low risk; 0.50-5.0, higher risk; and greater than 5.0, even higher risk. Furthermore, this risk assessment can be supplemented with information regarding the deformation of the blood vessel wall. For instance, it is contemplated that the more cross-sectional rotational or radial deformation present in the vessel wall, the more likely the patient is at risk for developing vascular disease. The peak uptake of the contrast agent into the lumen of the blood vessel can also be utilized to assess the risk of the patient. For example, it is contemplated that a patient having a large uptake in contrast agent is more likely to have vaso vasorum present than a patient with less uptake. 
     The present invention has been described in terms of one or more preferred embodiments, and it should be appreciated that many equivalents, alternatives, variations, and modifications, aside from those expressly stated, are possible and within the scope of the invention.