Method and system for target angle heel effect compensation

Patent Center

The Netherlands

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© 1027643 © C PATENT²⁰ © Patent application: 1027643 &lt;© Filed: 02.12.2004 © lnt.CI .:

A61B6 / 03 (2006.01)

<td>© Priority:</td><td>© Patent holder (s):</td>

<td>05.12.2003 US 10/728611</td><td>GE Medical Systems Global Technology</td>

<td></td><td>Company, LLC of Waukesha, Wisconsin,</td>

<td>© Registered:</td><td>United States of America (US).</td>

<td>07.06.2005 IE 2005/08</td><td></td>

<td></td><td>© Inventor (s):</td>

<td>© Date:</td><td>Jiang Hsieh of Brookfield, Wisconsin (US).</td>

<td> 27.06.2006</td><td></td>

<td></td><td>© Authorized representative:</td>

<td>© Published:</td><td>Ir. HV Mertens et al. At 2280 GE Rijswijk.</td>

<td>01.09.2006 IE 2006/09</td><td></td>

© Target angle heel effect compensation method and system.

An imaging system (10) for scanning an object (22) includes an X-ray source (14), an X-ray detector (18) with a plurality of detector rows (20) positioned to receive X-rays (16) from the source, an computer (36) operatively coupled to the X-ray source and the X-ray detector and a filter positioned between the X-ray source and the X-ray detector, comprising a non-stationary filter, wherein the non-stationary filter is arranged to filter different detector rows differently based on the detector row number.

NLC1027643

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The contents of this patent correspond to the original filed description with claim (s) and any drawing (s).

The Netherlands Patent Office is the Office for Industrial Property, an agency of the Ministry of Economic Affairs

Short designation: Method and system for target angle heel effect compensation.

The invention generally relates to computed tomography (CT) imaging systems, and more particularly to target angle effect compensation.

In at least some known imaging systems, an X-ray tube source projects an X-ray beam that passes through an object to be imaged, such as a patient, and strikes an array of X-ray detector rows. This technique is quite effective in medical CT scanners, but has some drawbacks when the detector coverage area becomes large, such as in the case of multi-slice CT. With the advent of multi-slice CT imaging systems, which include a number of detector rows, there are at least two major drawbacks, a non-uniform X-ray flux and a non-uniform slice thickness. The non-uniform X-ray flux can result in a bead effect and a non-uniform slice thickness can result in spatial resolution variations.

The heel effect can produce image quality differences across the detector rows. For example, a 40 mm volumetric computed tomography (CVT) detector with a nominal 7 ° target angle has an effective target angle of 5 ° at the outer anode silk yard and 9 ° at the outer cathode silk yard, resulting in an intensity variation of roughly 20% of one end of the detector to another. This variation in radiation intensity due to the heel effect reduces the image quality across the X-ray detector rows and thereby reduces the image quality of the X-rays.

A non-uniform slice thickness results when a first projected focal spot height is significantly greater than a second projected focal spot height. A non-uniform slice thickness translates into a spatial resolution in the z direction, which is a function of the detector array.

In multi-slice CT, it is desirable to design a system such that the X-ray flux and the spatial resolution do not change significantly from detector row to detector row.

In one aspect, a method of at least partially compensating an X-ray target angle heel effect is provided. The method includes providing an X-ray source, providing an X-ray detector with a plurality of detector rows positioned to receive X-rays from the source, and using a filter to ensure the uniformity of at least one of projection noise and spatial resolution where the projection noise and spatial resolution are nonuniform and a function of a target angle along a z axis.

In another aspect, an imaging system for scanning an object is provided. The imaging system includes an X-ray source, an X-ray detector with a plurality of detector rows positioned to receive X-rays from the source, an X-ray source and X-ray detector operatively coupled computer, which is configured to filter at least one of projection noise and spatial resolution to increase uniformity, and wherein the projection noise and spatial resolution are non-uniform and a function of the target angle along a z axis.

In yet another aspect, a computer readable medium is provided, which medium is encoded with a system executable program to at least partially compensate for an X-ray target angle heel effect. The program is arranged to instruct the computer to provide an X-ray source, to provide an X-ray detector with a plurality of detector rows positioned to receive X-rays from the source, and to use a filter to ensure the uniformity of at least one of projection noise and spatial resolution, where the projection noise and spatial resolution are non-uniform and a function of the target angle along a z axis.

Fig. 1 is an illustrative view of a CT imaging system.

Fig. 2 is a block diagram of the system shown in FIG. 1.

Fig. 3 shows an X-ray emission.

Fig. 4 shows the target angle heel effect.

Fig. 5 shows the variation of the projected focal spot height along a z axis.

Fig. 6 is a graph showing the calculated FWHM and FWZM.

Fig. 7 is a graph showing the calculated FWHM, FWZM and equivalent flux after adaptive filtering.

Fig. 8 is a block diagram of a bead effect compensation method.

0276433

In some known CT imaging system configurations, an X-ray source projects a fan-shaped beam collimated to lie within an XY plane of a Cartesian coordinate system and generally referred to as an imaging plane. The radiation beam passes through an object to be imaged, such as a patient. After being attenuated by the object, the beam strikes an array of radiation detectors. The intensity of the attenuated radiation beam received on the detector array depends on the attenuation of an X-ray beam caused by the object. Each detector10 element of the array produces a separate electrical signal, which is a measurement of the beam intensity at the detector location. The intensity measurements of all detectors are acquired separately to produce a transmission profile.

In third-generation CT systems, the X-ray source and detector array are rotated with a portal in the imaging plane and around the object to be imaged, so that the angle at which the X-ray beam intersects the object is constantly changing. A group of X-ray attenuation measurements, ie projection data, from the detector array at one gantry angle, is referred to as a view. A scan of the object includes a series of views taken at different gantry or viewing angles during one revolution of the X-ray source and detector.

In an axial scan, the projection data is processed to construct an image corresponding to a two-dimensional slice of the object. One method of reconstructing an image from a series of projection data is referred to in the art by the term filtered back projection technique. This process converts the scan attenuation measurements into integers, called CT numbers or Hounsfield units (HU), which are used to control the brightness of a corresponding pixel on a cathode ray tube display.

To reduce the total scan time, a helical scan can be performed. To perform a helical scan, the patient is moved while the data for the prescribed number of slices is acquired. Such a system generates a single helix from a fan beam helical scan. The helix shown by the fan beam provides projection data from which images can be reconstructed in any prescribed slice.

Helical scan reconstruction algorithms typically use helical weighting algorithms, which weigh the collected data as a function of the viewing angle and detector channel index. In particular, prior to a filtered back projection process, the data is weighted according to a helical weighting factor, which is a function of the gantry angle and the detector angle. The weighted data is then processed to generate CT numbers and to construct an image corresponding to a two-dimensional slice of the object.

To further reduce the total acquisition time, multiplak CT has been introduced. In multi-slice CT, multiple rows of projection data are acquired simultaneously at each time. In combination with a helical scan mode, the system generates a single helix of cone beam projection data. Similar to the single slice helical weighting scheme, a method can be derived to multiply the weighting factor by the projection data prior to the filtered backprojection algorithm.

As used herein, a singular and preceded word or step by the word should not be construed as plural forms thereof solely, unless such exclusion is exclusively stated. References to one embodiment of the invention are not intended to be construed as the existence of additional embodiments, which also include the stated features, only.

As used herein, the phrase reconstructing an image is not intended to exclude embodiments of the invention in which data representing an image is generated but not a visible image. However, many embodiments (or are arranged to generate) generate at least one visible image.

Reference is now made to Figs. 1 and 2, which show a multi-slice scanning imaging system, for example, a computed tomography (CT) imaging system 10, which system includes a portal 12 representative of a third generation CT imaging 35 imaging system. Portal 12 has an X-ray tube 14 (also referred to as X-ray source 14 herein) which projects a beam of X-rays 16 onto a detector array 18 on the opposite side of portal 12. The detector matrix 18 is formed by a number of detector rows (not shown in FIGS. 1 and 2), which comprise a number of detector

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include elements 20 which together detect the projected x-rays passing through an object, such as a medical patient 22, between array 18 and source 14. Each detector element 20 produces an electrical signal, representing the intensity of an incident X-ray beam, and therefore can be used to estimate attenuation of the beam as it passes through the object or patient 22. During a scan to acquire X-ray projection data, the portal 12 and the components mounted thereon rotate about a center of rotation 24. FIG. 2 shows only a single row of detector elements 20 (ie, a detector row). However, a multi-slice detector array 18 includes a plurality of parallel detector rows of detector elements 20, so that projection data corresponding to a number of quasi-parallel or parallel slices is acquired simultaneously during a scan.

The rotation of the components on the portal 12 and the operation of the X-ray source 14 are controlled by a control mechanism 26 of the CT system 10. The control mechanism 26 includes an X-ray control 28, which provides energy and timing signals to the X-ray source 14, and a gantry motor control 30, which controls the rotational speed and position of the components on the gantry 12. A data acquisition system (DAS) 32 in the control mechanism 26 samples the analog data from the detector elements 20 and converts the data into digital signals for subsequent processing. An image reconstruction element 34 receives the sampled and digitized X-ray data from DAS 32 and performs high-speed image reconstruction. The reconstructed image is supplied as an input to a computer 36, which stores the image in a storage device 38. The image reconstruction element may be specialized equipment or computer programs for execution on the computer 36.

The computer 36 also receives commands and scan parameters from an operator through a console 40, which has a keyboard. An associated cathode ray tube display 42 allows the operator to observe the reconstructed image and other data from computer 36. The commands and parameters provided by the operator are used by the computer 36 to provide control signals and information to DAS 32, the X-ray controller 28, and the portal engine controller 30. In addition, the computer 36 controls a table motor controller 44, which controller controls a motorized table 46 to position a patient 22 in the portal 12. In particular, the table 26 moves parts of the patient 22 through a portal opening 48.

In one embodiment, the computer 36 includes a device 50, for example, a flexible disk drive, CD-ROM drive, DVD drive, magnetic-optical disk (MOD) device, or any other digital device that includes a network connection device, such as an Ethernet device , contains for reading instructions and / or data from a computer readable medium 52, such as a flexible disk, a CD-ROM, a DVD or other digital source, such as a network or the

Internet, as well as digital resources yet to be developed. In another embodiment, the computer 36 executes stored instructions from the manufacturer installed software (not shown). The computer 36 is programmed to perform functions described herein, and accordingly the term computer used herein is not limited to only those integrated circuits referred to in the art as computers, but broadly refers to computers, processors, microcontrollers, microcomputers, programmable logic controllers, application specific integrated circuits and other programmable circuits, and these terms are used interchangeably herein. While the specific embodiment mentioned above refers to a third-generation CT system, the methods described herein apply equally to fourth-generation CT systems (stationary detector - rotating X-ray source) and fifth-generation CT systems (stationary detector and X-ray source) . In addition, it is contemplated that the advantages of the invention be valid for imaging modalities other than CT. In addition, while the methods and equipment described herein have been described in a medical arrangement, it is contemplated that the benefits of the invention be valid for non-medical imaging systems, such as those typically used in an industrial environment or transportation environment, such as for example, but not limited to, a baggage scanning system for an airport or other transportation center.

Described herein are methods and equipment for at least partially compensating for the bead effect by increasing the uniformity of the standard deviation of the radiation measurements performed by the X-ray detector rows using a non-stationary filter, in one embodiment.

W27 &amp; 437

Reference is now made to Figures 3-5, in which an X-ray tube 14 contains a cathode coil 102 and a target 104. Generally, electrons 106 move within the X-ray tube 14 from the coil 102 to the target 104.

The electrodes 106 contact the target 104 in different locations and cause emission of X-rays 16 and cause the emission of X-rays 16 at different angles to leave the target 14 towards the detector 18 and the detector rows 20. In FIG. electrons 106 in particular contact with a surface 10 of the target 14 at an acute angle α 110 with respect to a scanning plane 112 along an actual focal dot length L 114. A projected height of the focal dot h 116 is equal to L 114 multiplied by sin (a).

h = Lsin (a) (1)

The result is that the current focal spot length L 114 is significantly greater than the projected focal spot height h 116.

This has two main effects. One is a non-uniform X-ray flux across the slices in the z direction (shown in Figure 4) and the other is a non-uniform slice thickness (shown in Figure 5).

Fig. 4 shows a number of X-rays 16, a target angle ξ 120 and an average interaction depth 122. The number of X-rays 16 contains a first X-ray 124 and a second X-ray 126. The X-ray 124 emitted by the target 104 is a first distance d_X 130 within the target 104 before exiting the target 104. This is a shorter distance than a second distance d₂ 132, which is deposited by the X-ray 126 within the target 104 before leaving the target 104. Since d₂ 132 is greater than d_X 130, the X-ray 126 exits target 104 more attenuated than X-ray 124. This attenuation difference is the heel effect. As a result, the X-ray flux intensity changes as a function of the target angle ξ 120 formed by the X-rays 16 and the target surface 108. Given the one-to-one relationship between the target angle ξ 120 and the detector row 20, the X-ray intensity changes monotonically with each detector row 20. In general, a smaller target angle ξ 120 corresponds to a reduced X-ray flux intensity.

Fig. 5 shows the non-uniformity of the slice thickness. As explained above, electrons 106 contact target 104

I 027643s in different places and cause the X-ray emissions 16 at different angles to leave the target 14 towards the detector 18 and the detector rows 20. The non-uniformity of the slice thickness is caused by the variation of the projected focal spot height h 116 (shown in fig. 3). In particular, as shown in Fig. 5, a projected focal spot height 140 is significantly greater than the projected focal spot height h₂ 142. This translates into a spatial resolution variation in the z direction as a function of detector row 20.

In multi-slice CT, it is desirable to design a system such that both the X-ray flux strongly associated with the projection noise and the spatial resolution from detector row to detector row do not change significantly. To fully understand the impact of the heel effect, a simple model is used to estimate the spatial resolution of the system, although more complicated models can be used, the overall conclusion will not change. Generally, the spatial resolution of the system in z, s (z) is the convolution of the projected focal dot function, h (z), with the projected detector aperture function, d (z).

s (z) = h (z) ® d (z) (2)

The projected focal spot function, h (z), is obtained under the assumption of a point detector, while a projected detector aperture is assumed under the assumption of an X-ray source. The spatial resolution of the system in z, s (z) is often referred to as the slice sensitivity profile (SSP). If the focal dot function and the detector aperture function are approximated by means of rectangular functions, then a full width half maximum (FWHM) of the SSP is equal to the larger FWHM of the two projected functions. Similarly, the full-width-by-zero maximum (FWZM) is the sum of the widths of the two projected functions.

Fig. 6 is a graph 200 showing the FWHM 210 of the SSP and the FWZM 220 of the two projected functions. Graph 20 clearly shows the significant non-uniformity of SSP. FWHM 210 and FWZM 220 are calculated for a detector aperture of 0.525 mm in an iso center (0.625 mm detector pitch with 0.1 mm tungsten wire between the detector rows), a nominal focal spot height of 1.2 mm measured in ί

027643 “the central plane, a detector range of 40 mm in the iso center and a target angle of 7 °.

The examination of the non-uniformity of the X-ray flux is more complicated, as it requires the estimation of the mean interaction depth, the mean X-ray energy and the corresponding attenuation characteristics of the target. It has been determined that for a target angle of 7 ° and a coverage area in the iso center of 40 mm, the intensity variation from one end of the detector to the other is roughly 20%.

Our analysis indicates that SSP increases with the target angle ξ 120 (shown in Fig. 4), while the projection noise decreases with the target angle ξ 120. It is desirable to reduce both variations across the detector rows 20. One approach to reduce variation is to use adaptive filtering. The use of a low pass filter generally reduces the spatial resolution and at the same time reduces the noise. We can design a non-stationary filter (along z), the characteristics of which change as a function of the detector row 20. That is, the degree of smoothing decreases as a function of the target angle ξ 120. Since FWHM

220 and FWZM 230 do not change to the same extent, the shape of a filter kernel should also change so that the uniformity of both characteristics is met. The techniques for designing such filters are well known, since the original SSP, s (z), and the intended SSP function, t (z), are both known.

For example, we can set the original SSP function of the largest target angle ξ 120 as the intended response function. The filtering function f (z, can then be derived using known signal processing techniques to satisfy the following relationship:

t (z) = s (z) ® f (z) (3)

One of the most widely known techniques is, for example, the Weiner filter. The degree of smoothing increases with the reduction of the target angle ξ 120 (shown in Fig. 4). As a result, the amount of noise reduction also increases with the reduction of the target angle ξ 120. This translates into an enhancement of equivalent photon flux. Fig. 7 is a graph 300 showing the resulting FWHM 310 and FWZM 320 after filtering. Note that both parameters are much more uniform across z. At the same time, the equivalent photon flux magnification 330 from one end of the detector 20 (shown in FIG. 4) to the other end is roughly 20%. Given the previous research, that the flux drop is roughly 20%, the resulting noise in the projection with the filtering has become more uniform.

The target function, t (z), can be selected different from the SSP of the largest target angle ξ 120. For example, the target function can be chosen as the SSP of the central row. In this case, the filter functions for a target angle ξ 120 will be greater than the central row of high-pass filters. The filter functions for a target angle ξ

120 smaller than the central row will be low pass filters. As a result, the noise level for projections collected with rows of larger target angle ξ 120 is increased and the noise level for projections collected with rows of smaller target angle ξ 120 is reduced.

Alternatively, the filtering function can be determined based on the X-ray flux instead of the SSP. Based on previous analysis, it is known that the X-ray flux decreases as the target angle ξ 120 decreases. Therefore, a series of filtering functions is derived so that it smoothes the projection noise across all detector rows. For example, a 3-point filter can be used to produce the final projection based on the projections of the neighboring three detector rows. Filter coefficients for rows k-1, k and k + 1 are denoted by w_k_, w_k, w_{k +}, and r |_k is the normalized X-ray flux level for detector row (the reference row has a normalized flux level of one), where the filter must meet the following relationship:

w²k- + w²k, + w²k + = r |_k (4)

Optionally, the filter design may depend on the combined X-ray flux and SSP. For example, the filter is determined based on the X-ray flux for a target angle ξ 120 greater than the central row and the filter is determined based on the SSP function for a target angle ξ 120 smaller than the central row.

The application of the non-stationary filter can be combined with the non-uniform detector size. That is, the multi-slice detector can be designed so that the detector aperture changes from row to row. In this combination, the uniformity of the X-ray flux and the system response are obtained with both techniques. This thus allows for a more relaxed filter design.

Fig. 8 is a method 400 for promoting reduction of artifacts using the imaging system

10. The method 400 includes providing 410 an X-ray source, providing 420 an X-ray detector with a plurality of detector rows positioned to receive X-rays from the source, and using 430 a filter to ensure the uniformity of at least one of the projection noise and spatial resolution, wherein the X-ray flux and the spatial resolution are nonuniform and a function of the target angle along a z axis.

Previous attempts to reduce the bead effects of an X-ray tube have focused on modifying the detector angle of the X-ray tube. This approach exchanges uniformity of the X-ray flux and system response with the efficiency of the X-ray tube. The proposed approach eliminates such a compromise. In addition, redesign of the X-ray tube has been eliminated.

Exemplary embodiments of a bead effect compensation filter are described in detail above. The assemblies are not limited to the specific embodiments described herein, but instead components of each assembly can be used independently and separately from other components described herein.

Although the invention has been described in terms of various specific embodiments, those skilled in the art will recognize that the invention may be practiced with modifications within the spirit and scope of the claims.

027643e