Streak suppression filter for use in computed tomography systems

The disclosed streak suppression filter is for use with a Computed Tomography (CT) system. The CT system generates a plurality of projection data signals each of which is representative of the density of a portion of an object. The streak suppression filter includes a spatial filter for receiving the projection data signals and for generating therefrom a low frequency signal and a high frequency signal. The streak suppression filter further includes a non-linear filter for generating a filtered signal from the high frequency signal. The streak suppression filter further includes a device for combining the low frequency and filtered signals to generate a streak corrected signal. The streak suppression filter may form the low frequency signal by low pass filtering the raw density signals and may form the high frequency signal by subtracting the low frequency signal from the corresponding raw density signal. The streak suppression filter may form the filtered signal by clipping, or thresholding, the high frequency signal.

FIELD OF THE INVENTION 
The present invention relates generally to Computed Tomography (CT) systems 
used in the medical arts for generating CT images of, for example, human 
patients. More particularly, the invention relates to an improved streak 
suppression filter for reducing streaks in CT images. 
BACKGROUND OF THE INVENTION 
Computed Tomography (CT) systems of the third generation type include an 
X-ray source and an X-ray detector system secured respectively to 
diametrically opposite sides of an annular-shaped disk. The disk is 
rotatably mounted within a gantry support so that during a scan, the disk 
continuously rotates about a rotation axis while X-rays pass from the 
source through an object positioned within the opening of the disk to the 
detector system. 
The detector system typically includes an array of detectors disposed as a 
single row in the shape of an arc of a circle having a center of curvature 
at the point, referred to as the "focal spot", where the radiation 
emanates from the X-ray source. The X-ray source and the array of 
detectors are positioned so that the X-ray paths between the source and 
each detector all lie in the same plane (hereinafter the "slice plane" or 
"scanning plane") which is normal to the rotation axis of the disk. Since 
the X-ray paths originate from what is substantially a point source and 
extend at different angles to the detectors, the X-ray paths resemble a 
fan, and thus the term "fan beam" is frequently used to describe all of 
the X-ray paths at any one instant of time. The X-rays incident on a 
single detector at a measuring instant during a scan are commonly referred 
to as a "ray", and each detector generates an output signal indicative of 
the intensity of its corresponding ray. Since each ray is partially 
attenuated by all the mass in its path, the output signal generated by 
each detector is representative of the density of all the mass disposed 
between that detector and the X-ray source (i.e., the density of the mass 
lying in the detector's corresponding ray path). 
The output signals generated by the X-ray detectors are normally processed 
by a signal processing portion of the CT system. The signal processing 
portion generally includes a data acquisition system (DAS) which filters 
the output signals generated by the X-ray detectors to improve their 
signal-to-noise ratio. The filtered output signals generated by the DAS 
are commonly referred to as "raw data signals". The signal processing 
portion usually includes a projection filter which logarithmically 
processes the raw data signals to generate a set of projection data 
signals so that each projection data signal is representative of the 
density of the mass lying in a corresponding ray path. The collection of 
all the projection data signals at a measuring instant or interval is 
commonly referred to as a "projection" or a "view". During a single scan, 
as the disk rotates, a plurality of projections are generated such that 
each projection is generated at a different angular position of the disk. 
The angular orientation of the disk corresponding to a particular 
projection is referred to as the "projection angle". 
Using well known algorithms, such as the Radon algorithm, a CT image may be 
generated from all the projection data signals collected at each of the 
projection angles. A CT image is representative of the density of a two 
dimensional "slice", along the scanning plane, of the object being 
scanned. The process of generating a CT image from the projection data 
signals is commonly referred to as "filtered back projection" or 
"reconstruction", since the CT image may be thought of as being 
reconstructed from the projection data. The signal processing portion 
normally includes a back-projector for generating the reconstructed CT 
images from the projection data signals. 
One problem with CT systems is that a variety of noise and error sources 
may potentially contribute noise or artifacts to the reconstructed CT 
images. CT systems therefore typically employ a host of signal processing 
techniques to improve the signal-to-noise ratio and to reduce the presence 
of artifacts in the reconstructed CT images. 
One important type of noise in CT systems is manifested in the form of 
"streak like" artifacts, commonly known as "streaks", in the reconstructed 
CT images. FIG. 1 is an exemplary reconstructed CT image of a human head 
that is illustrative of the problems associated with streaks. In FIG. 1 
the white areas represent bone and the grey areas represent soft tissue. 
As those skilled in the art will appreciate, the soft tissue regions in 
FIG. 1 contain many streaks which interfere with interpretation of the 
image. 
One important factor which can give rise to streaks is aliasing caused by 
the finite size and spacing of the detectors in the detector array. 
Interfaces between bone and soft tissue of the patient, referred to as 
"bone-tissue interfaces", generate high frequency components in the 
projection data signals which are often under sampled due to the finite 
size and spacing of the detectors. Such under sampling gives rise to 
streaks in the reconstructed CT images. Streaks may also be caused by 
other factors such as movement of the patient or the gantry during a scan, 
or by the presence of metal implants or other high density prostheses in 
the patient. 
FIGS. 2A-B illustrate why bone-tissue interfaces generate high frequency 
components in the projection data. As those skilled in the art will 
appreciate, FIGS. 2A-B are not drawn to scale and are presented merely for 
illustrative purposes. FIG. 2A shows the spatial relationship between a 
cross section of a patient 50, an X-ray source 42, and a portion of a 
detector array 44 for a single projection angle. The cross section of 
patient 50 is disposed between source 42 and detector array 44 and 
contains a region of soft tissue 50:A and a region of bone 50:B. The 
detector array 44 shown includes seven individual detectors 44:1-44:7, and 
source 42 emits a fan beam 52 such that a ray 52:1 is incident on detector 
44:1, ray 52:2 is incident on detector 44:2, and so on. The output signals 
DET generated by the detector array 44 are filtered by DAS 45 which 
generates the corresponding raw data signals RDS. The DAS 45 as shown 
includes seven individual units 45:1-45:7, each unit corresponding to one 
detector. The raw data signals generated by DAS 45 are then filtered by an 
array 47 of projection filters which generate the projection data signals 
PDS. The array 47 as shown includes seven individual projection filters 
47:1-47:7, each filter corresponding to one detector. 
FIG. 2B is a graph of the amplitude of the projection data signals PDS 
generated by array 47. Since detectors 44:1, 44:2, 44:6 and 44:7 have only 
soft tissue (and air) disposed in their respective ray paths, the 
amplitude of the projection data signals generated by the corresponding 
projection filters 47:1, 47:2, 47:6 and 47:7 is relatively small. Since 
detectors 44:3, 44:4 and 44:5 have some bone disposed in their respective 
ray paths, and since bone is much denser than soft tissue, the amplitude 
of the projection data signals generated by the corresponding projection 
filters 47:3, 47:4 and 47:5 is relatively large. So the transition from 
soft tissue to bone (i.e., the bone-tissue interface) is represented in 
the projection data as a sharp "edge" or discontinuity, the edge being 
located where the amplitude changes rapidly (the edges being located in 
FIG. 2B at the transition between detectors 44:2 and 44:3, and at the 
transition between detectors 44:5 and 44:6). Such sharp edges represent 
high frequency components in the projection data. 
As shown in FIG. 2A, detector 44:3 straddles a bone-tissue interface (as 
does detector 44:5) meaning that a portion of the bone near the interface 
as well as a portion of the soft tissue near the interface are disposed in 
the ray path 52:3. In this position the intensity of ray 52:3 which is 
incident on detector 44:3 is representative of an average of the bone and 
tissue densities. Therefore, by straddling the bone-tissue interface, 
detector 44:3 "blurs", or under samples, the location of the interface. 
During a scan a single detector will typically straddle a bone-tissue 
interface for several projection angles before the disk rotates 
sufficiently to place a new detector under the interface. This new 
detector will then straddle the interface for several more projection 
angles. Having detectors straddle an interface in this manner for several 
projection angles makes it difficult for the scanner to precisely locate 
the interface and gives rise to streaks in the reconstructed CT images. 
One prior an method of reducing streaks is to decrease the size of the 
detectors and pack the detectors more closely together. Although simple 
and effective, this method raises the cost of the system because more 
detectors are required, and the manufacturing requirements for such small 
size detectors may exceed the limit of current technology. 
Another method of reducing streaks is to apply a linear low pass filter to 
the projection data and thereby remove the high frequency components that 
later give rise to streaks. Such a low pass filter is often incorporated 
into a convolution filter which ordinarily convolves the projection data 
signals with known convolution masks prior to applying these signals to 
the back-projector. The convolution filter may be thought of as being part 
of the back-projector. The low pass filter is also sometimes implemented 
by using a physical pre-filter, which averages or blurs the data, between 
the detectors and the X-ray source. One common and practical approach to 
achieving such a physical prefilter is to enlarge or oscillate the X-ray 
focal spot. Such methods do reduce streaks, however, they also have the 
disadvantage of removing high frequency components which are rich in 
information and thereby reduce the quality of the generated CT images. in 
general, such linear filters are not capable of sufficiently filtering the 
data to reduce streaks without also having the unwanted effect of reducing 
the quality of the CT images. 
There is therefore a need for improved methods of and apparatus for 
reducing streaks in CT images. 
OBJECTS OF THE INVENTION 
It is an object of the present invention to substantially reduce or 
overcome the above-identified problems of the prior art. 
Another object of the present invention is to provide an improved streak 
suppression filter. 
And another object of the present invention is to provide an improved 
non-linear streak suppression filter for reducing streaks in CT images. 
Yet another object of the invention is to provide an improved streak 
suppression filter for suppressing high amplitude, high frequency 
components of its input signals. 
And another object of the present invention is to provide an improved 
streak suppression filter including a high pass filter and a non-linear 
filter for filtering the output signals generated by the high pass filter. 
Still another object of the present invention is to provide an improved 
streak suppression filter including a high pass filter and a threshold 
device for clipping the output of the high pass filter. 
SUMMARY OF THE INVENTION 
These and other objects are provided by an improved streak suppression 
filter for use with a CT system. The CT system generates a plurality of 
projection data signals each of which is representative of the density of 
a portion of an object, and the CT system includes an image generating 
device for generating an image of the object from a plurality of 
projection signals. The streak suppression filter includes a spatial 
filter for receiving the projection data signals and for generating 
therefrom a plurality of low frequency signals and a plurality of high 
frequency signals. The streak suppression filter further includes a 
non-linear filter for filtering the high frequency signals to generate a 
plurality of filtered signals, and a device for combining corresponding 
low frequency and filtered signals to generate a plurality of streak 
corrected signals. The streak suppression filter then applies the streak 
corrected signals to the image generating device which generates therefrom 
a CT image of the object having reduced streaks. 
In one aspect, the streak suppression filter forms the low frequency 
signals by low pass filtering the projection data signals and forms the 
high frequency signals by subtracting the low frequency signals from their 
corresponding projection data signals. 
In another aspect, the non-linear filter includes a threshold device for 
generating the filtered signals by clipping the high frequency signals 
when they are greater than a threshold. 
Still other objects and advantages of the present invention will become 
readily apparent to those skilled in the art from the following derailed 
description wherein several embodiments are shown and described, simply by 
way of illustration of the best mode of the invention. As will be 
realized, the invention is capable of other and different embodiments, and 
its several details are capable of modifications in various respects, all 
without departing from the invention. Accordingly, the drawings and 
description are to be regarded as illustrative in nature, and not in a 
restrictive or limiting sense, with the scope of the application being 
indicated in the claims.

DETAILED DESCRIPTION OF THE DRAWINGS 
FIG. 3 shows an exemplary CT system, or scanner, 40 incorporating the 
principles of the present invention. Scanner 40 includes an X-ray source 
42 and a detector assembly 44 comprising an array of detectors mounted to 
a disk 46. Source 42 and detector assembly 44 are rotated about a rotation 
axis 48 (extending normal to the view shown in FIG. 3) so as to rotate 
around the object 50 that extends through the central opening of the disk 
46 during a CT scan. Object 50 may be a part of a live human patient, such 
as the head or torso. Source 42 emits within the scanning plane (normal to 
rotation axis 48) a continuous fan-shaped beam 52 of X-rays, which are 
sensed by the detectors of assembly 44 after passing through object 50. An 
array of anti-scatter plates 54 is preferably located between object 50 
and the detectors of assembly 44 to substantially prevent scattered rays 
from being sensed by the detectors. In a preferred embodiment the 
detectors number 384 and cover an arc of 48.degree., although the number 
and angle can vary. Disk 46, which may advantageously be of a light weight 
material, such as aluminum, is caused to rotate rapidly and smoothly 
around axis 48. The disk 46 is of an open frame construction so that 
object 50 can be positioned through the opening of the disk. Object 50 may 
be supported, for example, on a table 56, which is preferably as 
transparent as is practical to X-rays. 
The output signals generated by the detector assembly 44 are applied to DAS 
45 (shown in block diagram form) which generates therefrom the raw data 
signals. The raw data signals are applied to an array 47 of projection 
filters which generate the projection data signals. As disk 46 rotates, 
the projection data signals are used to provide projections from many 
projection angles. The projection data signals are applied to a streak 
suppression filter 70 which filters the projection data signals in 
accordance with the invention in a manner that reduces streaks in the 
reconstructed CT images. The output signals generated by streak 
suppression filter 70, referred to as "streak corrected projection data 
signals" or simply as "streak corrected signals", are then applied to a 
back-projector 72 which generates the CT images from the streak corrected 
signals. The back-projector 72 includes a convolution filter in the input 
stage to convolve the data for back projection. 
As will be discussed further below, streak suppression filter 70 is 
preferably a non-linear filter, and by using the streak corrected signals 
generated by streak suppression filter 70 rather than the projection data 
signals generated by projection filter 47, scanner 40 is able to generate 
improved, better quality CT images having fewer streaks and improved 
clarity. 
FIG. 4 is a block diagram of the signal processing portion of CT scanner 40 
showing streak suppression filter 70 in detail. Scanner 40 is an N channel 
device and detector array 44 includes N detectors 44:1-44:N. As stated 
above, in the preferred embodiment there are 384 detectors in array 44, so 
in the preferred embodiment N is equal to 384, however, other numbers of 
channels are of course possible. The N detectors in array 44 generate N 
detector output signals, DET:1-DET:N at every projection angle which are 
applied to DAS 45. DAS 45 filters the detector output signals and 
generates N corresponding raw data signals, RDS:1-RDS:N. The N raw data 
signals are applied to projection filter 47 which generates N 
corresponding projection data signals, PDS:1-PDS:N at every projection 
angle. The N projection data signals are applied to streak suppression 
filter 70 which generates N streak corrected signals, SCS:1-SCS:N for 
every projection angle. The N streak corrected signals from all of the 
projection angles are applied to back-projector 72 which generates the 
reconstructed CT images. 
Streak suppression filter 70 performs essentially the same functions for 
the data collected at each of the projection angles, and in general only 
the processing for the data collected at one of the projection angles will 
be discussed. Streak suppression filter 70 includes a spatial filter 110, 
a set of N threshold devices 120:1-120:N and a corresponding set of N 
adders 130:1-130:N. The N projection data signals PDS:1-PDS:N generated by 
projection filter 47 are applied to spatial filter 110 and in response 
spatial filter 110 generates N high frequency signals, PDSHF:1-PDSHF:N, 
and N low frequency signals PDSLF:1-PDSLF:N. The N high frequency signals 
PDSHF:1-PDSHF:N are applied to the N threshold devices 120:1-120:N which 
generate therefrom a set of N output signals referred to as "filtered 
signals" or "clipped signals". Each clipped signal and its corresponding 
low frequency signal are applied to the inputs of one of the N adders 
130:1-130:N which generates therefrom a corresponding streak corrected 
signal, so that the N adders 130:1-130:N generate the N streak corrected 
signals SCS:1-SCS:N. 
The data in each channel includes one detector output signal, one raw data 
signal, one projection data signal, one high frequency signal, one low 
frequency signal, and one streak corrected signal. So for example, the 
data in the third channel includes the third channel detector output 
signal DET:3 generated by the third channel detector 44:3, the third 
channel raw data signal RDS:3, the third channel projection data signal 
PDS:3, the third channel high and low frequency signals PDSHF:3, PDSLF:3, 
and the third channel streak corrected signal SCS:3. 
For purposes of generating the high and low frequency signals, each channel 
in scanner 40 is preferably associated with a neighborhood of channels. 
Spatial filter 110 preferably generates each channel's high frequency 
signal by performing a spatial high pass filtering operation on the 
projection data signals in that channel's neighborhood, and similarly, 
spatial filter 110 preferably generates each channel's low frequency 
signal by performing a spatial low pass filtering operation on the 
projection data signals in that channel's neighborhood. Each neighborhood 
preferably includes a group of adjacent channels, and one preferred 
neighborhood size is three, meaning that each neighborhood is formed by 
grouping a central channel with the two channels on either side of that 
central channel. For example, the third channel may be grouped with the 
second and fourth channel to form a neighborhood of three channels, and 
the spatial filter 110 then generates the high frequency signal in the 
third channel PDSHF:3 by combining the projection data signals in the 
second, third and fourth channels PDS:2, PDS:3, PDS:4 using an appropriate 
high pass filtering convolution mask. Similarly, spatial filter 110 
generates the low frequency signal in the third channel PDSLF:3 by 
combining the projection data signals in the second, third, and fourth 
channels PDS:2, PDS:3, PDS:4 using an appropriate low pass filtering 
convolution mask. Another preferred neighborhood size is five, and as 
those skilled in the art will appreciate, other neighborhood sizes will 
also function well with the invention. Further, the low and high pass 
filtering operations performed by spatial filter 110 preferably direct all 
of the energy of each projection data signal into either the corresponding 
low frequency signal or the corresponding high frequency signal such that 
a subsequent combination of corresponding high and low frequency signals 
will exactly reproduce the original projection data signal. 
Each of the N threshold devices 120:1-120:N performs a clipping operation 
on its input signal, i.e., each threshold device compares its input high 
frequency signal to a threshold, and generates a clipped signal that is 
equal to the high frequency signal when the magnitude of the high 
frequency signal is less than that of the threshold and is equal to the 
threshold when the high frequency signal is greater than the threshold and 
is equal to negative one times the threshold when the high frequency 
signal is less than negative one times the threshold. Alternatively, as 
will be discussed in greater detail below, each of the threshold devices 
may generate a clipped signal that is equal to a preselected value that is 
not equal to the threshold when the high frequency signal is greater than 
the threshold and is equal to negative one times the preselected value 
when the high frequency signal is less than negative one times the 
threshold. The values of the threshold and the preselected value may be 
permanently programmed into the threshold devices or may be selected by an 
operator. Each of the N adders 130:1-130:N then sums the clipped signal 
and the low frequency signal in its channel to generate that channel's 
streak corrected signal. 
FIG. 5 is a block diagram illustrating one preferred embodiment of streak 
suppression filter 70 in which spatial filter 110 is implemented using a 
set of N three-point, spatial low pass filters 210:1-210:N and a 
corresponding set of N subtractors 220:1-220:N. In each channel, the 
spatial low pass filter has three inputs coupled to receive the projection 
data signals from that channel's neighborhood and generates therefrom that 
channel's low frequency signal. The low frequency signal in each channel 
is applied to the negative input of that channel's subtractor and that 
channel's projection data signal is applied to the positive input of that 
channel's subtractor. In each channel, the subtractor then subtracts that 
channel's low frequency signal from that channel's projection data signal 
and thereby generates that channel's high frequency signal which is 
applied to that channel's threshold device. For example, the low frequency 
signal in the third channel PDSLF:3 generated by low pass filter 210:3 is 
applied to the negative input of the third channel's subtractor 220:3 and 
the projection data signal in the third channel PDS:3 is applied to the 
positive input of the third channel's subtractor 220:3. Subtractor 220:3 
subtracts the signal at its negative input from the signal at its positive 
input to generate the third channel's high frequency signal PDSHF:3 and 
applies this signal to the third channel's threshold device 120:3. Since 
the high frequency signal in each channel is generated simply by 
subtracting that channel's low frequency signal from that channel's 
projection data signal, a subsequent addition of a single channel's low 
and high frequency signals will reproduce exactly that channel's original 
projection data signal. This is preferable so that when a high frequency 
signal is below the threshold, the streak corrected signal is exactly 
equal to the original projection data signal and all high frequency 
information is preserved for use in reconstructing a CT image. In this 
embodiment, each streak corrected signal is equal to its corresponding 
original projection data signal when the amplitude of the high frequency 
signal is less than the threshold. However, when the amplitude of the high 
frequency signal exceeds the threshold, the streak corrected signal is 
essentially generated by removing some of the high frequency energy from 
its corresponding original projection data signal. 
In one preferred embodiment, the spatial low pass filters generate the low 
frequency signals according to Equation (1) 
EQU PDSLF:x=0.25PDS:(x-1)+0.50PDS:x+0.25PDS:(x+1) (1) 
in which x is an integer that may range from 2 to (N-1). As a specific 
example, spatial low pass filter 210:4 in the fourth channel generates the 
low frequency signal PDSLF:4 for the fourth channel by summing 0.25 times 
the projection data signal from the third channel PDS:3, 0.50 times the 
projection data signal from the fourth channel PDS:4, and 0.25 times the 
projection data signal from the fifth channel PDS:5. As those skilled in 
the art will appreciate, Equation (1) is equivalent to performing a 
convolution on the projection data signals with the three point 
convolution mask {0.25, 0.50, 0.25}. In this embodiment, the low pass 
filters are three point filters and use a neighborhood of three detectors. 
In another preferred embodiment, the low pass filters are five point 
filters and use a neighborhood of five detectors and use the mask {0.10, 
0.25, 0.30, 0.25, 0.10}. These masks are "averaging" or "blurring" masks, 
and are given by way of example. Other averaging masks and other 
neighborhood sizes will also function well with the invention. Further, 
the spatial low pass filters may be implemented using other finite impulse 
response (FIR) filters, infinite impulse response (IIR) filters, recursive 
or non-recursive filters, and by frequency domain filters using for 
example a Fourier transform technique. 
Spatial low pass filter 210:1 calculates the low frequency signal PDSLF:1 
for the first channel which is positioned at one end of the array of 
channels. Since this channel does not have adjacent channels on both 
sides, it is not possible for filter 210:1 to generate its low frequency 
output signal according to Equation (1), and as shown in FIG. 5, filter 
210:1 is only coupled to receive two projection data signals. This problem 
of computing neighborhood type processing near the end of an array is well 
understood in the image processing arts, and filter 210:1 may use any 
number of known methods for calculating its low frequency output signal 
PDSLF:1. For example, spatial low pass filter 210:1 may couple its third 
input (shown unconnected in FIG. 5) to a reference (e.g., ground) so that 
the third input is always interpreted as a zero and then convolve the 
three inputs with a scaled up version of the convolution mask used by the 
other filters. Alternatively, filter 210:1 may couple both the second and 
the third inputs to the second channel projection data signal PDS:2. Or, 
filter 210:1 may use a different convolution mask than the other filters. 
Similar options are of course available for filter 210:N located at the 
other end of the array. 
Streak suppression filter 70 has been discussed in connection with FIG. 5 
as including an array of low pass filters. As those skilled in the art 
will appreciate, steak suppression filter 70 may also be implemented using 
a set of high pass filters rather than low pass filters. In this 
embodiment, the high frequency signals are generated by the high pass 
filters, and the low frequency signals are generated by subtracting the 
high frequency signals from the corresponding projection data signals. 
Further, streak suppression filter 70 may be implemented using a set of 
high pass filters and a set of low pass filters for generating the high 
and low frequency signals, respectively. 
Streaks in reconstructed CT images are essentially caused by high 
amplitude, high frequency, components in the projection data signals. 
Therefore, CT images reconstructed from the streak corrected signals 
generated by streak suppression filter 70 (which have reduced high 
amplitude, high frequency components) contain fewer streaks. Further, the 
streaks that do remain in such CT images are of reduced amplitude. Streak 
suppression filter 70 therefore makes possible the reconstruction of clear 
CT images having increased utility. 
Preferably, the amplitude of the threshold is set so that the threshold 
devices remove only that portion of the high frequency signal which would 
generate a streak in the reconstructed CT image. If the threshold is set 
too low and the threshold devices thereby remove too much of the high 
frequency signal, the streak corrected signals will essentially contain 
only low frequency information. CT images reconstructed from such signals 
may not have any streaks but they will also be of poorer quality since all 
the high frequency information, which makes high resolution imaging 
possible, will have been removed from the streak corrected signals. 
Conversely, if the threshold is set too high, the quality of the resultant 
reconstructed CT images may be degraded by streaks. Since the regions of 
interest in a CT image are normally the soft tissue regions and not the 
regions containing bones, the threshold is preferably set high enough to 
preserve the integrity of the image information relating to the soft 
tissue regions and low enough to remove the high amplitude, high 
frequency, components generated by bone-tissue interfaces. Setting the 
threshold in this manner may degrade the resolution of the resultant CT 
images in the regions near bone-tissue interfaces, however since these 
regions are rarely of interest, it is acceptable to sacrifice some 
resolution In these regions to obtain high resolution streak free, or 
reduced streak, images in the soft tissue regions. The threshold is 
therefore preferably set higher than the amplitude of typical high 
frequency signals generated from regions free of bone-tissue interfaces 
and slightly lower than the amplitude of typical high frequency signals 
generated by bone-tissue interfaces. 
One method of tuning streak suppression filter 70 is to initially set the 
threshold to zero and generate some reconstructed CT images. These images 
will be derived only from the low frequency signals and therefore will be 
poorer quality. By viewing these images an operator may select an 
appropriate low pass filtering function (i.e., a convolution mask) for use 
in generating the low frequency signals. Preferably, a low pass filtering 
function is chosen that filters (or blurs) the raw data as little as is 
required to provide the desired amount of streak suppression. Once a low 
pass filtering function is chosen, the operator slowly increases the 
threshold from zero until streaks begin to appear in the reconstructed 
images. Since soft tissue regions contain high frequency components, 
setting the threshold too low will cause loss of information in soft 
tissue regions. Therefore, the threshold should be set as high as possible 
while maintaining the desired degree of streak suppression. Even when 
optimally tuned, streak suppression filter 70 may introduce small 
intensity errors into the reconstructed CT images. However, these errors 
are much smaller than the streaks and they are hardly visible to the human 
eye. The streak suppression achieved by filter 70 justifies the toleration 
of such errors. 
FIG. 6 is a CT image reconstructed using the same data that was used to 
reconstruct the CT image shown in FIG. 1. However, in FIG. 6, the 
projection data signals were first processed using streak suppression 
filter 70. For generation of this image, streak suppression filter 70 used 
a mask of {0.10, 0.25, 0.30, 0.25, 0.10} to generate the low frequency 
signals and a threshold of 0.006 while the maximum projection data signal 
was about 4.8. The image shown in FIG. 6 has far fewer streaks than the 
image shown in FIG. 1 and has increased clarity. 
Streak suppression filter 70 has been discussed in connection with using a 
threshold device 120 (shown in FIGS. 4-5) to clip the high amplitude 
portions of the high frequency signals. FIG. 7A is a graph of the transfer 
function of the threshold devices 120. The output signal generated by a 
threshold device 120 is equal to the input signal that is applied to the 
threshold device when the magnitude of the input signal is less than that 
of the threshold, and the output signal is limited to the threshold when 
the magnitude of the input signal is greater than that of the threshold. 
The threshold device 120 may be understood as generating an output signal 
that is a function of its input signal, a first threshold, and a second 
threshold so that the output signal is equal to the first threshold when 
the input signal is less than the first threshold, the output signal is 
equal to the input signal when the input signal is greater than the first 
threshold and less than the second threshold, and the output signal is 
equal to the second threshold when the input signal is greater than the 
second threshold. In the transfer function shown in FIG. 7A, the first and 
second thresholds are equal in magnitude and opposite in polarity, 
however, those skilled in the art will appreciate that in other 
embodiments the first and second thresholds need not be related in this 
manner. 
In still other embodiments, the threshold devices 120 may alternatively be 
implemented using filters having transfer functions that are different 
than the one shown in FIG. 7A. FIGS. 7B-E are examples of other transfer 
functions which will also function well with the invention. A preferred 
transfer function is shown in FIG. 7B, and when a threshold device 120 
uses this transfer function it generates an output signal as a function of 
its input signal, a positive threshold, a negative threshold, a positive 
preselected value PRESET, and a negative preselected value -PRESET, so 
that the output signal is equal to the negative preselected value -PRESET 
when the input signal is less than the negative threshold, the output 
signal is equal to the input signal when the input signal is greater than 
the negative threshold and less than the positive threshold, and the 
output signal is equal to the positive preselected value PRESET when the 
input signal is greater than the positive threshold. Preferably, the 
preselected value PRESET has a smaller magnitude than that of the 
threshold, and more preferably, the constant PRESET is set equal to zero. 
If the transfer function shown in FIG. 7C is used, then the devices 120 
generate an output signal that is equal to the input signal applied to the 
device 120 when the magnitude of the input signal is less than that of the 
threshold, and the output signal is equal to a linearly compressed version 
of the input signal when the magnitude of the input signal is greater than 
that of the threshold. Similarly, if the transfer function shown in FIG. 
7D is used, then the output signal generated by devices 120 is equal to a 
non-linearly compressed version of the input signal when the magnitude of 
the input signal is greater than that of the threshold. If the transfer 
function shown in FIG. 7E is used, then devices 120 apply some 
amplification to the input signal when the magnitude of the input signal 
is less than that of the threshold, and apply some compression to the 
input signal when the magnitude of the input signal is greater than that 
of the threshold. As those skilled in the art will appreciate, the 
transfer functions shown in FIGS. 7A-E are merely exemplary, and the 
invention will function well with any device 120 which compresses the 
streak forming, high amplitude, high frequency components of the 
projection data signals and which applies little or no filtering to the 
non-streak forming, low amplitude, high frequency components of the 
projection data signals. Such filters which selectively apply a relatively 
large degree of filtering to one portion of an input signal (i.e., high 
amplitude, high frequency components) and which selectively apply a 
relatively small degree of filtering to another portion of the input 
signal (i.e., low amplitude, high frequency components) are generally 
non-linear, and therefore, streak suppression filter 70 is preferably 
implemented as a non-linear filter. 
The invention has also been discussed in terms of applying a filter, such 
as threshold devices 120 to the high frequency signals. As those skilled 
in the art will appreciate, in other configurations the filter may also 
operate in an equivalent fashion directly on the projection data signals, 
rather than on low and high frequency signals, to generate the streak 
corrected signals. Also, streak suppression filter 70 has been discussed 
in connection with generating the high and low frequency signals so that 
when the high frequency signal is less than the threshold an addition of 
one channel's high and low frequency signals will exactly reproduce that 
channel's projection data signal. While it is preferred to generate the 
high and low frequency signals in this manner, those skilled in the art 
will appreciate that the invention will also function well if the high and 
low frequency signals are not generated in this manner. 
Streak suppression filter 70 has been discussed in connection with use in a 
CT system for generating the streak corrected signals from the projection 
data signals supplied from projection filter 47. However, CT systems 
constructed according to the invention may apply other types of filtering 
to the projection data signals prior to reconstructing a CT image. FIG. 8 
is a block diagram of the signal processing portion of a preferred CT 
system 900 constructed according to the invention. System 900, in addition 
to streak suppression filter 70, includes a parallel beam converter 910, 
and an interpolation filter 920. In system 900, the projection data 
signals generated by projection filter 47 are applied to parallel beam 
converter 910 which generates therefrom a set of parallel beam signals. 
The parallel beam signals are applied to streak suppression filter 70 
which generates therefrom the streak corrected signals. The streak 
corrected signals are then applied to interpolation filter 920, the output 
of which is applied to back-projector 72 which generates therefrom the 
reconstructed CT images. 
Parallel beam converter 910 includes a reorder converter 916 and an 
interleave converter 918. Reorder converter 916 receives the projection 
data signals from projection filter 47 and generates therefrom a set of 
reordered signals. The reordered signals are applied to interleave 
converter 918 which generates therefrom the parallel beam signals. The 
projections generated by projection filter 47 may be thought of as "fan 
beam" data, since all the projections are generated using fan beam 52 
(shown in FIG. 3). Parallel beam converter 910 re-organizes the 
projections to form parallel beam projections. 
The measurements of the projection data signals generated by projection 
filter 47 that are collected during a single scan (i.e., one rotation of 
the disk) may be organized in a matrix PDS as shown in Equation (2). 
##EQU1## 
Each element PDS(i,.theta.) in the PDS matrix represents a measurement of 
the projection data signal in the ith channel for a projection angle equal 
to .theta.. In Equation (2), N represents the number of channels in 
scanner 40. As stated above, in the preferred embodiment there are 384 
detectors in the array 44, so in the preferred embodiment there are 384 
channels in scanner 40 and N is equal to 384. .DELTA..theta. represents 
the mount of rotation of disk 46 between successive projections (i.e., the 
angular increment of the projection angle between successive projections). 
In the preferred embodiment, disk 46 rotates one eighth of a degree 
between each projection and scanner 40 generates 2880 projections in a 
single scan (i.e., eight projections per degree for 360 degrees), so in 
the preferred embodiment .DELTA..theta. is equal to 0.125 degrees. Each 
row of the PDS matrix represents all the measurements of the projection 
data signals collected at a single projection angle. In the preferred 
embodiment, there are 2880 rows in the PDS matrix. Each column of the PDS 
matrix represents all the measurements of one channel's projection data 
signal collected during one scan, and in the preferred embodiment there 
are 384 columns in the PDS matrix. The PDS matrix has cyclical nature in 
that the first row is the continuation of the last row, that is, PDS 
(i,0)=PDS (i,360). 
FIG. 9A illustrates a set of rays 1010 forming one portion of a single fan 
beam projection view of a cross section of patient 50. Since each of the 
rays emanates from X-ray source 42, which is essentially a point source, 
none of the rays 1010 are parallel, and the resulting projection is a fan 
beam projection. Each row of the PDS matrix corresponds to a single fan 
beam projection. Reorder converter 916 re-organizes the projection data 
signals so that each re-organized projection is formed by a set of 
parallel rays such as the rays 1020 shown in FIG. 9B. 
FIGS. 10A-B illustrate one preferred method, which may be used by reorder 
converter 916, for generating the re-organized projections. FIGS. 10A-B 
show the positions of X-ray source 42 and detector array 44 during 
generation of two successive projection views. During a scan, X-ray source 
42 and detector array 44 rotate in a counter clockwise direction about a 
center 1110 of circle 1120. During the first projection, shown in FIG. 
10A, a ray 1130 is incident on a detector 44:4 (i.e., the detector in the 
fourth channel of array 44). During the next projection, shown in FIG. 
10B, a ray 1132 is incident on detector 44:3 (i.e., the detector in the 
third channel of array 44). In the preferred embodiment, the spacing 
between the detectors is matched to the amount of rotation between 
generation of successive projections so that ray 1130 is parallel to, and 
slightly offset from, ray 1132. In the preferred embodiment, this basic 
relationship is true for all detectors so that any two rays incident on 
adjacent detectors during successive projections are parallel and are 
slightly offset from each other. As was stated above, in the preferred 
embodiment .DELTA..theta. is equal to 0.125 degrees, so in the preferred 
embodiment, each detector in array 44 is spaced apart from its adjacent 
detectors by 0.125 degrees. Reorder converter 916 uses this basic 
relationship to reorder the data and generate the re-organized 
projections. 
Reorder converter 916 preferably re-organizes the PDS matrix to form a 
matrix RE of reordered signals so that each row of the RE matrix is 
equivalent to a projection formed by a parallel beam. Reorder converter 
916 preferably generates the RE matrix so that each element RE(i,.theta.) 
of the RE matrix is chosen according to the formula shown in Equation (3). 
EQU RE(i,.theta.)=PDS(i,i-j!.DELTA..theta.!+.theta.) (3) 
where the jth channel is the channel nearest to the geometrical center of 
the detector array. Each element RE(i,.theta.) of the RE matrix represents 
a measurement of the reordered signal in the ith channel for a parallel 
beam projection angle of .theta.. Reorder converter 916 may also use a low 
pass filter to average the projections of adjacent angles for each 
channel. The averaged, or decimated, reordered matrix RE will have fewer 
numbers of rows at larger angular intervals .DELTA..theta.. Decimating the 
RE matrix in this fashion reduces the computations for subsequent 
operations. 
Interleave converter 918 (shown in FIG. 8) receives the reordered signals 
and generates therefrom the parallel beam signals. Interleave converter 
918 preferably combines pairs of parallel beam projections that are spaced 
apart by 180 degrees to form denser projections. FIGS. 11A and 11B 
illustrate the spatial relationship between X-ray source 42, a cross 
section of patient 50, and detector array 44 for projection angles of zero 
and 180 degrees, respectively. In FIGS. 11A-B, detector array 44 is shown 
containing seven detectors, and the detector in the fourth channel 44:4 is 
the central detector of the array 44. As was stated above, in the 
preferred embodiment, detector array 44 has 384 detectors, however, for 
convenience, the seven detector embodiment will now be discussed. In the 
preferred embodiment, detector array 44 is slightly offset from the center 
1210 of disk 46 such that a line 1234 intersecting the focal spot of 
source 42 and center 1210 does not intersect the center of the central 
detector 44:4. The arrangement of such a detector system is more fully 
described in U.S. patent application Ser. No. 08/191,428, entitled, X-RAY 
TOMOGRAPHY SYSTEM FOR AND METHOD OF IMPROVING THE QUALITY OF A SCANNED 
IMAGE, filed on Feb. 3, 1994, (Attorney Docket No. ANA-044) and assigned 
to the present assignee, which is hereby incorporated by reference. 
FIG. 12 illustrates the spatial relationship between detector array 44 at 
projection angles of zero and 180 degrees, and the rays 1310, 1312, 1314 
incident on three of the detectors. Because of the offset between array 44 
and the center 1210 of disk 46, the detector array 44 at a projection 
angle of zero degrees is slightly offset from the detector array 44 at 180 
degrees. Consequently, the ray 1310 that is incident on the sixth channel 
detector 44:6 for a projection angle of 180 degrees falls between the rays 
1312 and 1314 that are incident on detectors 44:2 and 44:3, respectively, 
for a projection angle of zero degrees. In this example, detector 44:6 may 
be thought of as a "central" detector and detectors 44:2 and 44:3 may be 
thought of as "opposite-adjacent" detectors. At each projection angle, 
each detector measures the density of a portion of the patient, and in 
general, the portions measured by the opposite-adjacent detectors are 
closer to the portion measured by the central detector than are the 
portions measured by any other detectors (e.g., the portions measured by 
detectors 44:2, 44:3 at a projection angle of zero degrees are closer to 
the portion measured by detector 44:6 at a projection angle of 180 degrees 
than are the portions measured by detectors 44:5, 44:7 at a projection 
angle of 180 degrees). Any two projections separated by 180 degrees may be 
interleaved using this relationship between central and opposite-adjacent 
detectors to form a single denser projection. For example, one such 
interleaved projection for the arrangement shown in FIG. 12 is composed of 
the quantities RE(1,0), RE(7,180), RE(2,0), RE(6,180), RE(3,0), 
RE(5,180), RE(4,0), RE(4,180), RE(5,0), RE(3,180), RE(6,0), RE(2,180), 
RE(7,0), RE(1,180)! in which RE(i,.theta.) is the reordered signal 
generated from the detector in the ith channel at a projection angle of 
.theta.. Interleave converter 918 interleaves the reordered signals in 
this manner to form denser projections. 
Interleave converter 918 preferably generates a matrix of measurements 
of the parallel beam signals, and each element (i,.theta.) of the 
matrix is a measurement of the parallel beam signal in the ith channel for 
a parallel beam projection angle equal to .theta.. The structure of the 
matrix is shown in Equation (4). 
##EQU2## 
As shown in Equation (4), the matrix has twice as many columns as the 
PDS matrix, and half as many rows. So each row of the matrix 
represents a parallel beam projection containing twice as much data as a 
row of the PDS matrix. So, each parallel beam projection may be thought of 
has having twice as many channels as a fan beam projection. Slightly 
different from the PDS matrix, the matrix has a cyclical property in 
which the last row continues into the first row in reverse order, that is, 
(0,180)= (2N-1,0); (1,180)= (2N-2,0) and so on. In the 
preferred embodiment, interleave converter 918 generates the elements of 
the matrix according to the formula shown in Equation (5). 
EQU (2i,.theta.)=RE(i,.theta.) 
EQU (2i+1,.theta.)=RE(N-1-i,.theta.+180) (5) 
EQU for 0&lt;i&lt;N 
As is well known, parallel beam converters such as converter 910 (shown in 
FIG. 8) for converting fan beam data to parallel beam data normally 
include an interpolation filter, such as filter 920. However, the 
interpolation filter is normally disposed immediately following the 
interleave converter 918. Since the detectors are generally spaced so that 
the angular offset between adjacent detectors, relative to the X-ray 
source, is equal for all detectors, the detectors are not spaced 
equidistantly in a linear sense. Therefore, each row of the matrix 
(i.e., each parallel beam projection) contains data points which are not 
spaced equidistantly. Rather, the elements near the middle of each 
projection are spaced further apart than are elements near the ends of 
each projection. The interpolation filter interpolates the data and 
generates a new matrix of parallel beam data such that all the elements of 
each projection are spaced equidistantly. In the preferred embodiment of 
system 900, the interpolation filter 920 uses known techniques to 
interpolate the data and generate projections containing equidistantly 
spaced elements, however, filter 920 is preferably disposed after streak 
suppression filter 70, rather than immediately following the interleave 
converter 918. However, the invention will also function well if the 
interpolation filter 920 is disposed immediately following the interleave 
converter 918 as is normally done in the prior art. 
Also, as is well known, converting fan beam data to parallel beam data 
generally introduces a slight rotation such that the parallel beam 
projection angle of zero degrees is not exactly coincident with the fan 
beam projection angle of zero degrees. If not corrected, this rotation 
results in generating a reconstructed image that is slightly rotated from 
the horizontal. This rotation is generally introduced because the center 
detector j as used in Equation (3) is generally not exactly in the center 
of the detector array. The amount of rotation is generally smaller than 
.DELTA..theta./2 and may be corrected using well known techniques by 
either the interpolation filter 920 or by the back-projector 72 or 
alternatively may simply be ignored. 
While streak suppression filter 70 may operate directly on the projection 
data signals generated by projection filter 47, the performance of streak 
suppression filter 70 may improve if the parallel beam signals generated 
by parallel beam converter 910 are applied to streak suppression filter 70 
rather than the projection data signals. 
When streak suppression filter 70 operates on the fan beam projection data 
generated by projection filter 47, those skilled in the art will 
appreciate that streak suppression filter 70 operates on one row of the 
PDS matrix at a time. For example, when streak suppression filter 70 is 
operating on the first row of the PDS matrix (i.e., the row for .theta. 
equal to zero degrees) to understand the processing of filter 70, the 
first element of the row PDS(0,0) may be thought of as being substituted 
for the previously described signal PDS:1, and the second element of the 
row PDS(1,0) may be thought of as being substituted for the previously 
described signal PDS:2, and so on. Similarly, when streak suppression 
filter 70 operates on the parallel beam signals generated by parallel beam 
converter 910, then filter 70 operates on one row of the matrix at a 
time. For example, when streak suppression filter 70 is operating on the 
first row of the matrix then the first element of the row (0,0) may 
be thought of as being substituted for the previously described signal 
PDS:1, and the second element of the matrix (1,0) may be thought of 
as being substituted for the previously described signal PDS:2, and so on. 
Since each row of the matrix has twice as many elements as a row of 
the PDS matrix, filter 70 preferably includes twice as many channels (or 
2N channels) when filter 70 operates on the parallel beam signals. 
The invention has been discussed in terms of suppressing streaks caused by 
undersampling of bone-soft tissue interfaces. However, as those skilled in 
the art will appreciate, streak suppression filter 70 may be used to 
suppress streaks that are caused by any type of irregularity or 
imperfection in high amplitude high frequency signals. Streak suppression 
filter 70 therefore effectively reduces streaks caused by motion of the 
patient and unwanted motion or vibration of the gantry during a scan, as 
well as streaks caused by the presence of high density fillings or 
implants in the patient. 
Streak suppression filter 70 has also been discussed in terms of being 
constructed from a group of components such as low pass filters, adders, 
and subtractors. As those skilled in the art will appreciate, the number 
of components used to implement streak suppression filter may be reduced 
by using multiplexing schemes. For example, with reference to FIG. 5, 
streak suppression filter 70 may be constructed using one low pass filter, 
one subtractor, one threshold device, one adder and two I:N multiplexers, 
rather than N of each component as shown. Further, streak suppression 
filter 70 may also be implemented by other means such as by a software 
program that is executed by a digital computer. 
Since certain changes may be made in the above apparatus without departing 
from the scope of the invention herein involved, it is intended that all 
matter contained in the above description or shown in the accompanying 
drawing shall be interpreted in an illustrative and not a limiting sense.