Abstract:
Harmonic motion is produced in a subject using vibro-acoustography. An ultrasonic imaging system repetitively interrogates the subject and the Doppler shift in the reflected echo signals is analyzed to measure the phase and amplitude of harmonic motion produced in the subject at different prescribed frequencies. Shear wave propagation through the subject is determined from this information and mechanical properties related to “stiffness” of the subject are determined. A Kalman filter is employed in the phase and amplitude measurement to extract the harmonic motion information from background noise.

Description:
CROSS-REFERENCE TO RELATED APPLICATIONS  
       [0001]     This application claims the benefit of U.S. Provisional patent application Ser. No. 60/508,371 filed on Oct. 3, 2003 and entitled “Motion Detection For Vibroacoustography”. 
     
    
     BACKGROUND OF THE INVENTION  
       [0002]     The field of the invention is coherent imaging using vibratory energy, such as ultrasound and, in particular, vibro-acoustography.  
         [0003]     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).  
         [0004]     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 either their amplitude or time delay is used to modulate the brightness of pixels on a display. With the B-scan method, enough data are acquired from which an image of the refractors can be reconstructed.  
         [0005]     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. A number of such ultrasonic transducer constructions are disclosed in U.S. Pat. Nos. 4,217,684; 4,425,525; 4,441,503; 4,470,305 and 4,569,231.  
         [0006]     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.  
         [0007]     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 signal from each transducer array element.  
         [0008]     This form of ultrasonic imaging is referred to as “phased array sector scanning”, or “PASS”. Such a scan is comprised of a series of measurements in which the steered ultrasonic wave is transmitted, the system switches to receive mode after a short time interval, and the reflected ultrasonic wave is received and stored. Typically, the transmission and reception are steered in the same direction (θ) during each measurement to acquire data from a series of points along a scan line. The receiver is dynamically focused at a succession of ranges (R) along the scan line as the reflected ultrasonic waves are received. The time required to conduct the entire scan is a function of the time required to make each measurement and the number of measurements required to cover the entire region of interest at the desired resolution and signal-to-noise ratio. For example, a total of 128 scan lines may be acquired over a 90° sector, with each scan line being steered in increments of 0.70. A number of such ultrasonic imaging systems are disclosed in U.S. Pat. Nos. 4,155,258; 4,155,260; 4,154,113; 4,155,259; 4,180,790; 4,470,303; 4,662,223; 4,669,314 and 4,809,184.  
         [0009]     Vibro-acoustography is an elasticity modality that vibrates tissue using ultrasound radiation force. The radiation force is generated by focusing two ultrasound beams on the object. These two ultrasound beams have slightly different frequencies and the tissue at the focal point vibrates at the difference frequency. The vibration frequency can be easily changed and the “stiffness” of the tissue at different frequencies can be measured. The tissue is scanned in a raster manner and its acoustic emission is detected by a hydrophone. The acquired emission data may be processed to reconstruct an image, which is related to the stiffness of the tissue. The details of the vibro-acoustography is described in U.S. Pat. No. 5,903,516 entitled “Acoustic force generator for detection, imaging and information transmission using the beat signal of multiple intersecting sonic beams” and U.S. Pat. No. 5,991,239 entitled “Confocal acoustic force generator”.  
         [0010]     A method was recently proposed to solve for the complex stiffness of a homogeneous medium or an artery by measuring shear wave speed dispersion. An oscillatory radiation force is applied to the subject using vibro-acoustography to generate shear waves of various frequencies. The speed of these shear waves is measured from shifts in phase detected over the distance propagated. Measurements of shear wave speed at multiple frequencies are then fit with appropriate theoretical models to solve for the shear elasticity and viscosity of the object as described in Ph.D. Thesis: Direct Methods for Dynamic Elastography Reconstruction: Optimal Inversion of the Interior Helmholtz Problem, Travis Oliphant, Ph.D. May 2001 and Ph.D. Thesis: Shear Property Characterization of Viscoelastic Media Using Vibrations Induced by Ultrasound Radiation, Shigao Chen, Ph.D. June 2002). Although the results are very promising, detection of the shear wave is achieved by an optical method, which limits its medical application because soft tissues are opaque.  
       SUMMARY OF THE INVENTION  
       [0011]     The present invention is a vibro-acoustic system for measuring mechanical properties of opaque subjects such as tissue which includes an acoustic force generator that imparts harmonic motion to the subject at a prescribed frequency, an ultrasonic system for interrogating points in the subject with ultrasonic pulses and receiving echo signals therefrom which indicate the amplitude and phase of the harmonic motion at the points. The echo signals are processed to extract the harmonic motion phase information, and from this a mechanical property of the subject is calculated.  
         [0012]     A more specific aspect of the invention is the method used to estimate the harmonic phase information in the echo signals. The echo signals are quadrature detected and the arctangent of the ratio of the Q and I components acquired at each point are calculated to produce a measured harmonic signal in slow time. The desired harmonic signal is modeled by a differential equation and the phase and amplitude parameters in this model are recursively estimated in a Kalman filter until the mean square error between the model and the measured harmonic signal is minimized.  
         [0013]     The foregoing and other objects 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 herein for interpreting the scope of the invention. 
     
    
     BRIEF DESCRIPTION OF THE DRAWINGS  
       [0014]      FIG. 1  is a block diagram of a vibro-acoustic system which employs the present invention;  
         [0015]      FIG. 2  is a block diagram of an ultrasound imaging system used in the system of  FIG. 1 ;  
         [0016]      FIG. 3  is a block diagram of a receiver which forms part of the ultrasound imaging system of  FIG. 1 ;  
         [0017]      FIG. 4  is a flow chart of the steps performed by a mid-processor which forms part of the receiver of  FIG. 3 . 
     
    
     GENERAL DESCRIPTION OF THE INVENTION  
       [0018]     The present invention is a vibro-acoustic system for measuring the mechanical properties of opaque subjects such as tissue. Where the subject tissue is buried deeply beneath other tissue, optical methods for measuring the very small harmonic motions of the subject tissue cannot be used. This problem is solved by employing an ultrasonic imaging system that interrogates the subject tissue with a pulsed ultrasound beam and examines the resulting echo signals to measure the phase and amplitude of the harmonic motion imparted to the subject tissue. The challenge is to extract this information from the echo signals where the amplitude of the harmonic motion is at the submicron level.  
         [0019]     The displacement of a point in harmonic motion of tissues can be represented in the form of, 
 
 D ( t )= D   0  sin(ω s   t +φ s )  (1) 
 
 The velocity of the motion is:  
               v   ⁡     (   t   )       =         ⅆ     D   ⁡     (   t   )           ⅆ   t       =       v   0     ⁢           ⁢     cos   ⁡     (         ω   s     ⁢   t     +     ϕ   s       )                   (   2   )             
 
 where v 0 =D 0 ω s . 
 
         [0020]     When a pulse echo ultrasound system is focused on the tissue motion, the tissue motion is represented in the echo signals as oscillatory Doppler shifts in the received signals.  
               r   ⁡     (       t   f     ,     t   s       )       =         A   ⁡     (       t   f     ,     t   s       )       ⁢           ⁢     cos   ⁡     (     ∫       (       ω   f     +         2   ⁢     v   ⁡     (   t   )         c     ⁢     ω   f     ⁢           ⁢   cos   ⁢           ⁢   θ       )     ⁢     ⅆ   t         )         ⁢     
     ⁢           =       A   ⁡     (       t   f     ,     t   s       )       ⁢     cos   ⁡     (         ω   f     ⁢     t   f       +     ϕ   f     +     β   ⁢           ⁢     sin   ⁡     (         ω   s     ⁢     t   s       +     ϕ   s       )           )                   (   3   )             
 
 where t f  is fast time representing depth, t s  is slow time representing repetitive pulses, ω f  is transmitting center frequency, ω s  is the tissue vibration frequency, φ s  is the tissue vibration phase, and θ is an angle between the ultrasound beam and direction of tissue motion. The modulation index is:  
             β   =       2   ⁢     v   0     ⁢     ω   f     ⁢           ⁢     cos   ⁡     (   θ   )             ω   s     ⁢   c               (   4   )             
 
 where c is the sound speed in the tissue. 
 
         [0021]     With quadrature demodulation of the received echo signal, we have in-phase and quadrature terms: 
 
 I ( t   f   ,t   s )= A ( t   f   ,t   s )cos(β sin(ω s   t   s +φ s )+φ f +φ 0 )  (5) 
 
 Q ( t   f   ,t   s )=− A ( t   f   ,t   s )sin(β sin(ω s   t   s +φ s )+φ f +φ 0 )  (6) 
 
 where φ 0  is a constant phase added during the quadrature demodulation to maintain I(t f ,t s ) to be either all positive or all negative in slow time at a location. Thus, 
 
 s ( t   f   ,t   s )=−tan −1 ( Q ( t   f   ,t   s )/ I ( t   f   ,t   s ))  (7) 
 
 y ( t   s )= s ( t   f   ,t   s )− {overscore (s)} ( t   f   ,t   s )=β sin(ω s   t   s +φ s )  (8) 
 
 where {overscore (s)}(t f ,t s ) is a mean value of s(t f ,t s ) in slow time. If the sampling frequency in fast time is high, I(t f ,t s ) and Q(t f ,t s ) can be averaged with a limited length in fast time to reduce noise before s(t f ,t s ) is calculated. 
 
         [0022]     A bandpass filter (BPF) centered at the vibration frequency can improve y(t s ) by reducing noise and distortions.  
         [0023]     The amplitude can be directly estimated from y(t s ), 
 
β={square root}{square root}{square root over (2)}σ y   (9) 
 
 where σ y  is a standard deviation of y(t s ). 
 
         [0024]     The phase and amplitude in Equation (8) can be directly obtained by another quadrature demodulation at the vibration frequency in the direction of the slow time, 
 
 I ( t   s )=β cos(φ s )  (10) 
 
 Q ( t   s )=β sin(φ s )  (11) 
 
β( t   s )={square root}{square root over ( I   2 ( t   s )+ Q   2 ( t   s ))}  (12) 
 
φ s ( t   s )=α tan( Q ( t   s )/ I ( t   s )  (13) 
 
         [0025]     The amplitude of the oscillatory Doppler shifts can also be directly measured by applying a turbulence estimation method to the r(t f ,t s ) to estimate the variance of motion velocity.  
         [0026]     In practice, the data will be noisy and have a stochastic nature. Therefore, a Kalman filter process is employed to recursively estimate the phase and amplitude. As described by R. G. Brown and P. Y. C. Hwang in “Introduction To Random Signals And Applied Kalman Filtering”, 3 rd  Edition, John Wily &amp; sons, 1997, a Kalman filter is a numerical method used to track a time-varying signal in the presence of noise. If the signal can be characterized by some number of parameters that vary slowly with time, then Kalman filtering can be used to tell how incoming raw measurements should be processed to best estimate those parameters as a function of time. In this application, a Kalman filter extracts the desired harmonic motion from random and noisy measurement data with known vibration frequency and unknown vibration amplitude and phase. Equation (8) can be represented by a 2 nd  order differential equation,  
                     ⅆ   2     ⁢     y   ⁡     (     t   s     )           ⅆ     t   s   2         +       ω   s   2     ⁢     y   ⁡     (     t   s     )           =   0           (   14   )             
 
 which can be transformed to a 2 nd  order state space form:  
               [             x   k     ⁢           ⁢     (   1   )                   x   k     ⁢           ⁢     (   2   )             ]     =           [         1       0           0       1         ]     ⁡     [             x   k     ⁢           ⁢     (   1   )                   x   k     ⁢           ⁢     (   2   )             ]       ⁢           ⁢   or   ⁢           ⁢     x   k       =     Φ   ⁢           ⁢       x   k     .                 (   15   )             
 
         [0027]     The measurement equation is:  
                 y   ⁡     (     t   s     )       =         β   ⁢           ⁢     sin   ⁡     (       ω   s     ⁢     t   s       )       ⁢   cos   ⁢           ⁢     ϕ   s       +     β   ⁢           ⁢     cos   ⁡     (       ω   s     ⁢     t   s       )       ⁢   sin   ⁢           ⁢     ϕ   s         ⁢     
     ⁢           =         [       sin   ⁡     (       ω   s     ⁢     t   s       )       ,     cos   ⁡     (       ω   s     ⁢     t   s       )         ]     ⁡     [       β   ⁢           ⁢   cos   ⁢           ⁢     ϕ   s       ,     β   ⁢           ⁢   sin   ⁢           ⁢     ϕ   s         ]       T         ⁢     
     ⁢         y   ⁡     (   k   )       =           [         h   k     ⁢           ⁢     (   1   )       ,       h   k     ⁢           ⁢     (   2   )         ]     ⁡     [         x   k     ⁢           ⁢     (   1   )       ,       x   k     ⁡     (   2   )         ]       T     ⁢           ⁢   or       ,     
     ⁢       y   ⁡     (   k   )       =         H   k     ⁢     x   k       +     n   ⁡     (   k   )                     (   16   )             
 
 where t s =k/f PRF  and n(k) is a white noise sequence having a variance R. f PRF  is the pulse repetition frequency. Thus the amplitude and phase parameters of the desired harmonic signal can be found: 
 
β( k )={square root}{square root over ( x   k   2 (1)+ x   k   2 (2))}
 
φ s ( k )=tan −1 ( x   k (2)/ x   k (1))  (17) 
 
         [0028]     The estimation {circumflex over (x)}(k) for x(k) is given by minimizing the mean square error: 
 
 P   k   =E [( x   k   −{circumflex over (x)}   k )( x   k   −{circumflex over (x)}   k ) T ]  (18) 
 
         [0029]     The estimated state variables in the final steps of Kalman filtering are averaged for the amplitude and phase. The filtering steps are listed below:  
         [0030]     1. Initializing: 
 
 P   1   −   =R,{circumflex over (x)}   1   −   =E[x   1 ]  (19) 
 
         [0031]     2. Calculating Kalman gain at k th  step: 
 
 G   k   =P   k   −   H   k   T ( H   k   P   k   −   H   k   T   +R ) −1   (20) 
 
         [0032]     3. Updating the state estimation: 
 
 {circumflex over (x)}   k   ={circumflex over (x)}   k−1   −   +G   k ( y ( k )−H k   {circumflex over (x)}   k−1   − )  (21) 
 
         [0033]     4. Calculating the error covariance matrix: 
 
 P   k =( I−G   k   H   k ) P   k   −   (22) 
 
         [0034]     5. Project ahead: 
 
 {circumflex over (x)}   k+1   − =Φ k   {circumflex over (x)}   k   (23) 
 
 P   k+1   − =Φ k   P   k Φ k   T   +Q   (24) 
 
 where Q is zero in this application and P is an estimation error covariance matrix that describes the estimation accuracy. 
 
         [0035]     The error covariance matrix provide by the Kalman filter is:  
             P   =       E   ⁡     [       (     x   -     x   ^       )     ⁢     (     x   -     x   ^       )       ]       =     (           σ   x1   2           σ   x2x1               σ   x1x2           σ   x2   2           )               (   25   )             
 
 where the diagonal elements represent variance of estimation errors for x, and x 2 . 
 
 x   1 =β cos φ s   (26) 
 
 x   2 =β sin φ s .  (27) 
 
         [0036]     The relations between the state variables and the amplitude β and estimated phase φ s  of the harmonic motion are,  
               β   ^     =         x   1   2     +     x   2   2                 (   28   )                   ϕ   ^     s     =     a   ⁢           ⁢   tan   ⁢       x   2       x   1                 (   29   )             
 
         [0037]     If y=f(x 1 ,x 2 ) and σ x1 , σ x2  are given, then the variance of y is,  
               σ   y     =             (       ∂   f       ∂     x   1         )     2     ⁢     σ   x1   2       +         (       ∂   f       ∂     x   2         )     2     ⁢     σ   x2   2                   (   30   )                 For   ⁢           ⁢       σ   β     :   f       =         x   1   2     +     x   2   2                                     ∂   f       ∂     x   1         =         ∂         x   1   2     +     x   2   2             ∂     x   1         =         1   2     ⁢     (     1         x   1   2     +     x   2   2           )     ⁢   2   ⁢     x   1       =       x   1           x   1   2     +     x   2   2                       (   31   )                   ∂   f       ∂     x   2         =         ∂         x   1   2     +     x   2   2             ∂     x   2         =         1   2     ⁢     (     1         x   1   2     +     x   2   2           )     ⁢   2   ⁢     x   2       =         x   2           x   1   2     +     x   2   2           .                 (   32   )             
 
         [0038]     Therefore, the standard deviation of estimation errors of amplitude estimates is:  
                     σ   β     =             (       ∂   f       ∂     x   1         )     2     ⁢     σ   x1   2       +         (       ∂   f       ∂     x   2         )     2     ⁢     σ   x2   2                       =             (       x   1           x   1   2     +     x   2   2           )     2     ⁢     σ   x1   2       +         (       x   2           x   1   2     +     x   2   2           )     2     ⁢     σ   x2   2                       =             x   1   2     ⁢     σ   x1   2       +       x   2   2     ⁢     σ   x2   2             x   1   2     +     x   2   2                         (   33   )                 For   ⁢           ⁢       σ     ϕ   ⁢           ⁢   s       :   f       =     a   ⁢           ⁢     tan   ⁡     (       x   2       x   1       )                                     ∂   f       ∂     x   1         =         1     1   +       (       x   2       x   1       )     2         ⁢     (     -       x   2       x   1   2         )       =     -       x   2         x   1   2     +     x   2   2                     (   34   )                   ∂   f       ∂     x   2         =         1     1   +       (       x   2       x   1       )     2         ⁢     (     1     x   1       )       =       x   1         x   1   2     +     x   2   2                   (   35   )                       σ   ϕs     =             (       ∂   f       ∂     x   1         )     2     ⁢     σ   x1   2       +         (       ∂   f       ∂     x   2         )     2     ⁢     σ   x2   2                       =             (     -       x   2         x   1   2     +     x   2   2           )     2     ⁢     σ   x1   2       +         (       x   1         x   1   2     +     x   2   2         )     2     ⁢     σ   x2   2                       =             x   2   2     ⁢     σ   x1   2       +       x   1   2     ⁢     σ   x2   2             (       x   1   2     +     x   2   2       )     2                     =             x   2   2     ⁢     σ   x1   2       +       x   1   2     ⁢     σ   x2   2               x   1   2     +     x   2   2                       (   36   )             
 
         [0039]     The above equation is for the standard deviation of estimation errors of phase estimates. Thus, the Kalman filter also provides a measure of estimation quality.  
         [0040]     The above method can be applied to estimate phase φ s  of tissue vibration propagating over a known distance Δr. Then, the shear wave speed can be estimated using the phase change Δφ s  over Δr: 
 
 c   s =ω s   Δr/Δφ   s   (37) 
 
 which can be used to characterize elasticity and viscosity of the artery. 
 
       DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT  
       [0041]     Referring particularly to  FIG. 1 , a vibro-acoustography system which employs the present invention employs an ultrasonic transducer having two elements  10  and  12  which produce two focused beams  14  and  16  that cross each other at their focal points as described in U.S. Pat. No. 5,991,239. The elements  10  and  12  are driven by respective continuous wave synthesizers  18  and  20  at ultrasonic frequencies ω 1  and ω 2  that differ by a desired beat frequency. The two focused beams  14  and  16  are aimed at target tissue  21  which is to be measured, and in response, the target tissue vibrates, or oscillates, at the difference frequency. These elements thus serve as a force generator which oscillates the target tissues  21  at a prescribed beat frequency.  
         [0042]     The vibrations of the target tissue  21  are measured by an ultrasound system  22 . As will be described in more detail below, the ultrasound system  22  drives an ultrasonic transducer  23  to apply a focused ultrasound beam to the target tissue  21  and to receive the echo signal reflected by the target tissue  21 . The phase and amplitude of these echo signals are processed as described below to measure mechanical properties of the target tissue  21 .  
         [0043]     Referring particularly to  FIG. 2 , a transducer array  23  is comprised of a plurality of separately driven elements  11  which each produce a burst of ultrasonic energy when energized by a pulse produced by a transmitter  13 . The ultrasonic energy reflected back to the transducer array  23  from the subject under study is converted to an electrical signal by each transducer element  11  and applied separately to a receiver  9  through a set of switches  15 . The transmitter  13 , receiver  9  and the switches  15  are operated under the control of a digital controller  19  responsive to the commands input by the human operator. A complete scan is performed by acquiring a series of echoes in which the switches  15  are set to their transmit position, the transmitter  13  is gated on momentarily to energize each transducer element  11 , the switches  15  are then set to their receive position, and the subsequent echo signals produced by each transducer element  11  are applied to the receiver  9 . The separate echo signals from each transducer element  11  are combined in the receiver  9  to produce a single echo signal which is employed to produce a line in an image on a display system  17 .  
         [0044]     The transmitter  13  drives the transducer array  23  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  23 . To accomplish this the transmitter  13  imparts a time delay (Ti) to the respective pulses  20  that are applied to successive transducer elements  11 . If the time delay is zero (Ti=0), all the transducer elements  11  are energized simultaneously and the resulting ultrasonic beam is directed along an axis  21  normal to the transducer face and originating from the center of the transducer array  23 . As the time delay (Ti) is increased, the ultrasonic beam is directed downward from the central axis  21  by an angle θ.  
         [0045]     A sector scan is performed by progressively changing the time delays Ti 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  21 , the timing of the pulses  7  is reversed.  
         [0046]     Referring still to  FIG. 2 , 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  11  of the transducer array  23  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  11 , however, these echo signals will not occur simultaneously and their amplitudes will not be equal. The function of the receiver  9  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 range R along the ultrasonic beam oriented at the angle θ.  
         [0047]     To simultaneously sum the electrical signals produced by the echoes from each transducer element  11 , time delays are introduced into each separate transducer element channel of the receiver  9 . In the case of the linear array  23 , the delay introduced in each channel may be divided into two components, one component is referred to as a beam steering time delay, and the other component is referred to as a beam focusing time delay. The beam steering and beam focusing time delays for reception are precisely the same delays (Ti) as the transmission delays described above. However, the focusing time delay component introduced into each receiver channel is continuously changing during reception of the echo to provide dynamic focusing of the received beam at the range R from which the echo signal emanates.  
         [0048]     Under the direction of the digital controller  19 , the receiver  9  provides delays during the scan such that the steering of the receiver  9  tracks with the direction of the beam steered by the transmitter  13  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.  
         [0049]     By selecting proper time delays, echoes from multiple focused locations can be simultaneously received to measure vibration information from several points of the tissue. The limitation of the lateral resolution of the transducer for two closely located points can be improved by assigning different transmitting codes for different locations.  
         [0050]     The display system  17  receives the series of data points produced by the receiver  9  and converts the data to a form producing the desired image. For example, if an A-scan is desired, the magnitude of the series of data points is merely graphed as a function of time. If a B-scan is desired, each data point in the series is used to control the brightness of a pixel in the image, and a scan comprised of a series of measurements at successive steering angles (θ) is performed to provide the data necessary for display of an image.  
         [0051]     Referring particularly to  FIG. 3 , the receiver  9  is comprised of three sections: a time-gain control section  100 , a beam forming section  101 , and a mid processor  102 . The time-gain control section  100  includes an amplifier  105  for each of the N=128 receiver channels and a time-gain control circuit  106 . The input of each amplifier  105  is connected to a respective one of the transducer elements  11  to receive and amplify the echo signal which it receives. The amount of amplification provided by the amplifiers  105  is controlled through a control line  107  that is driven by the time-gain control circuit  106 . 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  108  to values which provide a relatively uniform brightness over the entire range of the sector 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 by the TGC control circuit  106 . The settings of the potentiometers are employed to set the gain of the amplifiers  105  during each of the respective time intervals so that the echo signal is amplified in ever increasing amounts over the acquisition time interval.  
         [0052]     The beam forming section  101  of the receiver  9  includes N=128 separate receiver channels  110 . Each receiver channel  110  receives the analog echo signal from one of the TGC amplifiers  105  at an input  111 , and it produces a stream of digitized output values on an I bus  112  and a Q bus  113 . 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  114  and  115  with the I and Q samples from each of the other receiver channels  110 , they indicate the magnitude and phase of the echo signal reflected from a point P located at range R on the steered beam (θ).  
         [0053]     For a more detailed description of the receiver  9 , reference is made to U.S. Pat. No. 4,983,970 which issued on Jan. 8, 1991 and is entitled “Method And Apparatus for Digital Phase Array Imaging”, and which is incorporated herein by reference.  
         [0054]     Referring still to  FIG. 3 , the mid processor section  102  receives the beam samples from the summing points  114  and  115 . The I and Q values of each beam sample is a 16-bit digital number which represents the in-phase and quadrature components of the magnitude of the reflected sound from a point (R,θ). The mid processor  102  can perform a variety of calculations on these beam samples, where choice is determined by the type of image to be reconstructed.  
         [0055]     For example, a conventional ultrasound image may be produced by a detection processor  120  which calculates the magnitude of the echo signal from its I and Q components: 
 
 M={square root}{square root over (I     2     +Q     2     )}.  
 
 The resulting magnitude values output at  121  to the display system  17  result in an image in which the magnitude of the reflected echo at each image pixel is indicated. 
 
         [0056]     The present invention is implemented by a mechanical property processor  122  which forms part of the mid-processor  102 . As will be explained in detail below, this processor  102  receives the I and Q beam samples acquired during a sequence of measurements of the subject tissue  21  and calculates a mechanical property of the tissue  21 .  
         [0057]     Referring particularly to  FIG. 4 , the mechanical property processor  122  controls the measurements made by the ultrasound system  22 , the force generator elements, and it processes the resulting echo signals I and Q to satisfy equations (5) and (6) and to calculate a mechanical property of the target tissues. Such target tissues may be, for example, an artery and the mechanical property may be stiffness. The first step as indicated by process block  200  is to set the beat frequency of the force generator and excite the target tissues  21  with the force generator. As indicated at process block  202  the ultrasound system  22  is then operated to acquire echo signals from the subject tissues at a series of points. When measuring an artery, for example, 100 echoes sampled at a 40 MHz sample rate are acquired at each point, and 11 points spread evenly along 10 to 20 mm of the length of the artery are measured. Eight echo samples at the peak echo amplitude are used to obtain average I and Q values. As described above, it is necessary that all the I values remain either positive or negative in order to properly detect the harmonic signal. As indicated at process block  203 , the I values are checked and if a zero crossing occurs, all the I and Q values are reprocessed to add a constant phase φ 0  as indicated above in equations (5) and (6). Phase is added until no zero crossings are detected.  
         [0058]     As indicated at process block  204 , the amplitude and phase of the tissue motion at each point is then estimated from the acquired I and Q echo samples. As described above there are a number of different methods for accomplishing this, but in the preferred embodiment the arctangent of the ratio of the Q and I beam samples are calculated and the mean value is removed to obtain the harmonic motion in slow time as indicated above in Equations (7) and (8). The harmonic motion is modeled by a second order differential equation with random amplitude and phase and the known beat frequency. The amplitude and phase is then estimated in a recursive, Kalman filter process that minimizes the mean square error between the model and the measured tissue harmonic motion as indicated above in equations (14)-(18). As indicated by process block  205 , the change in tissue oscillation phase as a function of distance is then calculated for this beat frequency using the calculated phase values at the 11 points along the artery.  
         [0059]     The above process is repeated for each of the prescribed beat frequencies. When used for measuring artery stiffness, vibration frequencies of 100, 200, 300, 400 and 500 Hz are employed, and data acquisition continues until all frequencies have been acquired as determined at decision block  206 .  
         [0060]     As indicated at process block  208 , the next step is to calculate the shear wave speeds in the subject tissue  21  at the different beat frequencies. Linear regression is applied to the 11 phase changed measurements to yield a phase change over 10 mm distance along the artery. From this phase change over distance information, the shear wave speed at each beat frequency is estimated as described by equation (37).  
         [0061]     As indicated at process block  210 , the final step is to calculate a mechanical property of the tissue  21  from the shear wave speed information. In the preferred embodiment the shear elasticity and viscosity of the tissue  21  is estimated from the set of shear wave speeds. These mechanical properties indicate the stiffness of the artery which is a valuable clinical measurement. This calculation is based on shear wave dispersion, and as described by S. Chen et al “Complex Stiffness Quantification Using Ultrasound Stimulated Vibrometry”, 2003 IEEE Ultrasonics Symposium 941-944, the shear wave speeds at multiple frequencies are fit with appropriate theoretical models to solve for the shear elasticity and viscosity.  
         [0062]     While the analysis of the received echo signal is performed in the mid-processor section of an ultrasound receiver in the preferred embodiment described above, it should be apparent that these functions can also be performed in a separate processor or computer workstation.