Multi-element-amorphous-silicon-detector-array for real-time imaging and dosimetry of megavoltage photons and diagnostic X rays

A multi-element-amorphous-silicon detector-array real-time imager and dosimeter for diagnostic or megavoltage X rays having megavoltage photons having a plurality of photodiodes made of hydrogenated amorphous silicon arrayed in columns and rows upon a glass substrate. Each photodiode is connected to a thin film field effect transistor also located upon the glass substrate. Upper and lower metal contacts are located below and above the photodiodes to provide the photodiodes with a reverse bias. The capacitance of each photodiode when multiplied by the resistance of the field effect transistor to which it is connected yields an RC time constant sufficiently small to allow real time imaging.

BACKGROUND OF THE INVENTION 
1. Field of the Invention 
The present invention pertains generally to the field of radiation 
detecting devices and, more particularly, to the field of real-time 
radiation imaging devices. 
2. Discussion of the Background 
There are several instances in modern radiotherapy where real-time imaging 
of X rays is a highly useful and critically important technique. 
In external beam megavoltage photon radiation therapy, it is highly 
desirable that the maximum dose be delivered to the target volume and the 
minimum dose be delivered to the surrounding tissue. Prior to treatment, 
which typically consists of irradiating the patient on a daily basis for 
several weeks, the patient undergoes a number of preparatory steps in 
order to identify the region to be irradiated and to determine a 
"treatment plan" specifying exactly how this irradiation is to be 
performed. Often, one of these steps is to place the patient on a 
"treatment simulator", which simulates the motions and geometry of the 
therapy machine, and which makes diagnostic quality fluoroscopic and 
radiographic x-ray images. The fluoroscopic imaging allows a real-time 
means of simultaneously observing patient anatomy and manipulating the 
position of the patient so as to achieve a desired patient orientation 
with respect to the simulated treatment beam. Unfortunately, current 
fluoroscopic devices use large cumbersome image-intensifier tubes which 
restrict the possible motions of the simulator, thereby limiting the 
treatment positions that can be simulated. 
A permanent record of imaging information from the simulator is achieved by 
means of radiographic imaging with film as well as storage of the 
fluoroscopic images. These images are used to provide information which 
contributes to deciding what the target region should be and how the 
actual treatment is performed, i.e., what geometric and dosimetric 
combination of megavoltage beams to use to satisfactorily irradiate the 
target region but spare the surrounding normal tissues. Once a treatment 
plan has been determined, often with the assistance of a computer which 
allows, among other things, manipulation of the simulation information as 
well as CT or other imaging information, the patient is typically taken 
back to the simulator for a verification-simulation in order to verify the 
geometric correctness of the plan. 
When the patient is brought into the treatment room, it is highly 
desirable, prior to treatment, to verify that the orientation of the 
patient with respect to the treatment beam closely coincides with the 
setup achieved in the simulator room. Once verified, the prescription dose 
can be delivered to the target volume and surrounding tissues. The 
achievement of this goal is complicated by the fact that the patient 
anatomy moves due to both voluntary and involuntary patient motions. Such 
complications encourage the possibility of delivering too little dose to 
the target region and/or overdosing the surrounding tissues. In addition, 
for treatment machines which use a computer-controlled scanning treatment 
beam, there is the additional uncertainty of whether the beam is correctly 
directed on a burst by burst basis. 
The above problems can be overcome by real-time imaging. Several prototype 
real-time imagers are being developed around the world, but most have no 
practical applications to clinical use. A real-time clinical image 
detector has been developed by H. Meertens at the Netherlands Cancer 
Institute in Amsterdam which is disclosed in European Patent Application 
0196138. The Meertens' device operates on the principle of a scanning 
liquid ionization chamber. However, the Meertens' device is able to detect 
only a small fraction of the imaging signal. 
Radiation detecting devices are taught in Hynecek, U.S. Pat. No. 4,679,212; 
Luderer et al., U.S. Pat. No. 4,250,385; DiBianca, U.S. Pat. No. 
4,707,608; Haque, U.S. Pat. No. 4,288,264; Kruger, U.S. Pat. No. Re. 
32,164; Barnes, U.S. Pat. No. 4,626,688; and DiBianca et al., U.S. Pat. 
No. 4,525,628; however, these detectors do not make possible real-time 
imaging for megavoltage photons. 
Imaging equipment has been developed based on the use of photostimulable 
phosphors wherein an image receptor plate coated with such phosphors is 
exposed to a radiation beam and then "read out" by means of laser 
stimulated luminescence with direct conversion of the light to digital 
form. However, this technology appears to offer no possibility of 
real-time imaging. Efforts to develop imagers based on camera-fluoroscopy 
combinations have produced images of greatly varying quality at rates 
ranging from two images a second to one image every eight seconds. 
However, such a camera's expensive and delicate imaging electronics would 
be irreversibly damaged after approximately 10-130 kilorads of dose. 
Thus, mirrors are used to reflect the light image produced by a 
metal-phosphor screen combination to a camera sitting outside of the 
direct radiation field. This makes necessary the presence of a bulky light 
box located in the vicinity of the treatment table where such obstructions 
are highly undesirable. Furthermore, with the camera's imaging surface 2 
to 3 feet from the fluorescent screen, the solid angle subtended by the 
camera is small (less than 1%) and hence the image quality is limited by 
the light collection stage rather than by the available high-energy 
quanta. 
Recently, an imager consisting of tightly packed, tapered, optical fibers 
has been reported. The fibers make up a 40.times.40 cm.sup.2 surface, 12 
cm thick, which sits in the beam behind a metal-fluorescence screen and 
"pipes" the light to a video camera. The optical fibers are bunched 
together in bundles of 1.5.times.1.5 cm.sup.2 at the input end and the 
imager has a thickness of 12 cm. The optical fibers have to be bent to 
such an extent that light is lost due to the fact that the critical angle 
is exceeded. The system currently has a light collection efficiency no 
greater than that of the mirror-camera system and, like those systems, is 
rather bulky. 
In the optical imaging systems discussed above, considerably less than 1% 
of the visible light photons emitted by the scintillating layer are 
converted into signal. As a direct consequence, the quantum sink is the 
light collection stage rather than the stage where X rays are converted to 
high-energy electrons which enter into the phosphor. Thus, the quality and 
speed of imaging in the above systems are adversely affected. 
In selecting the materials for a real-time imager for megavoltage photon 
radiation therapy, care must be taken that the materials can withstand 
high levels of radiation exposure over long durations of time. Another 
consideration is that the radiation detecting elements be arranged over a 
relatively large surface area. For instance, a detection surface of at 
least 25.times.25 cm.sup.2 is necessary for head and neck portals. For 
pelvic, abdominal and thoracic portals, a surface area of 50.times.50 
cm.sup.2 is desirable. Though solid state imagers are highly desirable, 
the manufacture of crystalline semiconductor detectors over such an area 
is certainly prohibitively expensive. 
The development of a--Si:H (hydrogenated amorphous silicon) has resulted in 
the realization of a highly radiation resistant material which can be 
utilized over large surface areas at very economical cost. See V. 
Perez-Mendez, et al., "Signal, Recombination Effects and Noise in 
Amorphous Silicon Detectors", Nuclear Instrument and Methods in Physics 
Research A260 (1987) 195-200, Elsevier Science Publishers B. V.; and I. D. 
French et al., "The Effect of .gamma.-Irradiation on Amorphous Silicon 
Field Effect Transistors", Applied Physics A31, 19-22, 1983, 
Springer-Verlag. 
It is now realized that amorphous silicon thin film transistors have 
applications to large-area electronics, see H. C. Tuan, "Amorphous Silicon 
Thin Film Transistor and its Application to Large-Area Electronics," Mat. 
Res. Soc. Symp. Proc. Vol. 33 (1984) Elsevier Science Publishing Company, 
Inc. 
Amorphous silicon ionizing particle detectors made of hydrogenated 
amorphous silicon are known which can detect the presence, position and 
amount of high energy ionizing particles, see Street et al, U.S. Pat. No. 
4,785,186; however, the patent does not teach how a--Si:H photodiodes can 
be utilized in coordination with other elements to obtain a real-time 
imaging device. 
Rougeot, U.S. Pat. No. 4,799,094, teaches a photosensitive device having an 
array of p-doped floating grids which connect with a substrate of lightly 
n-doped hydrogenated amorphous silicon. Since Rougeot uses transistors as 
light detectors, the quantity of electron-hole pairs generated would 
appear quite insufficient to realize real-time imaging. 
SUMMARY OF THE INVENTION 
Accordingly, one object of the present invention is to achieve a 
pulse-by-pulse monitoring of the centroid of a transmitted megavoltage 
radiation beam. 
Another object is to obtain x-ray-like images of the patient for real-time 
portal localization and verification imaging using megavoltage beams. 
Yet another object is to measure the transmitted dose of the radiation 
field. 
A further object is to obtain real-time diagnostic quality images with a 
treatment simulator with far less restrictions of motion due to the 
presence of the imager. 
These and other objects are achieved by providing a real-time imaging 
device for storing and retrieving an imaging signal resulting from 
megavoltage radiation bursts in radiation treatment machines or from a 
diagnostic x-ray generator. The real-time megavoltage imaging device has a 
conversion layer for converting photons from the megavoltage radiation 
bursts into electrons, a scintillation layer in which the electrons 
created in the conversion layer create visible light photons, and an upper 
electrode layer which is transparent and allows the visible light from the 
scintillation layer to pass through. A plurality of a--Si:H sensors is 
arranged in rows and columns to form a radiation detecting surface area. 
Each a--Si:H sensor is paired with a thin film field effect transistor. 
The RC time constant of the imaging device is calculated by multiplying 
the resistance of the transistor by the capacitance of the a--Si:H sensor. 
The high energy electrons and visible light photons incident upon the 
sensors constitute an imaging signal. Sufficient amounts of this imaging 
signal can be detected, creating electron-hole pairs which are stored in 
the capacitance of the sensors. Thereafter this signal can be read out on 
a time scale determined by the RC time constant so that real-time imaging 
is made possible. 
For the application of real-time imaging of a diagnostic x-ray beam, given 
that the interaction probabilities are considerably higher and the range 
of the electrons created by the photons much shorter, a single layer for 
diagnostic-energy photon conversion and creation of the light photons by 
the resulting electrons suffices. 
In addition, given the thinness and uniformity of the amorphous silicon 
sensors and the substrates upon which they are deposited, it is 
furthermore possible to stack one imager upon another without significant 
degradation of the imaging information. For example, an array dedicated to 
determining the position of the scanning beam could be positioned under an 
array dedicated to imaging the megavoltage photon beam. Alternatively, an 
array dedicated to real-time diagnostic x-ray imaging could be positioned 
over an array dedicated to real-time megavoltage imaging. Such positioning 
would offer tremendous advantages to radiation therapy.

DESCRIPTION OF THE PREFERRED EMBODIMENTS 
Referring now to the drawings, wherein like reference numerals designate 
identical or corresponding parts throughout the several views and, more 
particularly, to FIG. 1 thereof, a sensor 30 and thin-film field effect 
transistor 52 are shown mounted upon a glass substrate 12. The gate 
contact region 14 of the thin-film field effect transistor is seen to be 
positioned atop the glass substrate 12. Surrounding the gate contact 
region is a gate dielectric layer 16 of silicon nitride Si.sub.3 N.sub.4 
which also makes contact with the glass substrate 12. Above gate 
dielectric layer 14 is an a--Si:H layer 18. 
Directly above the gate contact region 14 and making contact with the 
a--Si:H layer 18 is a second gate dielectric layer 24 made of silicon 
nitride. Adjacent to the sides of the lower portion of the second gate 
dielectric layer 24 are located n+ doped channel layers 25, 29 which are 
positioned so as to sandwich the lower portions of the second gate 
dielectric layer 24 therebetween. A drain contact region 26 and a source 
contact region 28 are positioned so as to sandwich the upper portion of 
the second gate dielectric layer 24 between them. These structured layers 
above and to the side of the gate contact region constitute a thin-film 
field effect transistor 52. Alternatively, microcrystalline silicon 
thin-film-transistors can be implemented to achieve faster read-out-speed. 
This thin-film field effect transistor 52 is connected to a sensor 30 which 
constitutes a p-i-n photodiode. The sensor 30 is connected to the source 
contact region 28 of the thin-film field effect transistor 52 by means of 
lower electrode layer 22. 
Above lower electrode layer 22 is p+ doped a--Si:H layer 36, layer 36 being 
approximately 400 .ANG. thick. Above layer 36 is an intrinsic a--Si:H 
layer 34 which is at least one micron thick and preferably being 
approximately 3 microns or more thick for reasons which will be 
subsequently discussed. Above the intrinsic layer 34 is an n+ doped layer 
of a--Si:H 32, which is approximately 100 .ANG. thick. 
Above n+ doped layer 32 lies the upper electrode 38 which is made of a 
material transparent to visible light. A material such as indium tin oxide 
(ITO) is a suitable material for the upper electrode 38. A scintillating 
layer 44 for purposes of converting electrons to visible light is located 
directly above and preferably makes contact with the transparent upper 
electrode 38. Scintillating layer 44 may be a CRONEX.TM. scintillating 
screen made by Dupont, a LANEX.TM. scintillating screen made by Kodak or 
other suitable material. 
For megavoltage beams, a photon-to-electron conversion layer 46 is located 
directly above and makes contact with scintillating layer 44. Conversion 
layer 46 is a 1 millimeter thick copper sheet; however, the thickness can 
be varied depending on the energy level of the radiation to which it is 
exposed. A 1 mm thick copper sheet when combined with a CRONEX.TM. 
scintillating screen and exposed to a megavoltage (.about.3-50 MV) photon 
beam produces pulses of light for about 10 microseconds. A 1 mm thick 
copper sheet when combined with a LANEX.TM. scintillating screen produces 
pulses of light for about 1 millisecond when exposed to a megavoltage 
beam. 
Optionally, a suitable phosphor material may be deposited directly onto the 
sensors or onto the photon-to-electron converter in order to optimize the 
spatial resolution of the imager. For diagnostic imaging, the interaction 
of X rays with the scintillation layer 44 is such that the conversion 
layer 46 is not necessary. 
As is shown in FIG. 1, polyimide 42 is placed over the field effect 
transistor 52 and between sensors 30 to provide insulation. 
FIG. 2 shows an array of sensors 50 according to the present invention. 
Biasing lines 40 are seen to connect the sensors 30 by being connected 
across the upper electrode layer 38 of each sensor in a given row. As will 
be appreciated by those skilled in the art, a metal layer (not shown), 
e.g. of aluminum, is coincident with those regions of biasing lines 40 
which are located between the sensors 30. Such a metal layer serves to 
shield the transistors 52 from light. Signal lines 54 are shown to connect 
to the drain of each field effect transistor 52 in a given row. 
From FIG. 1 it can be seen that a drain electrode layer 23 extends away 
from a side of drain contact region 26. Metallization (not shown) extends 
vertically downward from the end 27 of the drain electrode layer 23 away 
from transistor 52. This metallization is constructed to connect with a 
signal line 54 which lies on the substrate 12 in front of the sensor 30. 
Gate select lines 56 (FIG. 2) are seen to connect the gates of the 
thin-film field effect transistors located in a given column. 
Thus, the sensors and thin-film field effect transistors are arranged in 
rows and columns so as to form an array 50 mounted upon a glass substrate 
12 and form a detection panel. 
A 25.6.times.25.6 cm.sup.2 imaging panel is made of approximately 65,536 or 
more individual sensors. A typical imaging panel will have a 256.times.256 
array of sensors or more. Each sensor is approximately 0.9 millimeters 
long or smaller with a density of at least 1 sensor per square millimeter 
being desired. Four panels of the 25.6.times.25.6 cm.sup.2 sensor arrays 
can be combined to form a roughly 50.times.50 cm.sup.2 surface so that the 
invention can be utilized for virtually any imaging function. 
FIG. 4 shows the layout of the read-out electronics. Each gate select line 
56 is addressed sequentially with a shift register. The signal lines 54 
are connected to an amplifier and switching electronics. Either a charge 
or voltage amplifier may be used. If it is the latter, then a capacitor is 
included on the array at each signal line. The charge or voltage is sensed 
just after the end of each gate pulse, after which the signal line is 
reset to ground potential before the next gate pulse. 
As shown in FIG. 1, a radiation beam 10 is directed upon the 
photon-to-electron conversion layer 46 which converts the photons of the 
radiation beam to electrons, some of whose energy is absorbed in 
scintillating layer 44 and thereby converted to visible light. This 
visible light passes through the transparent upper electrode 38 and into 
the sensor 30 where electron-hole pairs are generated in the intrinsic 
layer 34. The present invention allows 70 to 95 percent of the visible 
light photons emitted from scintillating layer 44 to be converted to 
electron-hole pairs in the sensors. The sensor 30 has a capacitive effect 
when a reverse bias is applied to all of the sensors by means of the 
biasing lines 40. This reverse bias causes the electron-hole pairs to be 
attracted to the upper and lower electrodes 38 and 22 where the signal 
generated from the radiation bursts are thus stored. 
The thicker the intrinsic layer 34 is, the greater the ionization caused by 
the incident high energy and visible light radiation. However, as the 
thickness of layer 34 increases, the applied bias must increase 
proportionally in order to ensure that the electron-hole pairs generated 
throughout the intrinsic layer are efficiently collected and positioned at 
the upper and lower electrodes. A thickness of 1 micron is sufficient to 
absorb all the visible light from the phosphor. The choice of thickness is 
determined in part by the capacitance requirements of the sensor. 
The size of the signals stored by the capacitance of the sensors 
constitutes the information from which the images are produced. The speed 
at which the size of these signals may be sampled is characterized by a 
time constant, RC, which is given by the product of the sensor capacitance 
and the on-resistance of the transistor. In general, it is highly 
desirable to design the sensors and transistors so that this RC time 
constant is as small as possible since the contribution of noise to the 
signal, collected by interfacing electronics located off of the array, 
will be minimized when the sampling period is as short as possible. A 
second constraint upon the sampling period for certain applications of 
radiation therapy is the requirement that the array be capable of being 
read out after every burst of radiation. This latter constraint demands 
that the RC time constant be less than the time interval between radiation 
pulses divided by the number of columns in the array. The former 
considerations of noise would encourage even further reductions. 
The inverse of the on resistance of the transistor at a gate voltage 
V.sub.G, is given by: 
EQU 1/R=(W/L).mu.FE(V.sub.G -V.sub.T)C.sub.G 
where W and L are the width and length of the transistor, .mu.FE 
(.about.0.5 cm.sup.2 /Vsec) is the carrier mobility, V.sub.T (=1V) is the 
threshold voltage and C.sub.g (.about.10.sup.-8 F/cm.sup.2) is the gate 
capacitance. Typically, W/L is 10-100 so that R.about.0.1-1 Mohm. If the 
sensor capacitance is 50 pF, then the RC time constant can be as small as 
5 microseconds, thus satisfying the speed requirements. 
At present, a column of sensors can be read out approximately every 10 
microseconds. Therefore to read 256 columns would take 2.56 milliseconds. 
If the interval between radiation bursts is set at 2.56 milliseconds, this 
would give sufficient time to read out 256 columns in between radiation 
bursts. Thus, the time interval between radiation bursts, if set at 2.56 
milliseconds, when divided by 256 rows would yield a time of 10 
microseconds. Since the signal decays in an exponential fashion, the 
majority of the signal information is obtained during one time constant. 
Therefore, for this example, the desired time constant for the imaging 
device would be less than 10 microseconds. 
The capacitance of each sensor and the resistance of each transistor should 
be designed so that the product of their respective capacitances and 
resistances is equal to the time constant desired. With the signal 
information stored in the sensor being adequately sampled during a time 
scale determined by the time constant, real time imaging is made possible. 
The intrinsic layer 34 must be at least 1 micron thick or greater for 
purposes of converting and collecting photons over virtually the entire 
visible spectrum. As the intrinsic layer increases in thickness, 
capacitance diminishes and the ionization signal from high energy quanta 
increases. Furthermore, as the aspect ratio (W/L) of the thin-film field 
effect transistor is increased, resistance is reduced. Thus, by widening 
the channel of the field effect transistor, an increase in the aspect 
ratio and consequently a decrease in resistance will be realized. Also, 
resistance can be reduced by increasing the bias received by the gate of 
the field effect transistor. 
Once the signal is stored in the sensors, obtaining the signal is 
relatively easy. 
By applying a biasing voltage to the gate select lines the signal stored 
due to the capacitance of the sensor 30 is released from the source region 
to the drain region of the field effect transistor and is channeled 
through the signal lines 54 and on to interfacing electronics. 
The ultimate limitation to the signal-to-noise ratio is the quantum noise 
of (i.e., statistical variations in) the number of high energy photons 
converting to high energy electrons which deposit energy in the phosphor, 
and an ideal detector will introduce no significant additional noise. For 
megavoltage beams, a typical gamma-ray burst will present approximately 
5.times.10.sup.5 photons/mm-size pixel, of which approximately 1% or 
5.times.10.sup.3 are converted into electrons producing visible light in 
the phosphor. This results because of absorption in the patient and loss 
in the copper/phosphor converter. A signal-to-noise ratio of 1000:1 
requires 10.sup.6 of such converted photons for the quantum noise limit, 
and so requires approximately 200 bursts. Thus a high contrast image 
requires integration of the signal for 0.4-1 sec depending upon the dose 
rate and burst repetition frequency of the therapy machine. Such 
integration can be achieved either in the host computer or on the array. 
The inventors find that charge may be held in the array without 
significant loss for at least 1 second. 
The converted gammas from a single burst yields approximately 10.sup.7 
visible photons detected in each mm.sup.2 sensor of the array. The quantum 
noise of this signal is dominated by that of the converted gammas. The 
measured noise in the read-out of a single element in the array is 
10.sup.4 -10.sup.5 electrons. Therefore, no significant extra noise is 
introduced by the array, which will be capable of achieving the ultimate 
possible signal-to-noise ratio even when read out at every burst. Slower 
read out reduces even further the effect of read-out noise. 
As it is essential that the sensors and their electronics be adequately 
shielded from stray electromagnetic radiation, including radio frequency 
and ionizing, FIG. 3 shows a shielded housing 60 made of copper in which a 
panel of sensors 50 is enclosed. In the case of megavoltage imaging, the 
top of the shielded housing 60 is seen to comprise the photon to electron 
conversion layer 46 and scintillating layer 44. In the case of diagnostic 
x-ray imaging, the top shielded housing is a layer which acts as both the 
converter and the scintillator. In the case where it is desirable to 
combine several such images, such as one for megavoltage imaging and 
another for burst-by-burst determination of the centroid of a scanning 
beam, the arrays for these imagers would be stacked inside of the shielded 
housing along with their photon-electron converters and scintillating 
screens. 
FIG. 4 shows the array with known interfacing electronics. FIG. 4 serves to 
demonstrate how the gate select lines 56 can be activated by means of gate 
select electronics 70 connected to a microprocessor 72 which is connected 
to terminal 74. The signal lines 54 are seen to be interfaced with analog 
to digital converter 76 which is connected to microprocessor 72 and video 
monitor 78. 
FIG. 5 shows a radiation machine 80 and a patient 66 lying on table 82 
receiving treatment from the radiation beam 10. The shielded housing 60 
enclosing the array of sensors 50 is seen to lie below the patient 
underneath table 82. The arrangement would be similar in the case of a 
diagnostic imager located in the simulator room. 
The discussion which follows is intended to give the reader a firm 
understanding of what the term "real-time" means in regard to the present 
invention. 
If the invention is to be used for determining the centroid of a scanning 
megavoltage beam on a pulse-by-pulse basis then real-time operation 
requires that a very large fraction, preferably all, of the sensors be 
read out between bursts. This mode of operation is desirable with a 
scanning megavoltage beam machine. Such a machine typically has a variable 
pulse repetition rate ranging from 60 to 500 hertz. Thus, there are 2 to 
16.7 milliseconds between bursts. The speed at which a given row from the 
array must be read out to satisfy this real-time requirement will depend 
upon the number of rows per array and the pulse repetition rate. 
If the invention is to be used for imaging the megavoltage therapy beam, 
then there are two distinct modes, localization and verification imaging, 
in which real-time operation is required. In the case of localization 
imaging, just prior to the treatment, it is desired to give the patient a 
small fraction of his daily dose adequate to provide sufficient high 
energy photons to form an image. In this case, the signals stored in the 
sensors would be allowed to accumulate until the termination of the 
irradiation, at which time the sensors would be read out. The state of 
knowledge of the megavoltage beams indicates that acceptable images should 
be possible with the invention after periods of 0.1 seconds to several 
seconds depending upon the imaging situation and desired contrast. 
Real-time imaging would certainly be achieved in this instance if the 
final picture were available within several seconds or less after 
irradiation. As has been explained, it is essential that each column of 
sensors be read out as quickly as possible. 
In a second real-time imaging situation, it is desirable to produce images 
one after the other during the course of a treatment. Given that the 
treatment may last .about.10 to 60 seconds and given that there is imaging 
information after 0.1 seconds to several seconds, real-time operation in 
this case demands that the imager be read out as quickly as possible after 
sufficient information has accumulated in the sensors. 
In the case of radiographic imaging, as in localization imaging, the goal 
is to irradiate the patient sufficiently so as to produce a high quality 
image. The present invention achieves real-time operation by allowing a 
final picture to be available in several seconds or less. 
With regards to the RC time constant, the invention has been designed so 
that the columns of sensors can be read out as quickly as possible. This 
is a consequence of the fact that the external electronics which sample 
the signals from the sensors also sample noise from various sources, and 
this noise contribution increases with increasing sampling periods. Hence, 
it is highly desirable to keep this noise to a minimum by reducing the 
period during which the charge on the sensors is sampled. As has been 
mentioned, a major determinant of the speed at which this sampling can 
occur is given by the capacitance of the sensor times the on-resistance of 
the thin-film-transistor. Thus, by keeping the RC time constant to a 
minimum, the present invention achieves real-time imaging with a superior 
signal-to-noise ratio. 
The present invention makes possible the detection of the centroid of the 
megavoltage radiation beam as many times per second as there are radiation 
bursts and the determination of the transmitted radiation dose on a burst 
by burst basis. Further, the present invention verifies the radiation dose 
is directed upon the desired target area. The present invention achieves a 
superior signal-to-noise ratio and receives enough information for an 
image to be formed in 1/10 of a second, (10 images a second), the only 
limitation being the speed of the processing hardware. 
The present invention may be used for years at a time without a degradation 
in performance due to continued exposure to megavoltage radiation. When 
there is some radiation damage to the array, a simple heat treatment at 
130.degree.-150.degree. C. restores the original characteristics of the 
device. 
The present invention allows the replacement of the bulky image intensifier 
tube with a thin imaging system whose profile offers minimal obstruction 
to the motions of the simulator. 
Finally, the present invention allows the creation of combinations of 
imagers which are stacked one on top of the other. The various imagers in 
the stack may be optimized for various forms of imaging. 
Obviously, numerous modifications and variations of the present invention 
are possible in light of the above teachings. It is therefore to be 
understood that within the scope of the appended claims, the invention may 
be practiced otherwise than as specifically described herein.