This invention pertains to a method for improving the dynamic range
in computed tomography systems which incorporates an image intensifier detector
to aid in the formation of better images.
Background of the Invention
In recent years, much interest has been evidenced in a field now
widely known as computed tomography. In a typical procedure utilizing computed
tomography (or CT), an X-ray source and detector are physically coupled together
on opposite sides of the portion of a sample which is to be examined. The sample
can be a patient or phantom or other objects, for example. X-rays are made to
transit through the sample to be examined, while the detector measures the X-rays
which make it through the sample without being absorbed or deflected. Periodically,
the paired source and detector are rotated to differing angular orientations about
the sample, and the data collection process repeated.
A very high number of measurements of attenuation values may be obtained
by procedures of this type. The relatively massive amounts of data thus accumulated
are processed by a computer, which typically does a mathematical data reduction
to obtain attenuation values for a very high number of transmission valves (typically
in the hundreds of thousands) within the section of the sample being scanned.
This data may then be combined to enable reconstruction of a matrix (visual or
otherwise) which constitutes an accurate depiction of the density function of
the sample section examined.
By considering one or more of such sections, skilled medical diagnosticians
may diagnose various body elements such as tumors, blood clots, cysts, hemorrhages
and various abnormalities, which heretofore were detectable, if at all, only by
much more cumbersome and, in many instances, more hazardous-to-the-patient techniques.
While apparatus of the aforementioned type have represented powerful
diagnostic tools, and were deemed great advances in the radiography art, apparatus
of the first generation suffered from many shortcomings. Acquisition of the raw
data frequently entailed an undesirably long period, which among other things
subjected a patient to both inconvenience and stress. The patient's inability to
remain rigid for such a lengthy period, also could lead to blurring of the image
sought to be obtained.
In U.S. Patent US-A-4,149,248, to John M. Pavkovich, entitled
Apparatus and Method for Reconstructing Data, and assigned to the same
assignee as is the present patent, apparatus and methodology are disclosed which
alleviate a number of the prior art problems, most notably the lengthy period
that was previously required for computer processing of the raw data provided by
the detectors. The apparatus disclosed therein utilizes a fan beam source of radiation
coupled with application of a convolution method of data reduction, with no intervening
reordering of fan rays, to thereby eliminate the errors and delays in computation
time which would otherwise be involved in such reordering. The radiation source
and the detector means are positioned on opposite sides of the portion of the
patient to be examined and these elements are made to rotate through a revolution
or portion thereof about the patient. During such rotation, the detectors measure
the radiation absorption at the plurality of transmission paths defined during
the rotational process. In order to measure analog signals over a wide dynamic
range, application-specific conversion schemes are usually employed. That is, when
the signal-to-noise ratio of the input transducer exceeds that of the analog-to-digital
converter, then input signal preprocessing is typically used to compress the input
X-ray CT scanners are now a common tool for the diagnostic radiologist.
Typically these are expensive, i.e. greater than a million dollars. These systems
typically have scan times of 1 to 2 seconds with 0.3 mm spatial resolution. Density
resolution as low as 0.25%, with degraded spatial resolution, is achievable. The
technology of generator/detector design and the improvements in the microcomputer
area over the past 10 years have enabled image detection and processing to approach
Radiation Therapy Simulators
Radiation therapists often attempt to use scans from diagnostic CT
scanners in planning a radiation therapy treatment. Because high levels of radiation
are used during radiation therapy treatment it is important that the therapists
be able to precisely locate the site to be treated. However, the relative position
of organs within the body during a diagnostic CT scan are not the same as when
a patient is placed on a flat couch of the radiation therapy machine. This occurs
because the diagnostic CT scanner couch is more crescent shaped. Therefore, radiation
therapy simulators have come into use. These simulators have patient couches that
are identical to couches of radiation therapy machines. Also, in the simulator,
the X-ray focal spot for fluoroscopic/radiographic imaging is positioned to allow
the same target-to-patient isocenter as in the therapy machine. Beam shaping devices
and other accessories can be added which attempt to exactly duplicate the therapy
setup. Thus, simulators yield a projected planar image of the patient anatomy that
is much more geometrically compatible with the position of the radiation therapy
system. In addition to the properly oriented radiographic information, if cross-sectional
CT images could be obtained at the same time, then the therapist would be further
aided in planning the treatment.
A radiation therapy simulator is a diagnostic imaging X-ray machine
shaped to simulate the geometry of radiation therapy (or radiotherapy) treatment
units. Thus, a simulator includes an X-ray imaging source, a gantry to support
and position the X-ray imaging source, a couch to support the patient, and an
image forming system. The dimensions of the gantry are such that it positions the
x-ray imaging source relative to the couch in a geometry mathematically similar
to the geometry of the radiotherapy machine. Images formed on the imaging system
can then be interpreted in terms of the geometry of the radiotherapy machine. Images
can be taken from different angles to aid in the planning of how to form the radiotherapy
beam to maximize dose to the target and minimize damage to healthy organs.
In existing simulators, because the geometry of the simulator attempts
to very closely simulate that of the radiotherapy machine, the X-ray imaging source
and image forming system are limited to a configuration which is less than optimal
for the quality of the image. Both the source and the image-detector-part of the
image forming system are far from the patient. The image at the detector has been
recorded on film.
An image intensifier has been used to increase the brightness of
the image which can be used to produce a television image. A computer has been
used to process and enhance the television image.
Computed Tomography Simulators
In the prior art, it is known to form a computed tomography image
based on data obtained from a TV camera using an image intensifier tube (IIT) between
the patient and a television camera. The output signal from the television camera
is processed to form a digital signal which is further processed in a computer
to form a tomographic image. This prior art system employing the television camera
produces a noisy image of marginal value in simulation and planning.
Similar CT attempts using X-ray image intensifiers with video cameras
have been made in the past by various groups. However, from prior CT experience,
it is believed that the use of video camera signals based on data off the IIT was
one of the major limiting features in these designs. Compared to the IIT, conventional
video cameras have horizontal spatial resolution of 3-4 line pairs per mm over
a 30 cm field, but their intensity output is both limited and nonlinear. Typically,
tube video camera instantaneous signal dynamic range is limited to only two or
three orders of magnitude. Conventional solid state video cameras have good linearity
spatially and in intensity, but their signal dynamic range is also limited to about
1,000:1 at room temperature and with averaging lines possibly 4000:1.
In order to maintain X-ray photon statistics on a 16" (40 cm) diameter
body, a detector with a minimum signal to noise ratio (S/N) of at least 200,000:1
is necessary. This is assuming a typical surface dose of 2 rads/scan and no compensating
bolus around the patient. It is also necessary that the IIT, lens optics and photo
detector yield an X-ray to electron quantum efficiency of greater than unity.
Medical or industrial X-ray CT applications typically require a detector
system with millimeter spatial resolution and photon-limited intensity resolution.
The rate of photons emitted from an X-ray source is statistical and follows a Poisson
distribution. Thus, any ideal measurement of photon intensity has a root-mean-square
(rms) noise equal to the square root of the average number of photons detected.
Therefore, the detector system must have a total quantum detection efficiency (QDE)
greater than unity in order to maintain photon statistics. Also, since additional
random noise adds in quadrature, then the detector electronics must have a rms
input noise level below the photon noise.
US-A-4,647,975 discloses a method for improving the dynamic resolution
of an imaging system using a photo-detector array comprising a plurality of photodiodes
which are exposed to a visible light output of an image intensifier and cyclicly
sampled within successive sampling cycles with two different time intervals during
which photo-detector diodes are sampled for expanding the dynamic exposure range
thereof. The present invention improves on this arrangement according to claim
Examples of the invention will now be described with reference to
the accompanying drawings in which:
Fig. 1 shows a block diagram of a detector system.
Fig. 2 shows a diagram of the reconstruction diameter for a head
Fig. 3 shows a side view diagram of simulator geometry for a head
Fig. 4 shows a simplified perspective view of a radiation treatment
simulator in conjunction with which the detector of Fig. 1 can be used.
Fig. 5 is a block diagram which illustrates the elements (indicated
by a "*") which can be added to a radiation treatment simulator system to
obtain a CT simulator system in accordance with the present invention.
Fig. 6 shows a side view diagram of simulator geometry for a body
Fig. 7 shows a diagram of the reconstruction diameter for a body
Fig. 8 shows the addition of an imaging extension detector array
to the IIT which can be used to increase the scan circle diameter.
Fig. 9 provides a cross sectional view of a typical image intensifier
tube and associated optics.
FIG. 10 is a simplified schematic of the photodiode linear array,
preamplifier, and ADC of the present invention.
FIG. 11 provides a detailed schematic of the circuitry used to implement
one embodiment of the pre-amp, integrator, reset and clamp circuit.
FIG. 12 illustrates the relative positioning of the photodiodes of
the imaging extension detector array.
FIG. 13 illustrates a detector from the imaging extension detector
FIG. 14 illustrates the sensitivity profile along the detector face
of a detector from imaging extension detector array.
FIG. 15 is a detailed schematic of the sampling preamplifier circuitry
used with the imaging extension detector array.
FIG. 16 illustrates the relative timing involved in the data acquisition
from the photodiode linear array and the imaging extension detector array.
FIG. 17 shows a top view of the mirror alignment apparatus.
FIG. 18 shows a plot of detector response using the mirror alignment
apparatus of FIG. 17.
FIG. 19 shows the moving pin arrangement used in the center finding
procedure of the present invention.
FIG. 20 shows a plot of detector response in the center finding procedure
of the present invention.
FIG. 21 illustrates a table in which is tabulated the angle of the
occluded ray, and the detector which responded to the occluded ray.
FIG. 22 illustrates the manner in which the information tabulated
in FIG. 21 is used in connection with actual measurement data.
FIG. 23 shows the apparatus and response for determining the point
FIG. 24 illustrates the gathering and preparation of the point spread
FIG. 25 illustrates what the data might look like after the real
peaks have been zeroed out.
FIG. 26 shows the overall sequence of data collection and correction.
FIG. 27 illustrates a table tabulating photodiode number and coefficient
FIG. 28 shows a diagram of polynomial fitting of dose versus measured
FIG. 29 illustrates a procedure for determining the coefficients
for three 4th-order polynomials which correct for non-linearities in the photodiode
FIG. 30A illustrates the correction of the detector readings using
the three polynomials illustrated in FIG. 29.
FIG. 30B illustrates the correction of the detector readings according
to the preferred method of the present invention in which the results of each
polynomial are weighted, and the sum of the weighted polynomial then used as the
corrected detector reading.
FIG. 30C illustrates the weighing function used in FIG. 30B, for
one example of count ranges.
FIG. 31 illustrates the weight assigned to the data from each projection
for overlay projection corrections.
FIG. 32 illustrates the manner in which the E matrix, illustrated
in FIGS. 24 and 25, is used to implement the deconvolution which corrects for the
FIGS. 33A, 33B, and 33C illustrate how different positions of X-ray
source 30 cause different detectors in photodiode linear array to be affected
by a ray occluded by the calibration needle.
The following is a partial list of the lexicon used in the following
Automatic Brightness System
Analog to Digital Converter
Polynomial Coefficient for the ith polynomial and the jth order factor in the
Actual Detector Measurement
Actual Detector Measurement Corrected for Photodiode Non-linearities
Direct Memory Access
Focus to Axis (Isocenter) Distance
Focus to Intensifier Distance
Focal Spot to Access Distance
Image Intensifier Tube
Light Emitting Diode
Point Spread Function
Random Access Memory
Root Mean Square
Top Dead Center
Write Once Read Many
Description of the Preferred Embodiments
Referring to FIG. 1, in the simulator of the present invention an
object of interest is scanned to provide a series of projections at different selected
angles about a rotational axis of the object. The projections are formed by passing
a full or partial fan of radiation through the object of interest at each of these
selected angles. For each projection, the object-attenuated fan of radiation is
applied to an image intensifier tube which converts radiation photons to visible
light photons. The visible light photons are then detected using a photodiode linear
array. Signals from the photodiode linear array are conditioned and then converted
into digital form. The digital information is then processed under computer control
to correct the data for background noise, non-linearities in the image intensifier
tube and photodiode array, point spread in the imaging chain; and for other effects.
Thereafter, the corrected data from each of the projections is used to reconstruct
an image of the cross section of the object of interest for that scan. Multiple
scans can be taken to provide a three-dimensional view of the object of interest.
These scans can then be displayed on a monitor with variations in the image being
displayed representing different absorption coefficients or absorption densities.
The reconstructed image can also be stored in digital form for later viewing.
The use of a full or a partial fan-beam of x-ray radiation depends
upon the diameter of the object being scanned and the dimensions of the detector
electronics available. For example, where a 12" image intensifier tube is used,
and a scan of the head of a patient is desired, a full fan-beam is used. As such,
the centers of the fan-beam, the head of the patient, and the image intensifier
tube, are aligned along a common axis.
On the other hand, when the body of a patient is to be scanned, and
a 12" image intensifier tube is used, the diameter of the body is too large to
be fully contained within the 12" width of the tube. As such, a partial fan-beam
is used, with the image intensifier tube offset from the axis along which the
fan-beam and the patient centers lie.
Imaging Extension Embodiment
In a further embodiment of the present invention, an imaging extension
is provided by which the maximum patient scan circle diameter is increased, for
example, from 40 to 50cm when a 30cm (12") image intensifier tube is used. It allows
for an increase in patient to gantry clearance from 60cm to 77cm, and also allows
for a full 48cm wide patient couch to be used.
When a 12-inch image intensifier tube is used without the imaging
extension, the maximum size object which can be scanned is limited by the image
intensifier tube diameter and its distance from the x-ray source.
Full Fan Versus Partial Fan Beams
The 12-inch image intensifier tube active area varies from unit-to-unit,
but typically has been 28cm ±1cm across the face at 135cm from the x-ray source.
For a full-fan, head-scan centered at 100cm, as seen in FIG. 2, this geometry
limits the maximum size object to 21cm diameter. This covers approximately 95%
of the U.S. male population according to Diffrient's Humanscale publication. (See
N. Diffrient et al., Humanscale 11213 Manual, 1979, The MIT Press.) Most
diagnostic CT scanners have a 25cm maximum head scan circle which allows for approximately
100% population coverage and less critical patient positioning.
Both head and body scans can be increased in size by using an asymmetric
(or partial) x-ray fan-beam. The simulator head scan circle can be increased to
25cm by using an asymmetric verses full-fan approach. In this mode the image intensifier
tube is shifted a few cm off center and a 360 degree scan is performed. Thus,
a larger area is covered compared to the full-fan mode. Some loss in contrast
and spatial resolution results because the entire object is not viewed in each
The projection data is then reconstructed using a variation of the
Pavkovich fan-beam reconstruction method, described in copending patent application
entitled "Partial Fan-beam Tomographic Apparatus and Data Reconstruction Method",
filed even date herewith, and assigned to the assignee of the subject application.
In the case of body scans, the image intensifier tube is shifted
to the maximum extent. The largest body scan with the above arrangement is less
than 40cm diameter as depicted in FIG. 7. This covers approximately 95% of the
U.S. males across the chest, but less than 50% across the shoulders. The majority
of the diagnostic scanners have a 50cm maximum body scan which covers 97% of the
Radiation Treatment Simulator - Details
In one application, the present invention is used in conjunction
with a radiation treatment simulator and planning system.
Radiation treatment simulator and planning systems ("simulator")
simulate the geometry and movement of megavoltage radiation therapy equipment.
The following are the basic items into which the simulator system can be divided:
floor mounted drive unit with rotating arm, X-ray head and crosswire assembly,
detector including image intensifier, treatment couch, relay frame, and control
units. Many of the basic simulator system elements suitable for use with the present
invention can be found in the Ximatron CR Radiotherapy Simulator System, manufactured
by Varian, the assignee of the subject application.
Referring to FIG. 4, typically, the drive unit 10 comprises a welded
steel fabrication which is bolted on a plinth, which is preferably cast in the
floor, prior to the completion of the final floor finish. The drive structure
houses the variable speed electric drive unit and a high precision slewing ring
bearing on which is fitted the rotating arm 12. On the arm 12 are mounted the carriages
14 and 16 for the X-ray head assembly 18 and the image intensifier tube assembly
20, respectively. Attached to the front of the arm is a circular disk 22, the
circumference of which carries a scale mark in degrees from 0.0 to 360.0 degrees.
A screen wall (not shown) is supplied which carries the zero data mark for the
scale together with a small sub-scale for ease of reading if the zero data is visually
obstructed. The screen wall is built into a partition wall which seals off the
drive unit and control gear from the room, thus presenting a clean finish.
X-ray Head 18
Protruding through the top of the arm 12 is the X-ray head assembly
18 which is carried on a rigidly constructed steel fabrication. The X-ray system
on the simulator has a generator having an output of 125 kVp and 300 mA (radiographic
mode) or 125 kVp and 30 mA (fluoroscopic mode), in conjunction with a double focus
(0.6 mm and 1 mm) X-ray tube, with a permanent 2 mm element and filter. The X-ray
tube is mounted in a yoke on the end of the steel fabrication.
Mounted below the tube is a lead-bladed collimator which can be manually
set to give field sizes from 0 to 35 by 35 cm and 100 cm F.S.A.D. The collimator
also contains a lamp operated by a switch on the side of the housing, which will
define the area of the X-ray beam through the blades onto the patient's skin.
Mounted with and in front of the collimator is a crosswire assembly.
This is fitted with two pairs of motorized tungsten wires to give any square or
rectangular field from 4 x 4 cm to 30 x 30 cm at 100 cm F.S.A.D. Windows inside
of the crosswire housing have scales fitted indicating the field sizes at 100
cm. These are repeated on electrical indicators fitted in the remote control console.
The collimator and crosswire assembly have motorized and manual rotation over
the range ±45°. A suitable scale is provided for reading the angular position.
The complete head is capable of being electrically driven from its maximum F.S.A.D.
of 100 cm down to 60 cm.
Protruding through the front bottom of the arm 12 is the image intensifier
tube assembly 20. This unit is mounted on a double carriage to enable scanning
over an area of ±18 cm about the center of the X-ray beam both longitudinally and
laterally. The complete assembly is also capable of being electrically driven
from a maximum of 50 cm from the rotation axis to the image intensifier tube face
24, down to 10 cm. Anti-collision bars fitted to the image intensifier tube face
24 will, when operated, isolate the electrical supplies to the operating motors.
Provision is made to override the anti-collision interlocks in order to drive out
of the collision situation.
Treatment Couch 26
The couch assembly 26 includes a steel framework supported on a large
precision bearing ring. These are mounted in a pit cast in the floor. The framework
carries the telescopic ram assembly 28 for the couch 26 together with a circular
floor section. The bearing allows the couch 26 to be isocentered ±100° about the
X-ray beam, either electrically or manually. A scale is fitted around the edge
of the pit for positioning.
The telescopic ram assembly 28 provides vertical movement of the
couch top from a minimum height of 60 cm to a maximum of 120 cm.
Attached to the top of the telescopic ram is a sub-chassis which
provides for manually lateral movement. A manual brake is provided by levers on
either side of the couch, to lock the top in its set position. Fitted to this
sub-chassis is a side channel couch top having a width of 50cm and a length of
213cm. This is fitted for motorized longitudinal movement of 123cm and provision
for manual override to facilitate rapid setting. Manual rotation of the couch
top of the telescopic ram assembly is provided. The manual brake is provided to
lock the top to its desired position.
A cushion top is provided with three removable sections giving clear
openings of 43 x 31 cm. An overall transparent plastic film provides patient support
when the cushions are removed. A removable head cushion is provided to expose a
drill plate suitable for mounting head clamps, etc.
THE SYSTEM -- GENERALLY
Referring now to FIGS. 1 and 5, a computerized tomography system
according to the invention is shown. FIG. 1 is a block diagram which illustrates
the data gathering elements of the present invention in relation to an X-ray source
30 and patient 32. FIG. 5 is a block diagram which illustrates the elements (indicated
by a "*") of the present invention which are added to the above referenced
Ximatron simulator system to obtain a CT simulator system in accordance with the
In FIG. 5, it can be seen that the additional elements include: pre-patient
collimator 34, post-patient collimator 36, grid 38, image intensifier tube 40,
right angle flip mirror 42, photodiode linear array 44, imaging extension 45, 16-bit
ADC and interface electronics 46, data acquisition interface 48, and a processing
and display computer 50.
Data Acquisition Path
Referring now to FIG. 1, an x-ray source 30 passes radiation through
a pre-patient collimator 34, then through the patient 32, then through a post-patient
collimator 36 and anti-scatter grid 38 to an image intensifier tube 40 (IIT) and
imaging extension 45.
The image from image intensifier tube 40 is projected onto a photodiode
linear array 44 using first lens 52, right angle mirror 42 and second lens 54.
When the right angle mirror is 42 swung out of the way, the image from the image
intensifier tube 40 can be viewed with a television camera 56.
Signals from the photodiode array 44 and imaging extension 45 are
sent, on command, to a pre-amp, integrator, reset and clamp circuit 58. Control
logic circuitry 60 in the preamp, integrator, reset and clamp circuit 58 provides
timing signals, which are in turn derived from clocks provided by a phase locked
loop timing and control circuit 62. Circuit 62 is synchronized to the 50/60 Hz
line frequency. Control signals and light intensity signals are sent from the
pre-amp, integrator, reset and clamp circuit 58 through an amplifier and filter
64, multiplexed in multiplexer 65 with signals from an x-ray normalization detector
66, and signals from imaging extension detector array 45, to the analog-to-digital
converter and control circuit 68 (ADC). The ADC 68 sends an enable signal to a
gantry angle encoder and logic circuit 70, and both send signals to the data processing
computer 50 (FIG. 5), via an optically isolated data path 72 and multiplexer 67
in a time-multiplexed manner. The computer 50 returns handshake signals to the
Selected individual portions of the above system will now be discussed
in greater detail.
Pre-patient and Post-patient Collimators 34 and 36
Pre-patient collimator 34 provides the primary collimation of the
fan beam that is incident on the object of interest. These collimators are typically
constructed of lead ("Pb"), are configured to be removable, and typically provide
a beam width at the isocenter 74 of between 0.5cm and 1.0cm. Referring to FIGS.
2, 3, 6, and 7, typical dimensional relationships are illustrated (without the
imaging extension 45) between the X-ray source 30, the pre-patient collimator
34, the center of rotation of the patient 32 ("isocenter"), the post-patient collimator
36, and the image intensifier tube face 24. FIGS. 3 and 2 illustrate these dimensional
relationships for a full fan-beam "head" scan. FIGS. 6 and 7 illustrate these
dimensional relationships for a partial fan-beam "body" scan.
For a "head" scan, referring to FIGS. 3 and 2, it can be seen that
the focal spot of the X-ray source 30 is positioned approximately 100cm from the
isocenter 74, and the image intensifier tube face 24 is positioned approximately
32cm from isocenter 74.
FIG. 3 is taken transverse to the width of the beam, so that the
axis of rotation of the rotating arm 12 is in the plane of the paper. FIG. 2 is
taken looking across the width of the beam so that the axis of rotation of the
rotating arm 12 is coming out of the plane of the paper.
Positioned in the x-ray beam are beam dimension adjustment jaws 76,
followed by pre-patient collimator 34 at approximately 65 cm from the focal spot
of X-ray source 30. X-ray normalization detector 66 is positioned on the x-ray
source side of the pre-patient collimator 34. The post-patient collimator 36 is
positioned just above the image intensifier tube 40.
For a "head" scan, a full fan-beam is used with a beam thickness
of approximately 5mm (at isocenter 74). See Fig. 3. The image intensifier tube
face 24 is centered in the beam. Given the above separation of X-ray source 30,
isocenter 74, and image intensifier tube face 24, pre-patient collimator 34 has
a slit width of approximately 6mm. Post-patient collimator 36 has a slit width
of approximately 8mm. Beam dimension adjustment jaws 76 are set to provide a beam
thickness at pre-patient collimator 34 wide enough to illuminate x-ray normalization
detector 66, and to provide a beam width which is approximately 21.1cm at the
isocenter 74 (see FIG. 2). Further, beam dimension adjustment jaws 76 are positioned
close enough to pre-patient collimator 34 so that the latter is the primary collimator
of the beam.
For a "body" scan, a partial fan-beam is used, with a beam width
of approximately 1 cm used at the isocenter 74. As can be seen in FIG. 6, pre-patient
collimator 34 is positioned between 55-65cm from the focal spot of X-ray source
30, and has a slit width of approximately 6mm. The image intensifier tube face
24 is positioned approximately 35cm from the isocenter 74. Post-patient collimator
36 has a slit width of approximately 13mm. As can be seen in FIG. 7, with the
axis of rotation of the rotating arm 12 coming out of the plane of the paper, the
image intensifier tube face 24 is offset from center, and the beam dimension adjustment
jaws 76 are set so that a partial fan-beam is generated. For example, the beam
from the pre-patient collimator 34 would illuminate the image intensifier tube
face 24 edge-to edge, but the portion of the beam passing through the isocenter
74 would be incident approximately 3cm from one edge of the image intensifier tube
face 24. See FIG. 7.
In angular terms, the beam would have an outer edge approximately
1.27 degrees from a center line 75 running between the focal spot of the X-ray
source 30 and the isocenter 74, and its other outer edge at approximately 10.49
degrees from the center line 75.
In the current embodiment, the slice thickness for head scans is
5mm at isocenter, F.I.D. is 147cm. For body scans the current slice thickness is
1cm at isocenter, with an F.I.D. at 147cm. This provides additional patient scan
A 14:1 cylindrically focused grid is included in the post-patient
The pre-patient collimator 34 gives a well defined fan which helps
to reduce patient dose and scatter. It also carries the beam shaping filters which
attenuate the peripheral portions of the x-ray beam which pass through thinner
portions of the patient 32. This not only reduces patient dose but also reduces
the dynamic range over which the IIT 40 and photodiode linear array 44 must respond.
Dimensional Relationships With Imaging Extension 45
FIG. 8 illustrates the dimensional relationship when the imaging
extension 45 is used. In such a configuration, a partial fan beam can be used for
a 50cm patient scan circle. The pre-patient collimator 34 is positioned between
59-63 cm from the X-ray source 30; and the isocenter is approximately 100 cm from
the X-ray source 30. The image intensifier tube face 24 is positioned approximately
147 cm from the X-ray source 30; and the most vertical point of imaging extension
45 is approximately 8.5cm above the image intensifier tube face 24.
X-ray Normalization Detector 66
Returning to FIG. 1, the x-ray normalization detector 66 is mounted
to one side of the slit in the pre-patient collimator 34 and provides readings
of source intensity which are used to normalize x-ray tube output variations during
the scan. The x-ray normalization detector 66 measures the unattenuated x-ray
beam flux, allowing the detector to sample quite a large solid angle of the beam.
X-ray normalization detector 66 is formed by combining a scintillating cadmium-tungstate
(CdWO4) crystal with a silicon photodiode. Preferably, the crystal
measures 6x6mm by 3mm thick. Additional details of the crystal material, photodiode,
and construction are provided hereinbelow in the discussion of the imaging extension
detector array 45.
IIT 40 and Light Collection Optics 84
The 12-inch image intensifier tube (IIT) 40 is a conventional medical
image intensifier and serves as an X-ray photon to visible light photon converter.
FIG. 9 provides a cross sectional view of a typical image intensifier tube, and
the associated optics of the present invention. Incident X-ray photons are absorbed
by the thin 0.3 mm CsI (cesium-iodide) scintillator 78 on the image intensifier
tube face 24. The CsI crystal emits light photons which are converted to electrons
by the attached photocathode 80. The electrons are accelerated and focused by
focusing grids G1, G2 and G3, onto the output phosphor 82 for light conversion.
The quantum efficiency of this process is characterized by a 4 to 5 order of magnitude
(104 to 105) increase in light photons to incident X-ray
The CsI scintillator 78 is typically 12 mils thick. The output phosphor
82 is preferably "P20" type (ZnCdS). The accelerating voltage is typically between
30-35kV. A one inch diameter image is produced at the output of image intensifier
tube 40. Lead post-patient collimator 36 (FIG. 1) and anti-scatter grid 38 are
used to define the CT slice thickness and to reduce X-ray scatter. The post-patient
collimator 38 arrangement is mounted on a circular aluminum plate which is then
bolted onto the mounting ring of the IIT 40.
The overall QDE of the measurement system is dependent on the efficient
collection of these light photons at the IIT output phosphor 82. The light collection
optics 84 which views the IIT output is a lens system as shown in FIGS. 1 and 9.
The light collection efficiency for this lens geometry is proportional to the
transmission and the square of the numerical aperture. With both lenses 52 and
54 focused at infinity, their light collection efficiency is approximately 1%,
depending on the f-stop setting of the second lens. The IIT to light detector
QDE in this case is still 2 to 3 orders of magnitude (102 to 103)
In one embodiment of the present invention, first lens 52 is a conventional
82mm lens set at a 1.2 f-stop and focused at infinity; and second lens 54 is a
conventional 80mm lens set at a 5.6 f-stop and focused at infinity.
In order to allow both the fluoro TV camera 56 and the photodiode
linear array 44 to be permanently mounted, a dual port distributor with a motorized
45 degree flip mirror 42 is used, and is mounted on the IIT 40 in place of the
standard distributor. The flip mirror 44 normally rests in the fluoro position
and when the CT mode is selected, the mirror is flipped through 90 degrees so
that the IIT light output is directed onto the linear detector array 44 through
second lens 54.
Photodiode Linear Array 44
A variety of solid-state arrays have been evaluated for their suitability
as light detectors of image data from the IIT. The requirements for the light
detector in this system are: compatible spectral sensitivity, a wide signal dynamic
range, i.e. 100,000:1, and sufficient spatial resolution for image reproduction.
Compatible geometrical dimensions are also necessary to permit easy coupling of
the photodiode linear array 44 to the output of the IIT 40.
A commercially available 512 channel linear silicon diode array meets
these requirements and has yielded excellent results. This array is linear image
sensor number S2301, manufactured by Hamamatsu of Hamamatsu City, Japan. The array
is 1 inch (25.6 mm) in length and 2.5 mm in width. Each diode detector is 50 microns
by 2.5 mm with 72% active area. With the IIT output image at one inch in diameter,
a 1-to-1 arrangement light collection optics 84 is used between the IIT 40 and
the array 44. The photodiode linear detector array 44 is built into a camera housing
which is mounted at one of the exit windows of the right angled flip mirror 42.
The normalized photon response of the array is greater than 60% from
475 to 875 nanometers and overlaps the IIT output phosphor spectrum. The IIT light
output spectrum from the "P20" phosphor peaks at 532 nm. Thus, the silicon photodiode
spectral response is a reasonable match to the "P20" phosphor curve. The QDE for
silicon is approximately 0.6 - 0.7 electrons/photon. Therefore, from the above,
the overall QDE ratio of incident X-ray photon to electron-hole pairs in silicon
is still much greater than unity for the system as a whole.
To obtain a spatial resolution of 1 mm on an object, the detector
must have a sufficient number of channels to permit one to digitize the image.
The photodiode linear array 44 has 512 channels, which translates to 0.6 mm per
detector at the image intensifier tube face 24 for a 12 inch tube. Tests and specifications
on the image intensifier tube 40 have indicated that its spatial resolution is
approximately 3.5 line pairs/mm over the 12-inch diameter, which exceeds the diode
array's equivalent 0.9 line pairs/mm. Therefore, projected to the object, 1 mm
resolution is possible in the reconstructed image data.
The photodiode linear array 44 is a 512 channel linear device with
each channel capable of accumulating 22 pico coulombs of charge during an exposure.
The noise characteristics of the commercial array and preamp is specified as 3500
electrons rms. The saturation level to noise ratio yields a single measurement
maximum signal to noise ratio of 39,000:1.
The point spread response of the image intensifier tube 40 indicates
a dynamic signal range of at least 100,000:1. However, the single channel dynamic
range of the photodiode linear array 44 has been measured to be only 35,000:1 when
used with the manufacturer's preamp. Thus, using a single measurement, the photodiode
linear array 44 does not have sufficient dynamic range to match the IIT 40 output.
DYNAMIC RANGE IMPROVEMENTS
A dual exposure time scheme is employed, along with some additional
improvements in the charge preamplifier circuitry, to permit the 100,000:1 signal
range of IIT 40 to be utilized. A new preamplifier integrator design provides
a 50,000:1 range at room temperature. The combination of the dual exposure time
scheme and new preamplifier integrator design provides an array-preamp dynamic
range of 400,000:1, or 19 bits total, for each channel, over a measurement interval
of 100/83 msec (50/60Hz) while maintaining photon statistics.
The improvements in dynamic range are obtained in part by minimizing
of the effect of the charge amplifier reset noise in the preamplifier, by phase
locking the measurement to the line frequency, and by using an analog amplification
scheme to amplify the signal from photodiode linear array 44 prior to converting
it to digital form.
Phase Locking to the Line Frequency
One source of degradation in the dynamic range of the measurement
electronics is line frequency related ripple and harmonics in the x-ray source.
The ripple and harmonics are a by-product of the rectification of the line voltage
used in generating the high voltage CW for the x-ray source.
As can be seen in FIG. 1, phase locked loop timing and control circuit
62 provides a number of clocks which are synchronized to the line frequency. More
specifically, phase locked loop timing and control circuit 62 includes a voltage
controlled oscillator (not shown) which is operating at a preselected multiple
of the line frequency and which is synchronized to the line frequency. Within phase
locked loop timing and control circuit 62 are divider circuits which divide down
the voltage controlled oscillator signal into a sampling clock 86 and a start
frame clock 88. In the embodiment shown in FIG. 1, the sampling clock 86 is 262KHz
and the start frame clock includes a 13.3Hz component and a 120Hz component. These
clocks are applied to pre-amp integrator, reset and clamp circuit 58. As will be
described hereinbelow, these clocks are used in the sampling of the photodiode
linear array 44 and the dual exposure time scheme. Furthermore, phase locked timing
and control circuit 62 supplies a select signal to analog-to-digital converter
and control circuit 68 to synchronize its operation. When the timing in the measurement
electronics is phase-locked in the above manner, substantial rejection of the
line frequency ripple and harmonics can be obtained.
Pre-amp, Integrator, Reset and Clamp Circuit 58
Referring now to FIGS. 10 and 11, the pre-amp, integrator, reset
and clamp circuit 58 will be described in greater detail. In FIG. 10, a simplified
schematic of photodiode linear array 44 is shown. The anodes of the 512 photodiodes
are connected to a low noise reference, such as a bandgap voltage reference. The
cathode of each of the 512 diodes is coupled to a video line 90 by way of pass
transistors 92. The pass transistors are sequentially pulsed by an internal clock
which operates off of sample clock 86, FIGS. 1 and 11.
When a pass transistor 92 is pulsed, the charge which has accumulated
on its associated photodiode is placed on video line 90. This charge is transferred
onto capacitor 94 which is in the feedback loop of the input stage 96 of pre-amp,
integrator, reset and clamp circuit 58. Input stage 96 operates as a charge amplifier
and integrator, and provides at its output a voltage proportional to the amount
of charge present on capacitor 94. Capacitor 98 couples the voltage at the output
of input stage 96 to low pass filter 100, which provides a gain of approximately
four. Low pass filter 100 has a high impedance input and acts as a prefilter prior
to analog to digital conversion by ADC 68. ADC 68 is a single 16-bit linear analog-to-digital
Input stage 96 includes a low noise, high input-impedance amplifier
stage. In the preferred embodiment, a pair of discrete low noise field effect
transistors, such as type 2N5912, are used as a front end voltage follower of the
amplifier. See FIG. 11. The output of the stacked pair is then applied to the
inverting input of a high impedance operational amplifier, such as device number
OP-27, manufactured by PMI of Santa Clara, California.
Reset and Clamping
The pre-amp, integrator, reset and clamp circuit 58 includes a reset
transistor 102 connected in parallel with capacitor 94, and a clamp transistor
104 connected to the end of coupling capacitor 98 connected to low pass filter
100. Reset transistor 102 is pulsed to discharge capacitor 94 in preparation for
receipt of charge from the next photodiode being sampled.
It has been found that a random offset voltage is coupled into the
signal path by way of capacitive feedthrough from the gate of reset transistor
102. This offset can be on the order of one-half the control voltage being applied
to the gate of reset transistor 102. It has also been found that the addition
of clamp transistor 104 reduces the above offset by a factor of five.
In operation, reset transistor 102 is pulsed with a positive going
pulse for a predetermined time, such as two microseconds. At the same time, clamp
transistor 104 is pulsed with a negative going pulse, but for a period about twice
as long, such as four microseconds. During the time the clamp transistor 104 is
pulsed, coupling capacitor 98 charges to the offset voltage. When the negative
going pulse is completed, the end of coupling capacitor 98 connected to low pass
filter 100 follows the output of input stage 96, which will assume a voltage proportional
to the charge being transferred from the next photodiode being sampled in photodiode
linear array 44.
FIG. 11 provides a detailed schematic of the circuitry used to implement
one embodiment of the pre-amp, integrator, reset and clamp circuit 58. In this
embodiment, low pass filter 100 is implemented in three separate stages: 106, 108,
and 110, with coupling capacitor 98 and clamping transistor 104 being located
between stages 106 and 108. Stage 106 is non-inverting and provides a gain of 3.6
and low pass filter knee at 610 kHz; stage 108 operates as a follower; and stage
110 in inverting and provides a gain of 1.2 and a low pass filter knee at 220 kHz.
Also shown in FIG. 11 are circuits 112 for generating a three-phase
clock for use in sampling the photodiode linear array 44; the reset pulse 114
and clamp pulse 116 supplied to reset and clamp transistors, 102 and 104, respectively;
and also the convert signal 118 supplied to ADC 68.
Imaging Extension Detector 45
Referring to FIGS. 12, 13 and 14, the imaging extension detector
45 includes an array of 32 discrete detectors 200. Each detector 200 includes a
high density cadmium tungstate (CdWO4) scintillating crystal 202 mounted
and optically coupled to a UV-enhanced silicon photodiode 204. The scintillating
crystals 202 are 2mm wide by 12mm long by 3mm deep, and have the following characteristics:
Stopping power of 150keV gammas in 3mm90% Stopping power of 3MeV gammas in 12mm30% Light output relative to NaI (T1)40% Wavelength of maximum emission540nm Decay constant5µsec Afterglow at 3msec0.1% Index of refraction at 540nm2.2-2.3 Temp. coeff. of light out at 300K0%/deg.K Density7.9g/cc Melting point1598 K HygroscopicNo
These crystals are available from NKK of Tokyo, Japan and Harshaw Chemical, of
Each crystal is resin mounted to a photodiode 204, with the side
facing the photodiode being polished, painted with a white reflective coating,
and sealed with black epoxy. The photodiodes 204 are preferably model no. S1337-16Br
manufactured by Hamamatsu of Hamamatsu City, Japan, and have the following characteristics
at 25 degrees C:
Quantum efficiency at 540nm70% Radiant sensitivity at 540nm0.35A/W Noise equivalent power6x10-15W/root Hz Rise time0.2µsec Dynamic range10-12 to 10-4 A Dark current at 10mV rev. bias25 pA max. Junc. cap. at 10mV rev. bias65pF Dark current at 5V rev. bias60 pA typ. Junc. cap. at 5V rev. bias22pF
The photodiodes 204 have an active area of 1.1 by 5.9 mm. The photodiode
204 case of 2.7mm wide by 15mm long. This yields detector spacing of 2.85 and
the 9mm length of the imaging extension detector array 45. Each scintillating crystal
202 has an active face of 2mm by 12mm, which yields a slice thickness of the imaging
extension detector array 45, referred to the isocenter, of 8mm.
The limitation on photodetector length is due to the photodiode's
active length and case size. The x-ray signal is typically very high at the periphery
of the scan circle and thus the loss in signal is insignificant. More specifically,
the dynamic range requirements at the periphery of the scan circle are a factor
of 10 less as compared to detectors at the center. Detectors at the center typically
receive the fewest x-ray photons. Further, the width of the imaging extension
detector 45, referred to the isocenter, is approximately 1.9mm, compared to 0.37
mm for the photodiode linear array 44. This yields a spacial resolution approximately
a factor of 5 less than the photodiode linear array 44. However, this lower spacial
resolution is not significant because high spacial content objects are typically
not viewed at the periphery of a body scan circle.
The array is mounted at the edge of the existing 30cm image intensifier
tube face 24. A measured x-ray sensitivity profile along the detector face is
shown in FIG. 14. The combination of image intensifier tube 40 and imaging extension
detectors 45 forms an overlapping hybrid detector design.
X-ray photons, which have passed through the object being scanned,
are incident on the 2mm by 12mm face of the scintillating crystals 202. Each photodiode
204 is operated in the photoconductive mode with 5 volts of reverse bias applied
after each integration and readout. X-rays absorbed in the scintillator produce
light photons which are converted to electron-hole pairs in the diode. The resulting
current flow discharges the diode and the preamplifier measures this loss in charge
for each channel. This scheme is a discrete circuit implementation that is similar
to the approach used in large integrated circuit linear and 2-D photodiode arrays.
A multichannel scanning charge preamplifier is used to sample each
of the 32 detectors and multiplex this data with the existing IIT camera 512 channel
data. The sampling of these detectors operate similarly to that used in connection
with photodiode linear array 44, described hereinabove. Referring to FIGS. 1 and
15, a simplified block diagram and a detailed schematic of such a circuit are provided.
Imaging extension detector array 45 is connected to pre-amp, integrator,
reset and clamp circuit 58 by way of a multiwire harness 206. Signals from each
of the photodiodes 204 in the imaging extension detector array 45 are brought
out from, and a 5V reference and ground connections are provided to, imaging extension
detector array 45 over the multiwire harness 206. As was the case with the photodiodes
of the photodiode linear array 44, a low noise reference is used at the anode
of photodiodes 204. The signals from the photodiodes 204 are then applied in parallel
to a bank of select circuits 208. These select circuits 208 are controlled by
select logic circuit 210 to sequentially and serially place the signals from the
photodiodes 204 onto a video line 212.
In FIG. 15, it can be seen that select circuits 208 are gating circuits
which transfer data applied at their "D" inputs to the "S" output, under control
of signals applied to the "G" terminals. Select logic circuit 210 includes a bank
of parallel to serial shift registers 210-A, 210-B, 210-C and 210-D, which operate
to scan the bank of select circuits 208. Select logic circuit 210 can be implemented
using part number 74HC164.
The scan is begun upon receipt of a start extension pulse on line
213. This pulse is clocked along the parallel to serial shift registers 210-A
through 210-D, at a rate set by the extension clock supplied on line 214. As can
be seen from FIG. 1, the start extension pulse and extension clock are supplied
from control logic 60 on line 216.
As was the case with the photodiode linear array 44, the preamplifier
234 for the imaging extension detector array 45 uses a input amplifier 219 with
a capacitor 220 connected in a negative feedback configuration. Line 212 is tied
to one end of capacitor 220. A reset transistor 218 connected in parallel with
a capacitor 220 to reset it in preparation for receipt of the next sample. A coupling
capacitor 224 couples the output of input amplifier 219 to non-inverting amplifier
226. A clamp transistor 222 is connected to the end of coupling capacitor 224
that is connected to non-inverting amplifier 226. Finally, the output of non-inverting
amplifier 226 is connected to lowpass filter 228.
Reset transistor 218 is pulsed to discharge capacitor 94 in preparation
for receipt of charge from the next photodiode being sampled, and clamping transistor
224 is pulsed during this operation, as is the case with reset transistor 102 and
clamping transistor 104 in the charge amplifier for the photodiode linear array
44. The reset pulse for reset transistor 218 is provided on line 230, while the
clamping pulse is supplied on line 232 from the extension clock 216.
Data Acquisition Timing and Dual Exposure Time Scheme
Referring now to FIG. 16 the relative timing involved in the data
acquisition will be described in greater detail. A complete set of X-ray transmission
data is obtained by collecting 720 projections (sampling cycles) or 2 per degree
of rotation for 60 seconds. In practice, slightly more projections are collected
or slightly more than 360 degrees of rotation. However for purposes of this discussion,
720 projections and 360 degrees of rotation will be assumed.
In FIG. 16, line 120 illustrates the allocation of the 720 projections
across one revolution of the X-ray head assembly 18 and image intensifier tube
assembly 20 about the object of interest. Each such projection (sampling cycle)
takes up approximately 83.3msec (60Hz). Line 122 illustrates how each projection
can be viewed as a series of periods of duration T. The example illustrated in
FIG. 16 uses T equal to 8.33msec. Within a projection, the periods are grouped
to define a long time interval, 9T in length, and a short time interval, 1T in
It has been found that the dynamic range of the photodiode linear
array 44 can be greatly increased by utilizing a dual exposure time scheme in the
above manner. That is, by using a short time interval and a long time interval
over which the photodiodes are permitted to convert photons to electrons, the
most accurate measurement of the two can be selected for use.
As discussed briefly above, the point spread response of the IIT
20 indicates a dynamic signal range of at least 100,000:1. One the other hand,
the single channel dynamic range of the photodiode linear array 44 has been measured
to be 35,000:1 when used with the manufacturer's preamp. As such, the photodiodes
alone will saturate under high levels from IIT 40. With a two-interval sampling
approach, the short interval sample will be the most accurate for high intensity
levels from IIT 40, and the long interval sample will be the most accurate for
low intensity levels from the IIT 40. This effectively extends the dynamic range
of the photodiodes into the 100,000:1 range.
In practice, for the commercially available photodiode array identified
hereinabove, the saturation level is 22 picocoulombs. The f-stop of second lens
54 is set so that the photodiodes will not saturate over a short interval period
when there is no object in the x-ray beam; i.e. with the beam at 125kVp and 15mA.
Preferably, this f-stop setting will result in a light level at the photodiode
linear array 44 of about one-half to three-fourths the saturation level with no
object in the x-ray beam.
Returning now to FIG. 16, the timing for the long and short interval
sampling will be discussed in greater detail. Line 124 illustrates the points at
which the photodiodes of photodiode linear array 44 are sampled relative to one
another. In the leftmost part of line 124 there is shown a first series 126 of
544 sample points: 512 for the photodiode linear array 44, and 32 for the imaging
extension detector array 45. These occur during the left most T-period of line
122; i.e. during the last T-period of projection 1, line 120. In line 124, the
second series 128 of 544 sample points occurs during the first T-period in projection
2, line 122. It is to be noted that no sample points occur in line 124 until the
ninth T-period of projection 2, line 122. Thereafter, a third series 130 of 544
sample points are shown occurring during the tenth T-period. Finally, a fourth
series 132 of 544 sample points are shown occurring during the first T-period of
The long interval sample for projection 2 is taken during the third
series 130 of 544 sample points. The short interval sample for projection 2 is
taken during the fourth series 132 of 544 sample points. For example, it can be
seen from line 124 that the time period between the sampling of diode 1 in second
series 128 and third series 130 is nine T-periods. Diode 1 is therefore permitted
to integrate the incident photons for nine T-periods before it is again sampled.
The samples taken in third series 130 thus represent the long interval sample for
Conversely, the time period between the sampling of diode 1 in third
series 130 and in fourth series 132 is only one T-period long. Thus, samples taken
in the fourth series 132 represent the short interval sample for projection 2.
Line 134 in FIG. 16 illustrates the time for the sample points for
diodes 1-4 in series 126, line 124. As can be seen, the time between sample points
is approximately 15.3 microseconds. Within this 15.3 microsecond period, the input
stage 96 of pre-amp, integrator, reset and clamp circuit 58 is reset (line 136),
the clamp transistor 104 is pulsed (line 138), charge from the photodiode being
sampled (for example photodiode 1) is placed on the video line 90 (line 140),
and convert signal 114 is sent to ADC 68 (line 142). Lines 144 and 146 show the
relative timing of the sampling pulses for photodiodes 2 and 3.
It is to be noted that there are four base-clock periods in each
diode sample period. These base-clock periods are related to the 262KHz clock 86
from phase locked loop timing and control circuit 62, FIG. 1. Similarly, the 8.33
microsecond duration of each T-period is related to the 120Hz clock 88 from phase
locked loop timing and control circuit 62. Finally, the 13.3Hz clock 88 from phase
locked loop timing and control circuit 62 relates to the duration of nine T-periods.
X-ray Normalization Detector 66
As was discussed hereinabove, the intensity of the x-ray source is
monitored by x-ray normalization detector 66, FIG. 1. The signal from x-ray normalization
detector 66 is amplified and filtered and then applied to a multiplexer 65. Also
applied to the inputs of multiplexer 65 is the signal from pre-amp, integrator,
reset and clamp circuit 58. The output of multiplexer 65 is then applied to ADC
68. A select signal 61 is supplied from control logic circuit 60 to select between
the signal from x-ray normalization detector 66 or the from pre-amp, integrator,
reset and clamp circuit 58 for conversion by the pre-amp, integrator, reset and
clamp circuit 58.
A simple optical switch (not shown) is mounted on the drive stand
and a corresponding "finger" is attached to the main gantry gear wheel. This arrangement
generates a pulse when the gantry goes through zero degrees, indicating start of
scan. The next pulse from the optical switch (360 degrees of rotation later) signals
the end of a scan. When the system is ready for data collection, this trigger
initiates data acquisition, a pre-set number of projections are always collected,
and the end of scan pulse sets a flag in the projection data.
To allow for variations in gantry rotation speed the data collection
controller is set to acquire more projections than are required, so there is a
slight overscan (5-10 degrees) at the beginning and end of the gantry rotation.
This allows the x-ray generator output to settle and the gantry to reach constant
angular velocity before data acquisition begins. Data acquisition stops as soon
as the end of scan pulse is detected and the x-ray generator can be turned off.
A gantry angle encoder and logic circuit 70 is attached to the potentiometer
which measures the angle of the rotating arm 12. This encoder nominally gives
ten (10) pulses per degree of gantry rotation and hence allows the determination
of projection angles to 0.1 degree (12-bit counter).
Data Acquisition Interface 48
Data acquisition interface 48 is an optically isolated interface
which utilizes conventional photodiodes and a receiver link. Use of an optical
link greatly reduces electrical ground problems.
In addition to the digital data and handshake signals from ADC 68
and gantry angle encoder and logic circuit 70, an analog channel (not shown) is
brought out from amplifier and filter 64 via data acquisition interface 48 for
use in calibration and set up purposes.
Processing and Display Computer 50
Processing and display computer 50 is preferably a conventional 80286
based personal computer. A conventional 20MFlop array processor, a 250 MB WORM
(write-once-read-many) optical disk drive, a 4 MB RAM memory, 30 MB hard disk,
and an image display card, available from Matrox of Canada, are also used.
DATA CORRECTION, NORMALIZATION AND LINEARIZATION
Further improvements are made possible by correcting the data for
certain known error sources. The processing and display computer 50 corrects for
detector system spatial and intensity non-linearities and offsets. To minimize
the effects of the point spread response of the IIT 40, the data is preprocessed
in the processing and display computer 50 by the array processor. That is, after
background subtraction and normalization, the array data is convolved with an
empirical filter which compensates for the non-ideal point spread response. After
all the projection data is obtained, a 512 x 512 pixel image is then generated
using a convolution and backprojection technique. The resulting CT image has better
than 1 mm spatial resolution and 1% density resolution on a 20 cm water calibration
Sources of Error
Possible sources of error in acquired data lie in both the imaging
chain and in the mechanical system. Imaging chain sources of error (in no particular
priority order) are: 1) time varying x-ray flux from X-ray source 30; 2) photon
scatter; 3) image intensifier tube 40 (non-linearity across face, s-curve distortion,
EHT variation with current, center detector, edge effects, curved face, dark current);
4) photodiode linear array 44 (non-linear response, saturation, long versus short
integration values, dark current); and 5) optics (internal reflections, distortion,
Mechanical system sources of error are: 1) wandering isocenter 74;
2) mechanical flexing; 3) non-uniform rotational speed; 4) lack of stiffness in
IIT structure; and 5) non-repeatability of machine positioning.
The x-ray flux from the X-ray source 30 can vary with time (power
frequency fluctuations, photon statistics, etc.). This is measured directly by
means of x-ray normalization detector 66. The output from x-ray normalization
detector 66 provides a current proportional to the number of incident photons.
It is assumed that this device is perfectly linear and the readings from the other
detectors are normalized to it, i.e., detector elements are scaled as if the x-ray
flux was constant and at its peak value.
Scatter of x-ray photons during their passage through the body is
difficult to correct for. The approach used in the present invention is to attempt
to eliminate scatter problems by: 1) accurately collimating the fan beam, and
2) using a 14:1 cylindrically focused scatter suppression grid in front of the
face of the IIT, although this does cause loss of primary x-ray photons.
Errors and distortion from the image intensifier tube 40 can arise
for the following reasons: 1) uneven distribution of absorbing material/scintillator
(CsI) at the face; 2) curvature of the glass face and its increasing thickness
away from the center; 3) observed spatial non-linearity across the face of the
IIT due to electron focusing errors; 4) s-curve distortion which varies as the
IIT 40 is reoriented in the earth's magnetic field; 5) dark current (i.e., noise);
6) dynamic range (max signal:noise); and 7) finite point spread function across
the tube face due to internal optical light scatter at the input and output of
Distortions and errors in the optical path are generally due to internal
reflections and lens imperfections and can be summarized by the system point spread
For the system, overall light intensity is not a limitation, but
x-ray photons are. The second lens 54 f-stop is generally set to 5.6 so it could
easily be opened up to 4.0 to allow twice as many light photons through. The IIT
40 has a QDE of 1,000-10,000, therefore there can be a loss of light photons before
the QDE of the system reduces to unity. Potential sources of errors in the detector
array are: 1) non-linear detector/amplifier response with respect to number of
light photons detected; 2) dark current; 3) different response with changes in
integration period (for a given light input); 4) detector saturation; and 5) non-repeatability
of center detector position.
A calibration procedure has been implemented to quantify and correct
for data acquisition errors. The image intensifier tube 40, optics, and photodiode
linear array 44 chain are treated as a single unit for the purposes of calibration
and data correction. Information obtained as a result of these calibration steps
is used to correct the data collected during an actual scan.
The calibrations are performed in the following order: a) mirror
alignment; c) dark current (background); c) center detector and detector array
limits (fan angle limits); d) detector system spatial linearity; and e) system
point spread function. Additionally, every detector array is calibrated for response
non-linearities by using a calibrated light source.
Physical Alignment of the System
Physical alignment of right angle flip mirror 42 in this system is
important since a highly collimated beam of light is to be projected onto a long
narrow detector array. The adjustment is carried out when the system is installed
by carefully centering a specially drilled lead mask 300 over the face of the
image intensifier tube 40. See FIG. 17. The incident x-ray beam is collimated so
that only the drilled area of the mask 300 is illuminated and the x-ray flux is
adjusted so that none of the detectors saturate. The mirror is adjusted so that
the response (displayed on an oscilloscope) of the detectors in the photodiode
linear array 44 is symmetrical, flat, and with the correct number of detector
FIG. 18 illustrates the typical detector response pattern obtained
when using the mask 300. As can be seen from the figure, the magnitude of response
is greater for the detectors which receive signals from the large drilled holes.
The focus of first lens 52 and second lens 54 may be adjusted by
looking at the "sharpness" of the detected peaks. See FIG. 18. This adjustment
need not be performed again unless any of the elements in the optical chain are
Background Noise Measurement
The detector system dark current (noise) is determined by collecting
data with the x-ray beam off. The usual number of data projections are collected.
The readings for each detector are then summed and averaged to give an average
dark current (i.e., background) for each individual detector element in the array.
Dark current is a strong function of temperature, therefore calibration
values should be taken as a function of temperature, including temperatures during
warm-up and room ambient. The averaged background values are stored and subtracted
from any data collected when the beam is on.
Identifying Center Detector
Referring to FIG. 19, the center detector is identified as follows:
1) collect background data; 2) set the x-ray flux so that no detectors saturate;
3) collect data with nothing in the beam; 4) set a pin or needle 302 at the isocenter
74; 5) ignore detectors close to the ends of the photodiode linear array 44 where
readings fall off rapidly; 6) run a scan of the pin to collect a full set of readings.
FIG. 20 shows a typical set of readings for any particular projection.
The readings are then processed as follows: a) background correct
the data; b) select long/short integration values (the longs will always be selected
since the x-ray flux is set so that no detectors saturate); c) calculate ln(air
norm)-ln(data); and d) set non-useful detectors to zero.
The results of the calculations are 0 (zero) for most detectors,
with a positive attenuation value at the detectors which "saw" the pin 302.
The peak attenuation values for each projection are identified and
a corresponding interpolated detector number calculated, i.e., the center detector
for that projection.
All the center detector values from each of the projections are then
averaged to compensate for the possibility that the pin was not placed exactly
at the isocenter 74 and for the flexing of the mechanical structure. The result
is the identity of the center detector for the system, which is then saved for
Detector Spatial Non-Linearities
When a ruler with attenuating markers is placed over the face of
the IIT and irradiated, the markers will not be equally spaced when viewed at the
exit window of the IIT, if there are spatial non-linearities in the detector system.
In other words, although the photodetectors in photodiode linear array 44 are
uniformly spaced, various effects in the IIT 40 and the imaging path can cause
a detector, other than the predicted detector, to be affected when an object is
positioned in the fan beam. See FIG. 19.
Among the factors which influence spatial (or geometric) non-linearity
are the curvature of image intensifier tube face 24. As can be seen from FIG.
19, the tube face has a convex shape with respect to the fan-beam. This results
in rays at the outside of the beam striking the face 24 at a larger relative displacement
than rays near the center of the beam. Within the image intensifier tube 40, non-linearities
in the focusing grids G1, G2 and G3, can cause the trajectory of emitted electrons
to depart from the predicted path. Lens errors in first lens 52 and second lens
54, and mispositioning of the right angle flip mirror 42, are also a source of
In order to determine such system spatial non-linearity a second
pin 304 is placed at an offset from the center pin 302 and then moved slowly through
the beam. See FIG. 19. In practice, this effect is actually achieved by keeping
the pin fixed and rotating the gantry.
Referring to Figs. 33a through 33c, this effect is illustrated in
simplified form. In each of the figures, top dead center (TDC) of the gantry rotation
path is shown at the top of the figure. The circle illustrates the path of the
gantry rotation. X-ray source 30 is shown in a different position in each of the
figures. For these different positions of X-ray source 30 it can be seen that a
different detector in photodiode linear array 44 is affected by the occluded ray
The angle &thetas; indicating the angle between center line 75 and
TDC, is measured using gantry angle encoder and logic circuit 70. Angle α
is the angular position of the calibration needle 304 from top dead center. The
angle δ, can then be calculated from &thetas;, α, the distance from
X-ray source 30 to isocenter 74, and the distance between isocenter 74 and calibration
needle 304, using well known geometric techniques.
From this data the angle of the occluded ray 306, and the detector
which responded to the occluded ray 306, can be tabulated. See FIG. 21, where such
a table is shown with data selected to illustrate the concept. From the geometry
of the system, the detector which should have been affected for any position of
the "moving" pin 304 can be predicted. This information is also included in the
tabulation of FIG. 21. The total fan angle is calculated by examining when the
"moving" pin moves into the fan beam and then leaves it.
In the reconstruction method used in the CT simulation of this example,
data corresponding to rays of equal angular displacement are assumed. From the
tabulated data, FIG. 21, and corresponding detector readings taken during an actual
projection, an intensity reading can be determined for the specific angles desired.
The manner in which this information is used in connection with actual measurement
data is described in greater detail hereinbelow in connection with FIG. 22.
System Point Spread Correction
The system point spread function (PSF) is an inherent problem in
the imaging chain and is due to: 1) defocusing of electrons in the IIT 40, 2)
defocusing and scatter of light photons in the optics, and 3) internal reflections
within the optics. As illustrated in FIG. 23, the point spread function is measured
by using a lead slit 312 in conjunction with post patient collimator 36, to cause
a selected spot to be illuminated on image intensifier tube face 24. In the preferred
embodiment of the present invention, the PSF is determined by positioning the
lead slit at isocenter 74 and collecting data as the IIT 40 is moved laterally
by the simulator motor system on arm 16. The IIT 40 is moved to a position where
a detector is fully illuminated and the detector array readings for that position
are read out. This is repeated until all 512 detectors of the photodiode linear
array 44 have been illuminated and readings for the other 511 photodetectors in
the array taken.
The idealized response to this excitation is shown as curve 314 at
the bottom of FIG. 23 as being all zero except in the immediate vicinity of the
photodetector which receives the intensified photons from IIT 40. The dashed line
316 in the figure shows the expected response, taking into account the physical
limitations of the imaging system. Finally, curve 318 shows what has been actually
measured. It is to be noted that there is a distinct "tail" in the response, located
symmetrically opposite the real peak.
There is a different PSF for each slit position over the face of
the IIT, and a different tail. The magnitude of these "tails" increases as the
slit is moved towards the edge of the IIT 40. This has important implications
for body scans using a partial-fan beam. Typically, one side of the IIT 40 is
very brightly illuminated by x-rays after passing through the periphery of the
body, while the other side is dark because not many x-ray photons manage to get
through the center of the body. In this case, the "tails" can be so large as to
swamp out valid readings.
From the measured data a deconvolution function is obtained for each
slit position (and hence for every detector in the array) which then can be applied
to correct the actual detector readings to obtain the idealized response, untainted
by the PSF tails. This means that for every fan (i.e. projection) of data, there
can be 512 deconvolutions.
In practice, these "deconvolutions" are implemented according to
FIG. 24-26 in which amongst others the gathering and preparation of the PSF data
Merging Imaging Extension Detector Array 45
Measurements from imaging extension detector array 45 are combined
with those from photodiode linear array 44 to provide 544 measurements. When the
imaging extension detector array 45 is added to the image intensifier tube 40,
its first photodiode is positioned to overlap the last several photodiodes in
the photodiode linear array 44. The detector spacing of the photodetectors in the
imaging extension detector array 45 is approximately five times that of the effective
detector spacing in the photodiode linear array 44. However, this lower spatial
resolution is not significant because high spatial content objects are typically
not viewed at the periphery of a body scan circle.
As was the case with measurements from the photodiode linear array
44, values can be obtained which correspond to the desired uniform angular displacement
for measurements by the imaging extension detector array 45 by interpolating the
actual measurement data. These interpolated values are then moved into the correct
"expected detector" slots ready for further processing and back projection.
Dual Sample Interval Measurement Method
As is discussed hereinabove, to increase the dynamic range of the
detector system, two sets of readings are taken for each projection, a long integrate
and a short integrate set. The long integrate period is nine half-line cycles (T-periods)
long and the following short integrate period is one half-line cycle (T-period)
long. This approach gives an extended period in which to count low numbers of
photons accurately. If this reading saturates, then the short integrate value can
be used, multiplied by a scaling factor to scale it up to generate an equivalent
long integrate value. This approach assumes that the detector response is linear.
It also importantly preserves photon statistics. When the short integrate is used
the detected x-rays are of sufficient number to discard ninety percent of them
in the measurement.
This dual sample interval approach thus provides a substantial and
significant increase in the dynamic range of the imaging system. By way of example,
if the long interval samples have a 16-bit range, the use of the short sample
interval extends the measurement range to effectively 19-bits. In the context
of the specific example given, the long interval samples would be used for counts
up to about 62,000. The short samples would be used for counts between 62,000
and about 500,000. Thus a dynamic range increase of about 3-bits, or a factor of
about 9, is achieved by use of the dual sample interval approach of the present
Detector Non-Linearity Correcting Polynomials
In practice, it has been found that the response of the photodetectors
in photodiode linear array 44 are slightly non-linear. A simple apparatus for
calibrating the photodetector array is used. A photodiode calibration fixture holding
a single LED and a single normalization photodiode are employed and mounted over
the detector array in place of the second lens 54. The light output from the single
LED is directly proportional to the applied current, thus, the response curves
for the photodiodes in the linear array 44 can be determined by plotting photodiode
response versus current applied to the LED, and is normalized by curve fitting
the data against data from the normalization photodiode. The normalizing photodiode
can be the same photodiode as used in normalizing detector 66, but without the
In the preferred embodiment of the present invention, constants C0,
C1, C2, C3 and C4 for 4th- order polynomials
for each detector are determined and placed in memory during calibration. These
constants are used for example in a fourth order polynomial in which the data
obtained for each photodiode has been fit to the normalizing photodiode response
data using a least-squares curve fit approach. That is, calibration data is obtained
for the normalizing photodiode. This data is assumed linear. Then calibration
data for each of the photodiodes in linear array 44 is obtained. The calibration
data from the photodiodes in linear array 44 is then curve fit to the normalizing
photodiode calibration data using a least squares curve fit criteria. FIG. 27
illustrates a table tabulating photodiode number and coefficient values.
A polynomial rather than table lookup is used because the polynomial
is continuous and can be conveniently executed in the array processor. The use
of a lookup table lookup for the dynamic range found in the present invention would
require too much memory and be too slow. While fourth order polynomials are used
in the preferred embodiment of the present invention, it is to be understood that
polynomials of nth-order, where "n" is more or less than four, can be used within
the spirit of the invention.
In the preferred embodiment of the present invention, an adjustment
factor is determined prior to curve fitting the calibration data from the photodiodes
in the linear array 44 to the calibration data from the normalization photodiode.
This adjustment factor permits the non-linearities of the photodiode being curve-fit
to be more easily interpreted.
The adjustment factor is determined as follows. The normalization
factor is determined for the light intensity level that causes the linear array
photodiode to output a particular middle range level, e.g. 40,000 counts. Thus,
if a linear array photodiode output ("di") of 40,000 counts is produced
by a light intensity level of I0, and for that same light intensity
the normalizing photodetector output ("ni") is 36,000 counts, an adjustment
factor, g, is formed by taking the ratio of ni to di:
The normalizing photodetector output is always set to be lower than any detector
which is being calibrated so that the detector under calibration will staturate
before the normalizing detector.
The adjustment factor, g, is then used to multiply the calibration
data for the normalizing photodetector, and the calibration data for the linear
array photodiodes are curve fit against this adjusted normalizing photodetector
calibration data, ni*. This procedure effectively sets the coefficient
for the first order terms in the nth-order polynomial close to "1", thus better
revealing the higher order effects. The curve fitting preferably performed for
a 4th order polynomial is thus:
ni* = g ni = Co + C1di
+ C2di2 + C3di3
where ni = normalizing photodiode response for light intensity i, di
= linear array photodiode response for light intensity i, and g = the ratio of
ni for light intensity I0 to di for light intensity
In practice it has been found that use of more than one polynomial
is useful to describe the response curve of the linear array more accurately because
of the manner in which the response curve varies with dose, see FIG. 28. It has
been determined that it is more accurate and faster to use a number of higher
(e.g. fourth) order polynomials to model the response curve rather than to try
and find a single higher order polynomial.
In the current embodiment, one polynomial accurately describes the
curve below about 4000 counts and a second polynomial is used from about 2000
to about 62,000 counts. A third is used above about 44,000 counts. See FIG. 28.
Corrected values between 2000 and 4000 counts are obtained by applying both polynomials
1 and 2, then interpolating a final result. Corrected values between 44,000 and
62,000 counts are obtained by applying polynomials 2 and 3, then interpolating
a final result. In practice, polynomial 1 is used with data below 4,000 counts;
polynomial 2 is used with data between 2,000 and 62,000 counts; and polynomial
3 is used with data above 44,000 counts. The overlapping portion of these polynomials
is used to determine the values in the transitional ranges 2000 to 4000, and 44,000
to 62,000 by interpolation. This provides smooth transition between one polynomial
and the next.
Scaling Factor Determination
FIG. 29 illustrates a procedure for determining the coefficients
for the three polynomials. In step 332 the photodiode calibration fixture (not
shown) having the single LED light source and the single normalization photodiode
is substituted for lens 54, FIG 1. In step 336 the response of the photodiode
normalization detector and the photodiodes in the linear array 44 is determined
over the full range of the expected light intensity values. For each intensity
level used, a long and a short time interval is used to obtain a long and a short
Next, in step 337 the scaling factor (for multiplying the short interval)
is determined as follows using the data from the normalizing photodiode. The long
and short interval measurements are examined in the range between about 32,000
and 62,000 counts. The short measurements are multiplied by a scaling factor which
is optimized by least squares fit to obtain the best fit between the long sample
measurements and the scaled short sample measurements in that range. The optimized
scaling factor is then stored for use in multiplying the short sample measurements
for the normalizing photodiode.
In the preferred embodiment of the present invention, the scaling
factor is a two-part factor as follows. The relationship between the normalizing
photodiode long sample interval measurements ("Li") and short sample
interval measurements ("Si") is characterized by the equation:
Li (1 + αLi) = κ Si.
The constants α and κ are optimized for the best fit over the count
range from about 40,000 to 60,000. That is, for the light intensity values that
produce values for Li in the 40,000 to 60,000 range, the values for
Li and corresponding values for Si are applied to the above
equation and the constants α and κ are optimized for best least squares
Once the constants α and κ are determined, the calibration
values, ni, for the normalizing photodiode, which are to be used in
selecting coefficients for linearizing the photodiodes in linear array 44 (step
344), are defined as
ni = Li (1 + αLi), Li <
= κSi, Li ≥ 60,000.
Typical values for a are on the order of 10-7, and for
κ are approximately 9.
Thereafter, in steps 342, 344 and 346, the coefficients for three,
fourth-order polynomials are determined for the photodiode measurement values in
the linear array 44, one polynomial each for the ranges 0 to 4,000, 2,000 to 62,000,
and 44,000 to 500,000, respectively. A "least squares" curve fit is employed.
As discussed above, the curve fit is against the calibration data obtained from
the normalizing photodiode. It is to be noted that the calibration values, ni,
for the normalizing photodiode are provided over the range from 0 to 500,000 counts,
with the counts below 60,000 being provided using the long sample interval measurements,
Li, multiplied by (1 + αLi) and the counts above 60,000
being provided using the short sample interval measurements, Si, multiplied
by the constant κ.
The determination of the constants for the 4th- order polynomial
for the 44,000 to 500,000 count range is made using the normalizing photodetector
value κSi and the unscaled short sample interval measurement
for the particular photodiode from the linear array 44. Because of this, the coefficients
C30, C31, C32, C33 and C34
will effectively incorporate the scaling factor κ. Recall that the scaling
factor κ adjusts the magnitudes of the short time interval measurements
to the order of the long time interval measurement, and in doing so effectively
increases the dynamic range of the detection system by 3-bits. In the above manner,
the increase in dynamic range is passed on to the measurements made by the photodiodes
in the linear array 44. These coefficients are then saved for later use.
In step 344 the coefficients C10, C11, C12,
C13, and C14 correspond to polynomial 1; the coefficients
C20, C21, C22, C23, and C24
correspond to polynomial 2; and the coefficients C30, C31,
C32, C33, and C34
correspond to polynomial 3.
It is to understood that in determining these coefficients, the long
time interval measurements are used for polynomials 1 and 2. For polynomial 3
the short time interval measurements are used. These short measurements are fit
against the normalizing photodiode measurements which can be long interval measurements,
or short interval measurement which have been multiplied by the scaling factor.
SYSTEM OPERATIONData Collection and Correction
Referring to FIG. 26, the overall sequence of data collection and
correction is shown when the system is scanning a patient. After the system is
initialized, step 343, data is collected in step 345.
During data collection the readings obtained at every projection
are: Projection number; Short sampling interval value; Short normalizing detector
value; Long sampling interval value; Long normalizing detector value; and gantry
Corrected detector data is determined in step 347, FIG. 26. In accordance
with the preferred embodiment of the present invention, processing of the actual
detector readings is performed while a scan is in process. Because a scan typically
takes about a minute to complete, a significant amount of processing of the data
can be done during the scan.
As data is received from each projection during a scan the data is
converted to floating point and the background level is subtracted from each detector
reading. The coefficients for each of the three polynomials are then retrieved.
Prior to solving the three polynomials, projection averaging is run to obtain
one set of projection readings per degree of gantry rotation. This involves appropriately
weighting the projection data taken at angles close to the angle for which the
projected data is desired. As discussed earlier, data for two projections per
degree are taken; i.e., approximately 720 projections per 360 degrees of rotation
in a 60Hz system. The projection averaging reduces the number of projections to
about 360, and thus reduces the computational load. For example, projection for
gantry angle 321.5°, 322°, and 322.5° might be averaged together to obtain a set
of data for a projection 322°.
Once projection averaging is completed, polynomials 1 and 2 are solved
using the long sample interval measurement, and polynomial 3 is solved using the
short sample interval measurement.
DET.' = C10 + C11*DET. + C12*DET.2
where DET. = Long sampling interval value.
DET.' = C20 + C21*DET. + C22*DET.2
where DET. = Long sampling interval value.
DET.' = C30 + C31*DET. + C32*DET.2
where DET. = Short sampling interval value.
DET.' is the CORRECTED DETECTOR value being calculated. In practice, the coefficients
for polynomial 3 will incorporate the scaling factor so that the short sample
interval measurement need not be multiplied prior to being plugged into polynomial
Conceptually, referring to Fig. 30A, the magnitude of the measurement
DET. which is plugged into the polynomials will determine which polynomial result
will be actually used for DET'. Steps 364, 366, and 368 illustrate that the result
from polynomial 1 will be used for DET.' when DET. is less than 2,000. From steps
370 and 372, it can be seen that the interpolated valves from polynomials 1 and
2 will be used for DET.' when DET. is between 2,000 and 4,000. When DET. is less
than 44,000, but greater than 4,000, steps 366, 370, 374, and 376, the results
from polynomial 2 will be used for DET.'. For DET. values greater than 62,000,
steps 378 and 380, the results from the polynomial 3 will be used for DET.' Finally,
when DET. is between 44,000 and 62,000, steps 374, 378, and 382, the interpolated
values from polynomials 2 and 3 are used for DET.' This procedure is run for the
measurements from each photodiode in the linear array 44.
The particular order of processing can be selected to increase processing
speed by taking advantage of the array processor characteristics. Thus, in the
present embodiment, it is faster to run all three polynomials on the data, rather
than to screen the data first to determine the proper range and polynomial, then
run the polynomial.
In the present embodiment of the invention, where an array processor
is used, a different order of processing is used. See FIG. 30B. Because the IF-THEN
operations of steps 366, 370, 374, and 378, Fig. 30A, are operationally intensive,
resort has been made to a weighting scheme to increase computational speed. Referring
to FIGS. 30B and 30C, this weighting scheme will now be described.
In FIG. 30B, step 358, the coefficients for the three polynomials
are retrieved. In step 360, the polynomials 1 and 2 are run using the long sample
interval values for the photodiode from the linear array 44, and polynomial 3
is run using the short sample interval valve. In step 362, the "weights" W1,
W2, W3, for each of the three polynomials are determined
as a function of the magnitude of the long sample interval value. Then, in step
363, the results from each of the three polynomials, P1, P2,
and P3, are multiplied by their corresponding weights, then summed
to provide the linearized detector value, DET.':
DET.' = W1 P1 + W2 P2 + W3
In FIG. 30C, the determination of weights W1, W2,
and W3 is shown. The vertical axis represents the weight assigned, while
the horizontal axis represents the long sample interval count. In this particular
example, the count range is 0 to 500,000. The transitions between polynomials
occur at 2000 to 4000 and at 44,000 to 62,000 counts. The weighting factor, W1,
for polynomial 1 covers the count range from 0 to 4000 counts, with a break point
at 2000 counts. The weighting factor, W2, for polynomial 2, covers
the count range from 2,000 to 62,000, with breaks at 4,000 and at 44,000. Finally,
the weighting factor W3, for polynomial 3, covers the count range from
44,000 to 500,000. A curve, C1, is defined using the region of weighting
factor W1 from 2000 counts to 4000 counts:
C1 = 4000 - DET. / (2000)
where DET. equals the long sample interval value. A second curve, C3,
is defined. Using the region of weighting factor W3 from 44,000 to 62,000
C3 = DET. - 44000 / (18,000)
C1 and C3 are both solved for each value of DET. that is
processed, however, values of C1 and C3 above 1 and below
0 are clipped, i.e., ignored. The weights W1, W2, and W3
are then designated as follows using the values of C1 and C3
for the particular DET.:
The use of these weights, W1, W2, and W3,
in the above manner makes efficient use of the array processor, and increases
the speed by which the data can be processed.
Point Spread Function Correction
Returning to FIG. 26, following the correction of the data for background
noise and non-linearities, step 347, PSF corrections are made in step 390.
In FIG. 32, the ε matrix, discussed above in connection with
FIGS. 24 and 25, is used as follows to implement the deconvolution which corrects
for the PSF. Assume the following relationship:
[A][I] = [R]
where, [A] is a 512x512 matrix representing the point spread function, [I] is a
512 element vector representing the x-ray intensity incident on the image intensifier
tube face 24 for each of the 512 detectors, and [R] represents the actual measurements
from each of the 512 linear array photodiodes taken during a projection. The vector
[I] is the information that is being sought. To obtain [I], the vector [R] is
multiplied by the inverse of [A], [A]-1:
[A]-1[R] = [A]-1[A][I] = [I].
Note, however, that [A] can be expressed as identity matrix plus the ε matrix.
Note also that, since the E matrix is small, [A]-1 equals, to a first
order, the identity matrix minus the ε matrix.
Thus, in accordance with the preferred embodiment of the present
invention, the deconvolved values [I] for the measured data are determined by
[I] = ([IDENTITY MATRIX] - [ε MATRIX])[R] .
In step 392, FIG. 32, the ε matrix is retrieved from memory.
In step 394, the DET.' values ("[R]"), from step 347, FIG. 26, are retrieved. Then,
in step 396 the correction vector is determined by multiplying the DET.' values
by the ε matrix. Finally, in step 398, the correction vector is subtracted
from the DET.' values to obtain the vector [I], [DET. "O, DET." 1,...DET." 511].
In accordance with the preferred embodiment of the present invention,
the assumption that the PSF is a slowly varying function of position is further
exploited to speed up the above identified calculations. Instead of determining
the ε matrix for all 512 slit positions, values are collected for every fourth
or so position. Thus, the ε matrix might initially take the form of a 128x128
matrix. Further, actual measurements are taken for the corresponding 128 detectors,
and the vector [I] is then calculated from this more limited set of data. Because
the PSF is a slowly varying function of position, the resulting 128 element [I]
vector can be interpolated to a full 512 element vector with minimal loss of resolution.
Phantom Normalization and Line Integral Calculation
Returning to FIG. 26, step 400 is next processed. This step involves
a determination of the relationship:
Line Integral = In(Corrected DET.) - ln (normaliz. DET.) - In(phantom).
The line integral difference is conventional in the computerized
tomography scanner art, and involves taking the difference between the natural
logarithm of the intensity measured during an actual projection and the intensity
measured by the normalizing photodiode and the intensity using a phantom having
known absorption characteristics.
There can be approximately 800 projections (60Hz system) or 650 projections
(50Hz system) worth of data to store. In practice, the gantry is rotated 5 to
10 degrees beyond TDC at the end of a scan. This results in a slight overscan.
Projections are taken during this overscan region. Data from these projections
are blended with data from projections taken at the beginning of the scan. See
FIG. 26, step 401. Referring to FIG. 31, the weighting assigned to the data from
each projection is illustrated. From the figure it can be seen that the data taken
in the early projections, around zero degrees gantry angle, are lightly weighted,
while the data taken at the end of the scan, around 360 degrees, are weighted
more heavily and then decreasing in weight out to 370 degrees.
Geometric Non-Linearity Adjustment:
Next, step 402 is processed in which adjustments are made to compensate
for geometric or spatial non-linearites. As described in connection with FIGS.
33a, 33b, 33c, and 21, hereinabove, rays in the partial fan-beam which are separated
by uniform angles do not necessarily produce responses at detectors in the photodiode
linear array 44 which are spaced a correspondingly uniform distance apart.
FIG. 22 illustrates the averaging/interpolation technique employed
in step 402, FIG. 26, which corrects for these spatial non-linearites. The upper
section of axis 308 illustrates the desired uniform angular interval between measurements,
for example, a measurement every Δ degrees, between ± 12 degrees. The bottom
section of axis 308 illustrates the actual angular interval between the actual
measurements. Note that due to the spatial non-linearities in the imaging system
detector responses occur at other than the required angles.
As can be seen from portion 310 of FIG. 22, intensity values for
a desired angular position are determined by selecting a subset of the detector
measurements and interpolating those measurements. Thus, for example, the intensity
value for the angular position three-Δ intervals from the -12° point might
be determined by interpolating the measurements from detectors 1 and 2. Similarly,
the intensity value for the angular position two-Δ intervals to the left
of the 0° position might be determined by interpolating the measurements from
detectors 250-253. In the above manner, responding detector readings can be averaged/interpolated
together and then moved into the correct "required detector" slots ready for further
processing and back projection.
The corrected data is then written to a reconstructor input file
where it adjusted for partial fan reconstruction and is then ready for partial
fan reconstruction, step 404, FIG. 26. Reference is made to co-pending patent application,
entitled "Partial Fan-beam Tomographic Apparatus and Data Reconstruction Method",
in the names of John Pavkovich and Edward Seppi, and filed even date herewith,
in which the adjustment and partial fan reconstruction method are described in
One immediate result of the wide dynamic range provided by the present
invention is that a CT simulator system can be provided in which images are produced
which are calibrated to CT numbers. Unlike other previous CT simulator systems,
which produced signals calibrated to arbitrary numbers, the CT simulator system
according to the invention provides data which is calibrated to CT numbers covering
the scale of -1000 to +3000, as in conventional diagnostic CT scanners. To perform
such calibration, a phantom of known materials is scanned and the transmission
values obtained for each of the materials is stored. When transmission data from
an actual scan is obtained, such data is compared against the values obtained
for the phantom and appropriate adjustments are made to the data.
Verfahren zur Verbesserung der dynamischen Auflösung eines Bildgebungssystems,
bei dem ein Photodetektor-Array mit einer Vielzahl von Photodioden benutzt wird,
die dem sichtbaren Licht eines Ausgangs eines Bildverstärkers exponiert sind und
zyklisch innerhalb sukzessiver Abtastzyklen gleicher Dauer gemäß einer Zweifach-Technik
abgetastet werden, wobei jeder Abtastzyklus ein kurzes und ein langes Zeitintervall
die Kurzzeitintervall-Expositionsabtastwerte zur Weiterverarbeitung ausgewählt
werden, und diese Abtastwerte mit einem Skalierungsfaktor multipliziert werden,
wenn festgestellt wird, daß die Langzeitintervall-Expositionsabtastwerte oberhalb
eines Übergangsbereichs von Abtastwerten liegen; und
die Langzeitintervall-Expositionsabtastwerte zur Weiterverarbeitung ausgewählt
werden, wenn festgestellt wird, daß die Langzeitintervall-Expositionsabtastwerte
unterhalb des Übergangsbereichs von Abtastwerten liegen; und
eine gewichtete Kombination der Langzeitintervall- und der Kurzzeitintervall-Expositionsabtastwerte
ausgewählt wird, wenn die Langzeitintervall-Expositionsabtastwerte in dem Übergangsbereich
der Abtastwerte liegen,
wobei der Skalierungsfaktor in dem Selektionsschritt bestimmt wird durch:
(i) es wird über ein langes und ein kurzes Expositionsintervall die Antwort
einer von der Vielzahl der Photodioden verschiedenen Normierungs-Photodiode auf
Signale gemessen, die eine maximale Lichtintensität aufweisen, die derart gewählt
ist, daß die Normierungs-Photodiode während des langen Expositionsintervalls nicht
(ii) es werden die Meßwerte der langen Expositionsintervalle mit den Meßwerten
der kurzen Expositionsintervalle über einen vorgegebenen Bereich von Lichtintensitäten
verglichen, wobei die Meßwerte der kurzen Expositionsintervalle mit einem Skalierungsfaktor
multipliziert werden; und
(iii) es wird dieser Skalierungsfaktor angepaßt, um die beste Übereinstimmung
zu erhalten, wobei ein Kurvenanpassungskriterium nach den "kleinsten Quadraten"
zwischen den skalierten Meßwerten der kurzen Expositionsintervalle und den Meßwerten
der langen Expositionsintervalle benutzt wird.
Verfahren nach Anspruch 1,
dadurch gekennzeichnet, daß zusätzlich die Expositionsabtastwerte linearisiert
werden, die zur Weiterverarbeitung von Nichtlinearitäten in der Antwort des Photodetektor-Arrays
Verfahren nach Anspruch 2,
dadurch gekennzeichnet, daß der Linearisierungsschritt folgende Schritte umfaßt:
es wird eine Vielzahl von Polynomen n-ter Ordnung als Funktion der zu selektierenden
Expositionsabtastwerte eingerichtet, wobei diese Polynome n-ter Ordnung Koeffizienten
(C0-C4) aufweisen, welche als Funktion von ausgewählten Eigenschaften
des Photodioden-Arrays festgelegt werden, und wobei weiterhin verschiedene Polynome
n-ter Ordnung verschiedenen Bereichen der Abtastwerte zugeordnet sind;
es wird ein Polynom n-ter Ordnung ausgewählt, das demjenigen Bereich der Abtastwerte
entspricht, der die Werte der ausgewählten Expositionsabtastwerte einschließt;
es wird die Lösung des oder jedes ausgewählten Polynoms n-ter Ordnung als linearisierte
Verfahren nach Anspruch 3,
gekennzeichnet dadurch, daß die Polynome n-ter Ordnung Bereichen von Meßwerten
der Normierungsphotodiode zugeordnet sind, die sich überlappen, und wobei weiterhin
der Linerarisierungsschritt die folgenden Schritte umfaßt:
es wird bestimmt, ob die ausgewählten Expositionsabtastwerte in einen der Überlappungsbereiche
es werden die Lösungen der Polynome n-ter Ordnung interpoliert, die demjenigen
Überlappungsbereich entsprechen, in den die ausgewählten Expositionsabtastwerte
es wird die interpolierte Lösung als die linearisierten, ausgewählten Expositionsabtastwerte
zur Verfügung gestellt, wenn diese Werte in den Überlappungsbereich fallen.
Verfahren nach Anspruch 3,
gekennzeichnet dadurch, daß die Koeffizienten der Polynome n-ter Ordnung gemäß
der folgenden Schritte ausgewählt werden:
es werden kalibrierte Lichtintensitätssignale auf den Photodetektor-Array gestrahlt,
es wird die Antwort des Photodetektor-Arrays auf die Kalibriersignale gemessen;
es werden die Koeffizienten der Polynome n-ter Ordnung ausgewählt, um die beste
Anpassung des betreffenden Polynoms auf die eingestrahlten Kalibriersignale als
Funktion der Antwort des Photodioden-Arrays gemäß einem Kurvenanpassungskriterium
nach den kleinsten Quadraten zu erhalten.
Verfahren nach Anspruch 5,
gekennzeichnet dadurch, daß der Meßwertbereich der Antwort des Photodetektor-Arrays
Werte von 0 bis 500.000 umfaßt und daß der Lösungsschritt Lösungen von drei Polynomen
4. Ordnung benutzt.
Verfahren nach Anspruch 6,
dadurch gekennzeichnet, daß der Bereich von 0 bis 500.000 in drei Bereiche aufgeteilt
ist, und jeder der drei Polynome 4. Ordnung einem anderen der drei Bereiche entspricht.
Verfahren nach Anspruch 7,
dadurch gekennzeichnet, daß sich die drei Bereiche überlappen.
Verfahren nach einem der Ansprüche 1 bis 8,
dadurch gekennzeichnet, daß der Übergangsbereich der Abtastwerte einen einzelnen
ausgewählten Abtastwert abdeckt.
Verfahren nach einem der Ansprüche 1 bis 9,
dadurch gekennzeichnet, daß der Skalierungsfaktor aus zwei Teilen besteht und durch
folgende Gleichung charakterisiert ist:
Li (1 + α&peseta;Li) = k&peseta;Si,
wobei Li = Meßwerte der Langzeitexpositionsintervalle der Normierungs-Photodiode,
Si = Meßwerte der Kurzzeitexpositionsintervalle der Normierungs-Photodiode,
und wobei die Konstanten α und k durch Bestanpassung über den ausgewählten
Bereich optimiert sind.
A method for improving the dynamic resolution of an imaging system using a
photodetector array comprising a plurality of photodiodes which are exposed to
a visible light output of an image intensifier and cyclically sampled within successive
sampling cycles of equal duration according to a dual technique, wherein each sampling
cycle comprises a short time interval and a long time interval; and
selecting the short time interval exposure samples for further processing,
and multiplying the values of these samples by a scaling factor, when it is determined
that the values of the long time interval exposure samples fall above a transition
range of sample values;
selecting the long time interval exposure samples for further processing when
it is determined that the values of the long time interval exposure samples fall
below the said transition range of sample values; and
selecting a weighted combination of the long time and short time interval exposure
samples when the values of the long time interval exposure samples fall within
the said transition range of sample values,
wherein said scaling factor in said selecting step is determined
(i) measuring the response of a normalizing photodiode distinct from said plurality
of photodiodes over a long exposure interval and a short exposure interval to signals
having a maximum light intensity which is selected so that said normalizing photodiode
does not saturate during the long exposure interval;
(ii) comparing the long exposure interval measurements to the short exposure
interval measurements over a predetermined range of light intensities, wherein
the short exposure interval measurements are multiplied by a scaling factor; and
(iii) adjusting this scaling factor to obtain the best match using a "least
squares" curve fitting criteria between the scaled short exposure interval measurements
and the long exposure interval measurements.
The method of claim 1, further including the step of linearizing the exposure
sample values selected for further processing for non-linearities in the photodetector
The method of claim 2, wherein said linearizing step comprises the steps of:
establishing a plurality of nth-order polynomials as a function of the exposure
sample values to be selected, wherein these nth-order polynomials have coefficients
(Co-C4) which are determined as a function of selected characteristics
of said photodetector array, and further wherein different ones of the nth-order
polynomials apply to different ranges of sample values;
selecting an nth-order polynomial which corresponds to the range of sample
values which includes the values of said selected exposure samples; and
using the solution of the or each selected nth-order polynomial as the linearized
selected exposure sample values.
The method of claim 3, wherein the nth-order polynomials apply to ranges of
measurement values of the normalizing photodiode response which overlap, and further
wherein said linearizing step includes the further steps of:
determining whether the selected exposure sample values fall in any of said
interpolating the solutions from the nth-order polynomials which correspond
to the overlap range in which the selected exposure sample values fall; and
providing the interpolated solution as the linearized selected exposure sample
values when these values fall within that overlap range.
The method of claim 3, wherein said coefficients of said nth-order polynomials
are selected according to the steps of:
applying calibrating light intensity signals to said photodetector array;
measuring the photodetector array response to said calibrating signals;
selecting said coefficients for the nth-order polynomials to provide a best
fit of the respective polynomial to the applied calibration signals as a function
of the photodetector array response, according to a least squares curve fitting
The method of claim 5, wherein the range of measurement values of said photodetector
array response to said calibrating signals comprises values from zero to 500,000
and the solving step uses solutions from three fourth-order polynomials.
The method of claim 6, wherein said zero to 500,000 range is divided into three
ranges, and each of said three fourth-order polynomials corresponds to a different
one of said three ranges.
The method of claim 7, wherein said three ranges overlap.
The method of any one of claims 1 to 8, wherein said transition range of sample
values covers a single selected sample value.
The method of any one of claims 1 to 9, wherein said scaling factor is a two-part
factor characterised by the equation:
Li (1 + αLi) = k&peseta;Si,
wherein Li = the normalizing photodiode long exposure interval measurements,
Si = the normalizing photodiode short exposure interval measurements,
and the constants α and k are optimized for the best fit over the selected
Procédé pour améliorer la résolution dynamique d'un système d'imagerie utilisant
un réseau de photodétecteurs comprenant une pluralité de photodiodes, qui sont
exposées à une lumière visible délivrée par un amplificateur d'images et échantillonnées
cycliquement pendant des cycles d'échantillonnage successifs de même durée conformément
à une technique binaire, selon laquelle chaque cycle d'échantillonnage comprend
un bref intervalle de temps d'exposition et un long intervalle de temps d'exposition;
sélectionner des échantillons soumis à une exposition pendant le bref intervalle
de temps d'exposition, pour un traitement ultérieur, et multiplier les valeurs
de ces échantillons par un facteur d'échelle, lorsqu'il est établi que les valeurs
des échantillons d'exposition pendant un long intervalle de temps se situent au-dessus
d'une gamme de transition de valeurs d'échantillons;
sélectionner les échantillons d'exposition pendant un long intervalle de temps
pour un traitement ultérieur lorsqu'il est établi que les échantillons d'exposition
pendant un long intervalle de temps se situent au-dessous de la gamme de transition
des valeurs échantillons; et
sélectionner une combinaison pondérée des échantillons d'exposition pendant
un long intervalle de temps et pendant un bref intervalle de temps, lorsque les
valeurs d'échantillons d'exposition pendant un long intervalle de temps se situent
dans ladite gamme de transition de valeurs d'échantillons,
ledit facteur d'échelle utilisé lors de ladite étape de sélection
étant déterminé par :
(i) mesure de la réponse d'une photodiode de normalisation, distincte de ladite
pluralité de photodiodes pendant un long intervalle d'exposition et pendant un
bref intervalle d'exposition, à des signaux possédant une intensité de lumière
maximale, qui est choisie de telle sorte que ladite photodiode de normalisation
ne se sature pas pendant le long intervalle d'exposition;
(ii) comparaison des mesures effectuées pendant un long intervalle d'exposition
aux mesures exécutées pendant un bref intervalle d'exposition, dans une gamme prédéterminée
d'intensités de lumière, les mesures effectuées pendant le bref intervalle d'exposition
étant multipliées par un facteur d'échelle; et
(iii) ajuster ce facteur d'échelle pour obtenir la meilleure correspondance
en utilisant un critère d'ajustement de courbe "selon la méthode des moindres carrés"
entre les mesures effectuées pendant le bref intervalle d'exposition, qui ont
été soumises à un cadrage d'échelle, et les mesures effectuées pendant un long
Procédé selon la revendication 1, comprenant en outre les étapes consistant
à linéariser les valeurs d'échantillons d'exposition choisies pour un traitement
ultérieur dans le cas de non-linéarités dans la réponse du réseau de photodétecteurs.
Procédé selon la revendication 2, selon lequel ladite étape de linéarisation
comprend les étapes consistant à :
établir une pluralité de polynômes d'ordre n en fonction des valeurs d'échantillons
d'exposition devant être choisies, ces polynômes d'ordre n possédant des coefficients
(C0 - C4), qui sont déterminés en fonction de caractéristiques
sélectionnées du réseau de photodétecteurs, et en outre différents polynômes parmi
les polynômes d'ordre n s'appliquant à des gammes différentes de valeurs d'échantillons;
sélectionner un polynôme d'ordre n, qui correspond à la gamme de valeurs d'échantillons
qui inclut les valeurs desdits échantillons d'exposition sélectionnés; et
utiliser la solution du ou de chacun des polynômes sélectionnés d'ordre n en
tant que valeurs d'échantillons d'exposition sélectionnées et linéarisées.
Procédé selon la revendication 3, selon lequel les polynômes d'ordre n s'appliquent
à des gammes de valeurs de mesure de la réponse des photodiodes de normalisation,
qui se chevauchent, et en outre selon lequel ladite étape de linéarisation inclut
les étapes supplémentaires consistant à :
déterminer si les valeurs d'échantillons d'exposition sélectionnées se situent
dans l'une quelconque desdites gammes en chevauchement;
interpoler les solutions à partir des polynômes d'ordre n, qui correspondent
à la gamme de chevauchement, dans laquelle se situent les valeurs d'échantillons
d'exposition sélectionnées; et
délivrer la solution interpolée sous la forme des valeurs d'échantillons d'exposition
sélectionnées et linéarisées, lorsque ces valeurs se situent dans cette gamme de
Procédé selon la revendication 3, selon lequel lesdits coefficients desdits
polynômes d'ordre n sont choisis conformément aux étapes consistant à :
appliquer des signaux d'intensité lumineuse d'étalonnage audit réseau de photodétecteurs;
mesurer la réponse du réseau de photodétecteurs auxdits signaux d'étalonnage;
sélectionner lesdits coefficients pour les polynômes d'ordre pour obtenir le
meilleur ajustement du polynôme respectif aux signaux appliqués d'étalonnage en
fonction de la réponse du réseau de photodétecteurs, conformément à un critère
d'ajustement de courbes selon la méthode des moindres carrés.
Procédé selon la revendication 5, selon lequel la gamme des valeurs de mesure
de ladite réponse du réseau de photodétecteurs auxdits signaux d'étalonnage comprend
des valeurs allant de zéro à 500 000, et l'étape de résolution utilise des solutions
fournies par trois polynômes du quatrième ordre.
Procédé selon la revendication 6, selon lequel ladite gamme de zéro à 500 000
est subdivisée en trois gammes, et chacun desdits trois polynômes du quatrième
ordre correspond à l'une différente desdites trois gammes.
Procédé selon la revendication 7, selon lequel lesdites trois gammes se chevauchent.
Procédé selon l'une quelconque des revendications 1 à 8, selon lequel ladite
gamme de transition de valeurs d'échantillons englobe une seule valeur d'échantillon
Procédé selon l'une quelconque des revendications 1 à 9, selon lequel ledit
facteur d'échelle est un facteur en deux parties, caractérisé par la relation :
Li (1 + αLi) = k . Si,
dans laquelle Li = les mesures pendant un long intervalle d'exposition
de la photodiode de normalisation, Si = les mesures pendant un bref
intervalle d'exposition de la photodiode de normalisation, et les constantes α
et k sont optimisées pour le meilleur ajustement dans la gamme sélectionnée.