U.S. patent application number 12/366088 was filed with the patent office on 2009-07-02 for shielding an imaging array from x-ray noise.
This patent application is currently assigned to SCHICK TECHNOLOGIES, INC.. Invention is credited to Stan Mandelkern, Barmak Mansoorian, David Schick, Daniel Van Blerkom.
Application Number | 20090166545 12/366088 |
Document ID | / |
Family ID | 37910338 |
Filed Date | 2009-07-02 |
United States Patent
Application |
20090166545 |
Kind Code |
A1 |
Mandelkern; Stan ; et
al. |
July 2, 2009 |
SHIELDING AN IMAGING ARRAY FROM X-RAY NOISE
Abstract
An electronic imaging sensor, which has an improved immunity to
noise caused by unwanted x-rays, images an object by collecting
charge carriers produced in the sensor when the object is exposed
to x-rays. One or more shielding areas are formed proximate the
sensor to capture or sweep away any undesirable charge carriers
generated by the unwanted x-rays. The shielding areas extend deeper
beneath the surface of the sensor than the depth at which the
desired charge carriers corresponding to the object being imaged is
collected. The shielding areas capture charge carriers formed by
the unwanted x-rays, which penetrate into the sensor to a greater
depth than the depth at which the desired charge carriers are
collected. In this way, the undesirable charge carriers are
captured near the region where they are generated and before they
migrate towards the surface where they can be collected and
manifest as noise in the resulting image of the object.
Inventors: |
Mandelkern; Stan; (Teaneck,
NJ) ; Schick; David; (Kew Garden Hills, NY) ;
Mansoorian; Barmak; (San Diego, CA) ; Van Blerkom;
Daniel; (Altadena, CA) |
Correspondence
Address: |
FITZPATRICK CELLA HARPER & SCINTO
30 ROCKEFELLER PLAZA
NEW YORK
NY
10112
US
|
Assignee: |
SCHICK TECHNOLOGIES, INC.
Long Island City
NY
|
Family ID: |
37910338 |
Appl. No.: |
12/366088 |
Filed: |
February 5, 2009 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
|
11245519 |
Oct 7, 2005 |
7501631 |
|
|
12366088 |
|
|
|
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Current U.S.
Class: |
250/370.09 |
Current CPC
Class: |
H04N 5/32 20130101; G01T
1/2018 20130101; H01L 27/14663 20130101; G01T 1/24 20130101; H04N
5/357 20130101 |
Class at
Publication: |
250/370.09 |
International
Class: |
G01T 1/24 20060101
G01T001/24 |
Claims
1.-8. (canceled)
9. A radiation detector, comprising: a photoconductive portion that
generates charge carriers when irradiated with x-rays; a
semiconductor substrate incorporating dopants of a first
conductivity type; a sensing region positioned at a first depth
below a surface of the substrate to collect the charge carriers
generated in the photoconductive portion; a shielding portion
formed in the substrate and incorporating dopants of a second
conductivity type, wherein a shielding junction located at an
interface between the shielding portion and the substrate and
positioned at a second depth that is deeper below the surface of
the substrate than the first depth collects charge carriers
generated in the substrate by stray x-rays that pass through the
photoconductive portion into the substrate.
10. A radiation detector according to claim 9, further comprising a
plurality of shielding portions arranged around the sensing
region.
11. A radiation detector according to claim 9, wherein the
photoconductive portion is comprised of selenium, lead iodide or
mercuric iodide.
12. A radiation detector according to claim 9, wherein the depth of
the shielding junction correlates substantially with a penetration
depth of the stray x-rays in the substrate.
13. A radiation detector according to claim 9, wherein the
radiation detector is incorporated in a dental imaging device.
14. A radiation detector according to claim 9, wherein the
radiation detector is incorporated in a medical imaging device.
15. (canceled)
16. A radiation detector, comprising: photoconductor means for
generating charge carriers when irradiated with x-rays;
semiconductor substrate means incorporating dopants of a first
conductivity type; sensing means for collecting the charge carriers
generated by the photoconductor means, wherein the sensing means is
positioned at a first depth below a surface of the substrate means;
shielding means for collecting charge carriers generated in the
substrate means by stray x-rays that pass through the
photoconductor means into the substrate means, wherein the
shielding means incorporates dopants of a second conductivity type
and is formed in the substrate means such that an interface between
the shielding means and the substrate means is positioned at a
second depth that is deeper below the surface of the substrate than
the first depth.
Description
BACKGROUND OF THE INVENTION
[0001] 1. Field of the Invention
[0002] The present invention relates generally to the field of
radiation-sensitive imaging devices, and more specifically to
radiation-sensitive imaging devices that have improved immunity to
x-ray noise.
[0003] 2. Related Background Art
[0004] Increasingly, electronic imaging sensors have been replacing
film-based sensors in commercial, industrial, and medical imaging
applications. Examples of such electronic sensors include
charge-coupled device (CCD) sensors and complementary
metallic-oxide-semiconductor (CMOS) sensors, to name a few. CMOS
sensors, in particular, have emerged as a preferred candidate due
to advantages in manufacturing cost, integration of components,
charge efficiency, and low power consumption. An excellent example
of a detector system suitable for medical imaging applications that
uses a CMOS active pixel sensor (APS) array is provided in U.S.
Pat. No. 5,912,942 to David B. Schick et al., assigned to the
assignee of the present patent application. The Schick '942 patent
is incorporated herein by reference.
[0005] In the system of the Schick '942 patent, as in many medical
imaging applications, radiation or energy from x-ray photons,
commonly referred to as x-rays, is projected through a patient and
must be registered or detected by a sensor. However, conventional
sensors or imaging chips, which generally are fabricated from
silicon, are considerably more sensitive to photon energy in the
visible spectrum than to x-rays. Thus, a scintillator is disposed
on the sensors to convert the energy from the x-rays to visible
light. The scintillator is typically composed of gadolinium
oxysulphide or cesium iodide, although other alternative materials
may be used.
[0006] While it is desirous that the scintillator convert all of
the x-rays to visible light, in practice only a percentage is
actually converted, with the remaining x-rays passing through the
scintillator and reaching the silicon. A typical gadolinium
oxysulphide scintillator, for example, may have a so-called
stopping efficiency of 20-30%, meaning that only 20-30% of the
x-rays that impinge on the scintillator are converted to visible
light, with a significant amount of the x-rays (some 70-80%) being
transmitted through the scintillator into the surface of the
sensor. Although the sensor may be designed to be primarily
sensitive to visible light, it will nevertheless be secondarily
sensitive to effects from the transmitted x-rays. As a result, the
transmitted x-rays are registered as noise by the sensor, reducing
the overall quality of the captured image.
[0007] To reduce such noise, conventional systems typically shield
the sensor from transmitted (unconverted) x-rays. In the Schick
'942 patent, for example, such shielding is achieved by interposing
a fiber-optic plate (FOP) between the scintillator and the sensor.
The FOP allows light to pass, but blocks (i.e., absorbs) a large
amount of the unconverted x-rays. While generally good for its
intended application, this shielding approach has the drawback of
adding an undesirable thickness to the sensor, which compromises
patient comfort. FOPs also cause some degree of light signal loss
and light spreading, each of which reduces image quality. Also, as
FOP sizes increase, they become extremely expensive and fragile.
Accordingly, for many reasons, it is desirable to avoid using
FOPs.
[0008] Recently, advanced x-ray-sensitive photoconductive
materials, such as selenium, lead iodide (PbI.sub.2), and mercuric
iodide (HgI.sub.2), for example, have allowed designers to produce
sensors that can image x-rays directly, so that a conventional
scintillator is not needed. When x-rays strike the surface of such
a photoconductive material, electron-hole pairs are formed. These
charges are driven by an electrical potential or bias potential,
which causes charges of a selected polarity to be collected by a
storage element, such as a capacitor. This alternative design has
the benefit of limiting light-spreading, and may achieve preferred
spatial resolution as compared to typical scintillators. However,
although these photoconductive materials have superior stopping or
conversion efficiency, their efficiency nevertheless is imperfect.
And of course, because x-rays are directly imaged, an FOP generally
is not used in these arrangements. Accordingly, with these
constructions as well, an improved x-ray noise immunity is
desirable.
[0009] Other structures for providing radiation shielding have been
proposed. For example, U.S. Pat. No. 6,690,074 B1 to Dierickx is
aimed at providing a radiation-resistant semiconductor device, and
is particularly concerned with reducing the leakage current between
the source and drain electrodes in a MOS-type structure, resulting
from an overlap between the gate electrode and the field oxide.
Towards this end, the Dierickx device uses a doped guard ring
interrupted by an active area. This approach, however, while
perhaps adequate for its intended purpose of preventing or reducing
leakage currents, which typically occur near the surface of
semiconductor devices, is completely ineffective at shielding from
the deleterious effects of unconverted x-rays, which usually occur
well below the surface.
[0010] There is a great need, therefore, for a semiconductor x-ray
imaging chip that takes an entirely fresh approach, and provides
improved x-ray noise immunity, and at the same time a superior
image quality.
SUMMARY OF THE INVENTION
[0011] The present invention addresses the deficiencies in the
prior art by providing arrangements and designs of an electronic
imaging sensor or device with an improved immunity to noise caused
by unwanted x-rays. The sensor images an object by collecting
charge carriers produced in the sensor when the object is exposed
to radiation such as x-rays. The charge carriers may be generated
indirectly by the radiation, such as with the use of a scintillator
to convert x-rays into visible light, with a photocurrent produced
by the visible light being used to create an image of the object.
Alternatively, the charge carriers may be generated directly by the
radiation, such as with the use of selenium or other types
x-ray-sensitive photoconductive materials that enable a
photocurrent to be generated directly from exposure to x-rays.
[0012] One or more shielding areas are formed proximate the sensor,
to capture or sweep away any undesirable charge carriers generated
by stray or undesired radiation. The shielding areas extend deeper
beneath the surface of the sensor than the depth at which the
photocurrent corresponding to the object being imaged ("the desired
photocurrent") is collected. This is because the undesirable charge
carriers are formed deeper beneath the surface of the sensor.
Typically, the undesired charge carriers are generated by, for
example, unconverted x-rays or x-rays that are not absorbed by the
x-ray-sensitive photoconductive material, which penetrate into the
sensor to a greater depth than the depth at which the desired
photocurrent is collected. In this way, the undesirable charge
carriers are captured near the region where they are generated and
before they migrate towards the surface where they can be collected
and manifest as noise in the resulting image of the object.
[0013] According to an aspect of the invention, a radiation
detector is provided that possesses an improved immunity to x-ray
noise over conventional radiation detectors. The inventive
radiation detector includes:
[0014] a scintillator that converts radiation of a first energy to
radiation of a second energy;
[0015] a semiconductor substrate incorporating dopants of a first
conductivity type;
[0016] a sensing region formed in the substrate and incorporating
dopants of a second conductivity type, wherein a sensing junction
located at an interface between the sensing region and the
substrate and positioned at a first depth below a surface of the
substrate collects charge carriers generated in the substrate by
the radiation of the second energy; and
[0017] a shielding region formed in the substrate and incorporating
dopants of the second conductivity type, wherein a shielding
junction located at an interface between the shielding region and
the substrate and positioned at a deeper depth below the surface of
the substrate than the first depth collects charge carriers
generated in the substrate by radiation of the first energy.
[0018] According to another aspect of the present invention, a
radiation detector is provided that possesses an improved immunity
to x-ray noise over conventional radiation detectors, and that
utilizes a photocurrent generated directly from x-rays. The
inventive radiation detector includes:
[0019] a photoconductive portion that generates charge carriers
when irradiated with x-rays;
[0020] a semiconductor substrate incorporating dopants of a first
conductivity type;
[0021] a sensing region positioned at a first depth below a surface
of the substrate to collect the charge carriers generated in the
photoconductive portion;
[0022] a shielding portion formed in the substrate and
incorporating dopants of a second conductivity type, wherein a
shielding junction located at an interface between the shielding
portion and the substrate and positioned at a second depth that is
deeper below the surface of the substrate than the first depth
collects charge carriers generated in the substrate by stray x-rays
that pass through the photoconductive portion into the
substrate.
BRIEF DESCRIPTION OF THE DRAWINGS
[0023] The present invention will be better understood by a study
of the detailed description presented below considered in
conjunction with the attached drawings, of which:
[0024] FIG. 1 shows a cross-sectional view of a radiation detector
or sensor, according to an embodiment of the present invention;
[0025] FIG. 2 shows, in cross section, an arrangement of a
radiation detector that utilizes n-type regions in a p-type
substrate, according to an embodiment of the present invention;
[0026] FIG. 3 shows, in cross section, an arrangement of a
radiation detector that utilizes p-type regions in an n-type
substrate, according to an embodiment of the present invention;
and
[0027] FIG. 4 shows an arrangement of a large-area radiation
detector, according to an embodiment of the present invention.
[0028] It is to be understood that the attached drawings are
intended to schematically illustrate various aspects of the present
invention and may not be rendered to scale.
DETAILED DESCRIPTION OF THE INVENTION
[0029] In an embodiment of the present invention, the inventive
shielding arrangement is incorporated within a photosensitive CMOS
sensor containing an array of pixels. Each pixel includes a
photodiode with a p-n junction in a semiconductor substrate. The
photodiode capacitance is biased to a known or pre-set voltage and,
during exposure of the photodiode to photons, such as when an image
is being captured, electron-hole pairs are formed in the substrate
in proportion to the intensity of the photons. The electron-hole
pairs give rise to a photocurrent at the p-n junction, resulting
from charge carriers of the electron-hole pairs migrating through
the substrate and reaching the p-n junction. The photocurrent
causes a change in the voltage of the photodiode, and the voltage
is read out to obtain image data for that pixel. Image data
corresponding to multiple pixels are used to construct a captured
image.
[0030] FIG. 1 is a magnified cross-sectional view of an electronic
detector or sensor 1a for capturing an image by detecting
radiation, according to the embodiment of the present invention. As
shown in FIG. 1, the sensor 1a includes a scintillator 4 disposed
above a semiconductor 5, which is disposed above a tungsten layer
6. The scintillator 4, the semiconductor 5, and the tungsten layer
are supported on a passivated ceramic substrate 7. In general
terms, the scintillator 4 converts high-energy radiation, such as
x-rays, to lower-energy radiation, such as visible light. The
semiconductor 5, in turn, converts the lower-energy radiation into
electrical signals corresponding to image data representing the
captured image. The tungsten layer 6 absorbs any x-rays not
converted by the scintillator 4, to prevent those unconverted
x-rays from impinging on the patient, and absorbs any backscattered
radiation. The semiconductor 5 includes a CMOS APS array. According
to an aspect of the present invention, the image data produced by
the semiconductor 5 is conveyed to a cable 2 via a conductive lead
5'. The conductive lead 5' also may be used to convey electrical
power and/or control signals from a computer to the semiconductor
5. According to another aspect of the present invention, image data
and control signals are exchanged wirelessly, and in such a case
the cable 2 may be eliminated. The entire sensor 1a is enclosed in
a radiation housing 8, which protects the sensor 1a from shock and
enables the sensor 1a to be moisture resistant, and which is
transparent to x-rays.
[0031] The scintillator 4 is interposed between an x-ray source
(not shown) and the semiconductor 5, to both protect the
semiconductor 5 from unwanted exposure to x-rays and to convert the
x-rays to visible light, which can be directly detected by the
semiconductor 5. According to an aspect of the present invention,
the scintillator 4 is composed of gadolinium oxysulphate
(Gd.sub.2SO.sub.5) or thallium-doped cesium iodide (CsI(Tl)). Each
of these materials is sensitive to x-ray photons, and efficiently
converts them into visible photons in the 500-600 .mu.m wavelength
range. Other x-ray-to-light converting materials that may be used
for the scintillator 4 include: cadmium telluride; cadmium sulfide;
calcium tungstate (CaWO.sub.4); zinc sulfide; and zinc cadmium
sulfide. Scintillating glass, such as for example terbium glass, or
scintillating optical fibers may also be used for the scintillator
4.
[0032] In some types of x-ray imaging, as discussed above, x-rays
are first converted into visible light by the scintillator 4.
According to an aspect of the present invention, the scintillator 4
is formed of a material that converts x-rays to light to be
collected by the pixels in the APS array. However, as also
discussed above, the scintillator 4 will not convert 100% of the
x-rays that it receives; some x-rays inevitably will pass through
the scintillator 4 unconverted. Such unconverted x-rays can cause a
very large local charge to be generated in the sensor 1a, which, if
registered or collected by a pixel, can create noise and dark spots
in the resultant image.
[0033] For the purposes of the present discussion, the structure
and operation of a single pixel of the array of pixels is described
below. Adaptation of the below description to the entire array of
pixels is within the realm of a person of ordinary skill in imaging
device technology and therefore will not be discussed further.
[0034] In seeking to improve the x-ray noise immunity of a
semiconductor imaging chip or sensor, the present inventors
observed and exploited the phenomenon that energy from x-ray
photons is absorbed deeper in semiconductor material than energy
from visible light photons. The present invention exploits this
phenomenon to improve x-ray noise immunity by adding structure to
the semiconductor material of the sensor to redirect the
electron-hole pairs that are formed from x-ray photons absorbed
deep in the semiconductor material, thus preventing those
electron-hole pairs from being detected while allowing the
electron-hole pairs that are generated near the surface of the
semiconductor material to be detected, thereby providing a highly
effective shielding from x-ray noise.
[0035] An illustrative example of an aspect of the present
invention is provided in FIG. 2, which schematically shows a simple
illustration of a pixel 20 in a radiation sensor. The pixel 20
includes a photodiode consisting of a p-n junction 26 formed of an
n-type sensing region 22 in a p-type substrate or bulk region 21.
The sensing region 22 may be formed by known techniques, such as,
for example, ion implantation or diffusion of n-type dopants into
the bulk region 21. According to a preferred embodiment, the bulk
region 21 is formed of silicon, although other semiconductor
materials may be used and are within the scope of the present
invention. Preferably, the sensing region 22 is formed by ion
implantation and is heavily doped n-type (n+). The bulk region 21
may be comprised of, for example, silicon that is doped p-type
according to any of a number of techniques well known in the
art.
[0036] Prior to exposure to radiation, the photodiode is biased or
charged to a pre-set operating voltage. For example a voltage of 2
V may be used, although any biasing voltage sufficient for holes to
overcome the barrier voltage established by the p-n junction is
sufficient. When photons of visible light impinge on the pixel 20,
energy from the photons generate electron-hole pairs in the bulk
region 21. Positively charged holes from the electron-hole pairs
migrate through the bulk region 21 and, when the holes reach the
p-n junction 26, a photocurrent is produced, thus changing the
voltage of the photodiode.
[0037] That voltage change is representative of the captured image
at the location of the pixel 20. Following exposure to the photons
of visible light, the voltage of the pixel 20 is read out via a
line 24 to obtain image data for the pixel 20, and this data may be
used and combined with data from other pixels to construct the
captured image.
[0038] Also formed in the bulk region 21 are two n-type shielding
regions 23a, 23b. As schematically shown in FIG. 2, the shielding
regions 23a, 23b extend deeper into the bulk region 21 than does
the sensing region 22 (although their respective depths are not
necessarily illustrated to scale). The shielding regions 23a, 23b
form p-n junctions 27a, 27b with the bulk region 21 such that, when
radiation is incident on the sensor 20, electron-hole pairs
generated near the surface of the sensor 20 are detected through a
voltage change at the sensing region 22, while the electron-hole
pairs generated deeper in the bulk region 21, such as those
generated by unconverted x-rays, migrate to the shielding regions
23a, 23b and therefore are not detected or sensed at the sensing
region 22. In this manner, charge carriers from electron-hole pairs
corresponding to the captured image are detected and charge
carriers from electron-hole pairs formed by unwanted x-rays are not
detected. The shielding regions 23a, 23b are formed by known
techniques such as, for example, ion implantation. Preferably, the
shielding regions 23a, 23b are lightly doped (n-) relative to the
sensing region 22.
[0039] Optionally, instead of the two shielding regions 23a, 23b, a
single shielding or more than two shielding regions may be used in
the sensor 20.
[0040] While the embodiment discussed above refers to the bulk
region 21 or substrate as being p-type and refers to the sensing
region 22 and the shielding regions as being n-type, the present
invention may be implemented in other configurations. For example,
p-type sensing and shielding regions may be formed in an n-type
substrate or bulk region. The latter configuration can be
accomplished by forming an n-type epitaxial layer on top of a
p-type substrate or bulk region, and then forming p-type sensing
and shielding regions in the epitaxial layer.
[0041] FIG. 3 schematically shows a simple illustration of a pixel
20' in a radiation sensor, in which the polarity of the various
regions of the pixel 20' are opposite to the polarities of the
various regions in the pixel 20 of FIG. 2. The pixel 20' includes a
photodiode consisting of a p-n junction 26' formed of a p-type
sensing region 22' in an n-type substrate or bulk region 21'. Also
formed in the bulk region 21' are two p-type shielding regions
23a', 23b', which extend deeper into the bulk region 21' than does
the sensing region 22' (although their respective depths are not
necessarily illustrated to scale). The shielding regions 23a', 23b'
form p-n junctions 27a', 27b' with the bulk region 21'.
[0042] The specific depths of the sensing region 22 (22') and the
shielding regions 23a, 23b (23a', 23b'), i.e., the specific depths
of the p-n junctions 26, 27a, 27b (26', 27a', 27b') are dependent
on the nature and characteristics of the x-ray source. In a
photosensitive CMOS APS sensor, the p-n junctions 27a, 27b (27a',
27b') of the shielding regions 23a, 23b (23a', 23b') preferably are
formed at a depth that is greater than the penetration depth of
light emitted by the scintillator 4. More specifically,
scintillators fluoresce at a characteristic wavelength
corresponding to the color of the emitted light, such as light in
the blue-green portion of the visible-light spectrum. By forming
the shielding regions 23a, 23b (23a', 23b') such that their p-n
junctions 27a, 27b (26a', 27b') are positioned deeper in the bulk
region 21 (21') than the depth to which light generated by the
scintillator 4 penetrates, the shielding regions 23a, 23b (23a',
23b') effectively function to capture unwanted charge carriers from
electron-hole pairs that are generated deep in the bulk region 21
(21') by x-rays that pass through the scintillator 4 without being
converted to visible light. The shielding regions 23a, 23b (23a',
23b'), however, do not significantly affect the charge carriers
from electron-hole pairs generated by the visible light from the
scintillator 4, because these charge carriers are generated at a
relatively shallower depth in the bulk region 21 (21'). Thus,
charge carriers generated by the unconverted x-rays, which if left
to migrate to the p-n junction 26 (26') of the sensing region 22
(22') would result in noise in the captured image, are prevented
from being detected by the sensing region 22 (22').
[0043] Optionally, a line 25 may be use to apply a bias voltage to
the shielding regions 23a, 23b (23a', 23b').
[0044] In a preferred embodiment of the present invention, intended
for typical x-ray dental imaging applications, the sensing region
22 (22') is formed such that the p-n junction 26 (26') occurs at a
depth of about 0.2 microns, while the shielding regions 23a, 23b
(23a', 23b') are formed such that the p-n junctions 27a, 27b (27a',
27b') occur at a depth of about 4.5 microns, with the ratio of the
two depths being about 1:20. These depths and the depth ratio,
however, are exemplary only, and other depths and ratios are
possible and within the scope of the present invention.
[0045] In another embodiment of the present invention, a pixel of a
photoconductive sensor includes one or more shielding regions, such
as those discussed above. However, instead of using a scintillator
to convert x-rays to visible light, the sensor uses an
x-ray-sensitive photoconductive material, such as selenium,
PbI.sub.2, or HgI.sub.2, which directly converts x-rays to charge
carriers, which in turn produces an electrical signal. Optionally,
in such a sensor, the photodiode of the embodiment described above
is replaced by a capacitor, which is used to read out the charge
generated in the photoconductive material. In this case, the gate
of a read-out transistor is connected to a "floating" (capacitive)
node whose capacitance operates in a manner similar to the
photodiode capacitance discussed above.
[0046] Such a sensor also is similarly sensitive to spurious charge
carriers formed deep in the bulk region of the sensor and generated
by x-rays that penetrate beyond the photoconductive material (i.e.,
are not absorbed by the photoconductive material). By including
shielding regions in the sensor, such as the shielding regions
described above, sensitive charge-integration nodes in the pixel
are shielded from the spurious charge carriers.
[0047] A manifestation of the shielding arrangement of the present
invention is that a portion of the image signal may be diverted to
the shielding region(s) rather than to sensing region (or the
capacitive node). There is therefore a natural trade-off between
the loss of intended signal versus the desired shielding from noise
caused by unwanted x-rays. Depending upon the particular case, it
may be preferable to form an n-well that only partially surrounds
the node. By varying the geometry and the size of the shielding
region(s), an optimal image signal may be achieved. For example,
the sensing region (or the capacitive node) may be designed to have
a surface area that has a fixed proportion with respect to the
shielding region(s).
[0048] As mentioned above, the Schick '942 patent provides an
excellent example of a CMOS APS array for an x-ray sensor. The
x-ray sensor of the present invention has a number of differences
over the x-ray sensor of the '942 patent. For the x-ray sensor of
the '942 patent, the photodiode is formed in a relatively shallow
n-well or p-well, such that the detected signal comes from charge
carriers in the well. This shallow-well arrangement is based on the
principle that the desired signal comes from energy absorbed in the
thin depletion region near the surface of the photodiode.
Typically, the doping levels in the photodiode are such that there
is a high capacitance per unit area, which reduces its conversion
gain.
[0049] The shallow-well arrangement of the Schick '942 patent is
effective in blocking out unwanted charge carriers that diffuse
from the substrate. However, this arrangement also results in a
significant amount of desirable charge carriers being lost to the
junction between the well and the substrate instead of being
collected as part of the detected signal. This reduces the
responsivity of the photodiode.
[0050] In short, the shallow-well arrangement provides excellent
shielding from x-ray noise and cross talk, but suffers from a low
responsivity.
[0051] For the x-ray sensor of the present invention, the
photodiode is formed in the substrate or in an epitaxial layer
formed on the substrate. This arrangement does not shield as well
against x-ray noise and cross talk, compared with the arrangement
described in the Schick '942 patent, but instead is able to collect
a greater amount of desirable charge carriers and thus provide a
better responsivity. Incorporation of the shielding region(s) in
the x-ray sensor of the present invention reduces the amount of
x-ray noise in the detected signal by removing some of the unwanted
charge carriers generated deep in the substrate.
[0052] The photodiode arrangement of the present invention allows
for collection of desirable charge carriers that are present beyond
(below) the depletion region, by not confining the sensing region
to within a shallow well, as in the Schick '942 patent. That is,
the sensing region can extend deeper into the substrate than in the
arrangement of the '942 patent, and can be doped to a lower level
than in the arrangement of the '942 patent. This results in a lower
capacitance per unit area and a higher conversion gain of the
photodiode arrangement of the present invention.
[0053] The invention may be useful in a variety of applications,
such as intra-oral and large-area medical x-ray applications. In
dentistry, for instance, an FOP is commonly used to filter out
unconverted x-rays while directing light generated by a
scintillator onto a photosensitive array. The addition of an FOP,
however, adds to the overall sensor thickness and sensor weight,
and can result in a sensor that is uncomfortable for the patient,
which may then limit the medical practitioner's ability to
correctly position the sensor in the patient's mouth. By using a
shielding arrangement according to the present invention, the need
for a thick and heavy FOP is obviated. Further, the use of a
scintillator may be obviated by using an x-ray-sensitive
photoconductive material that allows for the direct imaging of
x-rays.
[0054] The present invention also is applicable to large-area
medical x-ray applications. Large format x-ray detectors may be
fabricated as large as 17''.times.17''. A suitable large-area FOP
is extremely expensive and fragile, as mentioned above, and the
inventive shielding arrangement eliminates the need for an FOP and
provides a cost-effective alternative approach to shield against
the effects of unwanted x-rays over a large area.
[0055] In a preferred embodiment, four wafer-scale CMOS dies are
butted together, end to end, to form a large-area detector 40 of
approximately 8''.times.10'' in size, as schematically shown in
FIG. 4 as dies I, II, III, and IV. Each die has two inactive
regions on adjacent sides that do not detect radiation (e.g., die I
has adjacent sides 41, 42 that do not detect radiation; die II has
adjacent sides 43, 44 that do not detect radiation; etc.). The
other two adjacent sides are adjoined to neighboring die. The array
is designed and manufactured so that pixels on adjacent sides are
approximately within two pixels of each other. To accommodate the
geometry, a left-sided and right-sided design are both required.
The aforementioned x-ray shielding routine allows the system to be
fabricated without a fiber optic plate, which would otherwise add
significant overall cost and fragility, as mentioned above.
[0056] The present invention has been described through the use of
illustrative examples. It is to be understood that the scope of the
present invention is not limited to the examples described herein,
and other structures or arrangements are encompassed by this
invention.
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