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ΤΜΗΜΑ ΙΑΤΡΙΚΗΣ
ΤΜΗΜΑ ΦΥΣΙΚΗΣ
ΔΙΑΤΜΗΜΑΤΙΚΟ ΜΕΤΑΠΤΥΧΙΑΚΟ ΠΡΟΓΡΑΜΜΑ ΣΠΟΥΔΩΝ ΣΤΗΝ
ΙΑΤΡΙΚΗ ΦΥΣΙΚΗ
ΔΙΕΡΕΥΝΗΣΗ ΤΩΝ ΑΠΕΙΚΟΝΙΣΤΙΚΩΝ ΧΑΡΑΚΤΗΡΙΣΤΙΚΩΝ
ΨΗΦΙΑΚΟΥ ΑΝΙΧΝΕΥΤΗ ΒΑΣΙΣΜΕΝΟΥ ΣΕ ΑΙΣΘΗΤΗΡΑ CMOS ΣΕ
ΣΥΖΕΥΞΗ ΜΕ ΦΘΟΡΙΖΟΥΣΑ ΟΘΟΝΗ ΕΡΥΘΡΑΣ ΕΚΠΟΜΠΗΣ
ΙΩΑΝΝΗΣ Ε. ΣΕΦΕΡΗΣ
ΔΙΠΛΩΜΑΤΙΚΗ ΕΡΓΑΣΙΑ
ΠΑΤΡΑ 2013
2
UNIVERSITY OF PATRAS
SCHOOL OF MEDICINE
DEPARTMENT OF PHYSICS
INTERDEPARTMENTAL PROGRAM OF POSTGRADUATE STUDIES IN
MEDICAL PHYSICS
INVESTIGATION AND IMAGING CHARACTERISTICS OF A CMOS
SENSOR BASED DIGITAL DETECTOR COUPLED TO A RED EMITTING
FLUORESCENT SCREEN
IOANNIS E. SEFERIS
MASTER THESIS
PATRA 2013
3
MEMBERS OF THE EXAMINATION COMMITTEE
Professor George Panayiotakis
Supervisor
Professor Ioannis Kandarakis
Member of the Examination Committee
Assistant Professor Eleni Costaridou
Member of the Examination Committee
4
ABSTRACT
ABSTRACT
The dominant powder scintillator in most medical imaging modalities for decades is
Gd2O2S:Tb due to the very good intrinsic properties and overall efficiency. Except for
Gd2O2S:Tb there are alternative powder phosphor scintillators like Lu2SiO5:Ce and
Gd2O2S:Eu that has been suggested for use in various medical imaging modalities.
Gd2O2S:Eu emits red light and can be combined mainly with digital imaging devices
like CCDs and CMOS based detectors. The purposes of the present thesis, is to
investigate the fundamental imaging performance of a high resolution CMOS based
imaging sensor combined with custom made Europium (Eu3+) activated Gd2O2S
screens in terms of Modulation Transfer Function (MTF), Normalized Noise Power
Spectrum (NNPS), Detective Quantum Efficiency (DQE), Noise Equivalent Quanta
(NEQ) and Information Capacity (IC) covering the mammography and general
radiography energy ranges.
The CMOS sensor was coupled to two Gd2O2S:Eu scintillator screens with
coating thicknesses of 33.3 and 65.1 mg/cm2, respectively, which were placed in
direct contact with the photodiode array. The CMOS photodiode array, featuring
1200x1600 pixels with a pixel pitch of 22.5  m , was used as an optical photon
detector. In addition to frequency dependent parameters (MTF, NPS, DQE)
characterizing image quality, image information content was assessed through the
application of information capacity (IC). The MTF was measured using the slantededge method to avoid aliasing while the Normalized NPS (NNPS) was determined by
two-dimensional (2D) Fourier transforming of uniformly exposed images. Both
parameters were assessed by irradiation under the RQA-5 protocol (70kVp digitalradiography) recommended by the International Electrotechnical Commission Reports
5
ABSTRACT
62220-1 and the W/Rh, W/Ag beam qualities (28kVp digital-mammography). The
DQE was assessed from the measured MTF, NNPS and the direct entrance surface
air-Kerma (ESAK) obtained from X-ray spectra measurement with a portable
cadmium telluride (CdTe) detector.
The spectral matching factor between the optical spectra emitted by the
Gd2O2S:Eu and the Gd2O2S:Tb screens and the CMOS optical sensor, evaluated in the
present study, was 1 and 0.95 respectively. The ESAK values ranged between 11.287.5 Gy , for RQA-5, and between 65.8-334 Gy , for W/Rh, W/Ag beam qualities.
It was found that the detector response function was linear for the exposure ranges
under investigation. Under radiographic conditions the MTF of the present system
was found higher than previously published MTF data for a 48  m CMOS sensor, in
the low up to medium frequency ranges. DQE was found comparable, while the
NNPS appeared to be higher in the frequency range under investigation (0–10
cycles/mm). NEQ reached a maximum (73563 mm-2) in the low frequency range (1.8
cycles/mm), under the RQA 5 (ESAK: 11.2 Gy ) conditions. IC values were found to
range between 1730-1851 bits/mm2. Under mammographic conditions MTF, NNPS
and NEQ were found comparable to data previously published for the 48  m CMOS
sensor while the DQE was found lower. The corresponding IC values were found
ranging between 2475 and 2821 bits/mm2.
The imaging performance of europium (Eu3+) activated Gd2O2S screens in
combination to the CMOS sensor, investigated in the present study, was found
comparable to those of Terbium (Tb) activated Gd2O2S screens (combined with the
CMOS sensor). It can be thus claimed that red emitting phosphors could be suitably
used in digital imaging systems, where the Silicon (Si) based photodetectors are more
sensitive to longer wavelength ranges, and particularly in the red wavelength range.
6
CONTENTS
CONTENTS
ABSTRACT
5
ACKNOWLEDGEMENTS
11
CHAPTER A
12
INTRDUCTION
A.1. The Problem
CHAPTER B
12
12
16
DIGITAL RADIOGRAPHY
16
B.1. Introduction
16
B.2. Advantage of Digital Radiography
19
B.3. Disadvantage of Digital Radiography
19
B.4. Digital Radiography and X-Ray Dose
20
B.5. Photostimulable Storage Phosphor Detectors (PSP)
21
B.6. Charge-Coupled Device Detectors (CCD)
23
B.7. Complementary Metal-Oxide Semiconductors Detectors
25
B.8. Active Matrix Flat Panel Imagers Detectors (AMFPI)
26
B.9. Image Quality
28
B.9.1. Pixel Size, Matrix and Detector Size
28
7
CONTENTS
B.9.2. Dynamic Range
28
B.9.3. Spatial Resolution
29
B.9.4. Modulation Transfer Function (MTF)
29
B.9.5. Noise Power Spectrum (NPS)
30
B.9.6. Detective Quantum Efficiency (DQE)
31
B.9.7. Information Capacity (IC)
31
CHAPTER C
33
MATERIALS AND METHODS
C.1. Experimental Setup
33
33
C.1.1. Phosphor Screen Preparation
33
C.1.2. CMOS Sensor
33
C.1.3. X-Ray Spectra Measurement
35
C.2. Image quality
37
C.2.1. Spectral Compatibility
36
C.2.2. Signal Transfer Property (STP)
37
C.2.3. Modulation Transfer Function (MTF)
37
C.2.4. Normalized Noise Power Spectrum (NNPS)
38
C.2.5. Detective Quantum Efficiency (DQE)
39
C.2.6. Noise Equivalent Quanta (NEQ)
39
8
CONTENTS
C.2.7. Information Capacity (IC)
CHAPTER D
40
41
RESULTS AND DISCUSSION
41
D.1. Optical Properties and MTF of Borosilicate Glass Substrate41
D.2. Spectra Compatibility
42
D.3. X-Ray Spectra
43
D.4. Signal Transfer Property (STP)
45
D.5. Modulation Transfer Function (MTF)
46
D.6. Normalized Noise Power Spectrum (NNPS)
49
D.7. Detective Quantum Efficiency (DQE)
51
D.8. Noise Equivalent Quanta (NEQ)
54
D.9. Information Capacity (IC)
56
CHAPTER E
59
CONCLUSIONS AND FUTURE WORK
59
E.1. Conclusions
59
E.2. Future Work
60
Appendix A: Information Capacity (IC)
61
Appendix B: Abbreviations
64
Appendix C: List of Figures
66
9
CONTENTS
References
68
10
ACKNOWLEDGEMENTS
ACKNOWLEDGEMENTS
I wish to express my appreciation to my supervisor Professor George Panayiotakis for
giving me the opportunity to carry out this thesis under his supervision, providing me
advises and support.
My gratitude goes to Professor Ioannis Kandarakis. His support and advice over this
project were valuable for me. His eagerness to help me whenever i need it played a
catalytic role for the completion of this work.
Special thanks to Assistant Professor George Fountos, Dr Christos Michail, Dr.
Nektarios Kalivas, Assistant Professor Ioannis Valais, Dr. Panagiotis Liaparinos, Dr
Stratos David, for their collaboration, assistance and suggestions throughout this work.
Last but not least, I want to deeply thank my parents for their continuous support and
love during my studies in Athens and Patra.
11
CHAPTER A
CHAPTER A
INTRODUCTION
A.1 The problem
It was recognized early on that the digitization of medical information would advance
the efficiency of diagnostic technology (Suzuki et al 2000, Liang et al 2008).
However, the digitization of image data, is dependent on advances in technologies
such as input (imaging conditions, powder scintillator/ digital sensor optimization),
processing, transmission, storage and display (Bruckmann and Uhl 2000, Suzuki et al
2000, Efstathopoulos et al 2001, Κoscis et al 2003). Insufficient advances in such
technologies have effectively limited the digitization of image data for medical use
(Suzuki et al 2000).
Scintillator materials are employed as radiation to light converters in detectors
of medical imaging systems (Johns and Cunningham 1983, Knoll 1989, Yaffe 2000).
In X-ray projection imaging, a large number of such materials have been employed in
the form of granular screens (Blasse and Grabmaier 1994). The X-ray detection and
imaging performance of phosphor screens are affected by intrinsic physical properties,
related to X-ray and light photon transport through the material. These properties have
been previously investigated by experimental (Alig and Bloom 1977, Venema 1979,
Dick and Motz 1981a, Dick and Motz 1981b, Derenzo et al 1990, Ginzburg and Dick
1993, Van Eijk 2001), theoretical (Venema 1979, Nielsen and Carlson 1984,
Nishikawa and Yaffe 1990, Kandarakis et al 1997, Michail et al 2007, Michail et al
2009) and Monte Carlo methods (Raeside 1976, Rubinstein 1981, Kalender 1981,
Aerts et al 1982, Morin 1988, Andreo 1991) and have been taken into account in the
design of commercial imaging systems (Yaffe and Rowlands 1997, Gruner et al 2002).
12
CHAPTER A
However the quest for the ideal scintillator compromising optimum imaging
performance with high efficiency and low cost is still in the front for various research
groups and the results are published in numerous works.
The development of new digital diagnostic imaging techniques and modalities,
such as digital breast tomosynthesis (DBT) (Dobbins and Godfrey 2003), dual energy
(DE) imaging (Alvarez et al 2004), scanning slot digital mammography detectors
(Mainprize et al 2002), energy selective techniques (Patatoukas et al 2006, Alvarez
2010) and digital fluoroscopy systems (Antonuk et al 2009), require efficient
scintillators to reduce patient dose. Digital imaging systems with smaller sizes for the
detector elements (pixels or dels) are also required to improve image resolution and
the visibility of microcalcifications, without noise increase and DQE degradation.
Furthermore, in many digital imaging systems based on crystalline Silicon (Si) optical
detectors such as CCDs and CMOS sensors (Nikl 2006, Chotas et al 1999, Kim et al
2001, Bigas et al 2006, Kim et al 2005) the green light emitted by Terbium-activated
phosphors is not very efficiently detected (Gurwich 1995). This is because a large
number of Si based devices, incorporated in X-ray imaging systems are not
adequately sensitive to these wavelengths (500-550 nm) (Gurwich 1995). Since most
Si based photodetectors are more sensitive to longer wavelength ranges, and
particularly in the red wavelength range, the properties of red emitting phosphors
(Rowlands and Yorkston 2000, Rodnyi et al 2001, Zych 2006, Nikl 2006) and in
particular, of the Europium (Eu3+) activated Gd2O2S has been investigated in terms of
optical output (Zych 2006), emission efficiency (Michail et al 2008, Michail et al
2010, Michail et al 2011, Wiatrowska et al 2010) and imaging performance (Gurwich
1995, Michail et al 2008, Michail et al 2010, Michail et al 2011). Such phosphors are
cost effective and can be easily prepared by inserting Europium ion activator (Eu3+) in
13
CHAPTER A
rare earth based host matrices. In addition Europium-activated phosphors have a
decay time of the order of one millisecond (slightly higher than Gd2O2S:Tb) which is
acceptable for applications that do not involve high framing rates ( Lempicki et al
2002, Nagarkar et al 2003). These include digital radiography, and mammography.
Gd2O2S:Eu has been previously employed in single pulse dual energy radiography
(Boone et al 1992), in digital mammography and diffraction enhanced breast imaging
with CCD arrays (Gambaccini et al 1996, Taibi et al 1997, Gambaccini et al 1998,
Harris et al 2003). Furthermore Gd2O2S:Eu has been previously used as phosphor
insert in phantoms for limited-angle x-ray luminescence tomography (XLT)
(Carpenter et al 2011) and hybrid x-ray luminescence/optical imaging (Carpenter et al
2010). Previously, the image quality of a detector based on a commercial Gd2O2S:Tb
screen coupled to an active CMOS sensor has been investigated by Michail and found
comparable with CCD and passive CMOS sensors (Michail et al 2011). In the present
study this CMOS sensor was coupled to two different custom made Gd2O2S:Eu
phosphor screens, prepared in our laboratory, and the whole detector was evaluated
under general radiography and mammography conditions.
The investigation of the imaging performance of this detector (i.e. CMOS
coupled to Gd2O2S:Eu phosphor screens) was achieved by experimental assessment of
the Signal Transfer Property (STP), the Modulation Transfer Function (MTF), the
Normalized Noise Power Spectrum (NNPS), the Detective Quantum Efficiency
(DQE), the Noise Equivalent Quanta (NEQ) and the Information Capacity (IC). The
experimental method was based on the guidelines published by the International
Electrotechnical Commission (IEC) (Medical Electrical Equipment-Characteristics of
Digital X-Ray Imaging Devices 2005), which has standardized the methodology for
measuring DQE in digital detectors. In addition the results of the present work have
14
CHAPTER A
been compared with the results of previously studied CMOS sensors coupled to
Gd2O2S:Tb
phosphor
screens
(Michail
et
al
2011,
Cho
et
al
2008).
15
CHAPTER B
CHAPTER B
DIGITAL RADIOGRAPHY
B.1. Introduction
Digital radiography detector systems were first implemented for medical applications
in the mid-1980s, but the promise of digital imaging was not realized until the early
1990s, in conjunction with the establishment of first generation picture archiving and
communications systems (PACS). At the time, film screen detectors have nearly
replaced by computed radiography systems (CR), charge coupled device (CCD) and
Active-matrix flat-panel imager (AMFPI) systems. These technologies for digital
image acquisition appeared in the mid-1990s with the use of large field of view (FOV)
X-ray phosphors and optical lens assemblies to focus the X-ray induced light output
onto a small-area charge coupled device (CCD) photodetector array, as well as
rectangular CCD arrays used with slot-scan geometries. Active-matrix flat-panel
imager (AMFPI) systems appeared in the late 1990s and early 2000s, employing
either an X-ray-to-light converter with photodiode array, or a semiconductor material
to directly convert incident X-rays into signals. Both CCD and flat-panel based
detectors use an "active" readout of the image following acquisition to present the
image immediately without further interaction by the technologist. Other technologies
such as complementary metal-oxide semiconductor (CMOS) detectors were
introduced in the same time frame as AMFPI’s, and so far are making great progress
in medical imaging, making it competitive with the other digital imaging
detectors.Beyond the digital detector characteristics are considerations for software
for pre- and post-processing of the digital image data, the user and modality interfaces,
display monitors and calibrations. Many unique acquisition capabilities, such as dual-
16
CHAPTER B
energy image tissue decomposition and limited angle digital tomosynthesis, are
important when considering future applications of specific importance (Seibert 2009).
Luminescent materials, either in the form of powder scintillators (phosphor
screens) or in the form of ceramics or single-crystals, are incorporated in many
medical imaging radiation detectors. Powder scintillators have been successfully
employed in a large number of X-ray radiography devices (from conventional and
digital X-ray radiography, mammography and X-ray computed tomography to digital
dental radiography and portal imaging) (Arnold 1979, Yaffe and Rowlands 1997,
Boone 2000, Van Eijk 2002, Nikl 2006). Among various materials, rare-earth-ion
doped Oxyorthosilicates provide excellent scintillating properties and thus they have
been extensively investigated for the development of new scintillators (Lee et al
2006). Terbium (Tb)-activated phosphors (i.e. Gd2O2S:Tb, La2O2S:Tb and Y2O2S:Tb)
have been up to now accepted to be of the most efficient X-ray-to-light converters
(Arnold 1979, Gurwich1995, Kandarakis et al 1996, Liaparinos et al 2007) employed
in mammography and radiography. Currently the most widely used phosphors are
Gd2O2S:Tb and CsI:Tl.
Gd2O2S:Tb has been proven very useful in conventional radiography screenfilm systems, where precise matching the spectral sensitivity of the X-ray film to the
emission of the phosphor is of primary consideration in order to obtain the highest
speed for the screen-film combination. However, in the past decade an increasing
tendency to introduce digital radiography and mammography systems has been shown
in various publications (Rodnyi et al 2001, Van Den Bergh and Leblans 2005).
In some digital imaging systems, based on crystalline Silicon (Si) optical detectors
(CCDs, photodiodes, novel CMOS sensors), the green light emitted by Terbiumactivated phosphors phosphors (as well as by Thalium activated CsI) is not very
17
CHAPTER B
efficiently detected (Gurwich 1995). This is because a large number of Si based
devices, incorporated in X-ray imaging systems are not adequately sensitive to these
wavelengths (500-550 nm); only 45-55 % of the light produced by Gd2O2S:Tb or
Y2O2S:Tb is registered by the Si photodiode (Gurwich 1995).
Since most Si and CMOS based photodetectors are more sensitive to longer
wavelengths, and particularly in the red wavelength range, it would be of interest to
investigate the emission efficiency of red emitting phosphors (Rodnyi et al 2001, Van
Den Bergh and Leblans 2005). For this purpose, Europium (Eu)-activated phosphors,
emitting at wavelengths towards the red region of the light spectrum, could be used
instead of green emitting Tb-activated ones (Gambaccini et al 1996). Furthermore the
performance of many Europium doped scintillators, and particularly Gd2O2S:Eu, have
been previously found comparable to Terbium-activated phosphors (both showing
190% of the CdWO4’s optical output). Gd2O2S:Eu has high light yield (60000
photons/MeV) and luminescence efficiency, high density (7.3 g/cm3), high radiation
4
detection index (  Z eff
=103×106) and finally decay time of the order of 1 ms (slightly
higher than Gd2O2S:Tb), which is acceptable for most X-ray radiography applications
that do not involve high framing rates (Lempicki et al 2002, Okumura et al 2002,
Nagarkar et al 2003, Michail et al 2008). These include stationary digital and
conventional general radiography and mammography. Gd2O2S:Eu has been
previously employed in single pulse dual energy radiography (Boone et al 1992), in
digital mammography and diffraction enhanced breast imaging with CCD arrays
(Gambaccini et al 1996, Taibi et al 1997, Gambaccini et al 1998, Harris et al 2003).
18
CHAPTER B
B.2. Advantage of Digital Radiography
An advantage of DR over film/screen (F/S) radiography is the increased dynamic
range and linear response of DR images (Artz 1997). According to Vuylsteke &
Schoeters (1994), F/S has an exposure range in the order of 100:1. In contrast, DR has
a dynamic range of nearly 1000:1 (Vuylsteke and Schoeters, 1994). These dynamic
range advantages also exist for direct and indirect digital radiography.
A further advantage of DR over F/S is that it provides the facility to perform
post-processing functions or manipulate the image after it has been acquired. Contrast
magnification and split screen viewing can be manipulated on the monitor screen to
aid in diagnosis. It has been noted that the value of contrast enhancement is actually
task-specific (Dunn and Kantor 1993).
B.3. Disadvantage of Digital Radiography
Several disadvantages of DR have been reported. The spatial resolution of an imaging
system is its ability to accurately represent small objects. Spatial frequencies in F/S
general radiography are limited by factors external to the image recording system.
Some external factors are the finite size of the focal spot where x-rays are produced
and the geometric relationship between the focal spot, the object being imaged and the
image recording system (Bushberg et al 2002, Bushong 2001, Curry et al 1990). A
digital image’s spatial resolution is determined by the number of rows and columns of
pixels in the image (Baxes, 1994). The spatial frequencies of a DR image are also
limited by the external factors discussed above, as well as the workstation capabilities
used for diagnosis. The spatial resolution of a DR image can be measured in pixels
per millimetre or in line pairs per millimetre (Bushberg et al 2002, Bushong 2001).
The spatial resolution of DR image recording systems is lower than that of F/S image
19
CHAPTER B
recording systems (Schaetzing et al 1990, Artz, 1997, Bushberg et al 2002, Bushong
2001, Weiser 1997). This lower spatial resolution impacts on the overall spatial
resolution of radiographic examinations. Specifically, the sampling frequency of the
DR recording system is the limiting factor in determining the spatial resolution of the
DR system (Bushberg et al 2002, Bushong 2001).
Direct radiographic image noise results from either internal system noise or
from scattered radiation. Inherent system noise is greater in DR than in F/S and is
reduced through software algorithms prior to the display of the image. CR and other
DR systems are more sensitive to scattered radiation than F/S. The amount of
scattered radiation is a function of the energy of the X-ray beam and the anatomy
which attenuates the x-ray beam. Scatter can be reduced prior to reaching the image
receptor in both F/S and DR systems. The effect of scattered radiation in DR can be
further reduced through the use of scatter correction algorithms (Bushberg et al 2002,
Bushong 2001).
Artifacts are a potential problem that can reduce the quality of images. DR has
artifacts that are not found in F/S systems. Artifacts in CR fall into three main
categories (Cesar et al 2001):
 Image plate artifacts
 Plate reading artifacts
 Image processing artifacts
B.4. Digital Radiography and X-Ray Dose
Digital radiography has been the subject of many studies evaluating the radiation
doses received by patients as compared to doses received in F/S radiographic
20
CHAPTER B
examinations. The consensus of the findings is that the dose received by the patient in
CR is similar to that of F/S radiographic examinations. In mammography especially
the imaging capabilities and system setups of DR are generally equivalent with the
corresponding CR [attached]. In paediatric radiographic examinations, doses received
by the child were lower in F/S than in CR imaging examinations according to
Nickoloff (2002).
A general consensus of recent articles is that the wide latitude of CR compared
to F/S has not provided the expected dose reductions for CR. It has been suggested
that quantum noise from low exposures used in CR systems limits its dose reduction
capabilities. However, dose advantages are gained in direct DR imaging examinations
over those of CR and F/S. The reduced dose of direct DR is attributable to the lower
noise levels of direct DR over CR and the wider latitude of direct DR compared to F/S
(Chotas et al 1993, Dobbins et al 2002, Marshall 2001, Marshall et al 1994).
B.5. Photostimulable Storage Phosphor Detectors (PSP)
More commonly recognized as CR, the PSP detector "system" is comprised of 2 main
components. The detector is usually a cassette based storage phosphor that absorbs Xray energy transmitted through the patient and temporarily stores the X-ray latent
image as a 2-dimensional array of electrons trapped. The imaging plate "reader" is
comprised of a scanning laser beam to stimulate the trapped electrons and produce
"photostimulated luminescence" of a different wavelength that is optically separable
from the stimulation wavelength. The reader also includes a light guide and
photomultiplier assembly that extracts and processes the stored X-ray content to a
sampling resolution on the order of 100 microns, digital electronics to create the
21
CHAPTER B
corresponding digital image, and an erasure stage to eliminate any residual signal and
prepare for the next exposure. From the reader, all images proceed to a quality control
workstation for image evaluation, annotation and transfer to PACS. Most often, the
storage phosphor is layered on a flexible or solid substrate in a cassette enclosure,
which allows for the ability to directly replace a screen-film cassette in a conventional
radiography room. Thus there is the flexibility and portability of a cassette with digital
radiography acquisition capability using existing X-ray equipment; this is CR's
greatest asset. CR cassettes of various size and number, together with a high-speed
imaging plate reader can service multiple X-ray rooms, resulting in a relatively low
initial acquisition cost. However, the technologist must handle the cassettes and
process the imaging plates in a manner similar to film, which can reduce patient
throughput in a busy clinical room and increase labor costs, as the time to handle the
exposed imaging plate to the reader and output of the X-ray image can often take
several minutes (about 45 to 60 seconds to "read" the plate with the moving laser
beam). Other expenses to consider are the need for high-frequency (e.g., 170 lines per
inch) antiscatter grids for stationary (non-Bucky) applications such as portable
bedside imaging and fixed grid cassette holders. Long-term costs include cassette and
imaging plate longevity, maintenance and cleaning of the imaging plates and reader
assembly, replacement of damaged detectors, and continuous oversight with a quality
control program (Seibert et al 2006)
22
CHAPTER B
FIG.1. Photostimulable storage phosphor system
B.6. Charge-Coupled Device Detectors (CCD)
The design of a charge-coupled device (CCD) based DR system is straightforward.
The detector is comprised of a large FOV (e.g., 43 cm by 43 cm) scintillator that
converts absorbed X-ray energy into light. It also includes an optical lens assembly to
focus the light onto the photosensitive CCD array, and a CCD camera to integrate,
scan and output the corresponding light image. While there were initially several
configurations in early systems, today's CCD based detector is typically comprised of
a single compound optical lens and a high resolution CCD camera comprised of 9
million pixels (3000× 3000 pixels) to 16 million pixels (4000 × 4000 pixels). When
referred back to the image plane, this results in image pixel sizes of ~0.10 to ~0.14
 m . The photosensitive area of the CCD chip is actually quite small, on the order of
2.5 cm × 2.5cm to 4.0 cm × 4.0 cm, which is required to maintain extremely high
charge coupling efficiency and low noise operation during the readout of the image.
Thus, there is a large optical demagnification that is necessary to focus the full FOV
light image onto the CCD sensor. One physical difficulty is the inefficiency of light
collection caused by the dispersed light emission from the phosphor, resulting in only
a small fraction that can be focused onto the CCD, thus potentially reducing the
23
CHAPTER B
statistical integrity of information carried by the X-ray photons and increasing overall
noise in the image. This is determined by the demagnification factor, conversion
efficiency, luminance and directionality of the light emission. A non-structured
phosphor such as gadolinium oxysulfide has a high light dispersion and corresponding
low fraction of light that can be focused on the CCD, while a structured phosphor
such as cesium iodide (CsI) produces a more forward-directed light output, so that the
lens-light collection efficiency, and thus the SNR in the output image, is better for a
given incident X-ray exposure. Newer, advanced CCD systems with a CsI phosphor
have proven to be reasonably efficient, particularly when using higher kilovolt peak
(kVp) techniques that produce more light photons per absorbed X-ray photon. One
minor disadvantage in some positioning situations is the relatively large and bulky
enclosure of a CCD based DR system, necessitated by placing the CCD out of the
direct X-ray beam and using mirror optics to reflect the light to the photosensor array
(Seibert 2009).
Linear CCD arrays optically coupled to a scintillator by fiberoptic channel
plates (often with a demagnification taper of 2:1 to 3:1) are used in slot-scan
geometries. A significant advantage is pre- and post-patient collimation that limits Xray scatter and allows grid free operation with equivalent image quality (in terms of
SNR) of a large area FOV at 2 to 4 times less patient dose. Disadvantages include the
extended exposure time required for image acquisition with potential motion artifacts
and reduced X-ray tube efficiency. Nevertheless, imaging systems based on slot-scan
acquisition have provided excellent clinical results for dedicated chest and full-body
trauma imaging (Seibert 2009).
24
CHAPTER B
FIG.2. Charged couple devise (CCD)
B.7. Complementary Metal-Oxide Semiconductors Detectors (CMOS)
CMOS light-sensitive arrays are based upon a crystalline silicon matrix and are
essentially random access memory "chips" with built-in photodiodes, storage
capacitors, and active readout electronics, operating at low voltage (3 to 5 volts) for
image acquisition and readout. The ability to randomly address any detector element
on the chip enables opportunities for automatic exposure control (AEC) capabilities
that are not easily performed with a CCD photo detector. However, electronic noise
has been a problem that has slowed the introduction of this technology (Seibert 2009).
In recent years, interest in CMOS sensors has increased rapidly, driven largely
by consumer electronics and the demand for sensors with lower cost, lower power
consumption and miniaturization (Fossum 1997, Lo 1998). Cost effectiveness is
considered by some to be the main advantage of CMOS detectors due to their
manufacturability (Zhang et al 2005, Elbakri et al 2009).
Baysal and Toker (2005) reported on the development of a CMOS
mammography cassette for specimen and biopsy imaging with MTF, DQE and ACR
phantom scores comparable to industry standards. Arvanitis et al (2007) performed a
25
CHAPTER B
detailed study of the characteristics of a CMOS sensor with 25  m dels and 13 cm ×
13 cm field of view. The study characterized various sources of noise in absolute units
of electrons and reported MTF, NPS and DQE results that showed the potential of
active-pixel sensor CMOS. Passive-pixel sensors (PPS) and active-pixel sensors (APS)
are the two broad classes of CMOS detectors. In the PPS design, originally invented
in 1967 (Weckler 1967), each detector element consists of a photodiode and an access
transistor. Activating the access transistor connects the photodiode to the data line for
readout and digitization (Elbakri et al 2009). A sensor with an active amplifier within
each pixel is referred to as an Active Pixel Sensor (APS). For such sensors power
dissipation is minimal and is generally less than a CCD, since each amplifier is only
active during readout (Billas 2011).
B.8. Active Matrix Flat Panel Imagers Detectors (AMFPI)
AMFPI technologies are based on thin-film-transistor (TFT) arrays, made from
amorphous silicon, upon which lithographic etching and material evaporation on the
micron scale produces the electronic components and connections necessary for
detector operation. The flat-panel substrate is divided into individual detector element
(del) compartments, arranged in a row and column matrix, typically with a spacing
dimension of 70 microns to 250 microns, depending on the detector specifications.
Components within each del include a thin-film-transistor (essentially an electronic
switch), a charge collection electrode and a storage capacitor. Electronic
interconnections including gate lines (rows) and drain lines (columns) are connected
to each of the TFTs to control the on/off status of each of the dels and to provide the
conduction paths to the charge amplifiers, respectively. All of the TFTs are closed
during an exposure to collect X-ray induced charge proportional to the incident X-ray
26
CHAPTER B
fluence in each del. After the exposure, active readout of the array occurs one row at a
time by activating the respective gate line, which turns on the TFTs and allows the
stored charge to flow along the columns from each del capacitor via drain lines to the
corresponding charge amplifier. Banks of amplifiers simultaneously amplify the
charge, convert to a proportional voltage, digitize signals in parallel from each row of
the detector matrix and produce a corresponding row of integer values in the digital
image matrix (Figure 4). This process is repeated for each row in the matrix. Detector
readout speed is governed by the intrinsic lag characteristics of the X-ray converter
material, switching speed of the TFT array electronics and the number of independent
charge amplifier arrays that can function in parallel. Limits to spatial resolution for
TFT arrays include fill factor, electronic interconnections and manufacturing yield.
Fill-factor is a term describing the active charge collection area to the total area of the
del, and ideally is 100% for most efficient collection of X-ray information. However,
because electronic components and connection lines occupy space on the substrate,
with conventional TFT designs the fill factor is less than ideal, this is more severe for
smaller del areas of ~0.1  m (e.g., <50%), and can ultimately limit the minimum size
of the del (highest spatial resolution achievable) (Seibert 2009).
27
CHAPTER B
FIG.3. Active matrix flat panel imager (AMFPI)
B.9. Image Quality
B.9.1. Pixel Size, Matrix and Detector Size
Digital images consist of an array of picture elements, often referred to as pixels. The
two dimensional collection of pixels in the image is called the matrix, which is
usually expressed as length (in pixels) by width (in pixels). Maximum achievable
spatial resolution (Nyquist frequency, given in cycles per millimeter) is defined by
pixel size and spacing. The smaller the pixel size (or the larger the matrix) is, the
higher the maximum achievable spatial resolution. The overall detector size
determines if the detector is suitable for all clinical applications. Larger detector areas
are needed for chest imaging than for imaging of the extremities. In cassette-based
systems, different sizes are available (Michail 2010).
B.9.2. Dynamic Range
Dynamic range is a measure of the signal response of a detector that is exposed to Xrays (Spahn 2005). In conventional screen-film combinations, the dynamic range
gradation curve is S shaped within a narrow exposure range for optimal film
28
CHAPTER B
blackening; thus, the film has a low tolerance for an exposure that is higher or lower
than required, resulting in failed exposures or insufficient image quality. For digital
detectors, dynamic range is the range of X-ray exposure over which a meaningful
image can be obtained. Digital detectors have a wider and linear dynamic range,
which, in clinical practice, virtually eliminates the risk of a failed exposure. Another
positive effect of a wide dynamic range is that differences between specific tissue
absorptions (e.g., bone vs. soft tissue) can be displayed in one image without the need
for additional images. On the other hand, because detector function improves as
radiation exposure increases, special care has to be taken not to overexpose the patient
by applying more radiation than is needed for a diagnostically sufficient image (Spahn
2005).
B.9.3. Spatial Resolution
Spatial resolution refers to the minimum resolvable separation between high-contrast
objects. In digital detectors, spatial resolution is defined and limited by the minimum
pixel size. Increasing the radiation applied to the detector will not improve the
maximum spatial resolution. On the other hand, scatter of X-ray quanta and light
photons within the detector influence spatial resolution. Therefore, the intrinsic spatial
resolution for selenium-based direct conversion detectors is higher than that for
indirect conversion detectors (Michail 2010).
B.9.4. Modulation Transfer Function (MTF)
Modulation transfer function (MTF) is the capacity of the detector to transfer the
modulation of the input signal at a given spatial frequency to its output (Hendee 1970).
At radiography, objects having different sizes and opacity are displayed with different
gray-scale values in an image. MTF has to do with the display of contrast and object
29
CHAPTER B
size. More specifically, MTF is responsible for converting contrast values of
different-sized objects (object contrast) into contrast intensity levels in the image
(image contrast). For general imaging, the relevant details are in a range between 0
and 2 cycles/mm, which demands high MTF values. MTF is a useful measure of true
or effective resolution, since it accounts for the amount of blur and contrast over a
range of spatial frequencies (Michail 2010).
B.9.5. Noise Power Spectrum (NPS)
The noise power spectrum (NPS) is a spectral decomposition of the variance. As such,
the NPS of a digital radiographic image provides an estimate of the spatial frequency
dependence of the pixel-to-pixel fluctuations present in the image (Williams et al
1999). Such fluctuations are due to the shot (quantum) noise in the X-ray quanta
incident on the detector, and any noise introduced by the series of conversions and
transmissions of quanta in the cascaded stages between detector input and output.
Examples of the latter are the signal and dark current shot noise, the reset noise, the
amplifier noise, the white noise of source follower, the quantization noise, the fixed
pattern noise due to pixel-to-pixel variation in sensitivity or photoresponse
nonuniformity, the dark current fixed pattern noise and the offset differences between
pixel and column parallel signal readout. The fixed pattern noise is the result of
nonuniformities in the process causing small variations in device physical dimensions,
and differences in pixel and column voltage thresholds (Williams et al 1999,
Arvanitis et al 2007). The NPS is a much more complete description of image noise
than is quantification of integrated (total) noise via simple measurement of the rms
pixel fluctuations, because it gives information on the distribution in frequency space
of the noise power. An understanding of the frequency content of image noise can
30
CHAPTER B
provide insight regarding its clinical impact. For example, in mammography, excess
high-frequency noise may render the detection of microcalcifications impossible. If
desired, the total variance can be obtained by integrating the NPS over spatial
frequency (Williams et al 1999).
B.9.6. Detective Quantum Efficiency (DQE)
Detective quantum efficiency (DQE) is one of the fundamental physical variables
related to image quality in radiography and refers to the efficiency of a detector in
converting incident X-ray energy into an image signal. DQE is calculated by
comparing the signal-to-noise ratio at the detector output with that at the detector
input as a function of spatial frequency (Dainty and Shaw 1974). DQE is dependent
on radiation exposure, spatial frequency, MTF, NPS and detector material. High DQE
values indicate that less radiation is needed to achieve identical image quality;
increasing the DQE and leaving radiation exposure constant will improve image
quality. The ideal detector would have a DQE of 1, meaning that all the radiation
energy is absorbed and converted into image information. During the past few years,
various methods of measuring DQE have been established, making the comparison of
DQE values difficult if not impossible. In 2003, the IEC62220-1 (Marshall 2006a)
standard was introduced to standardize DQE measurements and make them
comparable (Medical Electrical Equipment 2003, Saunders et al 2005).
B.9.7. Information Capacity (IC)
The concept of image information capacity (IC) has been introduced within the
context of Shannon's information theory, in order to assess image information content
(Shannon 1948, Wagner et al 1979, Panayiotakis 1984, Kandarakis et al 2001).
However, little work relevant to medical imaging has been published up to now
31
CHAPTER B
(Michail et al 2011, Kanamori 1968, Maiorchuk et al 1974, Brown et al 1979, Shaw
1979, Evans 1981, Kanamori and Matsuoto 1984, Michail et al 2009, Gregg 1967,
Cavouras et al 2000). According to Shannon’s theory the image information capacity,
per unit of image area, is given as a function of the output signal to noise ratio.
Information capacity shows an estimation of the system performance using a single
value index, instead to the spatial frequency dependent parameters. Information
capacity can be applied in newly developed technologies such as telemammography
(Mahesh 2004), which is an advancement of digital mammography, since it allows
underserved and geographically remote populations to access the latest in breast care.
32
CHAPTER C
CHAPTER C
MATERIALS AND METHODS
C.1. Experimental Setup
C.1.1. Phosphor Screen Preparation
The phosphor Gd2O2S:Eu for the experiments, were purchased in powder form
(Phosphors Technology Ltd, England, code UKL63/N-R1 with mean grain size of
approximately 8 μm and volume density of 7.3. The phosphor was used in the form of
thin layers (test screens) to simulate the intensifying screen employed in X-ray
mammography and radiography. The screens prepared by sedimentation of the
powder phosphors on Borosilicate glass substrate 22x22 mm2 with thickness of 0.14
mm (Waldemar Knittel- GmbH). Sodium Orthosilicate (Na2SiO3) was used as binding
material between the powder grains. The sedimentation procedure was achieved by
using a mixture consisting of 500 ml of de-ionized water, 10 ml of Na2SiO3, and the
appropriate amount of phosphor powder in a glass tube of 100 cm height. The
Borosilicate glass substrate was placed at the bottom of the tube (Giakoumakis et al
1990, Kandarakis et al 1997).
Sedimentation has been a widely accepted technique for preparation of
radiographic phosphor screens with good homogeneity in various dimensions (i.e.
area and thickness) and spatial resolution (Sadowsky 1949, Shigeo and William 1998).
The screens obtained, with a packing density of the order of 50% which is common in
commercial phosphor screens (Sadowsky 1949, Shigeo and William 1998).
C.1.2. CMOS Sensor
For the detector under investigation, two Gd2O2S:Eu phosphor screens, with coating
thicknesses of 33.3 mg/cm2, for mammography, and 65.1 mg/cm2, for radiography,
33
CHAPTER C
were manually coupled to an optical readout device including a CMOS Remote
RadEye HR photodiode pixel array (Remote RadEye Systems Rad-icon Imaging
Corporation a division of Dalsa). The CMOS photodiode array consists of 1200×1600
pixels with 22.5  m pixel spacing. The Gd2O2S:Eu screens were overlaid onto the
active area of the CMOS photodiode array, with the phosphor layer directly coupled
to the photodiode array to prevent image degradation caused by the glass substrate.
The photodiode array consists of an N-well diffusion on p-type epitaxial Silicon, and
held by using a thin Polyurethane foam layer for compression between the screen and
a 1-mm-thick Graphite cover. A component view is shown in Figure 5. Experiments,
on the CMOS optical sensor, were carried out in both X-ray mammography and
radiography energy ranges. Three beam qualities were used; 28 kV W/Rh and W/Ag
for mammography and 70 kV (RQA 5) for radiography according to the IEC
standards (Medical Electrical Equipment-Characteristics of Digital X-Ray Imaging
Devices 2005). IEC standards X-ray spectra were achieved by adding 2 mm Al and 21
mm Al filtration respectively. Half value layers were calculated and found 7.1 mm for
radiography and 0.77 and 0.89 mm for W/Rh and W/Ag mammographic conditions
respectively.
FIG. 4. Gd2O2S:Eu screen coupled to RadEye HR CMOS sensor.
34
CHAPTER C
A Hologic Selenia X-ray tube, available with two target/filter combinations (W\Rh
and W/Ag), was used for the mammographic beam qualities and a Del Medical
Eureka X-ray tube with a rotating Tungsten anode and 1.5 mm Aluminum (Al)
inherent filtration for the RQA 5 beam quality.
In this study the source-to-detector distances (SDD) between the X-ray focal
spot and the surface of the detector was set to 185 and 66 cm for the radiographic and
mammographic energies, respectively (Medical Electrical Equipment-Characteristics
of Digital X-Ray Imaging Devices 2005). The added filtration was placed as close as
possible to the source.
C.1.3. X-Ray Spectra Measurement
A portable Amptek XR-100T X-ray spectrometer (X-ray & Gamma Ray Detector,
XR-100T-CdTe, Amptek, Bedford), based on a cadmium telluride (CdTe) crystal
solid-state detector was used for direct diagnostic X-Ray spectra measurements (L.
Abbene et al 2007). According to manufacturer’s data, the energy resolution of the
CdTe detector at 122 keV, obtained with a source, was less than 1.2 keV at the full
width at half maximum (FWHM) (X-ray & Gamma Ray Detector, XR-100T-CdTe,
Amptek, Bedford, MA). The CdTe detector was calibrated for energy scales, linearity
checks and energy resolution by using  -ray calibration sources of
125
I and
99m
Tc.
The CdTe was placed at a focus to detector distance of 166 and 48 cm for the
radiographic and the mammographic energies respectively. X-ray spectra were
corrected by the inverse square law at the CMOS detector plane for all beam qualities.
In order to minimize pile-up distortions, dedicated collimation systems (1 mm thick
collimators with 400  m and 100  m diameters, respectively for the radiographic
and mammographic X-ray spectra) were used. In addition the measured X-ray spectra
35
CHAPTER C
were corrected for the efficiency of the CdTe detector. The air Kerma value K (in
mGy) at the surface of the detector was calculated according to (L. Abbene et al
2007):
E0
K  0.00869   (1.83 106   0 ()   ( en ( E ) /  ) air )
 min
(1)
Where Φ0 is the measured X-ray spectrum value (photons/mm2) at energy E (keV).
(en ( E ) /  )air ) (cm2/g) is the X-ray mass energy absorption coefficient of air at
energy E obtained from the literature (J. R. Greening 1985). The exposure rate at the
entrance surface of the CMOS photodiode array detector was measured for a range of
tube current-time products (mAs). The measured X-ray spectra were corrected for the
efficiency of the CdTe detector.
C.2. Image quality
C.2.1. Spectral Compatibility
The spectral compatibility between the emitted scintillator light and the spectral
sensitivity of the CMOS optical sensors was calculated by the spectral matching
factor (  S ) according to (1):
S   SP ( )SD ( )d  /  SP ()d 
(2)
where S P ( ) is the emitted light spectrum of the phosphor and S D ( ) is the spectral
sensitivity of the optical detector coupled to the phosphor (Michail et al 2010). The
S P ( ) of the powder phosphor screens was measured under X-ray excitation by an
36
CHAPTER C
optical spectrometer (Ocean Optics Inc., HR2000).
C.2.2. Signal Transfer Property (STP)
The Signal Transfer Property (STP), or detector response, states the relationship
between mean pixel value (MPV) and ESAK at the detector surface. This relationship
was obtained by plotting pixel values (PV) versus ESAK at the detector, as described
in the IEC method (Medical Electrical Equipment-Characteristics of Digital X-Ray
Imaging Devices 2005). A sequence of uniform images was acquired at different
exposure levels. MPV was evaluated in a 1 × 1 cm region of interest (ROI). The mean
pixel value (MPV) and the standard deviation within that region were measured. The
system’s response curve was fitted using a linear equation of the form:
MPV    b  K
(3)
Where  and b are fit parameters. From the slope of the system’s response curve, the
value of the gain factor (G) was obtained. This value is relating MPV to the incident
exposure at the detector (in digital units per Gy ) (Neitzel et al 2004). The
magnitude of the pixel offset at zero air Kerma was also estimated (Samei et al 1998).
C.2.3. Modulation Transfer Function (MTF)
The MTF was measured using the slanted-edge technique, following the procedures
described in IEC standard (Marshall 2006a, Dobbins 2000, Illers et al 2005). A PTW
Freiburg tungsten edge test device was used to obtain the slanted edge images in both
radiographic and mammographic energies. The edge test device consists of a 1 mm
thick W edge plate (100×75 mm2) fixed on a 3 mm thick lead plate. Images of the
edge, placed at a slight angle, were obtained under the radiographic and
mammographic imaging conditions. Three exposure levels have to be chosen for the
37
CHAPTER C
measurements. From these levels, the medium one should be the ’normal’ level
routinely employed in the clinical practice (Medical Electrical EquipmentCharacteristics of Digital X-Ray Imaging Devices 2005). The edge spread function
(ESF) was calculated by the extraction of an 1×1 cm2 ROI with the edge roughly at
the center. The angle of the edge was then determined using a simple linear least
squares fit and the 2D image data were re-projected around the angled edge (Marshall
2006a) to form an ESF with a bin spacing of 0.1 pixels. The ESF was smoothed with
a median filter of five bins to reduce high frequency noise. The median filter is much
less sensitive to extreme values (outliers) than averaging filters. Therefore it is more
efficient to remove these outliers without reducing the sharpness of the image. Then
the ESF was differentiated to obtain the line spread function (LSF) (Greening 1985,
Boone 2001). Finally, the normalized LSF was Fourier transformed to give the presampling MTF.
C.2.4. Normalized Noise Power Spectrum (NNPS)
The NPS was calculated according to IEC 62220-1-2. (Medical Electrical EquipmentCharacteristics of Digital X-Ray Imaging Devices 2005) For each ROI, the PV were
converted into air kerma units with equation (4). The slowly varying spatial
background effects including the heel effect were corrected by fitting and subtracting
a two-dimensional second-order polynomial to the original acquired image data. The
area of analysis was subsequently divided into sub-images of 1024×1024 pixels. Half
overlapping ROIs with a size of 128×128 pixels were then taken from the sub-images
(Medical Electrical Equipment-Characteristics of Digital X-Ray Imaging Devices
2005). A total of 128 ROIs were taken from each flood image. For all the ROIs taken
from each image 2D fast Fourier transform (FFT) of each ROI was calculated and
38
CHAPTER C
added to the NPS ensemble. NNPS was obtained by dividing NPS by the square of the
corresponding K and afterwards the ensemble average was obtained.
C.2.5. Detective Quantum Efficiency (DQE)
The DQE of the system was calculated by the following equation:
DQE (u ) 
MTF 2 (u )
K  q  NNPS (u )
(4)
Where q is the number of photons per unit Kerma ( Gy ) per mm2, determined by
dividing the number of photons per mm2 (measured with the portable X-ray
spectrometer) with the corresponding air Kerma value ( Gy ) (Abbene et al 2007).
Values of 21738 photons  Gy 1  mm2 , for the RQA 5, and 6161 and 5422
photons  Gy 1  mm2 , for mammographic W/Ag and W/Rh beam qualities
respectively were calculated from our direct X-ray spectra measurements, instead of
using tabulated data.
C.2.6. Noise Equivalent Quanta (NEQ)
A measure of the lower number of exposure quanta required by an ideal detector to
2
yield the same SNRout
as the practical detector working at a given exposure level can
be expressed by the noise equivalent quanta (NEQ) (Shaw 1978). The latter is the
combined effects of signal and noise, in terms of spatial frequency. Also provides an
index of the signal-to-noise ratio (SNR) associated with the diagnostic value of a
medical image. NEQ can be written as follows (Dobbins 2000):
NEQ(u)  DQE(u) 0
(5)
39
CHAPTER C
C.2.7. Information Capacity (IC)
The concept of image Information Capacity (IC) has been introduced within the
context of Shannon's information theory, in order to assess image information content
(see Appendix) (Wagner et al 1979, Dainty and R. Shaw 1974, Shannon 1948, Jones
1961, Jones 1962, Kanamori 1968, Maiorchuk et al 1974, Brown et al 1979, Shaw
1979, Evans 1981, Kanamori and Matsuoto 1984). However, little work relevant to
medical imaging has been published up to now (Michail et al 2011, Kanamori 1968,
Maiorchuk et al 1974, Brown et al 1979, Shaw 1979, Evans 1981, Kanamori and
Matsuoto 1984, Michail et al 2009, Gregg 1967, Cavouras et al 2000). According to
Shannon’s theory the image information capacity, per unit of image area, is given by
(6) (see Appendix) (Kanamori 1968, Brown et al 1979, Shaw 1979, Evans 1981,
Kanamori and Matsuoto 1984, Shannon 1998, Kandarakis et al 2001c Cavouras et al
2000).

IC  π  log 2 1  NEQ( u ) udu
0
(6)
IC shows an estimation of the system performance using a single value index, instead
of the spatial frequency dependent parameters (Kandarakis et al 2001c).
40
CHAPTER D
CHAPTER D
RESULTS AND DISCUSSION
D.1. Optical Properties and MTF of Borosilicate Glass Substrate
The optical properties of the Borosilicate glass substrate have been measured via a
Perkin-Elmer Lambda 15 UV/VIS Spectrophotometer. In the visible range the optical
transparency was found very high, of the order of 93%, while absorption was found
very low (0.029). These values show that the substrate has very high transparency,
low reflectivity and low absorption in the wavelength range of the phosphor emission
In addition the MTF of the thin glass substrate was experimentally
investigated. The MTF was measured: a) with the glass substrate directly coupled to
the photodiode array (MTF1) and b) with the phosphor directly coupled to the
photodiode array (MTF2). In both cases the slanted-edge technique was used (Michail
et al 2011). By taking into account that the MTF of an imaging system is the product
of separate MTF components of an imaging system, thus the MTF of the glass
substrate (MTFglass) is (Michail 2010):
MTFglass (u ) 
MTF1 (u )
MTF2 (u )
(7)
Figure 5 shows the MTF1 and the MTF2 of the detector. From this figure can be
depicted an obvious degradation of the MTF due to thin glass substrate. MTF1 is 40%
lower than MTF2 on average, in the whole spatial frequency range. When the glass
substrate is directly coupled to the photodiode array (MTF1), the light photons emitted
41
CHAPTER D
by the phosphor interact with the glass substrate leading to a subsequent degradation
of the MTF. When the phosphor is directly coupled to the photodiode array (MTF2)
the light photons are directly detected by the photodiode array without interaction
with the glass substrate. In this case there is no degradation of the MTF caused by the
glass substrate. Furthermore figure 4 shows the MTF of the glass substrate (MTFglass)
calculated by (1). The glass substrate has a low impact on resolution of the image at
low frequencies (12% on average, up to 2.7 cycles/mm) and a higher impact at
medium and high frequencies (70% on average).
FIG.5. Comparison of the MTFs curves for both cases. MTF of the glass substrate is also shown.
D.2. Spectra Compatibility
The spectral compatibilities between the optical spectra emitted by Gd2O2S:Eu and
Gd2O2S:Tb screens and the sensitivity of the CMOS optical sensor were calculated for
the wavelengths range from 375 nm up to 675 nm by evaluating the spectral matching
factor (relation (3)). In particular, it was found that this factor is higher in the case of
Gd2O2S:Eu screens coupled to the CMOS sensor (matching factor: 1), than in the case
of the Gd2O2S:Tb screens (matching factor:0.95). This shows that red emitting
42
CHAPTER D
phosphors have better compatibility with the CMOS sensors and, apparently, with the
rest of Silicon based sensors. Such results should be taken into account in evaluating
phosphor and scintillator materials since their effective efficiency can be significantly
altered depending on the optical sensor incorporated in the detector.
D.3. X-Ray Spectra
Figures 6 and 7 show measured spectra produced by the X-ray tubes (radiographic
with W/Al target-filter combination and mammographic with W/Ag and W/Rh targetfilter combinations), used in the present study. The tube settings were: 20, 63 and
157.5 mAs at 70 kVp (RQA 5) and 20 mAs, 60 mAs and 120 mAs at 28 kV
respectively. The system reproducibility was verified by measuring X-ray spectra
several times. Entrance surface air Kerma was calculated by relation (2) using the
Amptek XR-100T X-ray spectrometer measurements and was found 11.2, 34.1 and
87.5 Gy for the RQA 5 and 65.8, 197.4 and 394.1 Gy and 55.7, 167.0 and 334.0
Gy for mammographic conditions W/Ag and W/Rh , respectively.
43
CHAPTER D
Table I. Parameters of the mammographic and radiographic beam qualities.
Radiation Beam
Mammographic
Mammographic
Radiographic
Parameters
quality (W/Rh)
quality (W/Ag)
quality (RQA 5)
W/Rh
W/Ag
W/Al
Tube voltage (KV)
28
28
70
Added filtration (mm Al)
2
2
21
Measured HVL (mm Al)
0.77
0.89
7.1
Q( photons  Gy 1  mm2 )
5422
6161
21738
Anode/filtration
combination
FIG. 6. Measured X-ray spectra for RQA 5 beam quality.
44
CHAPTER D
FIG. 7. Measured X-ray spectra for W/Rh and W/Ag beam qualities.
D.4. Signal Transfer Property (STP)
Figure 8 shows the detector response curve (STP) of the RadEye HR CMOS sensor
under RQA 5 (70kVp) and mammographic (28kVp) beam qualities. The detector was
found to have a linear response, covering the whole exposure range, with a pixel value
offset of 144.879 for the RQA 5 beam quality and 328.912 and 296.796 for the W/Ag
and W/Rh qualities, respectively. The linear no threshold fits gave correlation
coefficients (R2) greater than 0.9978, 0.9971 and 0.9989 for the RQA 5 and the Ag
and Rh filtered beam qualities, respectively. The gain factor G was determined as the
slope of the characteristic curve, relating the mean pixel value to the incident
exposure. Using flat-field images the gain factors were determined by linear
regression to be G=7.347 digital units per Gy for RQA 5 and 2.635 and 2.400
digital units per Gy for Ag and Rh filtered beams, respectively.
45
CHAPTER D
FIG. 8. STP curves for the RQA 5, W/Rh and W/Ag beam qualities.
D.5. Modulation Transfer Function (MTF)
Figure 9 shows the modulation transfer function (MTF) curves for three exposure
levels under the RQA5 beam quality. In this figure the MTF of a CMOS sensor with
48  m pixel pitch coupled to a Gd2O2S:Tb screen of 59.2 mg/cm2 coating thickness
(Cho et al 2008), is also shown for comparison purposes. This MTF has been obtained
under RQA 5 beam quality at 43.5 Gy . As it can be depicted from figure 9, the
MTF of the 48  m CMOS sensor (Cho et al 2008) is 13% lower, on average, than
the MTF of the 22.5  m CMOS sensor (investigated in the present study) in the
frequency range from 0 to 5 cycles/mm. In the higher spatial frequency range (5-10
cycles/mm), the MTF values differ by 3% on average. To explain these differences
the combined effects of screen thickness/sensor pixel size, as well as the screen
optical properties must be taken into account. The screen of our system is slightly
thicker (65.1 mg/cm2 instead of 59.2 mg/cm2) and, in addition, it has not been
optically optimized (i.e. incorporation of special light absorbing dyes) like
commercially available screens. However, taking into account our results, it could be
46
CHAPTER D
claimed that the red phosphor based system can show improved or acceptable signal
transfer properties in the whole spatial frequency range.
FIG. 9. Comparison of the MTFs of the 65.1 mg/cm2 Gd2O2S:Eu/CMOS detector with a 59.2 mg/cm2
Gd2O2S:Tb/CMOS detector under RQA 5 beam quality.
Figure 10 shows MTF curves for mammographic beam qualities with two different
anode/filter combinations (W/Ag and W/Rh). MTFs obtained with the Rhodium
filters are higher than those obtained with the Silver filter by approximately 5, 6 and 5
%, for the three exposure levels respectively. Figure 11 shows a comparison between
the MTFs of the present CMOS sensor combined with the 33.3 mg/cm2 Gd2O2S:Eu
screen and the, previously published, CMOS sensor coupled to a 33.91 mg/cm2
Gd2O2S:Tb screen (Michail et al 2011). The MTFs of the present CMOS/33.3 mg/cm2
Gd2O2S:Eu screen are comparable to those of the CMOS/33.91 mg/cm2 Gd2O2S:Eu
screen (Michail et al 2011) (deviations lower than 5, 2 and 1 % for the three exposure
levels). However a point by point comparison is not possible, since exposure levels
and beam qualities were clearly different in the two studies, e.g. W/Rh and W/Ag
47
CHAPTER D
anode/filters combinations in the present study instead of Mo/Mo (RQA M2) of the
previous study.
FIG. 10. MTF curve for W/Rh and W/Ag beam qualities.
FIG. 11. Comparison of the MTFs for the CMOS/Gd2O2S:Eu sensor under investigation and a
previously published CMOS/Gd2O2S:Tb sensor under W/Rh and RQA M2 beam qualities respectively.
48
CHAPTER D
D.6. Normalized Noise Power Spectrum (NNPS)
Figures 12 and 13 show extracted 1D normalized noise-power spectra (NNPS), for the
u direction, obtained from uniformly exposed images under general radiography
(RQA 5) and mammographic conditions with two different anode/filters combinations
(Ag and Rh). NNPS shows variation over a large spatial frequency range. For
example, a decrease from 3.54×10-6 mm2 down to 2.77×10-6 mm2 (i.e. 22%) was
observed in NNPS values in the range from 2 to 5 cycles/mm (exposure level 11.2
Gy for the RQA 5 beam quality). In the W/Ag and the W/Rh beam qualities
(exposure levels 197.4 Gy and 167.0 Gy respectively) a decrease of 58% and 50%
was observed. This is attributed to the fact that in the CMOS system, a 2D correction
is applied reducing the structured noise in each image. In the medium to high
frequency range, the noise levels of the image are related to the MTF of the system.
The rather slow variation of the NNPS may be attributed to the fact that MTF remains
high up and in the higher frequency range. For comparison purposes, NNPS of a
CMOS sensor measured in a previous study is also shown in Figure 12. (Cho et al
2008) This curve was obtained from a 48  m passive pixel sensor under the
representative RQA 5 beam quality, at an exposure level of 43.5 Gy . Figure shows
that the noise levels of the CMOS sensor under investigation is higher than that of the
previously published study (Cho et al 2008), due to the larger effective pixel size of
the passive CMOS senso (Cho et al 2008) leading to higher photon collection
efficiency and to a subsequent increase in the signal to noise ratio (SNR) (Chen et al
2000, Farrell et al 2006). Furthermore in the fabrication process of the custom made
Gd2O2S:Eu screens which has not been optically optimized (i.e. incorporation of
49
CHAPTER D
special light absorbing dyes) and are possible to be subjected to more structural nonuniformities than the commercial Gd2O2S:Tb screen used in the previous study.
FIG. 12. Comparison of the NNPS curves between the CMOS/Gd2O2S:Eu sensor under investigation
and a previously published CMOS/Gd2O2S:Tb sensor under RQA 5 beam quality.
FIG. 13. NNPS curves for W/Rh and W/Ag beam qualities.
An additional comparison for the mammographic beam quality is shown in Fig. 14 in
which the NNPS of a CMOS sensor measured in a previous publication of our group
50
CHAPTER D
is also shown for comparison purposes (Michail et al 2011). This curve was obtained
for the CMOS devise under investigation, coupled to a commercial Gd2O2S:Tb
scintillation screen under the representative RQA M2 (Mo/Mo at 28KV) beam quality
at exposure level of 40.1 Gy (Michail et al 2011). NNPS of CMOS
sensor/Gd2O2S:Eu screens combination shows higher noise in the whole spatial
frequency range. The reduced NNPS values of Michail et al (2011) can also be
attributed to the fabrication process of the Gd2O2S:Eu screen, as well as, in the lower
energy components of the RQA-M2 used in that study. The latter may enhance the Xray absorption properties of Gd2O2S, resulting in lower noise values.
FIG. 14. Comparison of the NNPS curves between the CMOS/Gd2O2S:Eu sensor under investigation
and a previously published CMOS/Gd2O2S:Tb sensor under W/Rh and RQA M2 beam qualities
respectively.
D.7. Detective Quantum Efficiency (DQE)
Figures 15 and 16 show DQE curves for the RQA 5, W/Ag and W/Rh qualities,
obtained according to (5). DQE was investigated at various air Kerma settings at the
detector surface. DQE decreases as the ESAK increases due to the influence of the
51
CHAPTER D
NNPS and MTF. All the DQE curves showed an increase in the spatial frequency
from 0 to 1.8 cycles/mm for the RQA 5 and from 0 to 5 cycles/mm for the
mammography beam qualities. This is due to the fact that NNPS falls off rapidly in
these spatial frequency ranges. Thereafter the reduction rate of the NNPS falls off in
contrast with the corresponding of the MTF curves, contributing to a decrease in the
DQE. For comparison purposes, DQE of the 48  m pixel size CMOS sensor (Cho et
al 2008) (RQA 5) at an exposure level of 43.5 Gy is also shown in Figure 15. The
DQE values of the previously published study (Cho et al 2008) were higher in the low
spatial frequency range (up to 1.8 cycles/mm). This is due to the higher NNPS values
of this spatial frequency range of the detector under investigation which contributes to
lower DQE values. Thereafter DQE curve is comparable in the whole spatial
frequency range due to the combined effects of MTF, NNPS and incident X-ray
spectra, as depicted in the previous Figures. These results show that a further
improvement on the DQE can be achieved with a commercialised Gd2O2S:Eu screen
preparation which will lead to an additional noise reduction.
FIG. 15. Comparison of the DQE curves between the CMOS/Gd2O2S:Eu sensor under investigation
and a previously published CMOS/Gd2O2S:Tb sensor under RQA 5 beam quality.
52
CHAPTER D
FIG. 16. DQE curve for W/Rh and W/Ag beam qualities.
In Fig. 17, DQE curves, of the CMOS sensor under investigation, coupled to a
commercial Gd2O2S:Tb screen, published in a previous study, is also shown for
comparison purposes, under mammography conditions, at an exposure level of 40.1
Gy (Michail et al 2011). DQE values of the currently investigated CMOS
sensor/scintillating screen combination is lower than the corresponding of the
CMOS/Gd2O2S:Tb screen combination (Michail et al 2011), due to the higher noise
performance of the current CMOS/Gd2O2S:Eu for the reasons previously reported.
53
CHAPTER D
FIG. 17. Comparison of the DQE curves between the CMOS/Gd2O2S:Eu sensor under investigation
and a previously published CMOS/Gd2O2S:Tb sensor under W/Rh and RQA M2 beam qualities
respectively.
D.8. Noise Equivalent Quanta (NEQ)
Figures 18 and 19 show noise equivalent quanta (NEQ) of the RQA 5 and
mammographic beam qualities, calculated from the MTF, NNPS and the measured Xray photon distribution at the detector surface. The shapes of the NEQ curves are
affected by both NNPS and MTF. The amplitude of NEQ is affected by the number of
X-ray photons at the surface of the detector (photons/mm2). As shown in these
Figures, NEQ reaches a maximum (73563 mm-2) in the low frequency range (1.8
cycles/mm), under the RQA 5 (ESAK: 11.2 Gy ) conditions and in the medium
frequency range (183074 mm-2 at 3.6 cycles/mm) under mammographic conditions
(ESAK: 167.0 Gy W/Rh).
54
CHAPTER D
FIG. 18. NEQ curves for RQA 5 beam qualities.
FIG. 19. NEQ curves for W/Ag and W/Rh beam qualities.
In Fig. 20 NEQ values of a CMOS sensor is also shown for comparison purposes
(Michail et al 2011). NEQ values of this CMOS sensor is higher in the spatial
frequency range up to 1 cycles/mm. NEQ curves of both CMOS sensors are
comparable thereafter.
55
CHAPTER D
FIG. 20. Comparison of the NEQ curves between the CMOS/Gd2O2S:Eu sensor under investigation
and a previously published CMOS/Gd2O2S:Tb sensor under W/Rh and RQA M2 beam qualities
respectively.
D.9. Information Capacity (IC)
In Table II the information capacity values for the RQA 5 and mammographic beam
qualities are shown. CMOS sensor/Gd2O2S:Eu screens combinations show higher IC
values in the mammography energy range (2475-2821 bits/mm2), in comparison to the
radiography energy range (1730-1851 bits/mm2) due to the higher MTF and incident
air Kerma values. These data show that, for a given level of incident X-ray fluence,
information capacity is mainly determined by the intrinsic phosphor material
properties, while the role of the screen thickness is of low significance. Thin screens
exhibit low levels of brightness (Michail et al 2010), which is compensated for by
their higher spatial resolution (MTF), whereas thick screens show poor resolution
properties but with high brightness gain (Michail et al 2010). The practical value of
the IC is that it defines an imaging performance index that evaluates image
information quantity by a single numerical value. Information capacity is not
56
CHAPTER D
expressed for specific frequency values since it is the outcome of integrating over the
spatial frequency bandwidth (Michail 2010). IC may be considered as being roughly
proportional to the area under the curve of the frequency-dependent image SNR.
Thus, IC mainly reflects the quantity and not the quality of image information. Image
quality is better described via the frequency-dependent signal parameters. It must be
noted, however, that IC is significantly affected by the maximum frequency contained
in the signal, i.e. the frequency bandwidth over which integration in (6) is performed.
A practical problem is to properly select this maximum frequency, which may
significantly affect the final IC result. In addition to this, information capacity is
calculated using spatial frequency as a factor, which multiplies the logarithmic term
containing SNR. Hence, the contribution of high spatial frequency SNR components
to the final result may be relatively enhanced. This affects IC in a sense that it may be
preferably emphasized by high frequency information content, i.e. the kind of
information that is expressed by spatial resolution and sharpness, which is better
displayed by thin layers. This may explain the higher IC values obtained for thin
screens.
Table II. Information capacity and ESAK values for the mammographic and radiographic beam
qualities.
Entrance Surface Air-
Information capacity
Kerma (ESAK Gy )
(bits/mm2)
(W/Ag)
65.8±0.036
2475±26
(W/Ag)
197.4±0.071
2678±41
(W/Ag)
394.1±0.084
2733±39
(W/Rh)
55.7±0.031
2535±42
(W/Rh)
167.0±0.066
2821±34
Beam Quality
57
CHAPTER D
(W/Rh)
334.0±0.073
2747±25
RQA 5
11.2±0.024
1730±31
RQA 5
34.1±0.032
1826±29
RQA 5
87.5±0.059
1851±34
Figures 21 show an initial image of a simple electronic device obtained from the
CMOS/Gd2O2S:Eu sensor screen combination under the RQA 5 beam quality
(obtained at 74 KVp with 21 mm Al filtration). This image demonstrates, that the use
of a CMOS/Gd2O2S:Eu sensor screen combination may be practically feasible.
FIG. 21. Image obtained from the CMOS/Gd2O2S:Eu combination at 74 KVp with 21 mm Al filtration.
58
CHAPTER E
CHAPTER E
CONCLUSIONS AND FUTURE WORK
E.1. Conclusions
In the present study the imaging performance of a high resolution CMOS based
imaging sensor combined with custom made Europium (Eu3+) activated Gd2O2S
screens was investigated in terms of Modulation Transfer Function (MTF),
Normalized Noise Power Spectrum (NNPS), Detective Quantum Efficiency (DQE),
Noise equivalent Quanta (NEQ) and Information Capacity (IC) covering the
mammography and general radiography energy ranges.
The spectral matching factor between the optical spectra emitted by the
Gd2O2S:Eu and the Gd2O2S:Tb screens and the CMOS optical sensor, evaluated in the
present study, was 1 and 0.95 respectively. The ESAK values ranged between 11.287.5 Gy , for RQA-5, and between 65.8-334 Gy , for W/Rh, W/Ag beam qualities.
It was found that the detector response function was linear for the exposure ranges
under investigation. Under radiographic conditions the MTF of the present system
was found higher than previously published MTF data for a 48  m CMOS sensor, in
the low up to medium frequency ranges. DQE was found comparable, while the
NNPS appeared to be higher in the frequency range under investigation (0–10
cycles/mm). NEQ reached a maximum (73563 mm-2) in the low frequency range (1.8
cycles/mm), under the RQA 5 (ESAK: 11.2 Gy ) conditions. IC values were found to
range between 1730-1851 bits/mm2. Under mammographic conditions MTF, NNPS
and NEQ were found comparable to data previously published for the 48  m CMOS
sensor while the DQE was found lower. The corresponding IC values were found
ranging between 2475 and 2821 bits/mm2.
59
CHAPTER E
Results showed that the red emitting phosphor/CMOS sensor combination has
comparable image quality parameters in terms of MTF, NNPS, DQE and NEQ
compared to previously published data for Terbium activated Gd2O2S/CMOS sensor
combinations. Since the imaging performance of Europium (Eu3+) activated Gd2O2S
screens, combined to CMOS sensors was found comparable to that of Gadolinium
(Gd) activated Gd2O2S screens, red emitting phosphors could be used in digital
imaging systems, where the Si based photodetectors are more sensitive to longer
wavelength ranges, and particularly in the red wavelength range.
E.2. Future Work
The aim of the future work is the experimental and theoretical evaluation of the
imaging performance and light emission efficiency of nanophosphor materials in the
mammography and radiography energy ranges for application in a novel high
resolution digital CMOS based X-ray sensor. The optimum CMOS/powder
nanophosphor combination for digital imaging applications as well as for fast data
acquisition (digital breast tomosynthesis, dual energy imaging, scanning slot
mammography) techniques will be determined. Also the application of the
information theory through the information capacity (IC) for the evaluation of the
digital imaging CMOS based system will be used.
60
APPENDIX A
APPENDIX A: INFORMATION CAPACITY (IC)
According to Shannon’s information theory, if it is possible to distinguish Ns different
signal intensity levels of duration T on a channel, we can say that the channel can
transmit log 2 N s bits in time T. The rate of transmission is then log 2 N s / T . More
precisely, the image information capacity, per unit of image area, may be defined by
(5) as follows (Shannon 1998):
IC  lim log 2 N S / T  n p log 2 N S
T 
(A.1)
Where n p is the number of image elements (pixels) per unit of area that can be
registered in an image element. If the transmitter has a dynamic range (D) and the
noise in the transmission channel is assumed to be Gaussian (Normal (0, σ 2 ) and
independent of the signal amplitude, then an expression may be derived for Ns in
terms of D and σ . For this case, the levels are equally spaced between zero and D
with a spacing determined by the acceptable error rate for signal transmission. If each
level have an interval of ± kσ , according to Wagner (Wagner et al 1979), then the
spacing becomes 2kσ , and one obtains: N S  1  D / 2kσ . (A.1) may now be written
as follows: IC  nlog 2 ( 1  D / 2kσ ) . Shannon showed that in the frequency domain
there is a definition of information capacity that is consistent with the equations
described above and which serves as a unique upper limit for the rate of information
transmission (Shannon 1998). This upper limit for the number of distinguishable
signals
in
communication
channel
of
bandwidth
Δf
is
defined
as:
61
APPENDIX A
NS 

( SPS  NPS ) / NPS

2T f
. Where SPS is the signal power spectrum which can
defined as: SPS  ( MTF  G )2 , where G is the gain factor (in digital units per Gy )
(Wagner et al 1979) and NPS is the noise power spectrum. Consequently, the channel
capacity is limited by(Shannon 1998):
IC  log2 NS / T  f log2 ((SPS  NPS ) / NPS )
(A.2)
Where SPS and NPS are assumed constants over the frequency interval Δf . Wagner
has expressed this equation (Wagner et al 1979) under the following form:
ΙC( Δu )  2 Δu log 2 ( SPS  NPS ) / NPS 
1/ 2
(A.3)
Where the exponent of ½ is explained since the ratio of amplitudes is replaced by a
ratio of powers, and the factor 2Δu results directly from the sampling theorem. For
any function having a bandwidth of Δu , sampling at intervals 1 / 2 Δu completely
specifies the function. If the function is assumed to be limited to a time interval T,
then for a real function there are just 2ΔuΤ independent values of 2Δu values per
unit time. The dimensions of IC are bits per unit time (Wagner et al 1979). If (A.2) is
integrated over the whole frequency range then information capacity is given by:
ΙC  0u log 2 1  SPS( u ) / NPS( u )du
(A.4)
62
APPENDIX A
Where u is greater than or equal to the band limit of the signal power. The 2D form of
(A.3) is given by IC  ( 2 Δux )( 2 Δu y )log ( SPS  NPS ) / NPS 1/ 2 , written in terms of
2
spatial frequency, leading similarly to an integral over the upper right-hand quadrant
in u x , u y ( u  ( ux 2  u y 2 )1/ 2 ) space or using polar frequency coordinates and assuming
cylindrical symmetry of SPS and NPS,
ΙC  1 / 20 log 2 1  SPS( u ) / NPS( u )2πudu
u
(A.5)
expressed in bits per unit area (Wagner et al 1979). For the detectors used in both
conventional and digital X-ray imaging, power spectra can be represented as onedimensional functions of spatial frequency u, SPS(u) and NPS(u) due to rotational
symmetry (Wagner et al 1979). (A.5) now becomes (Kanamori 1968, Brown et al
1979, Shaw 1979, Evans 1981, Kanamori and Matsuoto 1984, Shannon 1998,
Kandarakis et al 2001c, Cavouras et al 2000)

IC  π  log 2 1  ( SPS( u ) / NPS( u )) udu
0
(A.6)
Where SPS( u ) / NPS( u )  NEQ , therefore (A.6) becomes(A.7):

IC  π  log 2 1  NEQ( u ) udu
0
(A.7)
63
APPENDIX B
APPENDIX B: ABBREVIATIONS
AMFPI
Active Matrix Flat Panel Imagers
a-Si:H
Amorphous Silicon
CCD
Charge-coupled device
CR
Computed Radiography
CMOS
Complementary metal Oxide semiconductor
Si
Crystalline Silicon
DQE
Detective Quantum Efficiency
ESAK
Entrance surface air-Kerma
FT
Fourier Transform
HVL
Half-value layer
IEC
International Electrotechnical Commission
IC
Information capacity
nI
Informational efficiency
0
Incident X-ray photon fluence
FFT
Fast Fourier Transform
Gd2O2S:Tb
Gadolinium Oxysulphide
LSF
Line Spread Function
Lu2SiO5:Ce (LSO)
Lutetium Oxyorthosilicate
MPV
Mean pixel value
Mo
Molybdenum
MTF
Modulation transfer function
NEQ
Noise equivalent quanta
NPS
Noise power spectrum
NNPS
Normalized noise power spectra
64
APPENDIX B
PPS
Passive-pixel sensors
PSP
Photostimulable Storage Phosphor
Rh
Rhodium
SEM
Scanning electron microscope
SNR
Signal to noise ratio
Ag
Silver
SDD
Source to detector distance
as
Spectral matching factor
CsI:Tl
Thallium-doped Cesium Iodide
TFT
Thin film field-effect transistors
W
Tungsten
X
X-ray exposure
65
APPENDIX C
APPENDIX C: LIST OF FIGURES
FIG.1. Photostimulable storage phosphor system
FIG.2. Charged couple devise (CCD)
FIG.3. Active matrix flat panel imager (AMFPI)
FIG.4. Gd2O2S:Eu screen coupled to RadEye HR CMOS sensor.
FIG. 5. Comparison of the MTFs curves for both cases. MTF of the glass substrate is
also shown.
FIG. 6. Measured X-ray spectra for RQA 5 beam quality.
FIG. 7. Measured X-ray spectra for W/Rh and W/Ag beam qualities.
FIG. 8. STP curves for the RQA 5, W/Rh and W/Ag beam qualities.
FIG. 9. Comparison of the MTFs of the
CMOS /Gd2O2S:Eu sensor under
investigation and a previously published CMOS /Gd2O2S:Tb sensor under RQA 5
beam quality.
FIG. 10. MTF curve for W/Rh and W/Ag beam qualities.
FIG. 11. Comparison of the MTFs for the CMOS/Gd2O2S:Eu sensor under
investigation and a previously published CMOS/Gd2O2S:Tb sensor under W/Rh and
RQA M2 beam qualities respectively.
FIG. 12. Comparison of the NNPS curves between the CMOS/Gd2O2S:Eu sensor
under investigation and a previously published CMOS/Gd2O2S:Tb sensor under RQA
5 beam quality.
FIG. 13. NNPS curves for W/Rh and W/Ag beam qualities.
FIG. 14. Comparison of the NNPS curves between the CMOS/Gd2O2S:Eu sensor
under investigation and a previously published CMOS/Gd2O2S:Tb sensor under W/Rh
and RQA M2 beam qualities respectively.
66
APPENDIX C
FIG. 15. Comparison of the DQE curves between the CMOS/Gd2O2S:Eu sensor under
investigation and a previously published CMOS/Gd2O2S:Tb sensor under RQA 5
beam quality.
FIG. 16. DQE curve for W/Rh and W/Ag beam qualities.
FIG. 17. Comparison of the DQE curves between the CMOS/Gd2O2S:Eu sensor under
investigation and a previously published CMOS/Gd2O2S:Tb sensor under W/Rh and
RQA M2 beam qualities respectively.
FIG. 18. NEQ curves for RQA 5 beam qualities.
FIG. 19. NEQ curves for W/Ag and W/Rh beam qualities.
FIG. 20. Comparison of the NEQ curves between the CMOS/Gd2O2S:Eu sensor under
investigation and a previously published CMOS/Gd2O2S:Tb sensor under W/Rh and
RQA M2 beam qualities respectively.
FIG. 21. Image obtained from the CMOS/Gd2O2S:Eu combination at 74 KVp with 21
mm Al filtration.
67
REFERENCES
REFERENCES
Abbene L, Manna A. L, Fauci F, Gerardi G, Stumbo S, and. Raso G, (2007),“X-ray spectroscopy and
dosimetry with a portable CdTe device,” Nucl. Instrum. Methods Phys. A571, 373–377.
Aerts W, De Meester P J and Bollen R, (1982),“A Monte Carlo simulation of complete systems for
radiographic image registration,” Mater. Eval. 40, 1071-75.
Alig R C and Bloom S, (1977) “Cathodoluminescent Efficiency” J. Electrochem. Soc. 124 1136-38.
Alvarez R E, (2010) “Near optimal energy selective x-ray imaging system performance with simple
detectors,” Med. Phys. 37, 822-841.
Alvarez R E, Seibert J A, Thomson S K, (2004),“Comparison of dual energy detector system
performance,” Med. Phys. 31(3), 556-565.
Andreo P, (1991),“Monte Carlo techniques in medical radiation physics”, Phys. Med. Biol. 36, 861920.
Antonuk L E, Zhao Q, Mohri Y El,. Du H, Wang Y, Street R A, Ho J, Weisfield R and. Yao W,
(2009),“An investigation of signal performance enhancements achieved through innovative
pixel design across several generations of indirect detection, active matrix, flat-panel arrays,”
Med. Phys. 36(7), 3322-3339.
Arnold B A, (1979),“Physical characteristics of screen-film combinations”. In: The Physics of Medical
Imaging: Recording System, Measurements and Techniques ed A G Haus (New York:
American Association of Physicists in Medicine) p 30.
Artz DS, (1997),“Computed radiography for the radiological technologist, ”Seminars in
Roentgenology, vol. 32, no. 1, pp. 12-24.
Arvanitis C D, Bohndiek S E, and Royle G, Blue A, Liang H X, Clark A, Prydderch M, and Turchetta
R,Speller R (2007) “Empirical electro-optical and x-ray performance evaluation of CMOS
active pixels sensor for low dose, high resolution x-ray medical imaging,” Med. Phys. 34(12)
4612-25.
Baxes, GA (1994),“Digital image processing: Principles and applications,” John Wiley & Sons, New
York.
Baysal M A and Toker E (2005),“A new CMOS-based digital imaging detector for applications in
mammography,” Proc. SPIE 5923 1–11.
68
REFERENCES
Bigas M,. Cabruja E,. Forest J, and. Salvi J, (2006),“Review of CMOS image sensors,” Microelectron.
J. 37, 433-451.
Billas I, (2011) “Angiogenesis measurements in mammography using time-resolved dual energy
analysis,” Master thesis, university of Patras
Black G and Linfoot E H (1957),“Spherical aberrations and the information capacity of optical
images,” Phil. Trans. Roy. Soc. A239 522-40.
Blasse G and Grabmaier B C (1994),“Luminescent materials,” (Berlin: Spinger) pp.21-9.
Boone J M (2000),“X-ray production, interaction, and detection in diagnostic imaging,” in Handbook
of Medical Imaging vol. 1 Physics and Psycophysics ed J Beutel, H L Kundel and R L Van
Metter (Bellingham, WA: SPIE pp 3-77 chapter 2.
Boone J. M, (2001),“Determination of the presampled MTF in computed tomography,” Med. Phys.
28(3), 356-360.
Boone J. M, Tecotzky M, Alexander G. M, (1992),“Binary screen detector system for single-pulse
dual-energy radiography,” Radiology 183, 863-870.
Brown D.G, Anderson M.P and. Wagner R.F, (1979),“Information capacity considerations in medical
imaging,” SPIE 206, 77-82.
Bruckmann A. and Uhl A, (2000),“Selective medical image compression techniques for telemedical
and archiving applications,” Computers in Biology and Medicine, 30, 153–69.
Bushberg JT, Seibert JA, Leidholdt EM & Boone JM, (2002),“The essential physics of medical
imaging,” 2nd edn, Lippincott, Williams and Wilkins, Philadelphia.
Bushong SC, (2001),“Radiological science for technologists: Physics, biology and protection,”, 7th
edn, Mosby, St Louis.
Carpenter C M, Pratx G, Sun C and Xing L, (2011),“Limited-angle x-ray luminescence tomography:
methodology and feasibility study,” Phys. Med. Biol. 56, 3487-3502.
Carpenter CM, Sun C, Pratx G, Rao R, Xing L, (2010),“Hybrid x-ray/optical luminescence imaging:
characterization of experimental conditions,” Med Phys. 37(8), 4011-4018.
Cesar LJ, Schueler BA, Zink FE, Daly TR, Taubal JP & Jorgenson LL, (2001),“Artefacts found in
computed radiography,” British Journal of Radiology, vol. 74, pp. 195-202.
Chen T, Catrysse P. B, El Gamal A, and Wandell Brian A, (2000),“How small should pixel size be?,”
Proc. SPIE, 3965, 451-459.
69
REFERENCES
Cho M. K, Kim H. K, Graeve T, Yun S. M, Lim C. H, Cho H, and. Kim J. M, (2008),“Measurements
of X-ray Imaging, Performance of Granular Phosphors With Direct-Coupled CMOS Sensors,”
IEEE Trans. Nucl. Sci. 55(3), 1338-1343.
Chotas H. G, Dobbins III J T, and Ravin C E, (1999),“Principles of digital radiography with large-area,
electronically readable detectors: A review of the basics,” Radiology, 210, 595-599.
Chotas HG, Floyd CE Jr, Dobbins JT 3d & Ravin CE, (1993),“Digital chest radiography with
photostimulable storage phosphors: signal-to-noise ratio as a function of kilovoltage with
matched exposure risk”, Radiology, vol. 186, no. 2, pp. 395-8.
Curry TS, Dowdey JE & Murray RC, (1990),“Christensens physics of diagnostic radiology,” 4th edn,
Lea and Febiger, Philadelphia.
Dainty J C and Shaw R, (1974),“Image Science,” (New York:Academic Press) pp. 232.
Derenzo S E, Moses W W, Cahoon J L, Perera R C C, and Litton J E, (1990),“Prospects for new
inorganic scintillators”, IEEE Trans. Nucl. Sci. NS-37, 203-08.
Dick C E and Motz J W (1981),“Image information transfer properties of X-ray fluorescent screens,”
Med. Phys. 8, 337-46.
Dick C E and Motz J W, (1981),“Utilization of monoenergetic X-ray beams to examine the properties
of radiographic intensifying screens,” IEEE Trans. Nucl. Sci. NS-28, 1554-58.
Dobbins III J T and Godfrey D J, (2003),“Digital x-ray tomosynthesis: current state of the art and
clinical potential,” Phys. Med. Biol. 48, R65-R106.
Dobbins J. T III, (2000),“Image quality metrics for digital systems,” in Handbook of Medical Imaging
Physics and Psychophysics, edited by J Beutel, H L Kundel and R L Van Metter (Bellingham,
WA: SPIE) vol. 1, pp 161–222.
Dobbins JT 3rd, Samei E, Chotas HG, Warp RJ, Baydush AH, Floyd CE Jr & Ravin CE, (2003),“Chest
radiography: optimization of x-ray spectrum for cesium iodide-amorphous silicon flat-panel
detector,” Radiology, vol. 226, no. 1, pp. 221-30.
Dunn SM and Kantor ML(1993),“Digital radiology. Facts and fictions,” JADA 124(12): 38-47
Efstathopoulos E P, Costaridou L and Panayiotakis G, (2001),“A Protocol-based Evaluation of Medical
Image Digitizers,” The British Journal of Radiology, 74, 841-6.
70
REFERENCES
Elbakri I A, McIntosh B J and Rickey D W, (2009),“Physical characterization and performance
comparison of active- and passive-pixel CMOS detectors for mammography,” Phys. Med.
Biol. 54(6) 1743-55.
Elias P, (1953),“Optics and communication theory,” J. Opt. Soc. Amer. 43 229-32.
Evans A L, 1981,“The evaluation of medical images,” edited by Adam Hilger, (Bristol), pp 45-46.
Farrell J, Xiao F, Kavusi S, (2006),“Resolution and light sensitivity tradeoff with pixel size,” Proc.
SPIE, 6069, (2006) Digital Photography II, 60690N.
Fellgett P B and Linfoot E H (1955),“On the assessment of optical images,” Phil. Trans. Roy. Soc.
A247 369- 407.
Fellgett P B, (1953),“Concerning photographic grain, signal to noise ratio and information,” J. Opt.
Soc. Amer. 43 271-81.
Fossum E R, (1997),“CMOS image sensors: electronic camera-on-a-chip,” IEEE Trans. E lectron
Devices 44 1689-98.
Gambaccini M, Taibi A, Baldelli P, Del Guerra A, Tuffanelli A, (1998),“Image noise properties of a
phosphor-coated CCD X-ray detector,” Nucl. Instr. Meth. Phys. Res. A 409, 508-510.
Gambaccini M, Taibi A, Del Guerra A, Marziani M and Tuffanelli A, (1996),“MTF evaluation of a
phosphorcoated CCD for x-ray imaging,” Phys. Med. Biol. 41, 2799-2806
Giakoumakis G E, Nomicos C D, Yiakoumakis E N and Evangelou E K, (1990),“Absolute efficiency
of rare earth oxysulphide screens in reflection mode observation,” Phys. Med. Biol. 35 101723.
Ginzburg A and Dick C E (1993),“Image information transfer properties of X-ray intensifying screens
in the energy range from 17 to 320 keV,” Med. Phys. 20, 1013-21.
Greening J R, (1985),“Fundamentals of Radiation Dosimetry,” London, U.K.: Institute of Physics, pp.
56-64.
Gregg E C, (1967),“Modulation transfer function, information capacity and performance criteria of
scintiscans,” J.Nucl. Med. 9, 116-127).
Gruner S M, Tate M W and Eikenberry E F, (2002),“Charge-coupled device area X-ray detectors,”
Rev. Sci. Instur. 73, 2816-42.
Gurwich A. M, (1995),“Luminescent screens for mammography,” Radiat. Mes. 24, 325-25.
71
REFERENCES
Harris E. J, Royle G. J, Speller R. D, Griffiths J. A, Kidane G and. Hanby A.M, (2003),“Evaluation of
a novel low light level (L3Vision) CCD technology for application to diffraction enhanced
breast imaging,” Nucl. Instrum. Methods Phys. Res. A 513, 27-31.
Hartley R V L, (1928),“Transmission of information,” Bell. Syst. Tech. J. 7 535-63.
Hendee W R, (1970),“Medical radiation physics,” (Year Book Medical Publishers, Chicago) 145-48.
Illers H, Buhr E, and. Hoeschen C, (2005),“Measurement of the detective quantum efficiency (DQE) of
digital x-ray detectors according to the novel standard IEC 62220-1,” Radiat. Prot. Dosim.
114(1-3), 39-44.
Johns H E and Cunningham J R, (1983),“The physics of radiology,” (Springfield IL: Thomas) pp. 557–
669.
Jones R C, (1961),“Information capacity of radiation detectors I,” J. Opt. Soc. Amer. 50, 1166-1170.
Kalender W, (1981),“Monte Carlo calculations of X-ray scatter data for diagnostic radiology,” Phys.
Med. Biol. 26, 835-49.
Kanamori H and. Matsuoto M, (1984),“The information spectrum as a measure of radiographic image
quality and system performance,” Phys. Med. Biol. 29, 303-313.
Kanamori H, (1968),“Information capacity of radiographic films,” Jpn. J. App. Phys. 7, 414-421.
Kandarakis I, Cavouras D, Panayiotakis G S and Nomicos C D, (1997),“Evaluating x-ray detectors for
radiographic applications: A comparison of ZnSCdS:Ag with Gd2O2S:Tb and Y2O2S:Tb
screens,” Phys. Med. Biol. 42 1351-73.
Kandarakis I, Cavouras D, Panayiotakis G S, Agelis T, Nomicos C D and Giakoumakis G S,
(1996),“X-ray induced luminescence and spatial resolution of La 2O2S:Tb phosphor screens,”
Phys. Med. Biol. 41 297–307.
Kandarakis I, Cavouras D, Panayiotakis G S, Triantis D, Nomicos C D (1997),“An experimental
method for the determination of spatial-frequency-dependent detective quantum efficiency
(DQE) of scintillators used in X-ray imaging detectors,” Nucl. Instr. Meth. Phys. Res. A 399.
335-42.
Kandarakis I, Cavouras D, Panayiotakis G.S, Nomicos C.D, (2001c),“Experimental investigation of the
optical signal, gain, signal to noise ratio and information content of x-ray phosphor screens,”
Appl. Phys. B. 72, 887-883.
72
REFERENCES
Kim H K, Ahn J K, and. Cho G, (2005),“Development of a lens-coupled CMOS detector for an X-ray
inspection system,” Nucl. Instr. Meth. A 545, 210-216.
Kim H K, Cho G, Lee S W, Shin Y H, and. Cho H S, (2001),“Development and evaluation of a digital
radiographic system based on CMOS image sensor,” IEEE Trans. Nucl. Sci. 48(3), 662-666.
Knoll G F, (1989),“Radiation Detection and Measurement,” 2nd ed. (New York: Wiley) pp.237.
Koscis O, Costaridou L, Mandelos G, Lymberopoulos D and Panayiotakis G (2003), “Compression
assessment based on medical image quality concepts using computer generated test images”,
Computer Methods and Programs in Biomedicine 71, 105-15.
Lee J K, Muenchausen R E, Lee J S, Jia Q X, Nastasi M, Valdez J A, Bennett B L and Cooke D W,
(2006),”Structure and optical properties of Lu2SiO5:Ce phosphor thin films,” Appl. Phys. Lett.
89 101905.
Lempicki A, Brecher C, Szupryczynski P, Lingertat H, Nagarkar V, Tipnis S V,. Miller S. R,
(2002),“A new lutetia-based ceramic scintillator for X-ray imaging,” Nucl. Instr. and Meth. A.
488, 579-590.
Liang Z, Du X, Liu J, Yao X, Yang Y and Li K, (2008),“Comparison of Diagnostic Accuracy of Breast
Masses Using Digitized Images Versus Screen-Film Mammography,” Acta Radiologica 49(6),
618-22
Liaparinos P, Kandarakis I, Cavouras D, Delis H, and Panayiotakis G S, (2007),”Monte Carlo study on
the imaging performance of powder Lu2SiO5:Ce phosphor screens under x-ray excitation:
Comparison with Gd2O2S:Tb screens” Med. Phys. 34 1724-33.
Linfoot E H, (1959),“information theory and photographic images,” J. Photogr. Sci. 7 148.
Lo Y (1998) “Solid-state image sensor: technologies and applications,” Proc. S PIE 3422 70–80.
Mahesh M, (2004),“AAPM/RSNA Physics Tutorial for Residents Digital Mammography: An
Overview,” RadioGraphics 24 1747-60.
Maimon S R, (1971),“The role of technology in forecasting and assessment in planning and resource
allocation,” Abstract of the 3rd SPSE Symposium on Unconventional Photographic Systems
(Washington D.C.) pp.107.
Mainprize J G, Ford N L, Yin S, Gordon E E, Hamilton W J, Tumer T O, Yaffe M J, (2002),“A
CdZnTe slotscanned detector for digital mammography,” Med. Phys. 29(12), 2767-2781.
73
REFERENCES
Maiorchuk M.A,. Nikitin V.V and. Samoilov V.D, (1974),“Possible recording of Fourier holograms
with an information capacity of the order of 107 bit/mm2,” Sov. J. Quantum Electron. 4(2),
168-171.
Marshall N W (2006a) “A comparison between objective and subjective image quality measurements
for a full field digital mammography system,” Phys. Med. Biol. 51 2441-63.
Marshall, NW, Faulkner, K, Busch, HP, Marsh, DM & Pfenning, H (1994), “A comparison of radiation
dose in examination of the abdomen using different radiological imaging techniques”, British
Journal of Radiology, vol. 67, no. 797, pp. 478-84.
Medical Electrical Equipment (2003) “Characteristics of Digital X-ray Imaging Devices Part 1:
Determination of the Detective Quantum Efficiency,” IEC, International Electrotechnical
Commission, Geneva, Switzerland, IEC 62220-1.
Medical
Electrical
Equipment-Characteristics
of
Digital
X-Ray Imaging
Devices,
(2005)
“Determination of the Detective Quantum Efficiency-Mammography Detectors” IEC,
International Electrotechnical Commission, Geneva, Switzerland, , IEC 62220-1-2.
Michail C M, Fountos G P, David S L, Valais I G, Toutountzis A E, Kalyvas N E, Kandarakis I S and
Panayiotakis G S, (2009) “A comparative investigation of Lu2SiO5:Ce and Gd2O2S:Eu
powder scintillators for use in x-ray mammography detectors,” Meas. Sci. Technol. 20,
104008.
Michail C, David S, Liaparinos P, Valais I, Nikolopoulos D, Kalivas N, Toutountzis A, Cavouras D,
Kandarakis I and Panayiotakis G S (2007), “Evaluation of the imaging performance of LSO
powder scintillator for use in x-ray mammography” Nucl. Instrum. Methods Phys. Res. A.
580, 558-61.
Michail C, Fountos G, David S, Valais I, Toutountzis A, Kalyvas N, Kandarakis I, Panayiotakis G
(2009) A comparative investigation of Lu2SiO5:Ce and Gd2O2S:Eu powder scintillators for
use in x-ray mammography detectors” Meas. Sci. Technol. 20, 104008.
Michail C, Toutountzis A, David S, Kalivas N, Valais I, Kandarakis I, Panayiotakis G S (2009)
“Imaging performance and light emission efficiency of Lu2SiO5:Ce (LSO:Ce) powder
scintillator under X-ray mammographic conditions”, Appl. Phys. B 95, 131-39.
74
REFERENCES
Michail C, Valais I, Toutountzis A, Kalyvas N, Fountos G, David S, Kandarakis I, and Panayiotakis G
(2008), “Light Emission Efficiency of Gd2O2S:Eu (GOS:Eu) Powder Screens Under X-Ray
Mammography Conditions”, IEEE Trans. Nucl. Sci. 55(6), 3703-09.
Michail C. M, (2010) “Investigation of optical and imaging characteristics of fluorescent screens for
use in digital imaging detectors suitable for telemedicine,” PH.D. thesis, University of Patras,.
Michail C. M,. Fountos G. P,. Liaparinos P. F,. Kalyvas N. E, Valais I., and. Kandarakis I. S,.
Panayiotakis G. S, (2010) “Light emission efficiency and imaging performance of
Gd2O2S:Eu powder scintillator under x-ray radiography conditions,” Med. Phys. 37, 36943703.
Michail C. M,. Fountos G. P,. Valais I. G,. Kalyvas N. I,. Liaparinos P. F,. Kandarakis I. S, and.
Panayiotakis G. S, (2011) “Evaluation of the Red Emitting Gd2O2S:Eu Powder Scintillator
for Use in Indirect X-Ray Digital Mammography Detectors,” IEEE Trans. Nucl. Sci. 58,
2503-2511.
Michail C. M,. Spyropoulou V. A,. Fountos G. P,. Kalyvas N. I,. Valais I. G, Kandarakis I. S., and.
Panayiotakis G. S, (2011) “Experimental and Theoretical Evaluation of a High Resolution
CMOS Based Detector Under X-Ray Imaging Conditions,” IEEE Trans. Nucl. Sci. 58(1),
314-322.
Morin R L (1988), “Monte Carlo simulation in the radiological science” (Florida: CRC Press) pp.108109.
Nagarkar V. V,. Tipnis S. V,. Miller S. R,. Lempicki A,. Brecher C,. Szupryczynski P, and. Lingertat
H, (2003) “A New X-Ray Scintillator for Digital Radiography,” IEEE Trans. Nucl. Sci. 50,
297-300.
Neitzel U,. Gunther-Kohfahl S,. Borasi G, and. Samei E, (2004) “Determination of the detective
quantum efficiency of a digital x-ray detector: Comparison of three evaluations using a
common image data set,” Med. Phys. 31, 2205-2211.
Neofotistou, V, Tsapaki, V, Kottou, S, Schreiner-Karoussou, A & Vano, E (2005), “Does digital
imaging decrease patient dose? A pilot study and review of the literature”, Radiation
Protection Dosimetry, vol 117, no. 3, pp. 204-10.
75
REFERENCES
Nickoloff, EL, Berdon, WE, Lu, ZF, Ruzal-Shapiro, CB, So, JC & Dutta, AK (2002), “Pediatric high
kv/filtered airway radiographs: comparison of CR and filmscreen systems”, Pediatric
Radiology, vol. 32, no. 7, pp. 476-84.
Nielsen B and Carlson C A (1984), “Energy imparted to fluorescent screens from primary and scattered
radiation. Variations with atomic composition and screen thickness”, Phys. Med. Biol. 29,
315-28.
Nikl M. (2006) “Scintillation detectors for x-rays,” Meas. Sci. Technol. 17, R37-R54.
Nishikawa R. M. and Yaffe M J. (1990), “Model of the spatial-frequency dependent detective quantum
efficiency of phosphor screens”, Med. Phys. 17, 894-904.
Okumura M, Tamatani M, Matsuda N, Takahara T and Fukuta Y (2002) “Ceramic scintillator, method
for producing same, and X-ray detector and X-ray CT imaging equipment using same” United
States Patent patent number 6,384,417.
Panayiotakis G (1984) “Διερεύνηση φυσικογεωμετρικών παραμέτρων για βελτιστοποίηση της
ακτινογραφίας,” Ph.D. thesis, University of Patras.
Patatoukas G,. Gaitanis A,. Kalivas N, Liaparinos P, Nikolopoulos D.,. Konstandinidis A,. Kandarakis
I,. Cavouras D,. Panayiotakis G, (2006) “The effect of energy weighting on SNR under the
influence of non-ideal detectors in mammographic applications,” Nucl. Instr. Meth. Phys. Res.
A 569, 260-263.
Pierce J R (1972) “Communication,” Scient. Amer. 227(3) 31-41.
Raeside D E. (1976), “Monte Carlo principles and applications” Phys. Med. Biol. 21, 181-97.
Remote RadEye Systems Rad-icon Imaging Corporation a division of Dalsa, Sunnyvale, CA [Online].
Available: http://www.rad-icon.com/ products-remote.php
Rodnyi P. A,. Mikhrin S. B,. Mishin A. N and. Sidorenko A. V, (2001) “Small-size pulsed x-ray source
for measurements of scintillator decay time constants,” IEEE Trans. Nucl. Sci. 48, 2340-2343.
Rowlands J. A and. Yorkston J, (2000) “Flat Panel Detectors for Digital Radiography,” in Handbook of
Medical Imaging, Physics and Psycophysics edited by J Beutel, H L Kundel and R L Van
Metter (Bellingham, WA: SPIE) vol. 1 chapter 4 pp.223-313.
Rubinstein R Y (1981) “Simulation and the Monte Carlo method” (New York: Wiley) pp. 64-79.
Sadowsky M (1949) “The preparation of luminescent screens,” Trans. Electrochem. Soc. 95 112-28.
76
REFERENCES
Samei E,. Flynn M. J, and. Reimann D. A, (1998) “A method for measuring the presampled MTF of
digital radiographic systems using an edge test device,” Med. Phys. 25, 102-113.
Saunders R S Jr, Samei E, Jesneck J L and Lo J Y (2005) “Physical characterization of a prototype
seleniumbased full field digital mammography detector,” Med. Phys. 32 588–99.
Schaetzing, R, Whiting, BR, Lubinsky, AR & Owen, JF (1990), “Digital radiography using storage
phosphors,” in Digital imaging in diagnostic radiography, JD Newell & CA Kelsey (Eds),
Churchill Livingston, New York, pp. 107-36.
Seibert JA, (2006),“Acceptance Testing and Quality Control of Photostimulable Storage Phosphor
Imaging Systems,” AAPM Report 93.
Shannon C E and Weaver (1949) “W The mathematical Theory of communication,” (Urbana IL:
Univesrity of Illinois Press) pp.1-28.
Shannon C.E, (1948) “A mathematical theory of communication,” Bell. Syst. Tech. J. 27, 379-423.
Shannon C.E, (1998) “Communication in the Presence of Noise,” in Proc. IEEE, 447-457.
Shaw R, (1978) “Evaluating the efficiency of imaging processes,” Rep. Prog. Phys. 41, 1103-1155.
Shaw R, (1979) “Some modem aspects of image evaluation,” in The physics of medical imaging:
recording system, measurements and techniques, edited by A.G. Hags, (New York: American
Association of Physicists in Medicine), pp 515-523.
Shigeo S William M (1998) “Phosphor handbook,” (CRC Press) pp. 576.
Spahn M (2005) “Flat panel detectors and their clinical applications,” Eur. Radiol. 15 1934-47.
Suzuki J, Furukawa I, Fujii T, Ono S, Ashihara T, Hata J-I, Ando Y, (2000) “Application of Super
High Definition Images in Telemedicine: System Requirements and Technologies for
Teleradiology and Telepathology”, Multimedia Tools and Applications 12 7–38.
Taibi A,. Del Guerra A,. Gambaccini M,. Marziani M,. Tuffanelli A, (1997) “Evaluation of a digital xray detector based on a phosphor-coated CCD for mammography,” Nucl. Instr. Meth. Phys.
Res. A 392, 210-213.
Van Den Bergh R and Leblans P (2005) “Rare earth activated lutetium oxysulfide phosphor for direct
x-ray detection,” European Patent EP1493798.
Van Eijk C W E (2001), “Inorganic-scintillator development” Nucl. Instrum. Meth. A 460, 1-14.
Van Eijk C W E (2002) “Inorganic scintillators in medical imaging” Phys.Med.Biol. (47) R85-106.
77
REFERENCES
Venema H W (1979), “X-ray absorption, speed and luminescent efficiency of rare earth and other
intensifying screens”, Radiology. 130 765-71.
Vuylsteke, P & Schoeters, E (1994), “Multiscale image contrast amplificiation (MUSICA™),”, SPIE
Image Processing, vol. 2167, pp. 551-60.
Wagner R.F,. Brown D.G and. Paster M.S, (1979) “Application of information theory to the assessment
of computed tomography,” Med. Phys. 6, 83-94.
Weckler G P (1967) “Operation of p-n junction photodetectors in a photon flux integration mode,”
IEEE J. Solid-State Circuits SC-2 65–73.
Weiser, JC (1997), “Digital radiography using storage phosphor technology: how computed
radiography acquires data”, Seminars in Roentgenology, vol. 32, no. 1, pp. 7-11.
Wiatrowska A,.Zych E, ˛epinskiL.K, (2010) “Monoclinic HfO2:Eu X-ray phosphor,” Radiation
Measurements 45, 493-496.
Williams M B, Mangiafico P A, Simoni P U (1999) “Noise power spectra of images from digital
mammography detectors,” Med. Phys. 26(7) 1279-93.
X-ray & Gamma Ray Detector, XR-100T-CdTe, Amptek, Bedford, MA [Online]. Available:
http://www.amptek.com/xr100cdt.html
X-ray & Gamma Ray Detector, XR-100T-CdTe, Amptek, Bedford, MA [Online]. Available:
http://www.amptek.com/xr100cdt.html
Yaffe M J (2000) “Digital Mammography” in Handbook of Medical Imaging vol. 1 Physics and
Psycophysics ed J Beutel, H L Kundel and R L Van Metter (Bellingham, WA: SPIE), Chap. 5.
pp. 329-72.
Yaffe M J and Rowlands J A. (1997) “X-ray detectors for digital radiography,” Phys. Med. Biol. 42, 139.
Zhang Q, Li Y, Steele B, Wu X, Chen W R, Rong J and Lie H (2005) “Comparison of a CMOS-based
and a CCD-based digital imaging system: observer studies,” J. Electron. Imag ing 14 023002.
Zych E, (2006) “Spectroscopy of Eu-activated Lu2O3 X-ray phosphors,” in Frontal semiconductor
research, edited by O.T Chang, (Nova Science, New York) pp. 1-24.
78