text: Conventional ultrasound scanners create two-dimensional B-mode images of tissue in which the brightness of a pixel is based on the intensity of the echo return. In color flow imaging, the flow of blood or movement of tissue can be imaged. Measurement of blood flow in the heart and vessels using the Doppler effect is well known. The frequency shift of backscattered ultrasound waves may be used to measure the velocity of the backscatterers from tissue or blood. The change or shift in backscattered frequency increases when blood flows toward the transducer and decreases when blood flows away from the transducer. The Doppler shift may be displayed using different colors to represent speed and direction of flow. The color flow velocity mode displays hundreds of adjacent sample volumes simultaneously, all color-coded to represent each sample volume's velocity. Power Doppler imaging (PDI) is a color flow mode in which the amplitude of the flow signal, rather than the velocity, is displayed. The color flow image may be superimposed on the B-mode image.
  The present invention is incorporated in an ultra-sound imaging system consisting of four main subsystems: a beamformer 2 (see FIG. 1), processors 4 (including a separate processor for each different mode), a scan converter/display controller 6 and a kernel 8. System control is centered in the kernel 8, which accepts operator inputs through an operator interface 10 and in turn controls the various subsystems. The master controller 12 performs system level control functions. It accepts inputs from the operator via the operator interface 10 as well as system status changes (e.g., mode changes) and makes appropriate system changes either directly or via the scan controller. The system control bus 14 provides the interface from the master controller to the subsystems. The scan control sequencer 16 provides real-time (acoustic vector rate) control inputs to the beamformer 2, system timing generator 24, processors 4 and scan converter 6. The scan control sequencer 16 is programmed by the host with the vector sequences and synchronization options for acoustic frame acquisitions. Thus, the scan control sequencer controls the beam distribution and the beam density. The scan converter broadcasts the beam parameters defined by the host to the subsystems via scan control bus 18.
  The main data path begins with the digitized RF inputs to the beamformer from the transducer. Referring to FIG. 2, a conventional ultrasound imaging system includes a transducer array 36 comprised of a plurality of separately driven transducer elements 38, each of which produces a burst of ultrasonic energy when energized by a pulsed waveform produced by a transmitter (not shown). The ultrasonic energy reflected back to transducer array 36 from the object under study is converted to an electrical signal by each receiving transducer element 38 and applied separately to the beamformer 2.
  The echo signals produced by each burst of ultrasonic energy reflect from objects located at successive ranges along the ultrasonic beam. The echo signals are sensed separately by each transducer element 38 and the magnitude of the echo signal at a particular point in time represents the amount of reflection occurring at a specific range. Due to the differences in the propagation paths between an ultrasound-scattering sample volume and each transducer element 38, however, these echo signals will not be detected simultaneously and their amplitudes will not be equal. Beamformer 2 amplifies the separate echo signals, imparts the proper time delay to each, and sums them to provide a single echo signal which accurately indicates the total ultrasonic energy reflected from the sample volume. Each beamformer channel 40 receives the echo signal from a respective transducer element 38.
  To simultaneously sum the electrical signals produced by the echoes impinging on each transducer element 38, time delays are introduced into each separate beamformer channel 40 by a beamformer controller 42. The beam time delays for reception are the same delays as the transmission delays. However, the time delay of each beamformer channel is continuously changing during reception of the echo to provide dynamic focusing of the received beam at the range from which the echo signal emanates. The beamformer channels also have circuitry (not shown) for apodizing and filtering the received pulses.
  The signals entering the summer 44 are delayed so that they are summed with delayed signals from each of the other beamformer channels 40. The summed signals indicate the magnitude and phase of the echo signal reflected from a sample volume located along the steered beam. A signal processor or detector 4 converts the received signal to display data.
  The beamformer outputs two summed digital baseband receive beams. The baseband data is input to B-mode processor 4A and color flow processor 4B, where it is processed according to the acquisition mode and output as processed acoustic vector (beam) data to the scan converter/display processor 6. The scan converter/display processor 6 accepts the processed acoustic data and outputs the video display signals for the image in a raster scan format to a color monitor 22.
  The B-mode processor converts the baseband data from the beamformer into a log-compressed version of the signal envelope. The B function images the time-varying amplitude of the envelope of the signal as a grey scale using an 8-bit output for each pixel. The envelope of a baseband signal is the magnitude of the vector which the baseband data represent.
  The frequency of sound waves reflecting from the inside of blood vessels, heart cavities, etc. is shifted in proportion to the velocity of the blood cells: positively shifted for cells moving towards the transducer and negatively for those moving away. The color flow (CF) processor is used to provide a real-time two-dimensional image of blood velocity in the imaging plane. The blood velocity is calculated by measuring the phase shift from firing to firing at a specific range gate. Instead of measuring the Doppler spectrum at one range gate in the image, mean blood velocity from multiple vector positions and multiple range gates along each vector are calculated, and a two-dimensional image is made from this information. The structure and operation of a color flow processor are disclosed in U.S. Pat. No. 5,524,629, the contents of which are incorporated by reference herein.
  The color flow processor produces velocity (8 bits), variance (turbulence) (4 bits) and power (8 bits) signals. The operator selects whether the velocity and variance or the power are output to the scan converter. The output signal is input to a chrominance control lookup table which resides in the video processor 22. Each address in the lookup table stores 24 bits. For each pixel in the image to be produced, 8 bits control the intensity of red, 8 bits control the intensity of green and 8 bits control the intensity of blue. These bit patterns are preselected such that as the flow velocity changes in direction or magnitude, the color of the pixel at each location is changed. For example, flow toward the transducer is typically indicated as red and flow away from the transducer is typically indicated as blue. The faster the flow, the brighter the color.
  In conventional ultrasound imaging systems, the array of ultrasonic transducers transmit an ultrasound beam and then receive the reflected beam from the object being studied. The array typically has a multiplicity of transducers arranged in a line and driven with separate voltages. By selecting the time delay (or phase) and amplitude of the applied voltages, the individual transducers can be controlled to produce ultrasonic waves which combine to form a net ultrasonic wave that travels along a preferred beam direction and is focused at a selected range along the beam. Multiple firings may be used to acquire data representing the desired anatomical information along a multiplicity of scan lines. The beamforming parameters of each of the firings may be varied to provide a change in the position of focus or otherwise change the spatial position of the received data. By changing the time delay and amplitude of the applied voltages, the beam with its focal point can be moved in a plane to scan the object.
  The same principles apply when the transducer is employed to receive the reflected sound (receiver mode). The voltages produced at the receiving transducers are summed so that the net signal is indicative of the ultrasound reflected from a single focal point in the object. As with the transmission mode, this focused reception of the ultrasonic energy is achieved by imparting separate time delay (and/or phase shifts) and gains to the signal from each receiving transducer.
  Such scanning comprises a series of measurements in which the steered ultrasonic wave is transmitted, the system switches to receive mode after a short time interval, and the reflected ultrasonic wave is received and stored. Typically, transmission and reception are steered in the same direction during each measurement to acquire data from a series of points along a scan line. The receiver is dynamically focused at a succession of ranges or depths along the scan line as the reflected ultrasonic waves are received.
  In an ultrasound imaging system, the beam spacing for an optimum image is determined by the beam width or lateral point spread function. The lateral point spread function is determined by the product of the wavelength and the f-number. The wavelength is in turn a function of the transmit waveform center frequency and the receiver demodulation frequency. The f-number equals the focal depth divided by the aperture.
  The number of beams fired is determined by the spatial sampling requirements and the desired frame rate. Frame rate is inversely proportional to the time taken to transmit and receive all the beams required to form a complete frame of data. High frame rates are required to minimize the possible motion-induced errors in the image. In order to maintain a high frame rate the number of beams is kept to the minimum which would satisfy the Nyquist spatial sampling requirement. When fewer beams are fired than minimum spatial sampling requirements, spatial aliasing occurs. At the optimum spatial sampling the highest resolution is obtained together with the highest frame rate.