AU8237391A

AU8237391A – Electronically enhanced x-ray detector apparatus
– Google Patents

AU8237391A – Electronically enhanced x-ray detector apparatus
– Google Patents
Electronically enhanced x-ray detector apparatus

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Publication number
AU8237391A

AU8237391A
AU82373/91A
AU8237391A
AU8237391A
AU 8237391 A
AU8237391 A
AU 8237391A
AU 82373/91 A
AU82373/91 A
AU 82373/91A
AU 8237391 A
AU8237391 A
AU 8237391A
AU 8237391 A
AU8237391 A
AU 8237391A
Authority
AU
Australia
Prior art keywords
detector
linear array
range
signal
ray
Prior art date
1990-07-02
Legal status (The legal status is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the status listed.)

Granted

Application number
AU82373/91A
Other versions

AU644670B2
(en

Inventor
John M Pavkovich
Edward J Seppi
Edward G Shapiro
Current Assignee (The listed assignees may be inaccurate. Google has not performed a legal analysis and makes no representation or warranty as to the accuracy of the list.)

Varian Medical Systems Inc

Original Assignee
Varian Associates Inc
Priority date (The priority date is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the date listed.)
1990-07-02
Filing date
1991-07-02
Publication date
1992-01-23

1991-07-02
Application filed by Varian Associates Inc
filed
Critical
Varian Associates Inc

1992-01-23
Publication of AU8237391A
publication
Critical
patent/AU8237391A/en

1993-12-16
Application granted
granted
Critical

1993-12-16
Publication of AU644670B2
publication
Critical
patent/AU644670B2/en

1999-07-08
Assigned to VARIAN MEDICAL SYSTEMS, INC.
reassignment
VARIAN MEDICAL SYSTEMS, INC.
Request to Amend Deed and Register
Assignors: VARIAN ASSOCIATES, INC.

2011-07-02
Anticipated expiration
legal-status
Critical

Status
Ceased
legal-status
Critical
Current

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Classifications

H—ELECTRICITY

H05—ELECTRIC TECHNIQUES NOT OTHERWISE PROVIDED FOR

H05G—X-RAY TECHNIQUE

H05G1/00—X-ray apparatus involving X-ray tubes; Circuits therefor

H05G1/08—Electrical details

H05G1/64—Circuit arrangements for X-ray apparatus incorporating image intensifiers

A—HUMAN NECESSITIES

A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE

A61B—DIAGNOSIS; SURGERY; IDENTIFICATION

A61B6/00—Apparatus for radiation diagnosis, e.g. combined with radiation therapy equipment

A61B6/02—Devices for diagnosis sequentially in different planes; Stereoscopic radiation diagnosis

A61B6/03—Computerised tomographs

A61B6/032—Transmission computed tomography [CT]

A—HUMAN NECESSITIES

A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE

A61N—ELECTROTHERAPY; MAGNETOTHERAPY; RADIATION THERAPY; ULTRASOUND THERAPY

A61N5/00—Radiation therapy

A61N5/10—X-ray therapy; Gamma-ray therapy; Particle-irradiation therapy

A61N5/1048—Monitoring, verifying, controlling systems and methods

A—HUMAN NECESSITIES

A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE

A61B—DIAGNOSIS; SURGERY; IDENTIFICATION

A61B6/00—Apparatus for radiation diagnosis, e.g. combined with radiation therapy equipment

A61B6/54—Control of apparatus or devices for radiation diagnosis

A61B6/548—Remote control of the apparatus or devices

A—HUMAN NECESSITIES

A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE

A61B—DIAGNOSIS; SURGERY; IDENTIFICATION

A61B6/00—Apparatus for radiation diagnosis, e.g. combined with radiation therapy equipment

A61B6/58—Testing, adjusting or calibrating apparatus or devices for radiation diagnosis

A61B6/582—Calibration

A61B6/583—Calibration using calibration phantoms

Description

60,000.
Typical values for α are on the order of 10″7, and for j 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, n , 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, Lif multiplied by (1 + αLA) and the counts above 60,000 being provided using the short sample interval measurements, Sif multiplied by the constant j.
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 jSA and the unsealed 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 j . Recall that the scaling factor j adjusts the magnitudes of the short sample interval measurements to the order of the long sample 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, Cχι, 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 sample interval measurements are used for polynomials l and 2. For polynomial 3 the short sample 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 OPERATION Data 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 angular position.
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.
POLYNOMIAL 1;
DET.’ = C10 + Clχ*DET. + C12*DET.2 + C13*DET.3 + C14*DET.4, where DET. = Long sampling interval value.
POLYNOMIAL 2:
DET.’ = C20 + C21*DET. + C22*DET.2 + C23*DET.3 + C24*DET.4, where DET. = Long sampling interval value.
POLYNOMIAL 3:

DET. ‘ = C30 + C31*DET. + C32*DET.2 + C33*DET.3 + C34*DET.4, 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 3.
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. 3OB. 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. 3OB and 30C, this weighting scheme will now be described.
In FIG. 3OB, 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” Wχ, 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, Pχ, P2, and P3, are multiplied by their corresponding weights, then summed to provide the linearized detector value, DET. ‘ :
DET. « = W, W„ W3 P3*
In FIG. 30C, the determination of weights 1# W2, and 3 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, j_, 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, Cχ, is defined using the region of weighting factor Vll from 2000 counts to 4000 counts: r m 4000 – DET. Cl 2 2000000 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 counts: c „ PET. – 44,000 3 18,000
C2 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 2,
W2, and W3 are then designated as follows using the values of Cχ and C3 for the particular DET. :
0, Cχ < 0; Wj = l, Cχ > 1; c,, 0 < C. < 1. o, o; w3 = i, c > i;
‘3’ o < c, < 1. w2 = l - w3| - v \ DET. DET. The use of these weights, VI l f 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 e 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!.*] = A]"1[A] [I] = [I]- Note, however, that [A] can be expressed as identity matrix plus the e matrix. Note also that, since the e matrix is small, [A]""1 equals, to a first order, the identity matrix minus the e matrix. Thus, in accordance with the preferred embodiment of the present invention, the deconvolved values [I] for the measured data are determined by the equation: [I] = ([IDENTITY MATRIX] - [€ MATRIX] ) [R] . In step 392, FIG. 32, the e matrix is retrieved from memory. In step 394, the DET.1 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 e matrix. Finally, in step 398, the correction vector is subtracted from the DET. ' values to obtain the vector [I], [DET. "0, 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 € matrix for all 512 slit positions, values are collected for every fourth or so position. Thus, the e 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.) - In (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. Overlap Correction 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 D 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-D 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-D 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 detail. This copending application is hereby incorporated herein by reference. A full print-out of a computer program which implements the present invention in connection with a CT simulator system is provided in Microfiche Appendix A herein. 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. This invention is not limited to the preferred embodiment and alternatives heretofore described, to which variations and improvements may be made, without departing form the scope of protection of the present patent and true spirit of the invention, the characteristics of which are summarized in the following claims. Claims (41) ClaimsWhat is claimed is: 1. An electronically enhanced detector for use in a radiation therapy simulator machine, wherein the radiation therapy simulator machine has a source of x-ray photons, comprising: an x-ray image intensifier tube, said intensifier tube providing a visible output intensity spectrum in response to x-ray photons impinging on said intensifier tube input; a linear array of a multiplicity of photodiodes; optical means for coupling the visible output of said intensifier tube to the linear array; and electronic signal processing means for processing detector signals from said photodiodes to condition said detector signals for use in constructing a tomographic x-ray image corresponding to the geometry of the radiation therapy simulator machine. 2. The apparatus of claim 1 wherein said signal processing means includes means for linearizing the detector signals from the linear array. 3. The apparatus of claim 2 wherein said signal processing means includes means for accommodating a wide range of detector signal magnitudes from the linear array. 4. The apparatus of claim 3 wherein said means for accommodating a wide range of detector signal magnitudes includes means for measuring the detector signal over a long sample interval and a short sample interval. 5. The apparatus of claim 1 where said electronic signal processing means includes low-noise wide-dynamic-range analog electronics. 6. The apparatus of claim 5 including digitizing means for providing an effective 19 bit signal representative of the conditioned detector signal. 7. The apparatus of claim 5 where said digitizing means includes a 16 bit analog-to-digital converter and means for transforming the 16 bits of data from the 16 bit analog-to-digital converter into an effective 19 bit signal when a predetermined sample interval is used to sample the diode photodetector means. 8. The apparatus of claim 1 where said electronic signal processing means includes means for background noise correction. 9. The apparatus of claim 1 where said electronic signal processing means includes means for providing correction for spatial non-linearity of said image intensifier tube. 10. The apparatus of claim 1, wherein an object to be scanned is positionable between the source of x-ray photons and the image intensifier tube, further including pre-object means positionable between the source of x-ray photons and the object being scanned for collimating and filtering the x-ray photons from the source of x-ray photons; and post-object means positioned between the object being scanned and the image intensifier tube for collimating x-ray photons. 11. The apparatus of claim 1 where said linear array includes a main array positioned to receive the visible output being coupled by the optical means, and an extension array positioned to receive x-ray photons. 12. The apparatus of claim 2 wherein the means for linearizing the detector signals from the linear array comprise means for solving a plurality of nth-order polynomials as a function of the detector signals from the linear array, wherein the nth- order polynomials have coefficients which are a function of selected characteristics of the photodiodes in the linear array, and further wherein the different ones of the nth-order polynomials are valid over different detector signal magnitude ranges; and means responsive to the detector signals from the linear array and communicating with the solving means for providing linearized versions of the detector signals from the linear array, wherein the linearized version of any particular detector signal is the solution from a selected one of the plurality of nth-order polynomials, and further wherein the selected polynomial is chosen by matching the magnitude of the detector signal being linearized to the range of magnitudes for which the selected n-th order polynomial is valid. 13. The apparatus of claim 12, wherein the nth- order polynomials are valid for ranges of detector signal magnitudes which overlap, and further wherein the means for providing linearized detector signals interpolates the solutions from polynomials which are valid for overlapping detector signal magnitude ranges when the detector signal being linearized falls within that overlapping range. 14. The apparatus of claim 13, wherein the coefficients of the nth-order polynomials are selected as a best fit of the polynomial to the applied visible intensity spectrum magnitude as a function of the photodiode output signal magnitude, using a least squares curve fitting criteria. 15. The apparatus of claim 14, wherein the range of detector signal magnitudes is from zero counts to 500,000 counts and the means for solving provides solutions for three fourth-order polynomials. 16. The apparatus of claim 15, wherein the zero to 500,000 count range is divided into three ranges, and each of the three fourth-order polynomials is valid for a different one of the three ranges. 17. The apparatus of claim 16, wherein the three ranges overlap. 18. The apparatus of claim 4, wherein the photodiodes of the linear array accumulate charge when the visible intensity spectrum impinges upon them, and further wherein the means for measuring a long and a short sample interval are in communication with the photodiodes of the linear array and control a short time interval and a long time interval over which photodiodes are permitted to accumulate charge. 19. The apparatus of claim 18, wherein the sum of the short and long time intervals for accumulating charge is a predetermined time interval, and further wherein the long time interval is approximately nine times longer than the short time interval. 20. The apparatus of claim 19, wherein the simulator machine is operable to obtain a plurality of scans, and a plurality of projections within each of the plurality of scans, and further wherein all photodiodes in the linear array are each sampled over a long time interval and over a short time interval during each of the plurality of projections. 21. The apparatus of claim 19, wherein the detector signal has an expected range of magnitudes, and further wherein the long time interval samples are selected as the detector signal to be used in constructing the tomographic x-ray image when the long time interval sample magnitude is below a transition range of magnitudes; and further wherein the short time interval samples, multiplied by a scaling factor, are selected as the detector signal to be used in constructing the tomographic x-ray image when the long time interval sample magnitude is above the transition range of magnitudes, and a weighted combination of the long interval samples and the scaled short interval samples are selected as the detector signal to be used in constructing the tomographic x-ray image when the long time interval sample magnitude falls within the transition range of magnitudes. 22. The apparatus of claim 21, wherein the photodiodes of the linear array can be characterized by a series of calibrating signals which are measured using a long sample interval and a short sample interval, and further wherein the scaling factor is determined by comparing the long sample interval calibrating signal measurements to the short sample interval calibrating signal measurements, multiplied by the scaling factor, over a predetermined range of magnitudes, and adjusting the scaling factor to obtain the best match using a "least squares" curve fitting criteria. 23. The apparatus of claim 22, wherein the calibrating signals measurements have a range from zero counts to 500,000 counts, and the predetermined range of magnitudes over which the long and short sample interval calibrating signal measurements are compared is approximately 44,000 to 62,000 counts. 24. The apparatus of claim 23, wherein the scaling factor is approximately equal to nine. 25. The apparatus of claim 5, wherein the detector signal from the photodiodes of the linear array is in the form of a quantity of charge, and further wherein the low-noise wide-dynamic range analog electronics includes means for integrating the quantity of charge in a detector signal, wherein the integrating means has low noise and high input impedance; and means for resetting the integrating means in preparation for receipt of a subsequent detector signal. 26. The apparatus of claim 25, wherein the means for integrating includes a low noise, high input-impedance amplifier having an output, an inverting input to which the detector signal is applied, and a non- inverting input which is coupled to a system common point; capacitor means coupled between the output and the inverting input of the amplifier for accumulating the quantity of charge in the applied detector signal; and further wherein the resetting means includes first switch means coupled in parallel with the capacitor means for controllably discharging the capacitor means in preparation for receipt of the next detector signal; and second switch means coupled between the output of the amplifier and the system common point for clamping the output of the amplifier to system common whenever the first switch means is controlled to discharge the capacitor means. 27. The apparatus of claim 26, wherein the output of the amplifier is digitized by analog to digital converter means. 28. The apparatus of claim 27, wherein the low- noise wide-dynamic-range analog electronics include means for phase locking the operation of the integrating means, the resetting means, and the analog to digital converter means to a designated reference signal. 29. The apparatus of claim 28, wherein the radiation therapy simulator machine is powered by a line voltage, and the designated reference signal is the line voltage. 30. The apparatus of claim 11, wherein extension detector array comprises a plurality of detectors, each of which comprises a body of scintillating crystal material providing visible light photons over a range of the visible light spectrum in proportion to a number of impinging x-ray photons; and a photodiode optically coupled to the body, having a response characteristic which is compatible with the range of the visible light spectrum provided by the body, and providing a current which is proportional to the visible light photons received from the body. 31. The apparatus of claim 30 wherein said signal processing means includes means for linearizing the detector signals from the linear array. 32. The apparatus of claim 31 wherein said signal processing means includes means for accommodating a wide range of detector signal magnitudes from the linear array. 33. The apparatus of claim 32 wherein said means for accommodating a wide range of detector signal magnitudes includes means for measuring the detector signal over a long sample interval and a short sample interval. 34. The apparatus of claim 30 where said electronic signal processing means includes low-noise wide-dynamic-range analog electronics. 35. The apparatus of claim 34 including digitizing means for providing an effective 19 bit signal representative of the conditioned detector signal. 36. The apparatus of claim 34 where said digitizing means includes a 16 bit analog-to-digital converter and circuitry for combining 16 bits of data from the 16 bit analog-to-digital converter representing the conditioned detector signal, with 3 additional bits from a gantry angle encoder. 37. The apparatus of claim 30 where said electronic signal processing means includes means for background noise correction. 38. The apparatus of claim 1, wherein optical means comprise lens means for receiving the visible light photons from the intensifier tube and projecting the visible light onto the linear array with an overall selectable f-stop number. 39. The apparatus of claim 38, wherein the lens means comprise a first lens, focused at infinity and receiving the visible light photons from the intensifier tube, for projecting the visible light photons at a first selectable f-stop number; a mirror positioned to receive the projected visible light photons from the first lens and to deflect the received light photons at a first predetermined angle; and a second lens, positioned to receive the deflected light photons from the mirror at an infinity focus, for focusing the received light photons onto the linear array at a second selectable f-stop number, wherein the combined first and second f-stop numbers equal the overall f-stop number. 40. The apparatus of claim 39, wherein the mirror is repositionable to a second predetermined angle different from the first predetermined angle. 41. The apparatus of claim 1, wherein the image intensifier tube has a dynamic range of at least 100,000:1, and the processing means has a dynamic range of at least 100,000:1. 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