U.S. patent application number 11/090123 was filed with the patent office on 2009-11-26 for high resolution imaging system.
This patent application is currently assigned to NOVA R & D, INC.. Invention is credited to Martin Clajus, Tumay O. Tumer.
Application Number | 20090290680 11/090123 |
Document ID | / |
Family ID | 41342113 |
Filed Date | 2009-11-26 |
United States Patent
Application |
20090290680 |
Kind Code |
A1 |
Tumer; Tumay O. ; et
al. |
November 26, 2009 |
HIGH RESOLUTION IMAGING SYSTEM
Abstract
New sensors, pixel detectors and different embodiments of
multi-channel integrated circuit are disclosed. The new high energy
and spatial resolution sensors use solid state detectors. Each
channel or pixel of the readout chip employs low noise preamplifier
at its input followed by other circuitry. The different embodiments
of the sensors, detectors and the integrated circuit are designed
to produce high energy and/or spatial resolution two-dimensional
and three-dimensional imaging for different applications. Some of
these applications may require fast data acquisition, some others
may need ultra high energy resolution, and a separate portion may
require very high contrast. The embodiments described herein
addresses these issues and also other issues that may be useful in
two and three dimensional medical and industrial imaging. The
applications of the new sensors, detectors and integrated circuits
addresses a broad range of applications such as medical and
industrial imaging, NDE and NDI, security, baggage scanning,
astrophysics, nuclear physics and medicine.
Inventors: |
Tumer; Tumay O.; (Riverside,
CA) ; Clajus; Martin; (Riverside, CA) |
Correspondence
Address: |
FISH & ASSOCIATES, PC;ROBERT D. FISH
2603 Main Street, Suite 1000
Irvine
CA
92614-6232
US
|
Assignee: |
NOVA R & D, INC.
Riverside
CA
|
Family ID: |
41342113 |
Appl. No.: |
11/090123 |
Filed: |
March 28, 2005 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
|
60556507 |
Mar 26, 2004 |
|
|
|
Current U.S.
Class: |
378/62 ; 250/307;
250/311; 250/370.09; 250/370.14; 257/E21.002; 438/56; 438/64 |
Current CPC
Class: |
G01T 1/247 20130101;
H01L 2224/4847 20130101; Y10T 29/49126 20150115; H01L 2224/16145
20130101; G01T 1/249 20130101 |
Class at
Publication: |
378/62 ; 250/307;
250/311; 250/370.09; 438/56; 438/64; 250/370.14; 257/E21.002 |
International
Class: |
G01N 23/04 20060101
G01N023/04; G01N 23/00 20060101 G01N023/00; G21K 7/00 20060101
G21K007/00; G01T 1/24 20060101 G01T001/24; H01L 21/00 20060101
H01L021/00 |
Goverment Interests
STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT
[0002] Some parts of the invention were made with U.S. Government
support under Contract Numbers DAAE 30-02-C-1015, DAAE
30-03-C-1074, and DAMD 17-01-1-0356, which are awarded by the
Department of Defense. The U.S. Government has certain rights on
parts of the invention.
Claims
1-45. (canceled)
46. A method for producing a plurality of images, comprising the
steps of: using at least one source of radiation; constructing at
least one position sensitive direct conversion detector; producing
at least one charge signal from at least one portion of radiation
emitted from said at least one source of radiation in said at least
one position sensitive direct conversion detector; coupling at
least one integrated circuit to said at least one position
sensitive direct conversion detector, wherein said at least one
integrated circuit produces at least one signal from said at least
one charge signal; directing said at least one source of radiation
to said at least one position sensitive direct conversion detector;
placing at least one object between said at least one source of
radiation and said at least one position sensitive direct
conversion detector; scanning at least one portion of said at least
one object by said at least one position sensitive direct
conversion detector; and producing a plurality of images
simultaneously for each scan of said at least one portion of said
at least one object, wherein each of the plurality of images
corresponds to a different depth within said at least one portion
of said at least one object, wherein at least two of said plurality
of images are produced for different energy ranges.
47. The method for producing the plurality of images of claim 46,
wherein said at least one position sensitive direct conversion
detector has a plurality of pixels.
48. The method for producing the plurality of images of claim 46,
wherein said at least one portion of radiation emitted from said at
least one source of radiation is selected from the group consisting
of positrons, electrons and photons.
49. The method for producing the plurality of images of claim 47,
wherein said at least one signal is transferred from pixel to pixel
within said plurality of pixels timed to a scan speed of said at
least one position sensitive direct conversion detector.
50. The method for producing the plurality of images of claim 46,
wherein said at least one position sensitive direct conversion
detector is made to form an abuttable design wherein the abuttable
design is selected from the group consisting of two-side abuttable,
three-side abuttable and four-side abuttable design.
51. The method for producing the plurality of images of claim 49,
wherein said at least one charge signal is accumulated across at
least one column of pixels in the opposite direction of said scan
motion and output at least at one position of said at least one
column of pixels.
52. The method for producing the plurality of images of claim 46,
wherein said at least one charge signal is accumulated over a
plurality sections of said at least one position sensitive direct
conversion detector using said at least one integrated circuit and
output at multiple points across said at least one position
sensitive direct conversion detector using said at least one
integrated circuit.
53. The method for producing the plurality of images of claim 46,
wherein said at least one integrated circuit is disposed under said
at least one position sensitive direct conversion detector.
54. (canceled)
55. The method for producing the plurality of images of claim 46,
wherein said at least one signal is at least one count signal used
to produce said at least one image.
56. The method for producing the plurality of images of claim 55,
wherein said at least one count signal is used to produce at least
two images for different energy ranges.
57. The method for producing the plurality of images of claim 46,
wherein TDI technique is used.
58-65. (canceled)
66. A method for producing at least two images, comprising the
steps of: using at least one source of radiation; constructing at
least one four-side abuttable position sensitive direct conversion
detector; producing at least one charge signal from at least one
portion of radiation emitted from said at least one source of
radiation in said at least one position sensitive direct conversion
detector; coupling at least one integrated circuit disposed under
said at least one position sensitive direct conversion detector,
wherein said at least one integrated circuit produces at least one
signal from said at least one charge signal and wherein said at
least one integrated circuit is a mixed signal integrated circuit
having at least one analog circuit and at least one digital
circuit; directing said at least one source of radiation to said at
least one position sensitive direct conversion detector; placing at
least one object between said at least one source of radiation and
said at least one position sensitive direct conversion detector;
processing said at least one signal; and producing at least two
images of at least one portion of said at least one object. wherein
at least two of said at least two of images are produced for
different energy ranges.
67. The method for producing the at least one image of claim 66,
wherein said at least one position sensitive direct conversion
detector has a plurality of pixels.
68. The method for producing the at least one image of claim 66,
wherein said at least one portion of radiation emitted from said at
least one source of radiation is selected from the group consisting
of positrons, electrons and photons.
69. The method for producing the at least one image of claim 67,
wherein said at least one signal is transferred from pixel to pixel
within said plurality of pixels.
70. The method for producing the at least one image of claim 66,
wherein said at least one four-side abuttable position sensitive
direct conversion detector comprises at least two four-side
abuttable position sensitive direct conversation detectors
positioned to abut one another.
71. The method for producing the at least one image of claim 69,
wherein said at least one charge signal is accumulated across at
least one column of.
72. The method for producing the at least one image of claim 66,
wherein said at least one charge signal is accumulated over a
plurality sections of said at least one position sensitive direct
conversion detector using said at least one integrated circuit and
output at multiple points across said at least one position
sensitive direct conversion detector using said at least one
integrated circuit.
73. The method for producing the at least one image of claim 66,
wherein said at least one image consists of a plurality of images,
wherein each image corresponds to a different depth within said at
least one object.
74. (canceled)
75. The method for producing the at least one image of claim 66,
wherein said at least one signal is at least one count signal used
to produce said at least one image.
76. The method for producing the at least one image of claim 75,
wherein said at least one count signal used to produce at least two
images for different energy ranges.
77. The method for producing the at least one image of claim 66,
wherein TDI technique is used.
78. An imaging system comprising in combination: at least one
source of radiation; at least one position sensitive direct
conversion detector, wherein said at least one position sensitive
direct conversion detector produces at least one charge signal from
at least one portion of radiation emitted from said at least one
source of radiation and wherein said at least one source of
radiation is directed to said at least one position sensitive
direct conversion detector; at least one integrated circuit coupled
to said at least one position sensitive direct conversion detector,
wherein said at least one integrated circuit produces at least one
signal from said at least one charge signal; at least one object
placed between said at least one source of radiation and said at
least one position sensitive direct conversion detector, wherein
said at least one position sensitive direct conversion detector
scans said at least one object; at least one processing system
responsive to said at least one signal; and said at least one
processing system processes said at least one signal to produce
simultaneously at least two images for each scan of at least one
portion of said at least one object, wherein each of said at least
two images correspond to a different depth within said at least one
portion of said at least one object, wherein at least two of said
at least two of images are produced for different energy
ranges.
79. The imaging system of claim 78, wherein said at least one
position sensitive direct conversion detector has a plurality of
pixels.
80. The imaging system of claim 78, wherein said at least one
portion of radiation emitted from said at least one source of
radiation is selected from the group consisting of positrons,
electrons and photons.
81. The imaging system of claim 79, wherein said at least one
signal is transferred from pixel to pixel within said plurality of
pixels timed to a scan speed of said at least one position
sensitive direct conversion detector.
82. The imaging system of claim 78, wherein said at least one
position sensitive direct conversion detector is made to form an
abuttable design wherein the abuttable design is selected from the
group consisting of two-side abuttable, three-side abuttable and
four-side abuttable design.
83. The imaging system of claim 81, wherein said at least one
charge signal is accumulated across at least one column of pixels
in the opposite direction of said scan motion and output at least
at one position of said at least one column of pixels.
84. The imaging system of claim 78, wherein said at least one
charge signal is accumulated over a plurality sections of said at
least one position sensitive direct conversion detector using said
at least one integrated circuit and output at multiple points
across said at least one position sensitive direct conversion
detector using said at least one integrated circuit.
85. The imaging system of claim 78, further comprising sequentially
reading out alternating pixels of the detector, wherein scanning
said at least one object is performed linearly.
86. (canceled)
87. The imaging system of claim 78, wherein said at least one
signal is at least one count signal used to produce said at least
one image.
88. The imaging system of claim 87, wherein said at least one
processing system produces at least two images using said at least
one count signal for different energy ranges.
89. The imaging system of claim 78, wherein TDI technique is
used.
90. The method of claim 46, further comprising sequentially reading
out alternating pixels of the detector, wherein scanning said at
least one object is performed linearly.
91. The method of claim 46, further comprising: rotating said at
least one position sensitive direct conversion detector; and
rotating said at least one source of radiation.
92. An imaging system comprising in combination: at least one
source of radiation; at least one four-side abuttable position
sensitive direct conversion detector system, wherein said at least
one four-side abuttable position sensitive direct conversion
detector system produces at least one charge signal from at least
one portion of radiation emitted from said at least one source of
radiation and wherein said at least one source of radiation
directed to said at least one four-side abuttable position
sensitive direct conversion detector system; at least one
integrated circuit disposed under said at least one four-side
abuttable position sensitive direct conversion detector system,
wherein said at least one integrated circuit produces at least one
signal from said at least one charge signal and wherein said at
least one integrated circuit is a mixed signal integrated circuit
having at least one analog circuit and at least one digital
circuit; at least one object placed between said at least one
source of radiation and said at least one four-side abuttable
position sensitive direct conversion detector system; at least one
processing system responsive to said at least one signal; and said
at least one processing system that processes said at least one
signal to produce at least one image of at least one portion of
said at least one object, wherein said at least one processing
system produces at least two images each one for a different energy
range.
93. The imaging system of claim 92, wherein said at least one
position sensitive direct conversion detector has a plurality of
pixels.
94. The imaging system of claim 92, wherein said at least one
portion of radiation emitted from said at least one source of
radiation is selected from the group consisting of positrons,
electrons and photons.
95. The imaging system of claim 93, wherein said at least one
signal is transferred from pixel to pixel within said plurality of
pixels.
96. The imaging system of claim 92, wherein said at least one
four-side abuttable position sensitive direct conversion detector
is comprises at least two four-side abuttable position sensitive
direct conversation detectors positioned to abut one another.
97. The imaging system of claim 95, wherein said at least one
charge signal is accumulated across at least one column of
pixels.
98. The imaging system of claim 92, wherein said at least one
charge signal is accumulated over a plurality sections of said at
least one four-side abuttable position sensitive direct conversion
detector using said at least one integrated circuit and output at
multiple points across said at least one position sensitive direct
conversion detector using said at least one integrated circuit.
99. The imaging system of claim 92, wherein said at least one image
consists of plurality of images wherein each image correspond to
different depths within said at least one object.
100. The imaging system of claim 92, wherein said at least one
processing system produces at least two images each one for a
different energy range.
101. The imaging system of claim 92, wherein said at least one
signal is used to produce at least two images, wherein each image
is for a different energy range.
102. The imaging system of claim 101, wherein said at least one
processing system produces at least one image for said at least one
energy range using said at least one count signal.
103. The imaging system of claim 92, wherein TDI technique is
used.
104. The method for producing the plurality of images of claim 53,
wherein at least one mounting board is disposed in between said at
least one integrated circuit and said at least one position
sensitive direct conversion detector.
105. The method for producing the at least one image of claim 66,
wherein at least one mounting board is disposed in between said at
least one integrated circuit and said at least one position
sensitive direct conversion detector.
106. The imaging system of claim 78, wherein said at least one
integrated circuit is disposed under said at least one position
sensitive direct conversion detector.
107. The imaging system of claim 106, wherein at least one mounting
board is disposed in between said at least one integrated circuit
and said at least one position sensitive direct conversion
detector.
108. The imaging system of claim 92, wherein at least one mounting
board is disposed in between said at least one integrated circuit
and said at least one position sensitive direct conversion
detector.
109. A method for producing a plurality of images, comprising the
steps of: using at least one source of radiation; constructing at
least one position sensitive direct conversion detector; producing
at least one charge signal from at least one portion of radiation
emitted from said at least one source of radiation in said at least
one position sensitive direct conversion detector; coupling at
least one integrated circuit to said at least one position
sensitive direct conversion detector, wherein said at least one
integrated circuit produces at least one signal from said at least
one charge signal; partitioning at least one portion of said at
least one signal into at least two energy [[bins]] ranges;
directing said at least one source of radiation to at least one
portion of said at least one position sensitive direct conversion
detector; placing at least one portion of at least one object
between said at least one source of radiation and said at least one
position sensitive direct conversion detector; imaging at least one
portion of said at least one object by said at least one position
sensitive direct conversion detector; and producing a plurality of
images simultaneously of said at least one portion of said at least
one object, wherein at least two of said plurality of images
corresponds to said at least two energy ranges.
110. An imaging system comprising in combination: at least one
source of radiation; at least one position sensitive direct
conversion detector, wherein said at least one position sensitive
direct conversion detector produces at least one charge signal from
at least one portion of radiation emitted from said at least one
source of radiation and wherein said at least one source of
radiation is directed to at least one portion of said at least one
position sensitive direct conversion detector; at least one
integrated circuit coupled to said at least one position sensitive
direct conversion detector, wherein said at least one integrated
circuit produces at least one signal from said at least one charge
signal, wherein at least one portion of said at least one signal is
partitioned into at least two energy ranges; at least one portion
of at least one object placed between said at least one source of
radiation and said at least one position sensitive direct
conversion detector, wherein said at least one position sensitive
direct conversion detector images said at least one portion of said
object; and at least one processing system that processes said at
least one signal partitioned into said at least two energy ranges
to produce simultaneously at least two images for said at least one
portion of said at least one object, wherein each of said at least
two images corresponds to said at least two energy ranges.
111. The method of claim 109, wherein said at least one position
sensitive direct conversion detector has a plurality of pixels.
112. The imaging system of claim 110, wherein said at least one
position sensitive direct conversion detector has a plurality of
pixels.
113. The method of claim 109, wherein said radiation is selected
from the group consisting of positrons, electrons and photons.
114. The imaging system of claim 110, wherein said radiation is
selected from the group consisting of positrons, electrons and
photons.
115. A method for producing at least one four side abuttable
detector, comprising the steps of: producing at least one
integrated circuit with contact pads on at least one surface; using
at least one detector material sensitive to radiation with
electrodes on at least one surface with matching area to said at
least one integrated circuit, wherein said at least one detector
material has a two-dimensional surface area that covers at least
all of a surface of said at least one integrated circuit; making a
plurality of vias through said at least one integrated circuit to
establish an electrical contact between top and bottom surfaces of
the at least one integrated circuit; mounting said at least one
detector material on top of said at least one integrated circuit to
form at least one four side abuttable detector; processing a
plurality of signals produced inside said at least one detector
material by said at least one integrated circuit; and outputting
said processed plurality of signals from said at least one four
side abuttable detector through at least one portion of said
plurality of vias.
116. A method for producing at least one four side abuttable
detector, comprising the steps of: producing at least one
integrated circuit with contact pads on at least one surface;
making at least one connection board with connection pads on at
least one surface with matching area to said at least one
integrated circuit; using at least one detector material sensitive
to radiation with electrodes on at least one surface with matching
area to said at least one integrated circuit, wherein said at least
one detector material has a two-dimensional surface area that
covers at least all of a surface of said at least one integrated
circuit; making a plurality of vias through said at least one
integrated circuit and said at least one connection board to
establish electrical contact between top and bottom surfaces;
stacking said at least one detector material, said at least one
connection board and said at least one integrated circuit on top of
one another in any order to form said at least one four side
abuttable detector; processing a plurality of signals produced
inside said at least one detector material by said at least one
integrated circuit; and outputting said processed plurality of
signals from said at least one four side abuttable detector through
at least one portion of said plurality of vias.
117. A four side abuttable detector, comprising in combination: at
least one material sensitive to radiation with electrodes on at
least one surface with matching area to said at least one
integrated circuit; at least one integrated circuit with contact
pads on at least one surface, wherein said at least four side
abuttable integrated circuit processes a plurality of signals
produced inside said at least one detector material, and wherein
said at least one material has a two-dimensional surface area that
covers at least all of a surface of said at least one integrated
circuit; and a plurality of vias through said at least one
integrated circuit to establish electrical contact between top and
bottom surfaces of said at least one integrated circuit, connected
to at least one of said contact pads, wherein said at least one
detector material is stacked on top of said at least one integrated
circuit to form at least one four side abuttable detector.
118. A four side abuttable detector, comprising in combination: at
least one connection board with connection pads on at least one
surface; at least one detector material sensitive to radiation with
electrodes on at least one surface, wherein said at least one
detector material has a two-dimensional surface area that covers at
least all of a surface of said at least one connection board; at
least one integrated circuit with connection pads on at least one
surface, wherein said two-dimensional surface area of said at least
one detector material covers at least all of a surface of said at
least one integrated circuit, wherein said at least one detector
material, said at least one connection board and said at least one
integrated circuit are stacked on top of one another in any order,
and wherein said at least one integrated circuit processes a
plurality of signals produced inside said at least one detector
material, and; a plurality of vias through said at least one
integrated circuit and said at least one connection board that
establishes an electrical contact between top and bottom surfaces
of said at least one integrated circuit and said at least one
connection board, wherein said plurality of signals are transferred
through at least one portion of said plurality of vias.
119. The method of claim 115, wherein said radiation is selected
from the group consisting of positrons, electrons and photons.
120. The method of claim 116, wherein said radiation is selected
from the group consisting of positrons, electrons and photons.
121. The detector system of claim 117, wherein said radiation is
selected from the group consisting of positrons, electrons and
photons.
122. The detector system of claim 118, wherein said radiation is
selected from the group consisting of positrons, electrons and
photons.
123. The method of claim 115, wherein said at least one four side
abuttable detector mounted side by side to form at least one large
detector array.
124. The method of claim 116, wherein said at least one four side
abuttable detector mounted side by side to form at least one large
detector array.
125. The detector system of claim 117, wherein said at least one
four side abuttable detector mounted side by side to form at least
one large detector array.
126. The detector system of claim 118, wherein said at least one
four side abuttable detector mounted side by side to form at least
one large detector array.
127. The method of claim 115, wherein said at least one four side
abuttable detector connected to at least one circuit board that
collects data from said at least one four side abuttable
detector.
128. The method of claim 116, wherein said at least one four side
abuttable detector connected to at least one board that collects
data from said at least one four side abuttable detector.
129. The detector system of claim 117, wherein said at least one
four side abuttable detector connected to at least one board that
collects data from said at least one four side abuttable
detector.
130. The detector system of claim 118, wherein said at least one
four side abuttable detector connected to at least one board that
collects data from said at least one four side abuttable detector.
Description
CROSS-REFERENCE TO RELATED APPLICATION
[0001] This application claims the benefit of U.S. Provisional
Application Ser. No. 60/556,507 filed on Mar. 26, 2004, entitled
HIGH RESOLUTION IMAGING SYSTEM. The afore-mentioned application is
hereby incorporated by reference herein in its entirety, including
but not limited to those portions that specifically appear
hereinafter, the incorporation by reference being made with the
following exception: In the event that any portion of the
above-referenced application is inconsistent with this application,
this application supercedes said above-referenced applications.
BACKGROUND OF THE INVENTION
[0003] 1. The Field of the Disclosure
[0004] The present disclosure relates generally to radiographic
imaging, and more particularly, but not necessarily entirely, to
diagnostic x-ray imaging.
[0005] 2. Description of Related Art
[0006] Radiographic imaging systems are well known in the art and
in particular scanning radiographic systems are known in the art.
It is further known that CdZnTe or other solid state radiation
detectors can be used to detect x-rays. The prior art teaches
various other features that may be incorporated into a scanning
radiographic system. However, use of a scanning system coupled to
the disclosed position sensitive digital detection methods are
not.
[0007] The features and advantages of the disclosure will be set
forth in the description which follows, and in part will be
apparent from the description, or may be learned by the practice of
the disclosure without undue experimentation. The features and
advantages of the disclosure may be realized and obtained by means
of the apparatus and combinations particularly pointed out in the
appended claims.
BRIEF DESCRIPTION OF THE INVENTION
[0008] The disclosed fast full body digital radiography system is a
direct imaging digital radiography system and will not use
radiographic film. The solid state pixel detectors have high MTF,
about 5 to 8 line pairs per mm, approaching the resolution of a
screen-film system. The system has high DQE, about 50%-70%, and
high contrast, therefore low dose to the patient. The linear array
has a low projected area for x-rays scattered in the patient's body
and therefore low background and higher signal to noise ratio. The
images are taken in scanning mode using the Time Delayed
Integration (TDI) technique. This technique produces uniform images
without flawed or dead image pixels even if there are a moderate
number of random defective detector pixels.
[0009] This is a low cost, digital (no film), high resolution, low
dose and fast general whole or partial body radiography system that
will allow immediate and fast prescreening of patients arriving at
emergency rooms (ER). In human diagnostic applications the system
is capable of screening a whole body or parts of a body in several
seconds and displaying the image in real time on a high resolution
computer screen without the need for accurate patient positioning
and elaborate exposure setting. Small and portable machines can be
used in ambulances, special imaging vans, rural and remote
communities, during catastrophic events, on battlefields and in
field hospitals.
[0010] The high resolution imaging detectors described below can be
used for other applications, including security scanning for
explosives and contraband at airports and other installations,
computed tomography (CT) scanners for medicine, digital radiography
such as digital mammography, and industrial non destructive
inspection (NDI) and analysis (NDA) can be performed.
BRIEF DESCRIPTION OF THE DRAWINGS
[0011] The features and advantages of the disclosure will become
apparent from a consideration of the subsequent detailed
description presented in connection with the accompanying drawings
in which:
[0012] FIG. 1 is a drawing of the disclosed whole body fast digital
radiography scanning system;
[0013] FIG. 2 is a drawing demonstrating the Depth-of-field
considerations in TDI scanning;
[0014] FIGS. 3 A and B is a drawing of the hybrid CdZnTe pixel
detector with CdZnTe pixel array and detail of the indium bump
bonding onto the readout ASIC chip;
[0015] FIG. 4 is a drawing of the Schematic of the ASIC per-pixel
front-end analog chain;
[0016] FIG. 5 is a drawing of the Overall Geometric Concept for
DANA IC;
[0017] FIG. 6 is a drawing of the Signal Chain of DANA IC;
[0018] FIG. 7 is a top level diagram of a pixel with dual energy
channels;
[0019] FIG. 8 is a block diagram of a 160 cell DTDI circuit with
two counters in each cell;
[0020] FIG. 9 is a drawing of the staggered linear array formation
with tilted detector normal to the x-ray beam direction; and
[0021] FIG. 10 is a block diagram of a four side abuttable pixel
detector shown mounted onto a ceramic or printed circuit board.
[0022] FIG. 11 is a block diagram showing the geometry and
structure of the pixels on a solid state detector pixel array.
[0023] FIG. 12 is a block diagram showing a thick solid state pixel
detector for high energy photon detection and imaging.
[0024] FIG. 13 is a block diagram showing a single channel of a
chip which uses analog counters for accumulating multiple energy
images.
DETAILED DESCRIPTION
[0025] For the purposes of promoting an understanding of the
principles in accordance with the disclosure, reference will now be
made to the embodiments illustrated in the drawings and specific
language will be used to describe the same. It will nevertheless be
understood that no limitation of the scope of the disclosure is
thereby intended. Any alterations and further modifications of the
inventive features illustrated herein, and any additional
applications of the principles of the disclosure as illustrated
herein, which would normally occur to one skilled in the relevant
art and having possession of this disclosure, are to be considered
within the scope of the disclosure claimed. It is also noted that
the prior provisional application contains more detailed disclosure
of the invention, which can be referenced.
[0026] The publications and other reference materials referred to
herein to describe the background of the disclosure, and to provide
additional detail regarding its practice, are hereby incorporated
by reference herein in their entireties, with the following
exception: In the event that any portion of said reference
materials is inconsistent with this application, this application
supercedes said reference materials. The reference materials
discussed herein are provided solely for their disclosure prior to
the filing date of the present application. Nothing herein is to be
construed as a suggestion or admission that the inventors are not
entitled to antedate such disclosure by virtue of prior disclosure,
or to distinguish the present disclosure from the subject matter
disclosed in the reference materials.
[0027] The following publications are hereby incorporated by
reference herein in their entireties: Shi Yin, Tumay O. Tumer, Dale
Maeding, James Mainprize, Gord Mawdsley, Martin Yaffe and William
J. Hamilton "A Low-Dose High Contrast Digital Mammography System
(DigiMAM)", Presented in IEEE Medical Imaging Conference and
Submitted to IEEE Trans. Nucl Science (November 1998); J. G.
Mainprize, N. Ford, S. Yin, T. O. Tumer, M. J. Yaffe, "Image
Quality of a prototype Direct Conversion Detector for Digital
Mammography," Presented at SPIE's Inter. Sym. on Med. Imaging, (San
Diego February 1999); J. G. Mainprize, M. J. Yaffe, T. O. Tumer, S.
Yin and W. J. Hamilton, "Design Considerations for a CdZnTe digital
Mammography System," Presented at 1998 International Workshop for
Digital Mammography, Nijmegen, Netherlands (June 1998); Yin, S., T.
O. Tumer, D. Maeding, J. Mainprize, G. Mawdsley, M. J. Yaffe and W.
J. Hamilton, "Hybrid direct conversion detectors for digital
mammography", IEEE Tran. Nuc. Sci. Vol. 46, No. 6, 2093-97 (1999);
Yin, S., T. O. Tumer, D. Maeding, J. Mainprize, G. Mawdsley, M. J.
Yaffe, E. E. Gordon and W. J. Hamilton, "Direct conversion Si and
CdZnTe detectors for digital mammography", Nuc. Inst. & Met.
Phys. Res. A. 448 (2000) 591-97; Yin, S., T. O. Tumer, D. Maeding,
J. Mainprize, G. Mawdsley, M. J. Yaffe, E. Gordon and W. J.
Hamilton, "Direct conversion CdZnTe and CdTe detectors for digital
mammography", IEEE Tran. Nuc. Sci. Vol. 49, No. 1, 176-81 (2002);
T. O. Tumer et al., "New ASICs Specifically Developed for Position
Sensitive Solid State Detectors" Presented at the 10.sup.th
Symposium on Radiation Measurements and Applications, University of
Michigan (May 2002); Yaffe, M. J. and J. A. Rowlands, "X-ray
detectors for digital radiography", Phys. Med. Biol. 42, p. 1-39
(1997); Yin, S., T. O. Tumer, D. Maeding, J. Mainprize, G.
Mawdsley, M. J. Yaffe, E. Gorden and W. J. Hamilton, "Direct
conversion CdZnTe and CdTe detectors for digital mammography", IEEE
Tran. Nuc. Sci. Vol. 49, No. 1, 176-81 (2002); T. O. Tumer, S. Yin,
V. Cajipe, H. Flores, J. Mainprize, G. Mawdsley, J. A. Rowlands, M.
J. Yaffe, E. E. Gordon, W. J. Hamilton, D. Rhiger, S. O. Kasap, P.
Sellin and K. S. Shah "High-resolution pixel detectors for second
generation digital mammography", Nuc. Inst. & Met. Phys. Res.
A. 497 (2003) 21-29; Yaffe, M. J. and J. A. Rowlands, "X-ray
detectors for digital radiography", Phys. Med. Biol. 42, p. 1-39
(1997); T. Asaga, C. Masuzawa, A. Yoshida and H. Mattsuura,
"Dual-energy subtraction mammography", J. Digit. Imaging 8, 70-73
(1995); Lewin J M, D'Orsi C J, Hendrick R E, Moss L J, Isaacs P K,
Karellas A, Cutter G R. Clinical comparison of full-field digital
mammography and screen-film mammography for detection of breast
cancer. AJR Am J Roentgenol 2002; 179(3):671-7; Yin, S., T. O.
Tumer, D. Maeding, J. Mainprize, G. Mawdsley, M. J. Yaffe, E.
Gorden and W. J. Hamilton, "Direct conversion CdZnTe and CdTe
detectors for digital mammography", IEEE Tran. Nuc. Sci. Vol. 49,
No. 1, 176-81 (2002); T. O. Tumer, S. Yin, V. Cajipe, H. Flores, J.
Mainprize, G. Mawdsley, J. A. Rowlands, M. J. Yaffe, E. E. Gordon,
W. J. Hamilton, D. Rhiger, S. O. Kasap, P. Sellin and K. S. Shah
"High-resolution pixel detectors for second generation digital
mammography", Nuc. Inst. & Met. Phys. Res. A. 497 (2003)
21-29.
Scanning Detector for Digital Radiography
[0028] FIG. 1 shows the physical scanning motion of the digital
radiography system disclosed herein. The digital nature of the
imaging process permits the user to adjust the effective pixel size
by combining adjacent pixels, in order to reduce the patient dose,
reduce image size for data transmission and storage purposes, or
enhance image contrast. With the exception of the dose reduction,
these goals can be achieved after the image has been recorded, by
re-generating the image from the existing detector data with
different binning.
[0029] Digital radiography produces large image data sets. At 0.1
mm.times.0.1 mm pixel size, a typical radiograph may contain as
many as 7,000.times.20,000 pixels, and each image pixel may consist
of multiple TDI segments that will need to be retained if the
capability to generate arbitrary image slices is to be maintained.
The pixel sizes can range from <0.005.times.0.005 mm to
>10.times.10 mm. Also digital radiography may contain
<500.times.1,000 to >1,000,000.times.2,000,000 pixels. Before
data compression, which may considerably reduce the data size, this
can amount to gigabytes to terabytes of raw data, not including any
reconstructed images that may need to be archived separately.
[0030] The disclosed system has low noise, linear response and high
dynamic range of .about.14 to 16 bits. The dynamic range may be
extended to <8 to >28 bits for certain applications. Such a
detector can be produced using a hybrid technology consisting of a
converter detector layer bonded directly onto the readout
integrated circuit (IC). For one embodiment of the invention, the
detector layer is composed of high Z, high density photoconductive
solid state materials such as CdZnTe, CdTe, Silicon, GaAs,
PbI.sub.2, HgI.sub.2, CdWO.sub.4, Selenium, used to achieve high
x-ray quantum detection efficiency. Such materials directly and
precisely convert the absorbed energy to electric signals,
electron-hole (e-h) pairs 307. Each incoming x-ray photon produces
a pulse of e-h pairs 307. Each pulse is integrated and the charges
from different pulses detected by the same pixel during the same
time period are summed. In TDI operation, the signal-to-noise ratio
is further improved by shifting the charges from pixel to pixel.
This charge shifting is done synchronous with and in the opposite
direction of the scanning motion of the detector assembly,
resulting in an increased effective exposure time for each image
pixel and therefore improved contrast. Also, during the single
movement of the detector length in TDI only a single direction in
the object is viewed by each pixel of the detector because of the
transfer of charge in the reverse direction. Therefore, the number
of detected x-rays coming out for each pixel is in fact the
integrated sum of all the pixels along the length of the detector
observing the same direction inside the object imaged.
[0031] The tolerance for defective pixels will also increase
detector yield and lower the cost of production. There will be no
need for complicated patient positioning, because the patient will
be placed on the gantry and the scanning arm with the x-ray tube
will scan the whole or the selected part of the body of the
patient. The response of the direct conversion solid state
detectors is linear and there is no need for fine exposure control
or adjustment. The x-ray tube output is lower compared to standard
radiography systems and with shielding operator exposure will be
minimized.
[0032] We use 10-20 cm/s scan speeds. However, scan speeds can vary
from <0.1 cm/s to >200 cm/s. With 7,000 to 10,000 pixels for
the full length detector array, the image size is about 7-10
K.times.full length of a body or a part of the body as required.
Pixels may be combined into larger pixels such as 200.times.200
.mu.m.sup.2 or 400.times.400 .mu.m.sup.2 to reduce image size and
spatial resolution but in turn to increase contrast.
Detector Readout Chip with TDI
[0033] In one embodiment the detector readout IC for the disclosed
digital radiography system will utilize CCD-type architecture to
implement a time-delayed integration (TDI) technique. The charges
generated in a given detector pixel are collected in the
corresponding readout pixel's charge well for a preset duration,
about, but not limited to, 0.5 to 1 ms. In fact, there is no limit
to the duration. It can vary from <1 .mu.s to >1,000 s. The
charges are then transferred to the next adjacent pixel in the
direction opposite to the detector scanning motion, where
additional charges are collected before the transfer process is
repeated. At the end of each pixel column, the accumulated charges
are digitized (off-chip) and the results sent to the readout
computer for image reconstruction. By synchronizing the charge
transfer speed with the scanning speed, the individual charge
"buckets" remain stationary relative to the object being imaged and
the signal acquired for a given image pixel is enhanced in
proportion to the number of detector pixels per column. This
condition, together with the pixel size, related to a scanning
speed of 10 to 20 cm/s, not counting the small correction for the
magnification between the object plane and the detector. The
scanning speed is directly related to the rate of pixel-to-pixel
charge transfer.
[0034] The dynamic range that can be achieved in this manner is
limited by the capacity of the charge wells, about 3.times.10.sup.7
electrons and/or holes, depending on the specific CCD fabrication
process used, and the readout noise, .about.3.times.10.sup.3
electrons and/or holes. To overcome this limitation, NOVA has
developed a proprietary "early bypass readout" design. In this
design, the pixel columns are subdivided into short (four to eight
pixel) segments, each of which is read out and digitized
separately, boosting the effective well capacity by a factor 16 to
32 while increasing the noise by only the square root of that
factor. The charge well capacity can be designed to be anywhere in
the range of <10.sup.3 to >10.sup.10 electrons and/or
holes.
[0035] An added benefit of the TDI approach is that it
significantly increases the tolerance for low-performing or even
defective detector pixels. As such a pixel will only be one out of
more than one hundred that contribute to a given image pixel, its
contribution (or lack thereof) to the overall image quality will
typically not be significant, and it will take multiple bad pixels
in any single column to cause a noticeable deterioration. In turn,
this tolerance increases the yield of useable detectors and thus
contributes to achieving our goal of producing a low-cost
system.
[0036] To eliminate the effect of the non-working pixels and also
the detector pixel-to-pixel variations the pixel detector is
required to be calibrated. The calibration can be done in a variety
of ways. For example, a "flood" image can be taken using the x-ray
source without an object. The image then can be calibrated to be
uniform and then each pixel is assigned a weight factor required to
convert its contents to the set uniform image pixel value. After
this, the weight factor matrix is applied to all the pixels of the
images taken with the object in position to be uniform and without
any artifacts due to pixel response variations.
[0037] The averaging of the signal over a large number of pixels
makes it difficult to accurately assess detector quality.
Therefore, and for other test purposes and new applications such as
dental x-rays, the readout IC has the capability to operate in
staring (non-TDI) mode.
[0038] In the staring (non-TDI) mode each pixel observes a certain
direction in the object without any charge transfer. The pixels
accumulate the number of x-ray photons incident onto the
corresponding detector pixel. Then the charge accumulation is
stopped and the chip read out in series, pixel by pixel, or in
parallel. After the readout is complete the chip can be turned on
again to accumulate another image through the pixel detector.
[0039] The IC design is modeled after the MARY.TM. chip that NOVA
has developed for its digital mammography projects. The chip
architecture is largely the same but important design details will
be different, requiring a redesign of the chip. Table I compares a
few selected design parameters for the two chips.
TABLE-US-00001 TABLE I Selected design parameters for the disclosed
readout IC (preliminary) compared to those of NOVA'S MARY chip.
Parameter New readout IC MARY chip Type of charges Electrons Holes
collected Pixel size (.mu.m.sup.2) ~100 .times. 100 50 .times. 50
Size of pixel bonding Large, to support Minimal pads Alternative
bonding methods Pixel count (TDI 10 .times. 10 to 256 .times. 192
.times. 384 direction listed 1,024 first) Charge well capacity 3
.times. 10.sup.7 electrons 1.5 .times. 10.sup.7 holes Staring mode
Included Requires complex support circuitry
Depth-of-Field Correction Algorithm
[0040] The normal TDI scan requires a radial or circular motion. In
a straight linear scan as shown in FIG. 1 if the object to be
imaged has a very short thickness it can produce good images.
However, if the object has significant thickness then the images
will be blurred. A new embodiment described below solves this
problem.
[0041] One potential drawback of straightforward linear TDI
scanning is the limited depth of field that it can achieve. To
understand this, consider the sketch in FIG. 2. The fan beam from
the x-ray source 101 illuminates the detector 103 after passing
through the object under test 104. Consider a reference frame in
which the object moves in the direction indicated by the arrow 204
in FIG. 2 while the source and detector are at rest, rather than
the opposite situation that we outline here. The two frames are
equivalent in physics terms, though not necessarily in practical
terms. In order to obtain a sharp image of the features in the
middle of the object, at a distance y.sub.m from the source, the
object has to move by the length x.sub.m of the line 202 in the
time t that it takes to shift the charges from one end of the
detector to the other, so that the image of the line 202 moves by
the active width of the detector 205, x.sub.d. By simple geometric
considerations, x.sub.m is given as
x m = y m y d x d . ##EQU00001##
[0042] Obviously, in the same time t the top 201 and bottom 203 of
the object will move by the same distance. The projected images of
these bars, however, will move by
x b , t = y d y b , t x m = y m y b , t x d . ##EQU00002##
[0043] These values are different than x.sub.d and thus out of
synch with the movement of charges in the readout IC, causing
blurring of these out-of-focus regions of the object.
[0044] As a concrete example, with y.sub.b=150 cm, y.sub.t=120 cm
(that is, a 30 cm thick object), and x.sub.d=10 mm, x.sub.b and
x.sub.t would amount to 9 mm and 11.1 mm, respectively.
[0045] In digital mammography, this limitation is usually overcome
by moving the detectors in a circular arc centered on the x-ray
source spot. For a low-cost system of the proposed scale, however,
this does not appear to be a viable option; moreover, the
distortions that would result from projecting an extended but
essentially flat object onto a cylindrical surface might themselves
pose an image quality problem. Another solution is the use of a
parallel x-ray beam instead of a fan beam, but we are not aware of
any economically viable approach to producing a parallel beam of
the required width and intensity. Moving the source far enough away
that the difference between y.sub.b and y.sub.t is no longer
significant would require more space than is available in any
realistic hospital setting, much higher flux x-ray generator and
much higher cost system. This counteracts our goal to reduce the
overall x-ray flux compared to existing technology.
[0046] In one embodiment this may be overcome if the detector array
and the x-ray source are rotated 90.degree. together and the
detector system is pivoted around the x-ray generator focal spot.
Then the linear detector array and the source is moved similar to
the direction of the standard TDI systems in an arc centered at the
x-ray generator focal spot. In this case the scanning will be
carried out from side to side of the patient's body and not a
linear motion from head to toe.
[0047] Also, we have developed a completely different approach and
new embodiment to improving the depth of field, which takes
advantage of the early bypass readout design that was developed to
boost the readout IC's dynamic range. In this design, the
accumulated charges are not shifted through the entire length of
the chip but are read out every two, four, eight or sixteen pixels.
In fact, the number of pixels read out can be any number from 1 to
>1,000,000. The remainder of the TDI process, splicing together
the appropriate two-, four- (or eight-) pixel segments, is done
inside or outside the chip, either by electronic circuits on the
chip and/or outside the chip and/or the image reconstruction
software and/or by field-programmable gate arrays. In hardware
and/or software, the "appropriate" segments that are spliced
together to generate the TDI image can be selected to match the
speed with which the projection of a given plane of the object
moves across the detector, thus putting that plane into focus. Over
the individual segments, the blurring amounts to no more than half
a pixel for the geometric parameters discussed above. The system
will create multiple images, each of which focuses on a different
plane of the patient's body. By implementing this solution in
software, it will even be possible to make the imaging planes
user-selectable.
[0048] An added benefit of this approach is that it will provide
depth information to the user. For example, if an image shows a
bullet lodged in the patient's body, the treating physician can
vary the depth for which the image is reconstructed until the
bullet is in focus and then use the resulting depth information to
optimize the treatment plan, or the imaged body slices can be
viewed and the one that contains the bullet will give information
on the position of it inside the body. Therefore, this embodiment
will produce three-dimensional (3D) slices of the patient's body.
These slices will be horizontal and not vertical as in standard CT
scans. However, in a new embodiment the detector array and the
x-ray source is positioned at different angles to the body of the
patient and produce 3D image slices of the patients body at
different angles and orientations. This is a new 3D linear scan TDI
approach with slice reconstruction using hardware and/or software
for accumulating, processing and producing the images directly, in
slices, in 2D or 3D and/or in stereoscopic format.
System Design Details
[0049] The dimensions of the disclosed digital radiography system
will be controlled mainly by the length of the detector arm (70 to
100 cm, as mentioned before), the distance between the source and
the detector arm (.about.150 cm, possibly less for a mobile
system), and the length of the human body (.about.200 cm). By
adding room for mechanical support structures, acceleration and
deceleration of the scanner arm to and from the required constant
scanning speed of 10 to 20 cm/s, etc., we arrive at estimated
overall system dimensions (L.times.W.times.H) of 2.5 m.times.1
m.times.2.0 m for the full-size stationary system. However, these
dimensions may vary from <1 m.times.0.5 m.times.1.5 m for
children and small animals to >5 m.times.2 m.times.4 m for
larger objects such as patient's and animals.
[0050] Another embodiment is to use another detector arm on the
side illuminated by a second x-ray source. This will lead to
construct vertical image slices. It will also produce both images
simultaneously eliminating the need to turn the patient to his/her
side if necessary. The slice information together with the two
directional views will help to produce excellent 3D images.
[0051] In another embodiment, the two detector arrays can be placed
at a viewing angle to each other illuminated by two separate
sources. If the two detectors and the two sources are correctly
aligned to the viewing angle it will be possible to produce
stereoscopic imaging of the patient's body.
Another Embodiment
[0052] Because of the practical size limitations of the fabrication
technologies for both the IC readout chip 304 and solid state
detector such as CdZnTe detector 301, it is not feasible to
fabricate a continuous linear array detector of the required length
using silicon ICs and CdZnTe material. The solid state detector can
be selected from anyone or a combination of solid state detectors
such as Silicon, Selenium, Germanium, GaAs, CdTe, CdZnTe,
PbI.sub.2, HgI.sub.2, CdWO.sub.4 and other old or new detectors. In
one embodiment the array is assembled from smaller detector
hybrids. For best performance, neither the detectors nor the ICs
can have active pixels extending all the way to the edge of the
material, so a straight array would end up having inactive areas
that would lead to gaps in any image created by the system. It is
not a good idea either to use a flat linear array 901 because of
the significant thickness of the detectors, up to 2 mm, required to
achieve high DQE in CdZnTe for digital radiography at x-ray
energies up to .about.100 keV. Different applications may require
different thickness, width and/or length for the detector array.
All the different sizes can be designed and built using this
invention. Under a fan beam most x-rays would enter such a flat
detector array at an angle and have a good chance to penetrate into
adjacent pixels and produce blurring. To avoid this scenario, the
individual detectors need to be tilted 902 so that their cathode
surfaces are perpendicular to the incoming beam or they can be
mounted onto a curved surface such as an arc. Therefore, we use
detectors that are approximately 10 mm wide by 20 mm long, which is
the maximum size for fabricating both ICs and CdZnTe with good
yield and low cost, and place them into a staggered linear array
903 with overlaps. The overlapping is important so that the pixels
at the two edges of two adjacent pixel detectors image the same
section of the object. This will improve the images on the two ends
of the pixel detectors where one expects higher number of
non-working channels. In fact, in another embodiment the staggered
array detectors can be positioned to overlap totally or mostly.
This will provide improved imaging all around as the number of
pixels observing the patient is mostly doubled. The detector array
can be mounted on a detector arm 905 with a mounting bracket 902.
The detector signals can be processed with data acquisition
electronics 904.
[0053] Abutted flat or curved linear arrays can also be used. In
this case the distance between the two detectors is made as short
as possible so that the dead area where no image is produced is
minimized or reduced to zero. Another embodiment is to make the
array from a continuous detector without any gaps. The detector
array can also be made nonlinear depending on the application.
[0054] The solid state detector 301 has pixels 310 fabricated on
its surface as electrodes. The readout chip 304 also has pixels 303
fabricated to match the pixels of the detector 301. The detector
301 is bonded onto the chip 304 through bump bonds 302. The bump
bonds can be made from a conductive material such as indium, lead,
solder, tin, graphite, copper, aluminum, silver, gold, platinum,
and conductive epoxy. The bumps can be put only one surface either
the pixels on the chip or the detector. Or it can be put on both
surfaces in any combinations such as gold with conductive epoxy.
The pixels on the detector pixel array and the ASIC are then
connected together through these bump bonds 302 and hybridized. The
top of the detector 309 can also have a single or pixelated
electrode. A voltage source is connected to this electrode or
pixels to form an electric field 308, in a format that is required
by a specific application. The x-ray 306 and gamma rays then enter
the detector substrate and produce e-h pairs 307, which move in
opposite directions under the influence of the electric field 308
and the charge goes through the detector pixel 310 and bump bond(s)
302 and is collected by the ASIC's 304 pixel 305.
[0055] The time-delayed integration concept is used for this
technique to build a low-cost, high-resolution, low-dose, and fast
digital radiography system for full-body imaging with 3D image
construction capability. This technology will add the third
dimension to the images obtained, by letting the user focus on
different depths. In this sense, the system will then offer the
option of quasi-tomographic imaging.
[0056] A compact portable version of the detector with radio signal
communication capability can be made that can be mounted on a
conscious and mobile subject. The wireless communication with the
detector may also use other techniques such as different radio
wavelengths, visible, IR, UV, microwave, mm wave, sound,
ultrasound, etc. This detector can relay images of the distribution
of a radiotracer while the small animal is doing its normal
functions.
[0057] Some of the properties of the invention are listed below:
[0058] 1. Each pixel on the readout chip is multiplexed to a single
output and read out in a raster scan mode or other modes such as in
series per pixel row and/or per pixel column, and/or parallel and
or some parts in series and some parts in parallel. [0059] 2. Fast
trigger output is produced with low jitter for coincident detection
of annihilation photons for applications such as PET. [0060] 3.
Detector pixels DC or AC coupled to the readout ASIC pixels. [0061]
4. Use a two-dimensional (2D) array of resistance and/or
capacitance chip to achieve AC coupling. This chip may also be used
without capacitors and/or resistors, or other circuitry can be
placed onto this chip such as JFET transistors. This chip is
sandwiched between the 2D detector array and the 2D pixilated
readout chip. This chip is bump bonded onto the readout ASIC and
the detector array is bump bonded onto this chip. To achieve
sandwich type mounting, this chip requires deep vias, and/or both
surface processing, which connect the resistance and/or capacitance
inputs/outputs to both sides of this chip. Or a ceramic substrate
can be used with matching metal or gold pixels on both sides. The
pixels on both sides are connected to each other through plated
through holes or vias. Capacitors and/or resistors may be
fabricated into and/or onto this chip to achieve A/C coupling of
the detector's pixel output to chip's input. This chip is bump
bonded onto the ASIC or detector and then the other one is bump
bonded onto the other side. Therefore, this chip is sandwiched
between the ASIC 304 and the pixel detector array 301. It will use
similar type of bump bonding 302 to connect to the ASIC and the
detector from both sides. [0062] 5. Detector may be designed to
have a guard ring 1102 around the perimeter of the pixels 1101 as
shown in FIGS. 11A and 11B. The detector also contains alignment
marks 1103. Pixels are made from a blocking or unblocking type
conductive material such as conductive epoxy or paint, copper,
aluminum, indium, silver, gold, tungsten and/or platinum. FIG. 11B
shows a different pixel arrangement where the pixels 1101 are
surrounded by a grid 1104 which may be connected or unconnected to
the guard ring 1102. There may be more than one grid surrounding
the pixel area. One or more grids around each pixel steer the
charge onto the pixel pad 1105. This arrangement may work as a
Frisch Grid structure on each pixel. This grid 1104 and/or pixel
pad 1105 can have any type of geometry, number, material, size,
shape, symmetric, asymmetric depending on the application. Also any
size voltage(s) can be applied to the grid(s) 1104. The grid 1104
system improves energy resolution of each pixel 1101 especially
when a voltage applied to the grid(s) 1104 to steer electrons (or
holes) onto the pixel pad 1105. The indium bumps 1106 are also
shown. [0063] 6. The area inside the adjacent pixels may be used
for placing electronics circuitry of a pixel. [0064] 7. It may have
different pixel sizes and geometries on the same pixel detector and
ASIC as required by different applications. [0065] 8. It is
designed to work at, or above or below room temperature. [0066] 9.
It may be designed to image under low and/or high x-ray and/or
gamma-ray flux. [0067] 10. It may be designed to have wide energy
range of about 1 keV to over 1,000 keV. [0068] 11. Good tolerance
can be built in to accommodate high detector leakage current.
[0069] 12. CdZnTe detectors can be used with two thicknesses, such
as less than 1 to higher than 5 mm for Gamma Camera, CT & SPECT
and less than 5 to higher than 15 mm for PET, security, industrial,
NDE, NDI, and NDT imaging. Two or more pixel detectors can be
placed on top of each other to detect both very low energy and also
very high energy x-rays and/or gamma rays. [0070] 13. Fast trigger
output is provided with low jitter for coincident detection of
photons or particles such as the annihilation photons for
applications such as PET. [0071] 14. Two or more energy ranges may
be designed, such as below 10 to above 200 keV and below 100 to
above 1,000 keV, to accommodate different types of applications
such as Gamma Camera & SPECT, PET, CT, security, industrial,
NDE, NDI, and NDT imaging. Mounting Detector Pixel Arrays onto
Readout Chips
[0072] We have discussed above mounting detector pixel arrays onto
the readout chip using bump bonding technique. In another
embodiment the detector material can be directly deposited onto the
readout ASIC (pixelated surface) using techniques such as
evaporation, sputtering and electrochemical deposition. Another
technique is to grow the detector material on the chips pixelated
surface by using techniques such as liquid phase epitaxy, single or
amorphous crystal growth. It is important to produce sufficient
thickness quickly with good electrical contact between the detector
material and the ASIC pixel input pads. This technique can produce
low cost pixel detectors.
Detector Detailed Description
[0073] Such detectors can be produced using a hybrid technology
involving a pixilated converter layer in which high Z, high density
photoconductive materials are used to achieve high gamma ray
quantum efficiency. Such materials directly and precisely convert
the absorbed energy to an electric signal. The absorber layer is
hybridized to a readout device that allows the signal to be
collected directly and efficiently, digitized and stored in a
computer as an image matrix. Materials such as Silicon, GaAs,
Selenium, Germanium, PbI.sub.2, HgI.sub.2, CdTe, CdZnTe have
potentially excellent properties as the x-ray detector and have
demonstrated this potential in work to date. Since these detectors
are hybridized to Application Specific Integrated Circuits (ASICs),
they are very compact in size and ideal for imaging x-rays. We also
disclose a mixed signal readout ASIC for directly reading the 2-D
detector array pixels.
[0074] The key element in the detector must convert the x-rays into
electronic signals. Important detector properties are dynamic
range, quantum efficiency, sensitivity, noise characteristics,
linearity, uniformity and the ability to provide high spatial
resolution over the required area. First generation detectors
employ scintillators to convert x-rays into visible light, which is
subsequently converted into an electronic signal. Because of the
intermediate conversion stage, this process is inefficient, and
could lead to higher signal noise. In a pixel detector the light
produced in the scintillator propagates isotropically and, if the
detector thickness is not sufficiently small, it can diffuse into
adjacent pixels and cause blurring. Columnar scintillator structure
has alleviated this problem but it is costly to manufacture and
produces signal loss for crystals with very narrow aspect
ratio.
[0075] Each 80 keV x-ray photon produces about 2.1.times.10.sup.4
e-h (electron-hole) pairs in CdZnTe and a 1.5 mm thick detector
will stop about 99.8% and 90% of the 50 and 80 keV x-rays,
respectively. 1.5 mm to 2.0 mm thick CdZnTe detectors 301 will be
used for this application depending on the selected x-ray generator
peak working voltage. For a typical medical radiography machine the
x-ray tube voltage can be set at up to 180 kVp. We use a 3 mm Al
filtration and the intensity maximum will be between 30 to 50 keV
depending on the power supply. In one embodiment we select an x-ray
generator with 70 to 100 kVp supply.
Slot-Scan System Using the Time Delay Integration (TDI)
Technique
[0076] In one embodiment the individual detector unit has 96 to 128
stages in the TDI direction and 192 to 256 columns in the non-scan
direction. The detectors are abuttable up to the full body x-ray
image width (.about.70 cm) or it can be abutted to any required
length. The detector pixel size will be between about, but not
limited to, 50.times.50 to 100.times.100 .mu.m.sup.2 with about
100% fill factor and the overall size of about, but not limited to,
10 mm by 20 mm for each detector module.
[0077] The readout chip also operates in a flash mode where static
staring images are obtained for detector diagnostic purposes. High
resolution of 5 to 8 lp/mm MTF depending on the pixel size and 50%
to 70% DQE at zero spatial frequency is achieved. However, better
MTF and DQE approaching the theoretical limits is possible with the
detectors developed under this invention.
[0078] Each pixel holds 3.times.10.sup.7 electrons and the readout
noise at each output buffer is 3.times.10.sup.3 electrons. Because
of the 128 TDI stages and the early bypass readout design that
boosts the effective well capacity by up to a factor of 32, the
effective dynamic range can be increased to 6.times.10.sup.4 or 16
bits. These numbers are exemplary to the prototype detectors built.
Therefore, new versions of detectors developed using this invention
have the capability to increase these numbers significantly.
Hybrid Pixel Detector
[0079] FIG. 3 shows a drawing of the hybrid pixel detector and its
internal structure. It shows the pixilated CdZnTe detector 301 301
on top with HV bias electrode. Under the CdZnTe the pixilated array
ASIC 304 is shown with indium bumps 302 at each pixel input. CdZnTe
301 pixel array also has indium bumps on its pixels matching the
pixels of the ASIC. The detector is aligned on top of the ASIC and
pressed together to form a cold bond.
ASIC Design
[0080] A custom 16.times.16 ASIC for 300-.mu.m-pitch detectors
which is designed to be abatable on two or three sides to
facilitate tiling has been developed. It is also possible to
fabricate detectors, which are abuttable on one, two (FIG. 3),
three (FIG. 5) or four (FIG. 10) sides. If the ASIC is abuttable on
two sides 4 of them can be placed abutted at the two sides to
produce a 4 times the area without image artifacts. If the ASIC is
abuttable on three sides linear arrays with two rows can be
produced any length fully abutted. If the ASIC is abuttable on four
sides then 2D arrays can be fabricated any size area fully abutted
without producing significant loss of signal or image between the
pixel detectors.
[0081] To fabricate 4 sides abuttable pixel detector (FIG. 10) the
solid state detector pixel array 1001 is mounted on an ASIC 1002
using bump bonding 1005. The hybridized detector may include a
special connection or A/C coupling chip sandwiched in between the
detector 1001 and ASIC 1002 as described above. The hybridized
detector is then bump bonded 1006 onto another connection chip 1003
or a ceramic chip 1003 similar to the connection chip described
above. This chip has deep vias or plated through holes,
respectively. It is used as a Ball Grid Array (BGA) and mounted
onto a ceramic or printed circuit board 1004 using solder balls
1007 or bump bonding. In this technique the ASIC 1002 and
connection (BGA) 1003 dimensions are smaller or equal to the same
dimensions of the solid state pixel detector array 1001. Therefore,
the hybridized detectors using BGA technique can be placed side by
side in a 2D array fully abutted to build large area detectors.
[0082] In making abuttable pixel detectors it is also important
that the guard ring 1102 around the pixelated 1101 active area of
the solid state detector is also made thin to reduce the inactive
area on the perimeter of the pixel detector. This will help reduce
the loss of signal and/or image further. In another embodiment, the
guard ring can be built or fabricated around the sides of the
detector substrate allowing the pixels to come in touching distance
to the edges of the detector. This will allow the smallest possible
gap between the pixel detectors, improve abutting and will
eliminate artifacts in the images. In another embodiment the grid
1104 can be used instead of a guard ring. Since the grid structure
has the same pitch as the pixels it can come in touching distance
to the edges of the detector.
[0083] The ASIC uses identical analog processing chains for each of
the 256 pixels and is designed to be compatible with both CdZnTe
and silicon since it is externally configurable for detection of
positive or negative charges. The identical shaping time constant
for each chain can be selected from about 250 ns to 15 .mu.s to
accommodate different charge collection times.
[0084] The chip has three parallel logic inputs that select one of
eight functional modes. In normal operation the chip waits in Mode
0 for an x-ray event to occur. When this occurs, a trigger flag is
set in each pixel where the peak-detector voltage exceeding the
threshold voltage. When one or more trigger flags is set the ASIC's
trigger output is raised, alerting the processor controlling the
chip. To read-out the event the chip is put in Mode 1, where the
pixel number and held analog peak voltage are output for each
triggered pixel, in an order determined by the chip, with the
readout timing controlled by a clock input to the chip. After the
last triggered pixel is read out, the trigger output is lowered.
The ASIC may then be put in Mode 2, where the peak voltage for
additional pixels may be read out by inputting the pixel number on
the ASIC's address bus, and pulsing the chip's clock input. After
the readout is complete, the ASIC is put in Mode 3, where
capacitors in the analog electronics are discharged, and digital
signal paths reset. This refresh cycle, which requires about 50
.mu.s, is followed by a return to Mode 0, to await another trigger.
In the absence of any triggers the refresh cycle must still be
preformed periodically (.about. every 100 ms) because of charge
build-up resulting from detector leakage current.
[0085] Mode 4 and 5 are used for disabling and enabling the
triggering in individual pixels, allowing any noisy pixels to be
masked. Mode 6 and 7 are for enabling and disabling for individual
pixels the input of an external calibration pulse into the front
end of the pixel analog electronics.
[0086] The operation of the ASIC is also configurable through an
eight-bit control register, which is input serially. This allows
selection of positive or negative input from the detector;
selection of a global enable or disable of the test input; and
selection of a single external trigger enable/disable or the
individual pixel trigger enable/disable. In addition for testing of
individual pixels, voltage output from a single pixel can be
selected, with the voltage either from the output of the shaper
circuit, or the peak detector held voltage.
ASIC Performance
[0087] It is desirable that electronic noise be minimized. This
noise is a combination of chip noise coupled with the capacitance
of the CdZnTe pixel and its leakage current and noise injected from
peripheral circuits such and power supplies, voltage references and
noise pickup. Fortunately the CdZnTe pixel capacitance and its
leakage current are proportional to the area and the volume of the
pixel, so that the fine <10.times.10 .mu.m.sup.2 to
>10,000.times.10,000 .mu.m.sup.2 pixel pitch can be used. The
fine pitch reduces both the leakage current and the pixel
capacitance. The ASIC's input MOSFET transistor for the charge
sensitive amplifier has been optimized through SPICE simulations to
achieve minimum noise. In addition, the transistor gate size is
designed to match the detector input capacitance.
[0088] The input charge sensitive amplifier is designed for ultra
low noise operation. For example, a pure integrating charge
sensitive amplifier, a self resetting charge sensitive amplifier
and/or a transconductance amplifier, etc. can be used to achieve
low noise. However, pure integrating type amplifier requires very
low detector leakage current because the leakage current is
continuously integrated and the baseline increases requiring
periodic resetting of the whole chip to keep the energy resolution
high.
CdZnTe Pixel Array Fabrication
[0089] The electrodes are formed on the CdZnTe 301 after polishing
and pacification of the surface. The detector is then indium or
solder bump bonded 302 on to the electrodes 303 or pixels of the
CdZnTe crystals. Other bonding techniques such as gold or platinum
stud bonding, asymmetrically conductive epoxy or film bonding may
also be used. The pixel array can consist of 2.times.2 to over
1,000,000.times.1,000,000 array of less then 10.times.10
micron.sup.2 to over 10.times.10-mm.sup.2 pitch conductive such as
gold, platinum and/or indium pixels (FIG. 3). The pixel geometry
can be square, rectangular, circular, elliptical or any other
shape.
[0090] After the crystals had been prepared, indium bumps were
deposited both on the crystal readout pads and on the corresponding
ASIC pads. Using alignment marks, the two were then pressed
together to fuse the bumps, a process which takes place at room
temperature. If necessary, an underfill of insulating epoxy can be
used between the ASIC 304 and the CdZnTe 301 to provide additional
support and provide a more robust assembly. In practice, with a
large number of small pixels, this is not usually necessary and was
not done with these pixel detectors. Instead of underfill it is
also possible to epoxy the corners or the side edges to secure the
solid state detector pixel array onto the readout ASIC.
CdZnTe Material Quality and Leakage Current Compensation
[0091] It is important that the resistivity of the CdZnTe material
used for this pixel detector is high. Although the front end charge
sensitive amplifier of the new ASIC is designed to tolerate higher
leakage current lower leakage current will result in better energy
resolution. The front end amplifier of the readout chip has a reset
circuit which is a resistive feedback. This circuit cancels the
leakage current coming in and keeps the input amplifier's output or
offset from changing as the leakage currant from the detector is
integrated. The feedback circuit can be made in many different
ways, such as an active and/or passive resister with or without a
capacitor, one or more MOSFET or other type of transistors,
etc.
Pixel Detector and Readout ASIC Design Features
[0092] Staring array applications may need mounting of the
individual pixel detectors in a 2D array where each pixel detector
is abutted next to the other pixel detectors. If the two sides of
the readout chip then up to four chips can be abutted to produce an
active area by up to 4 times larger then the individual chip active
area. If the readout chip is abutable on three sides then linear
arrays of any length can be made with one or two chips wide. Or in
another embodiment any size 2D array can be formed by placing the
wire bonding section of the chip under the chip below it. This
requires that the pixel detectors are slightly tilted in order to
give room for the bonding of the chips before it. This tilt is
small and does not produce significant artifact in the image.
[0093] Scanning type applications may require the Time Delayed
Integration (TDI) technique to be used. The scanning requires a
linear array. The linear array may be based on a single chip
abutted into a linear array or two chips abutted on opposite ends
into a linear array. The linear array may also be formed by using a
staggered array 903 of chips and also may be formed by much wider
width using three to more chips along its width and/or TDI motion
direction mounted as described above.
Input Amplifier
[0094] The input amplifier 403 may have many different forms, such
as a charge sensitive amplifier 403, a transconductance amplifier,
an operational amplifier, current amplifier, voltage amplifier, low
noise amplifier, etc. The charge sensitive amplifier requires large
gain so that the small amount of charge created inside the detector
can be amplified and detected with good energy resolution. This
also means that it must have low noise. One embodiment of the
chip's input amplifier described is a pure integrator with only
capacitative 410 feedback. Although such a circuit has low noise it
is susceptible to continuous leakage current from the detector. The
leakage current is integrated continuously and the amplifier output
keeps rising and any signal received rides over this pedestal and
its pulse height, which is proportional to the x-ray photon energy,
is added to the pedestal. The solution to this problem is to use
low leakage current detector and/or to reset the amplifier
periodically to bring the pedestal to zero or use a charge
sensitive amplifier, which is not highly sensitive to the leakage
current.
[0095] One embodiment is a charge sensitive amplifier with a
resistive 603 feedback in parallel to the capacitative 604
feedback. The feedback capacitance is in the range of a few fF to
tens of fF depending on the required dynamic range. The presence of
the resistive feedback cancels the DC current component which is
due to the leakage current and does not allow the input amplifier
base line or pedestal, as sometimes called, to rise. In effect, the
base line signal level at the output 613 of the input amplifier
changes to compensate for the DC leakage current due to the
resistive feedback. This, of course, is only true within the design
range. Once we determine the CdZnTe resistivity range to expect we
are able to set the upper limit for the maximum leakage current
expected by designing the resistive feedback circuit to compensate
for it. There are several different embodiments to do the resistive
feedback circuit.
[0096] The first embodiment is to use a passive resistance in the
feedback circuit. This is limited on how much resistance can be
built into the chip at the feedback loop of the input amplifier at
the input of each channel.
[0097] The second embodiment is to use resistive multiplier
feedback circuit, which is an innovative way to create a linear
element (a resistor) with active MOSFET components. This circuit
uses current mirror divider circuits configured as an extremely
small transconductance amplifier to make an equivalent resistor
greater than 200 MOhm (2.times.10.sup.+8 Ohms). A small 100 KOhm
resistor is used as the reference device for this circuit. A
measurement of current going through the 100 KOhm resistor is made
and a new current level of 1/200th the magnitude is driven at the
output pin. The resistive multiplier circuit achieves high
resistances in a small area because it replaces the large area
required for a large linear resistor (>10.sup.8 Ohms) with a
circuit of area 0.042 mm.sup.2. A resistor of 200 MOhm made from 1
kOhm/square (Ohms cm length/width thickness) polysilicon would be 1
mm.sup.2 in size for a 0.5 .mu.m process. Also, a large resistor
element would suffer from RC delay effects, which the active
circuit is not subject to.
[0098] The third embodiment is to use a MOSFET transistor as a
resistive feedback circuit with its resistance is externally
controlled. Such circuitry can be used to produce a large
resistance feedback component, >10.sup.9 Ohms. It has a much
smaller area, about 0.001 mm.sup.2. Therefore, it is especially
appropriate to use for smaller pixel sizes such as 250.times.250
.mu.m.sup.2 pixel pitch. It can be added as an option to the
resistive multiplier circuit during the prototype fabrication so
that it can be tested and evaluated in comparison to the resistive
multiplier circuit. Unfortunately, MOSFET resistive element is
nonlinear and also it may not be used with the pole-zero
cancellation circuit.
[0099] The fourth embodiment is to use a combination of these
feedback resistance circuits. The capacitance feedback may also be
added to these different feedback circuits.
[0100] The fifth embodiment is to use a circuit to sense the
variations in the input amplifiers output and restore the output to
a preset value. In this case the sensor circuit need to respond
slower than the pulses coming from the detector so that it will
allow the detector pulses while restore the slow changes in the
output due to the leakage current coming from the detector.
[0101] The sixth embodiment is to use a circuit to anchor the
output of the input amplifier to a fixed value such as ground. This
circuit is made to allow the fast detector pulses to go through as
it will be designed to be slower to respond than the detector
signals.
[0102] The normal input amplifier size is expected to be about
0.023 mm.sup.2 for about 9 pF input capacitance. This size can be
reduced further. We expect this to be feasible, especially because
the input capacitance is <1 pF for CdZnTe pixel detector
application and the optimized input gate size to match the input
capacitance is, therefore, significantly smaller. The presence of a
resistive element in the feedback may increases noise.
Pole-Zero: Dc Coupled Feed Forward Circuit
[0103] A pole-zero cancellation circuit, well known in the art, is
built to improve the response to high count rate applications which
are expected for medical imaging. This circuit restores the
baseline and prevent pulse pileup when count rate approaches the
limitation imposed by the shaping time. The area of this circuit is
expected to be about 0.06 mm.sup.2 in general and can be reduced
further. However, a good place to put this circuit is after the
input amplifier 614.
Two Pole Shaper Circuit
[0104] A two pole shaper circuit 405, 606 is built. This shaper
circuit can be designed using many poles. This circuit can have 2
to more than 256 selectable different peaking times starting from
less than 0.001 .mu.s to over 100 .mu.s or more. The standard
circuit has 0.08 mm.sup.2 area with 8 selections. We reduced the
number of selections and the circuit area so that the pixel pitch
can be minimized. The aim for this circuit is to produce high
energy resolution. Therefore, speed is not an important issue.
However, the peaking time selection is important for compromising
between energy resolution vs. count rate. This circuit can be
placed at many different places.
Comparators
[0105] A comparator circuit 407, 607 is used to produce an event
trigger using the shaper output. This is essential as this trigger
signal notify external circuitry that an event has been received so
that an external process that reads out the ASIC can begin. It is
also recorded into the pixel address register and the sparse
readout circuit or hit register. The input amplifier/shaper base
line level is restored by a programmable 8-Bit DAC offset control
and the comparator threshold is adjusted by a programmable DAC so
that a uniform response can be achieved through calibration. The
expected area of the comparator circuit with adjustment circuits is
about 0.0095 mm.sup.2 for the comparator and 0.0427 mm.sup.2 for
the DAC if a 0.5 micron process is used. These circuits are
simulated, minimized and optimized for this application. The size
of the DAC can be also reduced by lowering it to 6-Bits or lower if
necessary.
[0106] A second or even more comparators for lower and/or upper
level discrimination may also be added to the circuit for each
channel. In this case the comparators are set, some below the peak
and some above the peak and some below and above background
region(s). When a trigger is received from the first comparator the
external circuit looks for a flag from the second or other
comparators if it is absent it expects that the event is within the
peak and reads it out. If there is a flag then it ignores the event
and resets the ASIC. This method is especially useful if the
radiopharmaceutical applied has several peaks and a peak below the
highest is being observed. Also it will allow easy background
subtraction if a background region and a peak region are observed
at the same time. This also allows increased data acquisition rate
capability as unwanted data can be ignored and the chip is reset.
If the second or any other comparators are deemed not necessary for
this application they may be omitted to keep pixel area small.
Fast Comparator
[0107] One or more fast comparators may be used. The circuit can be
a leading edge comparator and similar to the one that is used
above. It has programmable DAC based threshold adjustment. The
estimated area is similar to above. Improved fast comparator
circuits such as a constant fraction discriminator, zero crossing
discriminator, rise time compensated discriminator or comparator
can be incorporated into the chip. The fast comparator produces a
fast trigger signal, which may be used for the event arrival time
determination or coincident event recording.
Peak Detector
[0108] Peak detector 406, 609 circuits are well known in the art.
This circuit holds the signal in each pixel at its maximum value
until the pulse height is measured and the circuit is reset. This
circuit is critical for good energy resolution. Other circuits such
as track-and-hold, peak-hold, track-and-hold and/or sample-and-hold
circuits may also be used. The expected area of the circuit is
about 0.039 mm.sup.2.
Sparse Readout
[0109] The sparse readout capability is important for 2-D pixel
detector circuitry due to the large number of pixels. The sparse
readout system allows the external circuitry to readout only the
pixels which have data or produced a trigger. In this application
the majority of the events are expected to trigger a single pixel.
Therefore, sparse readout saves significant readout time and
software overhead later. In this circuit, the peak value of the
first pixel which has produced a trigger is connected to the output
and its address is also made available to the outside world. The
external circuitry downloads the address and digitize the pulse
height and inform the chip it is done. Then the chip automatically
drops to the second pixel, which has produced a trigger, if there
is one, and display or output it. The circuitry for this feature is
outside the pixel area.
Near Neighbor Readout
[0110] In a small pixel 2-D detector it is quite likely that the
charge is shared between adjacent pixels. Therefore, one or two or
three neighboring pixels may produce a trigger but the other pixels
surrounding them may also have some charge. Therefore, to get the
maximum energy resolution it is good to read out the charge
collected by the neighboring pixels. In this mode, the chip
produces a hit register, which shows which pixels produced a
trigger. The external circuitry downloads the hit register and then
it automatically reads the pixels, which produced a trigger and
also the neighboring pixels around them. This is only necessary if
probability for charge sharing is significant, which is true for
small pixel size and thick detectors. If charge sharing is not
significant the near neighbor mode can be turned off. The circuitry
for this mode is external to the pixel area. This circuit can also
be developed in hardware on the chip. It can have a lookup table,
which tells the circuit which pixel contents it should connect to
the outside together with its address.
Other Circuits
[0111] A test pulse circuit at the input of each channel or pixel
is implemented so that each pixel or channel can be tested by
applying an external pulse. Also the input of each pixel can be
turned on and off. This is useful during testing and calibration. A
polarity selection circuit 404 may be built in so that the input to
the amplifier can be selected to be electrons or holes. These
circuits are designed inside the pixel area but has minimal impact
on the pixel size. We also built two or more additional full
featured analog channels connected to the chip readout system on
the periphery of the ASIC outside the pixels with inputs connected
to wire bond pads of the ASIC. These analog channels are mainly
used to readout the cathode electrode deposited onto the top of the
detector to determine the depth of interaction of the x-ray or the
gamma ray using the pulse height ratio between the anode and the
cathode or the time difference between the pulses received between
the anode and the cathode on the detectors. Other channel(s) can be
used to process the signal from the guard ring and the
non-collecting grid. These peripheral channels may be connected to
pads on the perimeter of the chip. They may be also connected to
external detectors for testing. The circuitry for the depth of
interaction determination such as cathode/anode pulse height ratio
can be built onto the ASIC or it can be done externally.
[0112] Current integrating circuits may be used instead of or
combined with the pulse counting circuits on the readout chip at
each pixel. This will allow selection of either function or use
both functions at the same time. The current integrating circuitry
normally have a noise and/or other background signals that cannot
be subtracted, however, it can accept extremely high rates if
necessary. The photon counting detectors normally have non or very
low background as the noise or other unwanted external pulses can
be eliminated using the comparator(s) and/or discriminator(s).
[0113] Analog to digital converters (ADCs), digital to analog
converters (DACs), state machines, programmable gate arrays and
microprocessors may be included on the readout chip to carry out
functions that are done presently outside the chip. This will
reduce total size of the readout electronics and make very compact
detectors, such as detectors for in vivo molecular imaging of
mobile and conscious small animals.
[0114] Token Logic circuit 611 is included for each pixel cell.
Initially the token registers are reset. Then each of the cells
gets a pulse from the comparator at different times. The token
present signal is initially in the first cell. Upon receiving the
token clock, the token is passed from one cell to the next until
all cells have been read. Then the last token out signal goes to a
one and to an output signifying that all cells that received a hit
have been read.
[0115] A driver or a buffer circuit 610 is used to output the
analog peak detector signal to the output 612. Token logic 611
output can be used to control the data output 612 to a bus 615. The
data bus 615 can be a tri-state bus. The driver circuit can also be
a differential driver 408 with differential output 409.
A Two-Dimensional Pixel Readout Chip with Self Resetting Input
Amplifier
[0116] The DANA integrated circuit is designed for direct
connection to an array of detector diodes and/or pixels. Each diode
of the detector array is mated to a unit cell contained in the
integrated circuit through a bump connection 601 assembled in a
mechanical process. The resulting contact allows low parasitics and
high-density detectors for desired applications.
[0117] Each unit cell contains a preamplifier 614 with active or
passive feedback, gain stage 605, pulse shape processing circuits
606, a trigger event detector, comparator 607 and a peak height
measuring circuit 609. A token readout circuit 611 also contained
within the unit cell allows the readout of specific cells with
recorded events for a fast detection event duty cycle. The output
of the peak detector 610 goes into a tri-state analog buffer 610.
The output 612 of the buffer 610 goes into a tri-state bus 615.
[0118] Each pixel contains electronics to detect and process a
detection event associated with a single detector. Bond pads 305
are located on one side of the chip to accommodate multiple
detectors in the same system, one sided wire bonding 311 and tight
integration. This is forms a three sides abuttable pixel
detector.
[0119] Using deep vias it is possible to bring the electronics
inputs and outputs to the other side of the ASIC 304 so that a four
sides abuttable pixel detector (FIG. 10) can be fabricated. In this
case it will be important to integrate all the peripheral
electronics, not used inside the pixels to be placed in the
pixelated area so that the ASIC 304 area can be matched to the area
of the solid state detector pixel array 301.
[0120] Each unit cell contains all the necessary electronics to
sense, process, and readout events delivered from the attached
detectors.
[0121] The imaging using the disclosed detectors can be done in
different ways. The pixels can be read out in any desired sequence
as the detector and/or the object moves. The detector moves to a
new position stops and the pixels read out and the process is
repeated. The pixel detector moves continuously and the pixels read
out in columns using TDI technique in the TDI direction which is
the reverse direction to the movement of the object or the
detector. The detector is not moved and the images are read out in
the staring array mode in any desired sequence of time intervals.
The motion of the detector may need to be followed or matched by
the photon or particle generator. The motion of the detector and/or
the object may follow different type of paths.
Main Features:
[0122] 256 channels or pixels 303 arranged in a 16.times.16 matrix
at 500 um pitch. Other pixel ranges from <2.times.2 to
>4,096.times.4,096 can be used. [0123] Self-resetting charge
sensitive amplifiers 614. These amplifiers also have an active or
passive feedback that includes a passive or dynamic resistor 603
and/or capacitance 604, and or a MosFET transistor feedback
element. [0124] Count rate can range from 1 photon/sec/pixel to
1,000,000 photon/sec/pixel depending on the chip design. Therefore,
maximum count rate can be above 400E3 photons/s for all channels in
parallel. [0125] Gains and offsets of each channel are digitally
adjustable. [0126] Input energy range <1 keV to >150 keV.
Using thick detectors the input energy range can be >1,000 keV.
This is true for both input polarities. [0127] Selectable shaper
filter shaping time, <10 nsec to >40 usec in more than 16
steps for noise tradeoff versus detection rate selection. [0128]
Data readout is controlled by a programmable token logic 611.
[0129] Additional data read options are provided by an .gtoreq.8
bit data bus. The bus can be serial, or parallel. Number of bits
depends on the design and is not limited. [0130] Multiple chips may
be used in the same system by daisy-chaining up to 16 or more chips
that share a common analog output. The outputs use tri-state
buffers to avoid conflicts. [0131] A test signal input can be
connected or routed to the detector inputs or pixel inputs 601 of
any channel or combination of channels. [0132] An analog monitor
output can be connected to any channel or pixel to display its
analog signal. This system can look at the pre-amplifier 614, gain
stage 605, and/or shaper amplifier 606 outputs selectively or all
together. This is important for testing, characterization and
calibration. [0133] The cell token logic 611 ensures that only one
peak detect output is put on the analog bus at a time. [0134] Power
consumption: 1500 mW nominal. Much lower power is possible <1 mW
to >1500 mW. [0135] Input referred noise <200 e rms. [0136]
Die size: .apprxeq.8.575.times.9.535 mm.sup.2 sized to accommodate
tight integration of multiple detectors. The die size depends on
the pixel array size and the pixel pitch and has no other
limitation then design and application requirements. [0137] Input
pad spacing: 0.15 mm to allow 57 pads on one edge of the chip.
Dual, triple or higher pad arrays can be made for I/O connections.
Input pad size and dimensions has no limitations other than design
considerations. [0138] Active low (internally pulled down) chip
enable that allows for connecting the data busses of several IC's
in parallel. [0139] All digital inputs and outputs are compatible
with standard 3.3V logic (LV or LVDS). Different processes can be
used to design and fabricated the chips. These designs may have
different voltage requirements.
Sequence of Reading Out Events
[0140] The DANA chip contains logic for detection and readout of
events. The following sequence of events characterizes the
operation of the DANA chip. [0141] 1. The array is placed in
detection mode waiting for an event [0142] 2. An event occurs: an
analog level is triggered causing the wired-or trigger signal to
assert [0143] 3. External logic, decides to either read out the
array immediately or wait a predetermined amount of time for
required detection purposes. [0144] 4. The predetermined readout
operation is performed to collect the analog peak signals, with or
without the nearest neighbors, or in some other predetermined
pattern. [0145] 5. The array is then reset and the operation
repeats to #1
[0146] The above sequence should be completed in a relatively short
period of time so that as much as 500,000 loops can be completed
per second under random pulse conditions. The control logic present
within the unit cell was designed to try to achieve this fast frame
rate and also maintain simplicity and ease of use.
Multiple Energy Digital Radiography
[0147] FIG. 7 shows the electronics used in one of the channels of
a digital radiography system. The input comes from a detector pixel
into the input 701 of each channel. The input can be also turned on
and off so that the chip does not process any input signals until
the inputs are turned on. The signal is then amplified by an
amplifier 702, which is a charge sensitive type, it has feedback
element 703, a gain stage 704 may be used, and the output of the
amplifier and/or the gain stage goes to multiple comparators 705,
706 (two shown here). The comparators 705, 706 has thresholds 707,
708 controlled either by internal DACs or external circuitry, the
outputs of the comparators 705, 706 goes to counters 709, 710,
respectively. The counters are read out in sequence from previous
stage 711 to next stage 712, or in parallel. The counters in
different channels may also be linked in the same way for
readout.
[0148] The digital radiography system (FIG. 7) may also be used as
a multi energy imaging system. A comparator 705, 706 type
electronics readout system can be used as described above to divide
the incoming x-ray or gamma ray photons into different energy bins.
The is number of photons in each energy bin is then counted 709,
710 separately and separate images are produced using the photon
counts in different energy bins. The separation of the photons into
separate energy bins can be done in different ways, such as using
comparators as described above, pulse height information, overflow
wells or capacitors, leaky wells or capacitors, transistor ladder,
etc. The counting of the pulses from the detector pixels can be
also done in different ways such as analog counters, digital
counters, mixture of analog and digital counters, TDI for each
energy bin, etc. If small sections of TDI are used such as N
sections then the count rate at the beginning pixel of a TDI
channel with total TDI counts of C will have C*(1/N) counts. The
next channel will have C*(2/N), so on. Therefore, the digital or
analog counters at the beginning of each TDI section will use a
small number of bits. And more and more bit counters will be
required as the TDI section progresses down the TDI direction.
Since the counter sizes are different along the TDI section the
larger counters can share the area under the small counter pixels
to reduce the size of the pixels.
[0149] The analog counter system is shown in FIG. 13. The
charge-sensitive input amplifier 1301 of the cell shapes the charge
pulse received at its input with the help of the feedback circuit
1313. This amplifier 1301 is followed by an optional gain stage,
which is not shown. In order to determine the energy deposited by
the photon, the charge pulse is then discriminated by a number of
comparators 1302, 1312, 1314, each of which is set to a different
threshold and has associated with it counter circuitry to determine
the number of photons that exceed the threshold in a given time.
When the comparator output goes high a one shot circuit 1303
produces a small positive pulse, which is used by the analog
counter circuit proper. This circuit consists of a very simple
single stage amplifier 1307 with a large capacitive feedback 1310.
On the rising edge of the comparator pulse, charge flows through
the diode D1 1306 and is integrated on the capacitor C2 1310. The
capacitor current generated by the falling edge of the comparator
pulse is blocked by D1 1306 and instead flows to ground through D2
1305, allowing C2 to accumulate a charge that is proportional to
the number of pulses counted from each event. Periodically, the
system will poll the results of the ASIC. After readout through
Readout logic 1308, the array is reset and the timing is repeated.
The readout system, after reading the lowest threshold counter
result and find a very small value present, it may decide that
there is no need to read the rest of the counters in that cell to
save time.
[0150] We disclose a dual or multiple energy level detector system,
using high uniformity CdZnTe or other comparable materials for
digital mammography and radiography. Dual energy subtraction
techniques offer an alternative approach to conventional mammogram
images as they allow tissue images to be separated according to
their intrinsic attenuation of x-rays. In this approach, high and
low x-ray energy images can be acquired simultaneously without any
delay and using the same detector array without spatial differences
or variations, and the high and low energy images can be subtracted
from each other in a weighted fashion. This technology also allows
for multiple energy segmentations, which would be useful to make
multiple energy separated images. Using this technique it may be
possible, for example, to subtract the bones or certain kinds of
tissue out of a radiogram to produce higher quality images.
[0151] The disclosed system would have DQE >85%, limiting
resolution .gtoreq.8 lp/mm, ultra-low noise, linear response and
high dynamic range of >15 bits. This system's detector is
produced using a hybrid technology consisting of a converter layer
in which high Z, high density photoconductive materials such as
CdZnTe are used to achieve high x-ray quantum detection efficiency.
The detector electronics are hybridized to a CdZnTe pixel array 301
to fabricate a CdZnTe pixel detector. Each incoming pulse is
processed and separated based on its energy and then digitized and
stored in a computer as an image matrix.
[0152] Dual energy radiography is a technique that is able to
resolve tissue types, which vary in effective atomic number
(effective Z) by acquiring two images, each at a different x-ray
energy. Because tissues with substantially different effective Z's
(i.e. soft tissues, fat tissue, dense tissue, calcific tissues or
other contrast agents) demonstrate nonlinear energy dependencies in
terms of x-ray attenuation, images acquired at two different
energies allows the numerical elimination of one tissue type,
resulting in more contrast of the remaining tissue type due to the
elimination of structured noise. Dual energy x-ray mammography may
be performed subsequent to the injection of an iodine-based x-ray
contrast material. After some period of time (measured in minutes)
two images at two different energies are acquired in rapid
succession. These images are manipulated (using weighted
logarithmic subtraction techniques performed pixel by pixel), but
due to the slight time delay between the acquisition of the high
and the low energy images, motion artifacts can occur. Motion
artifacts can reduce the effectiveness of dual energy subtraction,
and be a significant source of noise which can obscure the presence
of lesions. Therefore, there is the need of an x-ray detector
system, which is capable of simultaneous dual energy imaging
capability which eliminates delays and motion artifacts.
[0153] An alternate single-shot technique for producing dual energy
images of the breast is to use stacked detector systems, in which
the detector immediately underneath the object detects the lower
energy x-rays which pass through the object, and the second x-ray
detector detects the higher energies because the lower x-ray
photons have been filtered out by the first detector and by the
placement of a metal foil between the two detectors. The problem
with the sandwich detector approach is that a large number of x-ray
photons, which have already passed through the object (and
therefore contributed to radiation dose of the breast) are absorbed
in the metal filter, which effectively wastes this radiation dose.
The use of sandwich detectors is also known to be less effective in
terms of energy separation, and thus images acquired in this
fashion generally suffer from lower signal-to-noise ratios in the
dual energy subtracted images. Since the two images are obtained by
separate detectors and there are detector response variations and a
magnification factor difference the two images may require image
manipulation for accurate image subtraction. This invention
eliminates the need for image manipulation and spatial variations
and differences.
[0154] We disclose a system and method, which is capable of
single-shot imaging (without motion artifacts and spatial
differences between two exposures) with superb energy
discrimination with externally adjustable threshold between the low
and high energy images. It also combines the benefits of two-shot
and one-shot dual energy imaging modalities with none of the
compromises mentioned above.
[0155] A high uniformity CdZnTe (or other suitable material)
detector with low leakage current is hybridized to an IC readout
chip developed specifically for dual or multiple energy digital
radiography, using a pulse processing, electronic energy
separation, digital counting and Digital Time Delayed Integration
(DTDI) technology.
Dual Energy Digital Mammography
[0156] Dual energy digital radiography is a subtraction technique
based on different attenuation characteristics of soft tissue and
calcifications. Dual energy mammography can provide an alternative
approach to the visualization of micro calcifications or other
contrast agents, which are the key features for identification of
breast cancer. With this technology, high and low energy x-ray
images are acquired and subtracted with weight so that the
background image clutter from other tissue structures such as
ducts, vessels and soft masses are suppressed so that the region of
interests can be better identified. The direct conversion detector
system, which is composed of a high-Z, high efficiency x-ray
absorption layer integrated onto an IC signal-processing chip where
each charge pulse converted from incoming an x-ray photon is
amplified, shaped, energy segmented and counted for. FIG. 3 shows
the schematic drawing of a direct conversion detector and how it is
connected to a readout chip through the indium bump 302 bonding
technology.
[0157] We use a photon counting technique with a digital counter
architecture, which is different from our present MARY readout IC,
which uses a current integrating and CCD based charge transport
architecture to achieve Time Delayed Integration (TDI). TDI
technology, compared to flat panel technology, has the advantage of
averaging contributions from many other pixels in the scanning
direction (192 in our MARY chip) and therefore, can tolerate more
bad pixels than the flat panel case. In FIG. 7 a top level
schematic diagram of a single pixel cell of the chip with dual
energy channels and digital counters 709, 710 is shown.
[0158] When an x-ray photon is captured by a solid state detector
301, a charge bundle 307 is generated. The charge bundle moves
under the effect of the bias voltage 308 applied between the anode
309 and the cathode pixels 310. The field may also be reversed so
that cathode 309 can be on the top and anode will be the pixels
310. In a CdZnTe detector normally electrons are collected but
alternatively holes could be collected, too. Therefore, the anode
is normally pixilated with same pixel array geometry as the readout
chip. The cathode is normally a single electrode where a negative
voltage is applied as the bias (this is opposite to what is shown
in FIG. 3B). When the electrons reach the anode they produce a tiny
charge pulse. The number of electrons in the pulse is directly
proportional to the energy of the incident x-ray photon. This pulse
is then coupled to the pixel cell of the readout chip through the
indium bump bond 302. The charge-sensitive input amplifier 702 of
the cell integrates and also shapes this pulse through the use of a
feedback circuit 703 with the right frequency response and low
noise in order to successfully detect the size of the charge
packet. The height of the shaped pulse is directly proportional to
the number of electrons and/or holes in the pulse and, therefore,
also directly proportional to the energy of the incident x-ray
photon. Therefore, the x-ray photons with different energies can be
discriminated and counted if the height of the pulse is selected or
determined. The input amplifier 702 is followed by an optional gain
stage 704. In order to determine the energy deposited by the
photon, the charge pulse is then discriminated by a number of is
comparators 705, 706, each of which is set to a different threshold
and has counter circuitry 709, 710 to accumulate the number of
photons that exceed the threshold in a given time.
[0159] The DTDI circuit will be achieved by transferring the counts
from pixel to pixel rather than charge as is done in the present
MARY chip. A readout logic circuit controls the DTDI circuit. In
DTDI mode, the counts in each counter is passed to its neighboring
cell in the opposite direction to the scanning motion, and only the
counts in the last cell of the DTDI chain is read out. The TDI
concept for charge transfer systems is described above.
[0160] We are using a slot-scanning technique, which is different
from a flat-panel approach. However, using these chips in the
staring array mode a panel type detector can be fabricated to image
the object without scanning similar to a flat-panel detector. To do
this a 2D abutted array of pixel detectors will be fabricated.
Tomography and Stereoscopic Imaging
[0161] These detector systems can also be used to produce
tomographic imaging by rotating it around the object or the
patient. The rotation can be helical or spiral. Or two detector
systems can be used at right angles to obtain images in orthogonal
directions. Another possible mode is to use three detector systems
to cover the three sides of the object under investigation. Or four
or more detectors can be used to cover all sides of the object.
This is another way of obtaining tomographic images. In this mode
the rotation may not be necessary or may be limited to an arc. One
or more x-ray generators may be used for these modes of operation.
Another possible mode of application can be scanning in the
direction of two visual angles using two detectors viewed buy two
x-ray sources. If the visual angles are set correctly a
stereoscopic images can be produced of the object or the patient
under investigation. Many different variations on positioning the
detector systems, scanning and obtaining different types of images
are possible.
Preamplifier, Shaper and Discriminator Stages
[0162] The input amplifier 702 consists of a common cascade single
ended preamplifier stage with a continuous-reset capacitive
feedback. This cell provides high gain and will be optimized to
match the input capacitance of the detector pixel. It also contains
a test input circuit, which is used only for testing and
characterization purposes.
[0163] A DC feedback 703 current proportional to the output pulse
height is provided by a linearized trans-conductor circuit
typically used for g.sub.m-C filters. This provides the DC feedback
703 current at a rate proportional to the preamplifier output. In
effect, it acts like a resistor but is referenced to an externally
supplied reference voltage. An external current input is used to
control the ratio of feedback current and thus the input amplifier
time constant.
[0164] The preamplifier is followed by a gain stage with
programmable gain and offset. It provides voltage gain from a
factor of about <4 to up to about >20. It is selectable
through control of three bits of programmability in each cell. The
large feedback resistor value is available in a 0.35 micron process
through the use of a high resistive poly option providing 1.2
kOhm/square. The baseline level at the amplifier output can be
adjusted via a six-bit offset DAC. The gain stage may be omitted if
necessary to reduce pixel area by careful design of the
preamplifier so that it incorporates the gain stage.
[0165] Disclosed unit cell block diagram (FIG. 7) is showing all
the required blocks except the offset adjustment circuit which is a
part of the gain stage. The offset adjustment circuit may not be
required if the signal path is capacitively coupled or the
comparators are designed to have built in offset adjustment
circuitry.
[0166] One embodiment uses a CMOS-inverter based high-speed
comparator. The threshold of a CMOS-inverter based comparator needs
to be set dynamically, but this can be accommodated by inserting a
small time gap between the count gate intervals used for successive
image frames. The benefit is a comparator of very high speed and
very small circuit area.
Digital Counter and DTDI Readout
[0167] The digital counter 709, 710 is made small due to a
flip-flop circuit design created specifically for use in a ripple
carry counter configuration. This flip flop circuit contains a
minimum number of transistors and is configured as a simple master
slave circuit configuration. However, in this configuration, the
slave latch is also able to drive the master latch of the next
stage even though both latches are identical. This produces a
compact design. Advantages of the new counter design are listed
below. [0168] 1. Requires a small number of transistors while still
retaining required functionality including the reset function and
the readout function. This is important to fabricate pixel
detectors with small pixel area or pitch. [0169] 2. Can be arrayed
easily due to the symmetrical use of the master and slave latches
in the flip-flop circuit. When the counter bit circuits are arrayed
in a ripple carry adder, the master latch drives the slave latch
while the slave latch drives the master latch of the second counter
bit. This is accomplished by creating a counter flip-flop that has
a high impedance input for both the master and slave latches and
has a low impedance output for both master and slave latches as
well. A standard flip-flop circuit has buffer stages at the output
to solve this problem but this solution is non-symmetrical,
non-differential, power hungry, and requires more circuit area.
[0170] 3. This circuit is differential. This improves noise
performance of the readout circuit as a whole due to the decrease
in transients within the array caused by the differential circuits
used therein.
[0171] FIG. 8 is a block diagram of a 160 cell DTDI circuit with
two counters in each cell. There is no limitation on the number of
cells and/or comparators and/or counters per cell. Therefore, same
design can be extended to include more than 2 counters per pixel or
channel. In Cell 0 802 the counter 1 803 and the counter 2 804
starts with 0 counts 801. Counter 1 803 counts the lower energy
x-ray photons and the Counter 2 804 the higher energy ones. The
counters increment every time an event is observed or the input
pulse causes the comparator to fire or turn on. In DTDI mode
Counter 1 803 in each cell is transferred in opposite direction 813
to the scan direction 814 into the Counter 1 806 of the adjacent
cell 1 805 by the use of the transfer clock 815. Counter 2 804 is
also transferred similarly into the Counter 2 807 of the
neighboring cell by the use of the transfer clock 815. The two
counters may have separate transfer clocks. This process continues
through all the following 2-158 channels 808 and comes out from the
last cell or channel 159 809 counter 1 810 and counter 2 811 at the
outputs 812. This system works without a need to have an adder
circuit because when the counts accumulated is transferred from
cell 1 802 Counter 1 803 to next cell's counter 1 it already
contains the counts from the first cell. Therefore, when the counts
are in the second cell's counter 1 806 this counter starts counting
from the counts accumulated inside counter 803. Therefore, in this
embodiment just transferring counts in sequence in the TDI
direction 813 opposite to the scan direction 814 is sufficient to
produce a natural DTDI process with minimum electronics to achieve
small pixel size.
[0172] In another embodiment counter 1 803 and counter 2 804 in
cell 0 802 start with the maximum counts and each event decrements
the counters. Then the counters are transferred to adjacent cell's
counters similarly. This allows the possibility of using the
opposite polarity signal or charge from the detector using the same
pixel detector and the ASIC. Since the counters are digital and the
noise and background signals are discriminated, this mode will be
as accurate as the incrementing system described above.
[0173] In a previous CCD based MARY readout chip, an 8 stage TDI
architecture was applied and summed up all the output signals off
the chip to achieve a total of 192 stages of charge integration in
the scanning direction. Limiting the TDI to 8 channels was due to
the fact that each x-ray photon producing several thousand
electrons is a strong signal unprecedented in the scintillator
based detectors. Therefore, CCD well size could not be made large
enough within the 50.times.50 micron.sup.2 pixel area to
accommodate so much charge. The 8 channel TDI required extra
peripheral electronics outside the chip to multiplex, sum,
transfer, digitize, etc.
[0174] The use of the DTDI technique eliminates most of these
problems. For example, a longer DTDI architecture can be used
reaching up to the whole DTDI length of .ltoreq.192 channels or
much more. This can be achieved because a counter is used. If the
total counts at the end of DTDI is over 16 bits then we have
several options: 1. Increase the counters to 17 bit or higher, 2.
Drop least significant bits, and/or 3. Use small number of bit
counters at the beginning cells and larger counters in the pixels
at the end of the DTDI chain. The last option is possible because
the beginning of the DTDI chain always have low counts and much
smaller counters such as 8 to 10 bit can be used for the pixels at
the beginning of the DTDI architecture. This will produce empty
area under these pixels. This extra real estate can be used for
larger counters such as 18 to 24 bit or more towards the end of the
DTDI chain, thus keeping the pixel pitch small, same size and
periodic. The DTDI circuitry will allow long DTDI chain reducing
multiplexing and summing, and eliminating digitization because it
produced an already digitized digital signal. This will reduce the
peripheral electronics, system size and cost.
[0175] In the case where the counters start with the maximum count
and events decrement the counters as described above, the beginning
pixel counters need to have the larger bit counters instead of the
smallest and the smallest bit counters instead of the largest at
the end of the DTDI chain.
[0176] Compared to Thin Film Transistor (TFT) technologies used in
flat-panel approaches, CMOS technology allows the implementation of
sophisticated circuitry inside the pixel area. Therefore,
interesting new designs can be implemented using CMOS technology
and/or some or most of the peripheral electronics needed outside
the ASIC can also be incorporated onto the ASIC. Thus it is
possible to reduce the area and the size of the circuits and the
detector array by such careful and innovative design and also using
finer foundry processes.
[0177] A properly selected detector thickness is essential for
highest conversion efficiency and lowest noise or charge sharing
among neighboring pixels. Compared to other solid detector
materials such as Silicon (Z=14), Se (Z=34) and GaAs (Z=31,33),
CdZnTe (Z=48, 30, 52) has much higher Z in average and therefore,
will attenuate much more higher energy x-rays than those lower Z
materials do. For CdZnTe, each 20 keV x-ray photon produces about
4.5.times.10.sup.3 electron-hole pairs, and detectors of thickness
0.2 mm and 0.5 mm will stop 95% and 99.8% of the 20 keV x-rays
respectively. Each 50 keV x-ray photon produces about
1.1.times.10.sup.4 electron-hole pairs and a 1.5 mm thick detector
will stop 99.8% of the x-rays and therefore, high efficiency or
high DQE is achievable. In our contrast imaging application, for
example, we may concentrate the x-ray energy around 33.2 keV, which
is the k-edge of iodine, as it has absorptions drastically
different before and after this edge. A proper energy separation at
this edge will give best contrast improvements for iodine based
dual energy imaging.
[0178] The charge spread among neighboring pixels should not be a
problem for our direct conversion detectors, as the electric field
established between the electrodes will naturally focus the charge
motion toward the pixel pad. For other detectors, such as the
scintillators, the light photons generated by incident x-ray
photons will propagate in all directions and thus causing low DQEs.
In our previous study on digital mammography, we found that we
could focus the signal charge to 0.05 mm pixels from a 2 mm thick
detector without spreading to neighboring pixels.
[0179] This system has potential for other medical applications
such as chest radiography and other suitable radiography by using
thicker detectors, where the bone shadows from the x-ray can be
removed so that better pulmonary nodule images are obtained. The
prostate imaging is another area that this dual energy approach can
help to remove the blocking bone shadows. It can also be applied to
arthritis joint examination, bone densitometry and specialized
medical radiology for a scanning type system. It may also be
possible to distinguish different tissue types such as fat, muscle,
liver, cancer tissues and image them with higher resolution using
dual or multiple energy imaging.
[0180] These detectors can also be used for industrial imaging that
require ultra high resolution inspection of high tech products and
safety-critical applications, such as security, inspection of
critical aircraft parts, corrosion and micro-fractures in aging
aircraft, and NDE and NDI of critical high tech military,
industrial and commercial products. It can be used to build very
high resolution baggage scanners where the finest detail of the
explosives such as sheet explosives, wires etc. And contraband can
be seen clearly. This will help in the identification process.
[0181] In another embodiment we can produce a staring array using
the same chip without changes. In the staring array mode, which is
also discussed in detail above, the pixel detector ASIC observes
single, dual or multiple energy images without scanning or any
motion. This mode of operation can be achieved by turning on the
inputs, accumulating an image, then turning off inputs and reading
out the counters as fast as possible in the same way it is
described above.
[0182] This process can be repeated as many times as needed to get
high quality and high statistics images. However, the external
circuitry needs to store the contents of each counter separately as
they come out of the chip to form the image.
[0183] These pixel detectors can be formed into many different
forms such as linear abutted arrays, linear staggered arrays,
curved arrays of any form, flat or curved 2D flat arrays, etc. The
arrays can also be put on top of each other to form 3D detector
array structures.
[0184] Those having ordinary skill in the relevant art will
appreciate the advantages provided by the features of the present
disclosure. In the foregoing Detailed Description, various features
of the present disclosure are grouped together in a single
embodiment for the purpose of streamlining the disclosure. This
method of disclosure is not to be interpreted as reflecting an
intention that the claimed disclosure requires more features than
are expressly recited in each claim. Rather, as the following
claims reflect, inventive aspects lie in less than all features of
a single foregoing disclosed embodiment. Thus, the following claims
are hereby incorporated into this Detailed Description of the
Disclosure by this reference, with each claim standing on its own
as a separate embodiment of the present disclosure.
[0185] It is to be understood that the above-described arrangements
are only illustrative of the application of the principles of the
present disclosure. Numerous modifications and alternative
arrangements may be devised by those skilled in the art without
departing from the spirit and scope of the present disclosure and
the appended claims are intended to cover such modifications and
arrangements. Thus, while the present disclosure has been shown in
the drawings and described above with particularity and detail, it
will be apparent to those of ordinary skill in the art that
numerous modifications, including, but not limited to, variations
in size, materials, shape, form, function and manner of operation,
assembly and use may be made without departing from the principles
and concepts set forth herein.
Fast Counting Imaging System
[0186] The pixel detectors and their readout chips can run in fast
mode for acquiring fast images. This is done with the circuits
shown in FIG. 7 and FIG. 13. In these cases the energy bins are
separated simultaneously by the comparators 705, 706, 1302, 1312,
1314 and the output of each comparator is counted by digital
counters 709, 710 and analog counters 1307, 1315. Therefore, each
channel or pixel accumulates multi energy data simultaneously
without intervention from external circuitry and/or software. Also
each channel or pixel is completely independent of each other and
accumulate data simultaneously without any dead time. This data can
also be collected in real time. After the data accumulation the
counting is stopped and the counts in each counter in the ASIC are
readout in parallel and/or in series as fast as possible to reduce
delay. The counters are then reset and the data accumulation can
start again. To reduce the delay to practically zero, the counts
inside the counter can be shifted into a register, counters reset
and the counting can start immediately. The contents of the
registers are than read out while the pixel detector is counting
the incident x-ray photons. The cycle can be repeated as many times
as necessary.
Imaging Emitted Radiation
[0187] On top of the pixel detectors and arrays a collimator can be
used to produce images of x-rays or gamma rays emitted by objects.
Since these photons can be emitted from anywhere inside or surface
of the object the photons does not follow well defined paths such
as generated by x-ray generators. Therefore, a collimator will be
necessary so that each pixel images photons coming from a certain
area on or inside the object emitting radiation. A collimator with
size, thickness, dimensions and geometry required for a specific
application can be placed on top the same detector systems
disclosed above. Images are also produced same way as described.
The collimator separates the particles, x-rays and gamma rays to
specific angles, which produces the image of the object under
investigation. The detectors developed in this embodiment may be
used for applications such as security (detecting radioactive
material and radioisotopes), gamma camera and SPECT (single photon
emission tomography) for medical imaging.
High Energy Imaging System
[0188] In another embodiment the two positron annihilation photons
with energy of 511 keV can be detected in or out of coincidence for
positron emission tomography (PET). This embodiment may also be
used for other high energy photon detections for energies reaching
above 1,000 keV. The fast trigger system required for the
coincident detection as discussed above. Compton scattering
dominates for CdZnTe detectors above 250 keV. This means that the
gamma ray most likely makes more than one interaction inside the
CdZnTe volume. At this energy level about three quarter of the
incident gamma rays undergo a Compton scatter. These interactions
can be detected, energy deposition determined and results used to
reconstruct the direction and energy of the incident gamma ray. In
PET application since the two simultaneous gamma rays produced
back-to-back are detected the direction measurement may not be
necessary. However, it will be important to determine the first
interaction position so that the path of the two simultaneous gamma
rays can be determined to construct the image.
[0189] Because the effective Z of CdZnTe is about 50 the quantum
efficiency is not as large as the typical PET scintillators (BGO
and LSO) and large thickness CdZnTe detectors are required. Or
higher Z solid state detector such as PbI.sub.2 may be used. This
system can also be used to fabricate small animal imagers for the
field of molecular imaging. Because the area of coverage for a
small animal is relatively small, one can create a lower cost
system with large solid angle coverage. A large solid angle
improves the sensitivity and ameliorates the low quantum
efficiency. If the CdZnTe detectors are developed for high
resolution human PET imaging, the detector layering or long CdZnTe
detectors or higher Z solid state detectors or higher Z
scintillators can be used.
[0190] FIG. 12 shows a diagram of a high energy imaging detector as
discussed above. The main features of this detector are:
1. The front-end readout electronics is designed onto the chip
1202. The CdZnTe pixel detector array 1201 is mounted on top of the
ASIC 1202 as shown in FIG. 12, using a bump bond technology. The
bias voltage 1205 is applied onto the top electrode 1204. 2. The
input pad 1203 at each pixel on the ASIC will be connected directly
to the pixel pad of the CdZnTe detector through a bump bond. This
will have small connection capacitance, <1 pF. The separation of
the bump bonds 1203 will be the same as the pixel pitch and thus
larger than the typical pitch of the wire-bond pads 1206 used. This
reduces the input capacitance and the chip will see only the true
detector pixel capacitance, which is small, <1 pF for a pixel
detector with thickness greater than 5 mm. 3. The peripheral
electronics required for the data acquisition, analysis and signal
generation is miniaturized to fit onto a small PCB. 4. Since the
new pixel detectors can be mounted at nearly touching distance to
each other, it is possible to make an array from 2 detectors up to
any size one-, two- or three-dimensional arrays to increase the
sensitivity and efficiency. This is especially important for
detecting gamma rays with high energy, >500 keV, because at such
high energies a large fraction of the gamma rays pass through a 1
cm thick CdZnTe crystal without an interaction or make one or even
two interactions and then escape from the crystal without
depositing all of their energy. To solve this issue there are two
options: 1. Significantly increase the detector volume, which is
costly; or 2. Mount several detectors in an array so that the
escaping gamma rays can be captured at the adjacent detectors. Thus
the abuttable pixel detectors are good candidates for the second
and more cost-effective solution. This option also makes the CdZnTe
pixel detectors modular, allowing the development of radiation
detectors of any size, geometry, thickness and volume with a
variety of sensitivities and efficiencies without loss of
performance. 5. The ASIC and the readout electronics can be
designed for low power consumption so that compact, portable and
battery operated radiation detection equipment can be developed. 6.
This pixel detector will have low energy threshold <20 keV. This
is because of the small detector capacitance and the low noise
ASIC. The small pixel effect will also play an important role in
improving the energy resolution. We expect an energy resolution of
about 0.5% to 0.8% FWHM @ 662 keV for single pixel events and a
threshold of .apprxeq.10 keV. Lower noise will also improve the
energy resolution of two- (or more-) pixel events. 7. The e-h pairs
created is directed into columnar motion under the applied bias
voltage, which allows excellent spatial resolution in the
determination of the photon interaction point (s). Normally a
photon will be absorbed by the photoelectric effect (PE). If a
photon makes one or more Compton scatters inside the detector each
interaction will produce a signal at the corresponding pixel. 8.
The pixel detector provides additional modes of operation beyond a
high energy resolution gamma ray detector. The x and y coordinates
of each absorption or Compton scatter event inside the pixel
detector will be known by determining the x and y coordinates of
the pixel(s) that have detected an interaction. The z coordinate of
each interaction can also be determined by either by pulse height
ratio between anode and cathode signals (good for single
interaction events) or by arrival time difference of each pixel
signal if photon makes more than one interaction. Since the x, y,
and z coordinates of each Compton interaction and the final PE
absorption are known, it is possible to calculate the direction of
the incident gamma rays using the Compton scatter formula through
overlapping Compton scatter cones. The overlapping cones produce an
image showing where the gamma rays are coming from. Determination
of the incident gamma ray direction can open the potential for new
applications for the pixel detector. 9. Due to its spatial
resolution, the new pixel detector also allows gamma ray imaging by
placing a collimator in front of the detector(s). The collimator
may be of any kind, such as pin hole, parallel, cone beam, or slot
type. 10. The efficiency can be further improved for high-energy
gamma rays >500 keV if they undergo three or more Compton
scatters. In this case, even if the scattered gamma ray escapes the
detector(s), the energy of the incident photon can be determined at
somewhat lower resolution by using the Compton formula. This will
increase the efficiency for detection of high-energy gamma rays in
small volume detectors. 11. This pixel detector can also be made to
have single or multiple energy ranges so that it can extend from
low energies to large energies. The multiple energy ranges make the
pixel detector to have good energy resolution and accuracy in each
range. The multiple energy ranges can be selectable. However, if
the photon source used has a large energy range than the detector
cannot be used effectively as selection of one of the ranges will
be required. To overcome this we can use a system that
automatically switches the energy range of the chip to the range of
the incident photon.
* * * * *