Method and apparatus for B-mode image banding suppression

Certain embodiments include a system and method for banding suppression in a B-mode ultrasound image. Banding occurs in a diagnostic image when two or more focal regions having different waveforms and/or frequencies, for example, are combined. The resulting difference in intensity between the focal regions produces banding in the final image. For purposes of illustration only, the method below will be described with two focal zones. Certain embodiments of the present invention reduce banding in a diagnostic image including a plurality of focal zones by processing the first focal zone, determining intensity around the border or “stitch line” between the two focal regions, and adjusting the second focal region based on the difference in intensity. Preferably, processing is done in real time as image frames are updated.

FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT

BACKGROUND OF THE INVENTION

The present invention generally relates to ultrasound imaging. In particular, the present invention relates to banding suppression in ultrasound imaging.

Ultrasound is sound having a frequency that is higher than a normal person may hear. Ultrasound imaging utilizes ultrasound waves or vibrations in the frequency spectrum above normal human hearing, such as the 2.5-MHz range. Ultrasound imaging systems transmit ultrasound into a subject, such as a patient, in short bursts. Echoes are reflected back to the system from the subject. Diagnostic images may be produced from the echoes. Ultrasound imaging techniques are similar to those used in SONAR and RADAR.

B-mode (brightness mode) imaging is a grayscale ultrasound imaging technique that constructs images based on echoes received from pulses transmitted through a cross-section of the subject scanned. In B-mode imaging, the brightness of a spot or pixel representing an echo in a grayscale image corresponds to the strength of the received echo. The voltage of an echo received at an ultrasound receiver is an indication of brightness. B-mode imaging may be used on its own or combined with Doppler imaging or another imaging technique.

Forming the best possible image at all times for different anatomies and patient types is important to diagnostic imaging systems. Poor image quality may prevent reliable analysis of the image. For example, a decrease in image contrast quality may yield an unreliable image that is not usable clinically. Additionally, the advent of real-time imaging systems has increased the importance of generating clear, high quality images. Differences between different body types may result in blurring, streaking, or introduction of ghost images or artifacts in a resulting image. Automatic optimization of diagnostic images helps to ensure consistent image quality over a wide range of patients.

Multiple focal zones are often used to improve resolution and/or penetration of an ultrasound image. A focal zone is a location within the body at which the transmitted ultrasound pulse is focused. Each focal zone has a corresponding focal region over which energy transmitted to that focal zone produces the best image. Typically, different waveforms and/or f-numbers (a ratio of lens focal length to lens aperture diameter) are used for different focal zones, and the focal region includes the focal zone. When multiple focal zones are used, an ultrasound image is formed by adjoining each focal region that corresponds to the focal zones. When two or more regions are joined together, the borders of the regions may be distinct and visible in the image. The artifactual edges are known as banding artifacts. Banding artifacts are caused by differences in speckle brightness or texture between two focal regions. Currently, transmit waveforms and depth-dependent gain curves are carefully selected to eliminate banding artifacts on the average patient. However, patient body types are diverse (particularly with pathology), and banding artifacts may occur despite the most carefully selected gain curves. Therefore, a real-time, adaptive band-suppression method is needed to reduce banding artifacts over a wide range of patient body types.

Thus, an ultrasound imaging system that automatically adjusts system parameters in real time to reduce banding on a wide range of patient body types would be highly desirable. Furthermore, a system that reduces banding in real time with faster processing than current systems would be highly desirable. A more accurate and efficient method for reducing banding would also be highly desirable. Furthermore, a method for reducing banding artifacts that is applicable to all B-mode applications would be highly desirable.

BRIEF SUMMARY OF THE INVENTION

Certain embodiments include a system and method for banding suppression in a B-mode ultrasound image. Banding occurs in a diagnostic image when two or more focal regions obtained using different waveforms and/or f-numbers, for example, are combined. The difference in image intensity between the focal regions produces banding in a composite image. For purposes of illustration only, the method below will be described with two focal zones. Certain embodiments of the present invention reduce banding in a diagnostic image including a plurality of focal zones by processing a first focal region corresponding to a first focal zone, determining intensity around a border or “stitch line” between the two focal regions, and adjusting a second focal region based on the difference in intensity. Processing is done in real time as image frames are updated.

Certain embodiments of the system include a transducer for transmitting ultrasound energy to a subject and receiving echoes from the subject, a front-end for producing data from the received echoes, and a processor for processing the data to produce an image, the processor combining data from at least a first focal zone and a second focal zone, wherein the processor determines a difference in intensity between data in first and second focal regions, the processor applying the difference to at least one of the first focal region and the second focal region.

The processor may include a B-mode processor, a control processor, and/or a display processor. The processor may accommodate a plurality of imaging and diagnostic modes. The front-end may include a receiver, a transmitter, and a beamformer. The system may also include a display for formatting and displaying the image data.

Certain embodiments of the method include examining a border between a first focal zone and a second focal zone in a combined ultrasound image, selecting a first subset of data adjacent to the border in the first focal zone and a second subset of data adjacent to the border in the second focal zone, determining a first intensity level for the first subset of data and a second intensity level for the second subset of data, determining a difference between the first intensity level and the second intensity level, and applying the difference to at least one of the first focal zone and the second focal zone.

The first and second intensity levels may be means or median intensity values for the first and second focal zones. The difference may be split equally and applied to the first and second focal zones. In an embodiment, the difference may not be applied if a certain threshold, such as a dark pixel threshold, a minimum intensity difference threshold, or a maximum intensity difference threshold, is exceeded. The method may also include applying a filter to the combined ultrasound image. The method may further include displaying the combing ultrasound image.

Certain embodiments of the present invention include a method for banding suppression in a B-mode ultrasound system. The method includes processing a first focal region to determine intensity of first focal region image data, determining a difference in intensity between first focal region image data and second focal region image data surrounding a border between the first focal region and a second focal region, and adjusting intensity of image data in the second focal region based on the difference in intensity. The method may include determining a difference in median or mean intensity between first focal region image data and second focal region image data surrounding a border between the first focal region and a second focal region. The method may also include splitting the difference equally and applying the split different to the first and second focal regions. The image data intensity may not be adjusted if at least one of a dark pixel threshold is exceeded, the difference is less than a minimum threshold, and the difference is greater than a maximum threshold. The method may further include applying a filter to the combined ultrasound image.

DETAILED DESCRIPTION OF THE INVENTION

FIG. 1illustrates a block diagram of an ultrasound imaging system5used in accordance with an embodiment of the present invention. A transducer10is used to transmit ultrasound waves into a subject by converting electrical analog signals to ultrasonic energy. The transducer10also is used to receive ultrasound waves that are backscattered from the subject by converting ultrasonic energy to analog electrical signals. A front-end20including a receiver, a transmitter, and a beamformer, is used to create transmitted waveforms, beam patterns, receiver filtering techniques, and demodulation schemes that are used for various imaging modes. The front-end20converts digital data to analog data and vice versa. The front-end20interfaces with the transducer10via an analog interface15. The front-end20interfaces with a B-mode processor30and a control processor50via a digital bus70. The digital bus70may include several digital sub-buses. The digital sub-bases may have separate configurations and provide digital data interfaces to various parts of the ultrasound imaging system5.

The B-mode processor30provides amplitude detection and data compression for B-mode imaging. In a certain embodiment, the B-mode processor30may also be used for other imaging modes, such as M-mode (motion mode), B+M-mode (both motion and brightness), harmonic imaging, and/or Doppler imaging. The B-mode processor30receives digital signal data from the front-end20. The B-mode processor30processes the received digital signal data to produce estimated parameter values. The estimated parameter values may be produced using the received digital signal data. The digital signal data may be analyzed in frequency bands centered at the fundamental, harmonics, or sub-harmonics of the transmitted signals to produce the estimated parameter values. The B-mode processor30passes the estimated parameter values to a control processor50over the digital bus70. The B-mode processor30may also pass the estimated parameter values to a display75via the digital bus70.

A user interface60allows user commands to be input by the operator to the ultrasound imaging system5through the control processor50. The user interface60may include a keyboard, mouse, switches, knobs, buttons, track ball, and/or on screen menus, for example (not shown).

The control processor50is the central processor of the ultrasound imaging system5. The control processor50interfaces to other components of the ultrasound imaging system5using the digital bus70. The control processor50executes various data algorithms and functions for various imaging and diagnostic modes. Digital data and commands may be transmitted and received between the control processor50and other components of the ultrasound imaging system5. In an alternative embodiment, functions performed by the control processor50may be performed by multiple processors and/or may be integrated into the B-mode processor30and/or the display processor80. In another embodiment, the functions of the processors30,50, and80may be integrated into a single personal computer (PC) backend.

FIG. 2illustrates a method200for ultrasound imaging in accordance with an embodiment of the present invention. First, at step210, the transducer10transmits ultrasound energy into a subject, such as a patient. Then, at step220, ultrasound energy or echoes backscattered from the subject are received at the transducer10. Signals are received at the front-end20in response to ultrasound waves backscattered from the subject.

Next, at step230, the received signals are transmitted from the front-end20to the B-mode processor30using the digital bus70. At step240, the B-mode processor30generates parameter values based on the received signals. Then, at step250, the parameter values are sent to the control processor50.

At step260, the control processor50processes the parameter values for use in display, storage, and diagnostics at the display75. The control processor50may combine a plurality of parameter value sets from a plurality of focal zones into a single image for display. The control processor50processes the image data parameter values to reduce banding artifacts in the resulting combined diagnostic image. Banding reduction will be described in further detail below.

Next, at step270, processed parameter values are transmitted to the display75. The display processor80may also process parameter values from a plurality of focal zone images to produce a combined image in conjunction with and/or in addition to the control processor50. Data processing to reduce banding resulting from differences between the focal regions will be discussed further below.

Finally, at step280, a diagnostic image is produced and output at the monitor90. The image may be stored, displayed, printed, and/or further transmitted, for example. The display processor80may produce the diagnostic image using the processed parameter values from the digital signal data.

Banding occurs in a diagnostic image when two or more focal regions having different waveforms and/or f-numbers, for example, are combined. The resulting difference in intensity between the focal regions produces banding in the final image. For purposes of illustration only, the method below will be described with two focal zones. Certain embodiments of the present invention reduce banding in a diagnostic image including a plurality of focal zones by processing the first focal zone, determining intensity around the border or “stitch line” between the two focal zones, and adjusting the second focal region based on the difference in intensity. Preferably, processing is done in real time as image frames are updated.

FIG. 3illustrates an image300with multiple focal zones and corresponding focal regions in accordance with au embodiment of the present invention. As indicated inFIG. 3, a splicer position is a position where a transition between the adjacent focal zones occurs. In a B-mode image, to avoid visible discontinuity at a transition region between adjacent focal zones, fading may also be applied to reduce image mismatch between adjacent zones.

FIG. 4illustrates a method400for banding suppression in accordance with an embodiment of the present invention. First, at step410, two kernels are selected along a splicer position between two focal zones in a middle portion of an image (a combined ultrasound image, for example). The two kernels are called “Up” and “Down” kernels. In an embodiment, the vertical width of a kernel is preferably approximately 3-5 pixels wide, and the lateral width is preferably approximately 80% of the entire image width, for example. However, the exact size of a kernel does not impact the operation of the method400. The splicer position may be included in either the “Down” or “Up” kernel, depending on configuration of the imaging scanner.

Then, at step415, representative image intensity levels are computed for both Up and Down kernels. A median or a mean value, for example, may be used as a representative image intensity level. Since the number of bits of incoming data may be limited (for example to 8bits), an efficient median-searching method may be used.

In certain embodiments, two thresholds may be applied to determine whether a data point may be used in calculating the median of the kernel. In certain embodiments, a bright vessel boundary, a dark cystic structure, and/or blood flow may be present inside a kernel, which may bias an estimation of a median or mean value. For example, for 8 bit data (256 intensity levels), a lower intensity threshold of 10 and an upper intensity threshold of 200 may be used to exclude extreme data points. In an embodiment, the entire image may be processed to determine thresholds before applying band suppression. Methods such as histogram analysis, for example may be used to find optimal thresholds.

Then, at step420, after representative image intensity values for both Up and Down kernels are found, a difference (or delta) between the representative image intensity value of the Up kernel and the representative image intensity value of the Down kernel is calculated.

Next, at step430, a number of dark pixels in a kernel and/or the difference between representative image intensities may be compared to certain threshold(s) to determine if compensation may be applied to the focal zone(s). For example, in an embodiment, if the Up kernel contains a bright vessel boundary and the Down kernel contains low-scattering blood flow (or vice versa), and neither dark nor bright intensity data has been excluded by lower or upper thresholds, the difference between the representative image intensity levels may exceed a threshold, such as 30 or 50, for example. In this case, the band suppression method may be turned off to avoid creating extra banding by biased values.

In another embodiment, a threshold may be used when a kernel contains a large number of dark pixels, such as when a splicer position is partially within or aligned with a blood vessel edge, for example. If a ratio between a number of dark pixels in a kernel to a number of total pixels in a kernel (referred to as a black ratio, for example) exceeds a black ratio limit (usually set to 1), banding compensation may be ignored for the splicer position. For example, a kernel with a black ratio above the defined threshold may contain noise, blood, or fluid.

In another embodiment, if the difference between representative intensity values from Up and Down kernels is less than a difference limit, compensation may not be performed in order to save computation time. For example, a gray level difference between two zones may be less than 50 but greater than an intensity difference corresponding to 4 dB. To avoid over-compensation, a 2 dB difference may be applied to each zone, so total banding compensation will not exceed 4 dB.

At step440, the difference in representative image intensity level may be applied to either focal zone. Preferably, the difference or compensation amount is split equally into two halves. That is, the entire compensation amount may not be applied to one focal zone. Half of the difference may be applied to each side or focal zone (one is positively compensated, and the other zone is negatively compensated, for example).

In an embodiment, compensation may be applied to an entire zone. However, banding may occur near the splicer position, and the compensation may not apply to an entire focal zone equally. Instead, in an embodiment, compensation may be applied in a linearly-decreasing fashion, as shown in FIG.5. In an alternative embodiment, first and last focal zones in an image may be compensated differently. For example, to prevent creation of excessive noise, the difference may be applied to half of the zone in a linearly-decreasing fashion.

If the method400is being used (“turned on”) and if the compensation fluctuates from frame to frame (caused by image noise or certain patient anatomy, for example), the image may be flickering. At step450, a filter, such as a temporal Infinite Impulse Response (IIR) filter, may be applied (similar to frame averaging) to the current and previous frames to avoid flickering in the image. An example of a filter is illustrated in equation (1):
Yout=0.5* Xcurrent+0.5*Yprevious  (1).

In certain embodiments, coefficients may have values other than (0.5, 0.5), such as (0.4, 0.6), etc.

In certain embodiments, when a certain condition occurs, the filter is initialized. For example, if image depth changes, focal zone position changes, number of focal zones changes, imaging frequency changes (from 4 MHz to 6 MHZ, for example), and/or application changes (from thyroid to breast imaging, for example), then the filter may be initialized. Compensation values may be initialized for each zone using values computed from the first focal region after a change occurs. However, if all values are initialized down to zero, initialization may cause some intermittent banding.

At step460, image data may be clipped if the image data exceeds or falls below a certain threshold. Then, at step470, the next splicer position at which focal zones are combined may be examined and the method400may be repeated.

Certain embodiments of the present invention suppress banding by processing the incoming detected and log-compressed data from the B-mode processor30in real-time to extract and compensate the median/mean image intensity difference between adjacent zones. Preferably, the method described in certain embodiments of the present invention is activated for images with at least two focal zones. Certain embodiments reduce banding in real-time using faster processing of signal and image data compared with current methods. Certain embodiments result in more accurate and efficient banding reduction for banding caused by waveforms/frequency/f-number and/or patient habitis (a patient's body build and health, particularly a patient's predisposition to disease), for example. Rather than applying a constant gain compensation across an entire zone, certain embodiments split the gain difference in half and apply ramp-compensation on the split gain difference. Additionally, in certain embodiments, data thresholds may be based on different applications, contrary to the prior art. In certain embodiments, banding suppression may be accomplished using a software-based implementation, for example, rather than by the addition of extra, more complicated, and slower hardware. Certain embodiments of the present invention may be applied to all B-mode applications.