Computed tomography scanners

Abstract: An x-ray source (12) of a CT scanner (10) has a focal spot (30) from which x-rays with an energy distribution (32) are generated. As the x-ray source rotates, a peak (34) of the radiation energy distribution seen by each detector (24) of a ring of detectors through a physical collimator (18) shifts. Each line of data generated by the detectors (24) has a shifting energy distribution across its data values which creates artifacts in reconstructed images. A filter (50) includes a convolver (52) which convolves the lines of data with deconvolution functions from a memory (54) in a recursive loop (56) in accordance with a relative angular position of the x-ray source and the corresponding detector. Filtered data lines are subtracted (56) from unconvolved data values to generate values with reduced off-focal radiation components. Filtered data is conveyed to a convolver (82) and a backprojector (84) which backprojects the data across an image memory (86).

Description: This invention relates to computed tomography (CT) scanners, more especially such scanners suitable for diagnostic imaging. It finds particular application in conjunction with volume imaging with CT scanners and will be described with particular reference thereto. However, it is to be appreciated that the invention also finds application in conjunction with other types of x-ray imaging for medical diagnostic purposes, and with x-ray imaging of manufactured articles to detect internal structures or flaws, and the like. In CT scanners, a fan-shaped plane of radiation about 1 to 10 mm thick is projected through the subject and detected by a plurality of detectors. The output of the detectors is reconstructed into cross-sectional images of a subject representing a series of adjacent planes about 1 to 10 mm thick or a volume with 1 to 10 mm resolution. In third and fourth generation CT scanners, an x-ray source produces radiation as the x-ray source rotates around the examination region. In a third generation scanner, the radiated detectors are rotated with the x-ray source and repeatedly sampled. In a fourth generation scanner, the examination region is surrounded by a stationary ring of detectors that are repeatedly sampled as the x-ray source rotates. The sampled data are then assembled in fans or views of data which are reconstructed, typically convolved and backprojected, to form a cross-sectional representation of the patient for display on a monitor. Prior art methods of reconstructing CT images assume that data inputted to convolvers and backprojectors are in the form of line integrals. The line integrals are each an integral of the attenuation of radiation of a characteristic uniform intensity along a line between the detector and the focal spot of the x-ray source. The line integrals are estimated from data sampled from the plurality of the detectors. Unfortunately, sampled data from conventional CT scanners do not yield proportional line integrals. The x-ray radiation emanating from the x-ray source is in the form of a bundle of rays having differing strengths rather than a single ray of uniform intensity. Rays from x-ray source are classified in two distinct groups. Rays from the center of the focal spot or on-focal radiation have a strong intensity and narrow base. Rays from regions of the anode surround the focal spot or off-focal radiation have a lower intensity but a broader base. Errors to line integral proportionality are introduced by the off-focal rays, i.e., those rays with a weak intensity origination over the broad base. The intensity errors and the broadened focal spot result in reconstructed images having major artifacts. Prior art CT scanners have attempted to correct the adverse effects created by off-focal radiation with physical collimation and electronic notch filters. In particular, these CT scanners have attempted to incorporate a correction factor in the reconstruction filter, known as an image algorithm. This correction involves convolving sampled data with a constant kernel using a fast fourier transformer. The prior art presumes that the off-focal radiation is static, i.e., the convolution kernel is invariant. The distribution of off-focal radiation as seen by the detector varies as a function of x-ray source position around the examination region. For instance, the detector at the center of a fan beam of radiation may see an intensity distribution having a large amplitude on-focal spike at the center with low amplitude off-focal shoulders to either side. As the x-ray source rotates, the detector may then see a large intensity on-focal spike at one side with an off-focal shoulder extending to the other side. Given the dynamic nature of the intensity distribution, convolving off-focal radiation using a constant deconvolution kernel will reduce off-focal effects for only a limited number of x-ray sourcew positions, i.e. those positions which generate intensity distribution with on-focal spikes at the center. From the document WO 87/02821 a collimator arrangement close to the x-ray source is known that is used to suppress off-focal radiation and to spatially shape the x-ray beam intensity. This collimator arrangement essentially consists of thin radiation absorbing sheets that are focused towards the focal spot area. From the European patent application EP 0 471 455 A2 a filter function for interpolating values from contiguous source-fan views acquired at the same angular position of the x-ray source is known. This filter function varies from ray to ray of a given source-fan view and is used in spiral scan acquisition performed with a 4 th generation CT scanner. This filter function allows for 180° plus fan angle reconstruction that yields a planar rather than a warped image representation. The present invention provides a CT scanner which alleviates the above described problems. In accordance with the invention there is provided a CT scanner including a radiation source which generates a beam of radiation, a gantry on which the radiation source is mounted for rotation around an examination region, and a multiplicity of radiation detectors which receive radiation from the source that has traversed the examination region and convert the received radiation into electronic data, and characterized by: an off-focal radiation filter which filters the electronic data with a filter function that changes in accordance with the relative position of the radiation source and the radiation detector at which the electronic data was detected. In one particular embodiment the filter includes a deconvolution memory for storing deconvolution functions, the deconvolution function memory being addressed in accordance with relative positions of the radiation source and the detectors to retrieve corresponding deconvolution functions; and a convolver for convolving the electronic data with the retrieved deconvolution functions. The method of use of the CT-scanner of the invention comprises the steps of generating a CT scanner image including directing a fan beam of radiation from a focal spot of a radiation source across an examination region, detecting radiation which has traversed the examination region with an arc of radiation detectors, repeatedly sampling each of the radiation detectors to convert each ray of the detected radiation into a corresponding electronic data value, filtering the electronic data values, and reconstructing the filtered electronic data values into an electronic image representation, and characterized by: determining a relative angle of each ray of radiation; and filtering the electronic data values with an off-focal radiation filter whose filtering changes in accordance with the relative angle of each ray. One advantage of the present invention is that it provides images having greater clarity. Another advantage of the present invention is that it reduces artifacts from displayed images. Yet another advantage of the present invention is that it removes radiation source position dependent off-focal radiation components from the data collected from the detectors. One CT scanner in accordance with the invention will now be described, by way of example, with reference to the accompanying drawings in which:- FIGURE 1 is a diagrammatic illustration of the CT scanner; FIGURE 2 is a diagram depicting a radiation beam energy distribution offset problem arising in the scanner and its correction in accordance with the invention; and FIGURE 3 is a flow chart diagramming data handling in a recursive loop of the scanner. With reference to Figure 1, the scanner 10 includes a radiation source 12, such as an x-ray tube, for projecting a beam of radiation through an examination region or scan circle 14. The x-ray source 12 is mounted on a rotatable gantry 16 to rotate the fan beam of radiation around the examination region 14. A collimator and shutter 18 collimates the beam of radiation into one or more narrow fans and selectively gates the beam on and off. A motor 20 provides motive power for rotating the gantry 16, preferably continuously, around the examination region 14. A rotational position encoder or tachometer 22 is connected with the motor 20 and gantry 16 to measure the angular position or relative angle of the gantry 16 and hence an apex of the fan beam of radiation. In the illustrated fourth generation CT scanner, a ring of radiation detectors 24, preferably about 4,000 detectors, are mounted peripherally around the examination region 14. For mechanical and mathematical conveniences, the detectors 24 are stationarily mounted around the gantry 16 in the same plane as the x-ray source 12. Alternately, an arc of detectors can be mounted to the gantry 16 across the examination region 14 from the x-ray source 12, i.e. a third generation scanner. Each of the radiation detectors 24 produces an output signal proportional to the intensity of radiation projected through the examination region 14. An optional reference detector (not shown) may detect radiation which has not traversed the subject (not shown) in the examination region 14. The difference between the magnitude of the radiation received by the reference detector and each detector 18 of the ring provides an indication of the amount of radiation attenuation along a corresponding ray of a sampled fan of radiation. With reference to FIGURE 2, the x-ray source 12 has a focal spot 30, across which there is an energy distribution 32. The energy distribution includes a peak 34 of higher intensity radiation and shoulders 36 of lower intensity radiation. The collimator 18 defines an aperture such that a central ray 38 from the center of the focal spot 30 traverses the center of the examination region 14. The edges of the collimator aperture limit the radiation to a fan with edge rays 40 that are tangent to opposite sides of the examination region. An exemplary one of the radiation detectors 24 shown in Figure 2 receives radiation as the x-ray source 12 moves in back of the examination region 14. That is, the detector first sees the radiation source 12 when it rotates to position A in which a leading edge tangential ray strikes the detector. In position A, parallax between the collimator 18 and the focal spot 30 causes the energy peak 34 to be offset as shown in distribution 42. As the x-ray source 12 continues to rotate, the energy peak 34 shifts. In x-ray source position B, in which the detector receives the central ray 38, the energy peak 34 is centered as shown in distribution 44. The peak continues to shift until as shown in position C only the trailing edge tangential ray strikes the detector. In the trailing edge detector ray, the energy peak 34 is shifted to a right side of the visible window as shown in distribution 46. The exemplary detector, like the other detectors 24, is sampled repeatedly as the x-ray source 12 rotates between positions A and C to generate the data values of a detector fan data line or view. It will be noted that the energy peak shifts with each point in the data line. Analogously, all detectors that are irradiated at one time by the fan of radiation can be sampled concurrently to generate data values of a source fan data line or view. Again, the position of the peak and the viewable portion of the energy distribution shift across the view. Of course, from the angular position of the focal spot 30 as determined by the rotational position encoder 22 and the detector as determined by the geometry of the system, the shift in the energy distribution for each data value of a corresponding line is known. Due to the circular symmetry, the shift in the energy distribution for all like sampled data lines is known. In a third generation scanner, the energy distribution is the same for all source fan data lines. In a fourth generation scanner, if the detectors are sampled M times before the leading (or trailing) ray moves a distance or angle equal to one detector, there are M energy distribution shift patterns. With reference again to FIGURE 1, the data lines are temporally stored and accumulated in a buffer memory 48 with an indication of their corresponding angular position around the examination region 14. As each line of digital data is completed, it is filtered by filter 50 to reduce effects of x-ray source position dependent off-focal radiation. It will be noted that in fourth generation scanners, as the x-ray source moves, each irradiated detector is concurrently generating intensity data. In order to accommodate this rapid flow of information, the filter preferably includes a plurality of parallel filters to process several data lines concurrently. Data d from the detectors can mathematically be represented by: d = (δ + e) * X where X represents desired on-focal data from the distribution peak 34 along a ray from the x-ray source 12, and e*X represents the error component sought to be filtered. The goal is to estimate X from the measured data d. From a theoretical point of view, X can be computed by convolving the data d with an estimated deconvolution function inversely related to (δ + e). Calculating X can be generally achieved using the following equation: where k i is an estimated deconvolution function whose values are dependent on x-ray source position and p j is related to X. The filter 50 implements Equation (2). In the preferred embodiment, each filter 50 includes a convolver 52 which convolves data d from a line of data with deconvolution function retrieved from a deconvolution function memory 54. The deconvolution function memory 54 holds a plurality of estimated deconvolution functions for a plurality of x-ray source positions. The filter 50 selects an appropriate deconvolution function in accordance with the relative angular position of the x-ray source and the detector for each data value of each data line to be convolved. A recursive loop 56 reduces the number of operations to produce p j , as is more fully explained below. The output of the filter 50 represents an approximation of X, the on-focal data. Error deviation for X is on the order of x*e 2 (power in convolution). A higher order of approximation can be mathematically derived. Second and third order corrections are provided as follows: X ≅ Y 2 - 3Y 1 + 3Y 0 X ≅ -Y 3 + 4Y 2 - 6Y 1 + 4Y 0 where Y 0 = d, and Y k+1 = (δ + e)*Y k , for k≥0. Radiation from the x-ray radiation source 12 can be described as a large amplitude spike at the center with low amplitude shoulders to either side as illustrated by distribution 32. At least, this is the distribution which the detector at the center of the fan beam sees. At one extreme of the fan, the radiation detector sees the spike offset completely to one side and a larger shoulder extending to the other side; at the other extreme ray of the fan, the spike is offset to the opposite side with the shoulder extending to the first side. The ray intensity from the x-ray radiation source is defined by: f(i) = w i if|i| ≤ N f(i) = w if|i| > N where N is a fixed number, e.g. N = 10. These predefined ray intensities f(i) provide the deconvolution functions in accordance with the following equation: Normally, the right side denominator of the immediately preceding equation varies as a function of the scanner geometry and beam limiters of the x-ray source. However, with certain CT scanner equipment the summation of f(m) is independent of scanner geometries and beam limiters. CT scanners manufactured by Picker International Inc. of Highland Heights, Ohio, USA exhibit these input characteristics. Using Picker CT scanners, the summation of f(m) can be replaced by a constant k. Accordingly, the deconvolution function is reduced to: k i = f(i) k Filtering off-focal radiation can also be implemented utilizing the equation as follows: where q is a deconvolution function and is defined by: q 0 = 2- f(0) k and q i = f(i) k for i ≠ 0. The deconvolution memory has at least enough memory for storing 2N + 2 values for deconvolution kernel q. Looking at the present filtering technique from a more mathematical perspective, the final result for each ray j of a detector fan can be expressed by: p j = I 1 + Q (I 2 + I 3 ) where the precomputed value Q is given by: Q = - w k and I 1 , I 2 and I 3 can be expressed with the following: for i = -N to i = N, and for j = S to j = -N -1 and j = n +1 to j = E, and I 3 = d j+s +S(d j+S-1 -d j+S ) + d j+E +e(d j+E+1 -d j+E ) Before the final result for ray j are computed, the starting and end limits s(j) and e(j) respectively, are calculated using: s(j) = a*j + b and e(j) = c*j + d Moreover, the computational limits for s and e are calculated by taking the fixed value of s(j) and e(j). In other words, S = S - s(j) e = e(j) - E S = fix (s(j)) and E = fix (e(j)) In implementing the filtering equation set forth in equation 12, the kernel q(i) is restricted to a constant Q for all i different than zero. This is the case when N=0, but this condition does not impose any practical restriction since in practice, N is smaller than |S(0)| and |e(1023)| wherein ray 0 and ray 1023 are the first and last rays of each view, respectively. Consequently, the deconvolution kernel Q from -N to N is ray independent. Assuming N=0, the deconvolution formula of equation 12 can be rearranged to become: Consequently, the following recursive formula can be implemented by the recursive loop 56. In this formula, each ray p j+1 is corrected based on the preceding ray p j . p j+1 = P j + (q(0)-Q) (d j+1 -d j ) + Q*(A - S) where A and S are calculated from the following expressions: for k= j + s(j) to j+1+s(j+1)and for k = j + e(j) to j+1+e(j+1). It should be noted that: j+1+s(j+1) = (j+s(j))+(1+a) and j+1+e(j+1) = (j+e(j))+(1+c) A and S are computed by: S=(1+a)d [j+s(j)] +FP(j+1+s(j+1))(d [j+1+s(j+1)] -d [j+s(j)] ) and A=(1+c)d[ j+e(j)] +FP(j+1+e(j+1))(d [j+1+e(j+1)] -d [j+e(j)] ) where [x] denotes the greatest integer ≤ x and FP(x) = x - [x). Within each view, these recursive relations reduce the number of operations to find p j . To implement the recursive relations, the initial values are set as: j = [-b/(1+a) + 1] s = (1+a)j + b m = [s] fs = s - m e = (1+c)j + d n = [e] fe = e - n and P j = (q(0) - Q)*d j + Q*(-fs*d m +d m+1 +...+d n-1 +fe*d n ) In each iteration of the recursive loop 56, the variables are indexed as follows: d0 = d m d1 = d n d = d j p = p j s = s + (1+a) m = [s] fs = s - m e = e + (1+c) n = [e] fe = e - n j = j + 1 and p j =p+(q(0)-Q)*(d j -d) + Q*((1+c)*d1+fe*(d n -d1)-(1+a)*d0-fs*(d m -d0)) The loop repeats until n reaches the last ray of the fan. With reference to FIGURE 3, the initial values of j, s, m, fs, e, n, fe, and p j are set in a step 60. The initial data value d j is read in a step 62. A processor or step 64 calculates p j in accordance with Equation 38. A memory 66 stores p j . A step 68 resets the values according to Equations 39a-48 including indexes j. The next data value d j is read at step 70. The next value of p j is calculated at step 72 and loaded into the memory 66. The loop is iteratively repeated for all d j . After each line of data is filtered, it is conveyed to a reconstruction processor 80. In the preferred embodiment, the reconstruction processor includes a convolver 82 which convolves each filtered data line with a convolution or filter function. Again, to accommodate the rapid flow of information in fourth generation scanners, the convolver preferably includes parallel convolvers for convolving several filtered data lines concurrently. The convolved data lines are conveyed to a backprojector 84 which backprojects each convolved data line across an image memory 86 to reconstruct an electronic image representation. A video processor 88 retrieves selected portions of the data in the image memory 86 and converts them to appropriate format for display on a video monitor 90. Exemplary images include single slice images, multiple slice images, surfaced rendered volume images, surface rendered volume images with cut planes, and the like. Reconstructed electronic images may also be sent to a central storage system for subsequent recall and viewing. Hard copy printers (not shown) convert the electronic images to human readable paper images.

Claims: A CT scanner including a radiation source (12) which generates a beam of radiation, a gantry (16) on which the radiation source (12) is mounted for rotation around an examination region (14), and a multiplicity of radiation detectors (24) which receive radiation from the source (12) that has traversed the examination region (14) and convert the received radiation into electronic data, and characterized by :an off-focal radiation filter (50) which filters the electronic data with a filter function that changes in accordance with the relative position of the radiation source (12) and the radiation detector (24) at which the electronic data was detected. A CT scanner according to claim 1 including an angular position encoder (22) for measuring the position of the radiation source (12) relative to the radiation detectors (24). A CT scanner according to claim 1 or claim 2 wherein the filter (50) includes a convolver (52) for convolving the electronic data. A CT scanner according to claim 3 including means (48) for processing the electronic data into sets of parallel ray data, the parallel ray data sets being filtered by the off-focal radiation filter (50). A CT scanner according to claim 3 or claim 4 wherein the off-focal radiation filter (50) includes a recursive loop (56)in which each electronic data value corresponding to a ray is filtered using a data value of a corresponding preceding ray. A CT scanner according to claim 1 or claim 2 wherein the filter (50) includes: a deconvolution memory (54) for storing deconvolution functions, the deconvolution function memory (54) being addressed in accordance with relative positions of the radiation source (12) and the detectors (24) to retrieve corresponding deconvolution functions; and a convolver (52) for convolving the electronic data with the retrieved deconvolution functions. A CT scanner according to claim 6 wherein by the convolver (52) includes an iterative loop (56).

Inventors: TUY HEANG K, TUY, HEANG K., TUY, HEANG K.,
Applicants: PHILIPS MED SYST INC, PHILIPS MEDICAL SYSTEMS (CLEVELAND), INC.
Application year: 1995
Main classification: A61B6
IPC codes: A61B6, G06T11