1. Višeenergetska Medicinska Radiografija
Teza submitovana Elektrotehničkom Fakultetu Unverziteta u Beogradu u cilju dobijanja
titule Master inženjer
Autor:
Nevena Damnjanović
Supervizori:
Asst. Prof. Dr. Miloš Vujisić,
Prof. Dr. Predrag Marinković,
Asst. Prof. Dr. Vladimir Petrović
Septembar 2016
3. Multi-energy Medical Radiography
A thesis submitted to the Faculty of Electrical Engineering, University of Belgrade in
partial fulfillment of the requirements for the master’s degree in Electrical Engineering
Author:
Nevena Damnjanović
Supervisors:
Asst. Prof. Dr. Miloš Vujisić,
Prof. Dr. Predrag Marinković,
Asst. Prof. Dr. Vladimir Petrović
September 2016
7. List of Figures
2.1 Photoelectric interaction of an incoming X-ray photon with atom [8]. . . . 11
2.2 Inelastic (Compton) scattering of an incoming X-ray photon with atom [8]. 13
2.3 The relative contribution of the photoelectric effect and of Compton scatter
to attenuation of X-rays in bone and muscle [8]. . . . . . . . . . . . . . . . 14
2.4 X-ray tube with a rotating anode and a heated filament [10]. . . . . . . . . 15
2.5 Bremsstrahlung and characteristic radiation spectra are shown for a Tung-
sten anode with X-ray tube operation at 80, 100, 120, and 140 kV p and
equal tube current. Adapted from [14]. . . . . . . . . . . . . . . . . . . . . 16
2.6 Attenuation µb in an organ thickness d inside of tissue with attenuation
coefficient of µa and thickness D. Adapted from [12]. . . . . . . . . . . . . 17
2.7 Direct readout (left) and an indirect readout in a detector system (right) [19]. 18
2.8 Flat Panel Detector showing the recording process [20]. . . . . . . . . . . . 18
2.9 Schematic of a DR system. . . . . . . . . . . . . . . . . . . . . . . . . . . . 19
2.10 Parallel (left) and focusing (right) anti-scattering grid. Red beam is noted
as primary radiation, and blue beam as scattered radiation. . . . . . . . . . 20
2.11 High-energy image (left), low-energy image (center), and soft-tissue image
(right) produced by subtraction with an appropriate ratio [23]. . . . . . . . 22
2.12 Schematic drawing of the photon energy spectra of 80 kV (—) and the
140 kV (- - -). The Sn filter increases energy separation by minimizing
the overlap of high and low kVp spectra and reduces dose by blocking low
energy photons from the high energy X-ray tube spectrum [25]. . . . . . . 22
3.1 Specimens prepared for X-ray radiation. Fat (upper left), Meat (upper
right), Bone (lower left) and water (lower right). . . . . . . . . . . . . . . . 26
3.2 The experimental setup on a DR system "Vision C". . . . . . . . . . . . . 26
3.3 Visaris Imaging Console. . . . . . . . . . . . . . . . . . . . . . . . . . . . . 27
3.4 kV-mAs for different filtering values. . . . . . . . . . . . . . . . . . . . . . 28
3.5 bolja slika . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 28
3.6 Graphical presentation of µ of different tissues at different kV . . . . . . . . 30
3.7 60 kV (left) and 120 kV (right) image of 4 samples without additional fil-
tering. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 31
3.8 60 kV (left) and 120 kV (right) image of 4 samples with light filtering. . . 31
3.9 60 kV (left) and 120 kV (right) image of 4 samples with hard filtering. . . . 32
3.10 Dual-energy subtraction of meat for α = 0.75 for an image without addi-
tional filtering (left), α = 0.67 for light filtering (middle) and α = 0.68 for
hard filtering (right). . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 32
5
8. 6 LIST OF FIGURES
3.11 Dual-energy subtraction of bone for α = 0.58 for an image without addi-
tional filtering (left), α = 0.46 for light filtering (middle) and α = 0.46 for
hard filtering (right). . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 33
9. List of Tables
2.1 Different K-edge values for different elements. . . . . . . . . . . . . . . . . 12
3.1 The thickness of the samples examined. . . . . . . . . . . . . . . . . . . . . 25
3.2 kV and mA s values for keeping the dose the same. . . . . . . . . . . . . . 27
3.3 Table of linear attenuation coefficient at different kV without using addi-
tional filtering. Values of µ are expressed in [cm−1
]. . . . . . . . . . . . . . 29
3.4 Table of linear attenuation coefficient at different kV with using 1 mm Al
with 0.1 mm of Cu additional filtering. Values of µ are expressed in [cm−1
]. 29
3.5 Table of linear attenuation coefficient at different kV using 2 mm Al with
0.3 mm of Cu additional filtering. Values of µ are expressed in [cm−1
]. . . . 29
3.6 Value of α for different tissue comination and for display of either soft or
hard image at 60-120 kV. For a certain α the tissue written in a table is
subtracted. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 30
3.7 Value of α for different tissue comination and for display of either soft or
hard image at 60-120 kV. For a certain α the tissue written in a table is
subtracted. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 33
7
11. Chapter 1
Introduction
1.1 Evolution to Digital Radiography
X-rays were discovered in November 1895 by Wilhelm Conrad Roentgen (1845-1923), a
Professor at Wurzburg University in Germany [1].
Roentgen observed a fluorescent glow from a barium platinocyanide-coated (BaPt(CN)4)
screen on a table near a cathode-ray tube in his laboratory [1].
He spent the next 6 weeks establishing the cause and applications of the observed
phenomena and gave his preliminary report to the Wurzburg Physical-medical society,
accompanied by experimental radiographs and images [2].
Medicinal application of x-rays began almost momentarily, as early radiologists used x-
rays to localise foreign bodies which helped surgeons remove them safely, without causing
excessive damage to the tissue [3].
What eluded the early pioneers was the harmful effective of radiation which led to
many deaths of practitioners and patients alike.
In 1913, William D. Coolidge (1873–1975) invented the Coolidge tube, which had a
cathode filament made of Tungsten. This was an improvement on the Crookes tube, used
by first radiologists [4].
The same year also saw the discovery of the anti-scatter grid by Gustave Bucky, which
helped reduce harmful radiation doses [4].
Next significant breakthrough in medicinal application of x-rays came in 1918 when
George Eastman, founder of Eastman-Kodak company, introduced roll film which replaced
glass photographic plates which were used until then [1].
X-rays were a field of significant scientific research in the ninety-twenties, resulting
in two Nobel prizes in the area – Karl Manne Siegbahn (1886-1978) for discoveries in
and research in the field of X-ray spectroscopy and Arthur Compton (1892-1962) for the
discovery of inelastic scattering of a photon by a charged particle [1].
X-ray image intensifier was originally introduced in 1948 and it enabled conversion of
x-rays into visible light at higher intensity than fluorescent screens do.
This led to higher image quality, as the structure of the object being imaged was easier
to see, and additionally required lowered absorption dose due to more efficient conversion
of X-ray quanta to visible light.
However, It must be noted that basic principles of radiography have not changed
significantly, although it is able to generate higher-quality images through the use of films
with larger variety of grain sizes [5].
Finally, the ability to capture and enhance images digitally have enabled higher-quality
9
12. 10 CHAPTER 1. INTRODUCTION
images that can be transferred across devices and stored indefinitely, without image qual-
ity deterioration [5].
1.2 The need for Dual-Energy subtraction radiography
In modern diagnostics Chest radiography represents one of the key tools for detecting
thoracic abnormalities. Detecting lung cancer in its early stages is an issue of profound
significance, but during a standard chest exam, small pulmonary nodules may be very
difficult to detect for a radiologist [6]. The reason lies in the overlapping of the structures
such as ribs and claviles.
Dual-energy has a higher sensitivity in detection of calcification within a pulmonary
nodule, bone abnormalities, recognitions of hilar and mediastinal masses, recognition of
tracheal narrowing, airway diseases and localization of indwelling devices [7].
1.3 Thesis outline
Chapter 1 starts with a short explanation laying out basic reasons behind the need for
Dual-energy imaging and with a further reflection on Digital Radiography evolution and
history.
Chapter 2 represent the theoretical background needed for a reader to get familiar
with the subject matter.
In the first section, the author emphasizes the importance of the effects that are
occurring during the interaction of X-ray with matter.
The second section explains the entire experimental setup introducing every part of
the X-ray chain that is making a difference to the end image. This chapter ends with the
mathematical explanation for obtaining Dual-energy subtracted image.
Chapter 3 presents the experiment and the results. Different specimens are compared
in order to verify the possibility of separation of tissues with different attenuation coeffi-
cients.
In Chapter 4 the results are discussed and conclusions are made, along with the sug-
gestions for further research.
13. Chapter 2
Theoretical Background
2.1 Interaction of X-ray with matter
Observing the X-ray energies applied in diagnostic imaging, one can distinguish two im-
portant types of interaction: photoelectric effect and Compton scattering.
2.1.1 Photoelectric effect
Incident X-ray photons interact with the matter by delivering its energy to an electron
which is ejected from its shell. As this photo-electron leaves a vacancy behind, one of
the electrons from the outer shells fills in the vacancy with an emission of a new X-ray
photon. This X-ray photon is emitted in a random direction. It may cause secondary
radiation, but it is mostly absorbed by new interactions.
A newly released photo-electron that was ejected at high speed further causes a large
number of secondary ionization. A graphical display of the photoelectric effect is shown
in Figure 2.1.
Figure 2.1: Photoelectric interaction of an incoming X-ray photon with atom [8].
The photoelectric effect is dominant when the incident X-ray photon energy is close
to the binding energy of an inner shell. Photon energy that is sufficient to release a
photo-electron from the K-shell is called "K-edge".
The K-edge is the minimum energy required for the photoelectric event to occur with
a K-shell electron. Below this level, no photoelectric events can occur with K electrons.
11
14. 12 CHAPTER 2. THEORETICAL BACKGROUND
K-shell electrons are important because the energy to dislodge them (produce photo-
electric absorption) is right in the middle of the diagnostic range.
Higher shells can also participate in the photoelectric effect, but they require much
lower energies that are typically below the diagnostic range. Those energies are filtered
out and/or absorbed by the body.
Photoelectric effect is proportional to:
PE =
Density ∗ Z3
E3
, (2.1)
where PE is photoelectric effect, Z is the atomic number and E is the X-ray photon
energy. This means that this effect strongly depends on the energy and a material that
it passes through.
Bone contains calcium, which has a somewhat higher Z and a K-edge at 4 keV (Table
2.1). While this K-edge is still of quite low value, the "tail" does reach into diagnostic
energies resulting in better absorption of bone than soft tissues at lower energies up to
some 50 keV [9].
At higher energies, the possibility for interaction of X-ray energy with the tissue drops
rapidly (E3
).
In Table 2.1, one can see the values of K-edge energies for different elements.
Element K-edge [keV]
Carbon 0.3
Nitrogen 0.4
Oxygen 0.5
Calcium 4.0
Lead 88.1
Table 2.1: Different K-edge values for different elements.
2.1.2 Compton effect
The inelastic (Compton) scatter (Figure 2.2) results from interaction of X-ray photons
with electrons in the outer shell. They are ejected leaving the atom ionized and the
incident X-ray photon with reduced energy and changed direction of motion.
At the end of its path, the incident X-ray photon becomes absorbed in the tissue.
The Compton scattering accounts for most of the scatter in diagnostic radiology which is
considered unwanted. It is proportional to:
Compton =
Density
Energy
(2.2)
Absorption due to Compton scattering is nearly independent of the X-ray energy.
Compton scattering is relatively constant for different energies although it slowly decreases
at higher energies. It does depend on the material (specifically, the electron density) but
still much less than the photoelectric absorption in diagnostic x-ray imaging.
As it creates scatter photons, they may hit the image detector at different angles and
contribute to background/noise. As a result, at higher energies, contrast between the
bone relative to the soft tissue decreases as photoelectric effect disappears [8].
15. 2.1. INTERACTION OF X-RAY WITH MATTER 13
Figure 2.2: Inelastic (Compton) scattering of an incoming X-ray photon with atom [8].
Each time a photon ejects an electron (i.e. ionizes an atom) this creates free radicals
that can damage DNA. The energy that a tissue absorbs from photon interactions is
referred to as dose.
2.1.3 Linear Attenuation Coefficient
The differential ability of various tissues to scatter and absorb X-ray photons, no matter
by which mechanisms, is given by their linear attenuation coefficient (µ expressed in
cm−1
).
It expresses the fractional reduction in beam intensity along a linear beam path after
passage through one centimeter of the tissue [8].
The linear attenuation coefficient of a given tissue varies with the X-ray photon energy.
It is high for lower energies where the photoelectric effect prevails and drops exponentially
for higher energies where Compton scatter dominates, and hence the mass density rather
than the atomic composition of the tissue becomes the prime determinant of attenuation
(Figure 2.3).
The number of photons that passes through the tissue is given by Beer-Lambert Law:
I = I0e−µd
, (2.3)
where I is the intensity of photons transmitted across some distance d, I0 the initial inten-
sity of photons, µ linear attenuation coefficient and d distance traveled (tissue thickness).
Thus, the main contribution to differences in X-ray attenuation in diagnostic imaging
is photoelectric absorption. One can see in the Figure 2.3 that these differences are much
more apparent at low energies.
Lowering kV improves image contrast since the difference in attenuation coefficient
between bone and soft tissue is larger. This better contrast comes with a price of increased
dose as at lower energies [10].
Both the photoelectric effect and inelastic scatter result in a loss of electrons from
atoms. This may cause the breakage of chemical bonds, and because the ionized atoms
(notably those of C, N and O) are chemically highly reactive, new chemical bonds are
established that are alien to the tissue.
16. 14 CHAPTER 2. THEORETICAL BACKGROUND
Figure 2.3: The relative contribution of the photoelectric effect and of Compton scatter
to attenuation of X-rays in bone and muscle [8].
X-rays’ ability to cause ionization includes them in the family of ionizing radiation,
and it is these ionization and their derived chemical reactions that cause the biological
damage related with such radiation [8].
2.2 X-ray system setup
2.2.1 X-ray tube
X-rays are emitted when a high speed electron hits a metal target. In medical X-ray
tubes, the target is usually Tungsten, alloy of Rhenium and Tungsten (Rh0.5 %W95 %) or
molybdenum (used in mammography) [11]. Only 1 % of the energy in the electron beam
is converted to photons, while the rest is converted to heat.
Following a graphical presentation in Figure 2.4, a cathode is heated by applying a
small voltage to cause a release of electrons by thermionic emission. These are accelerated
by a high voltage towards the anode (target) which is positively charge relative to the
cathode.
This whole process needs to be conducted in a high vacuum (10−7
bar [12]) otherwise
the electrons will collide with molecules of air further loosing energy before they reach the
anode. When the electrons strike the anode, they release energy. Some of this energy is
released as X-ray, but most of it is converted to heat. To prevent the anode from melting,
it is continuously rotating so the hot spot moves away from the electron beam [13].
The maximum energy of the produced X-ray photon is limited by the energy of the
incident electron, which is equal to the voltage applied on the tube times the electron
charge. As an example, 100 kV applied on a tube can not create X-rays with a greater
17. 2.2. X-RAY SYSTEM SETUP 15
Figure 2.4: X-ray tube with a rotating anode and a heated filament [10].
energy than 100 keV.
Two different processes lead to an X-ray being released from the anode:
• X-ray fluorescence
• Bremsstrahlung
X-ray fluorescence
Considering the electron has enough energy, it can eject an electron out of the inner
electron shell of a metal target. Following this process, an electron from a higher energy
level fills up the vacancy with an emission of X-ray.
The emission spectrum of X-rays has few discrete frequencies (i.e. spectral lines) as
shown in Figure 2.5 for a Tungsten anode. These spectral lines are usually referred as
characteristic lines as they represent a characteristic of an element.
Bremsstrahlung radiation
"Bremsstrahlung" in German means "breaking radiation" which describes the radiation
emitted when electrons are decelerated by scattering of the metal nuclei.
It is characterized by a continuous distribution of radiation which becomes more in-
tense and shifts toward higher frequencies when the energy of the bombarding electrons
is increased (as shown in the Figure 2.5).
The resulting output of a tube consists of a continuous Bremsstrahlung spectrum
falling off to zero at the tube voltage, plus several spikes at the characteristic lines.
The voltages used in diagnostic X-ray tubes range from 20 kV to 150 kV and thus the
highest energies of the X-ray photons range from roughly 20 keV to 150 keV [15].
Note that the average energy of the beam is much less than the peak energy. A "rule
of thumb" is that it would equal 1
3
of the maximum energy [16].
18. 16 CHAPTER 2. THEORETICAL BACKGROUND
Figure 2.5: Bremsstrahlung and characteristic radiation spectra are shown for a Tungsten
anode with X-ray tube operation at 80, 100, 120, and 140 kV p and equal tube current.
Adapted from [14].
2.2.2 Contrast of radiographic image
The possibility of distinguishing different structures inside the radiographic image is lim-
ited in few ways.
To be able to distinguish between two adjacent structures, their transmission proper-
ties of X-rays must be different. As an example, in a chest exam, one can easily distinguish
lungs and ribs.
In order to quantify, one introduces a contrast of image CΦ.
Let Φ0 be the entering flux density of X-ray beam, as shown in Figure 2.6. This flux
passes through the thickness D and is being attenuated by µa. Another beam with the
same entering flux density is going though the tissue of linear attenuation coefficient µb
and thickness d, as shown in Figure 2.6.
Flux density of X-photons exiting the tissue with thickness D and linear attenuation
coefficient µa is:
Φ1(E) = Φ0(E)e−µa(E)D
, (2.4)
The exiting flux density of another X-ray beam with µa and µb linear attenuation
coefficient is:
Φ2(E) = Φ0(E)e−µa(E)(D−d)
e−µb(E)d
. (2.5)
Further, the contrast is defined (for Φ1 > Φ2) as:
CΦ(E) =
Φ1(E) − Φ2(E)
Φ1(E)
= 1 − e−[µb(E)−µa(E)]d
(2.6)
If (µa − µb)d << 1, then the exponent can be approximated as e−x
≈ 1 − x and the
contrast equation looks like:
CΦ(E) = [µb(E) − µa(E)]d. (2.7)
19. 2.2. X-RAY SYSTEM SETUP 17
As seen from the equations 2.6 and 2.7 the contrast of an object depends on the
difference of linear attenuation coefficients and on the thickness of an organ/specimen.
If the difference in attenuation is small the contrast will be small no matter the thick-
ness of an organ.
Contrast also depends on the energy because linear attenuation coefficient is energy
dependent. That is why usually the value of contrast is taken as a mean energy spectrum
of X-ray that leave the X-ray tube.
Figure 2.6: Attenuation µb in an organ thickness d inside of tissue with attenuation
coefficient of µa and thickness D. Adapted from [12].
2.2.3 Flat Panel Detector
DR systems can be distinguished by a method of conversion of X-ray photons to:
• Direct system
• Indirect system
Direct-conversion systems have detectors that consist of X-ray photo-conductor, such
as amorphous Selenium (a-Se) that converts X-ray photons into an electric charge directly,
without the intermediate stage. a-Se has excellent X-ray properties and extremely high
intrinsic spatial resolution (>20 lp/m at 100 keV [17]).
Before the exposure to the X-ray, an electric field is applied across its layer. After
absorption of X-ray, electrons and holes are released in the a-Se and further drawn directly
to the charge collecting electrodes due to an applied electric field. a-Se is technologically
well developed with its filling factor approaching 100 % [18].
In indirect systems X-ray photons are primarily converted to light using scintillators or
phosphorus, and the visible light is further converted into electrical signal. The scintilators
used in indirect-conversion detectors can be structured or unstructured.
Fluorescent screen, as an example of unstructured scintillator, can spread visible light
to adjacent pixels and thus reduce spatial resolution [19].
This problem is solved using structured scintillator of Cesium Iodide doped with Tal-
ium (CsI(Tl)) crystals. This crystalline structure, which consists of discrete and parallel
"needles", propagate light so most of the signal can be collected on photo-diodes.
Below the CsI layer, an amorphous silicon (a-Si) is used to transfer visible light to
electric signal. These photo-diodes are organized in a form of active matrix (Figure 2.8).
20. 18 CHAPTER 2. THEORETICAL BACKGROUND
Figure 2.7: Direct readout (left) and an indirect readout in a detector system (right) [19].
Figure 2.8: Flat Panel Detector showing the recording process [20].
2.2.4 Acquisition console
The experimental setup is organized as a standard DR system shown in Figure 2.9. The
generator supplies a voltage to the tube enabling the generation of X-ray photons.
The exiting X-rays are further collimated on a wanted area. The patient/specimen is
situated closer to the acquisitor of X-rays, the detector, to avoid magnification.
The X-ray bucky grid is used to collect scattered radiation and preserve image quality.
After the detector has collected the signal, the image is displayed on an acquisition console
as a digital image.
21. 2.2. X-RAY SYSTEM SETUP 19
Figure 2.9: Schematic of a DR system.
2.2.5 Influence of kV
We have already discussed the fact that high-energy photons pass through tissue with less
attenuation. While these photons create more damage when they do hit electrons in the
tissue, it should be said that this is a relatively small effect.
Thus, overall changing an X-ray beam to a higher energy will deposit less dose to the
patient. In addition, because more photons are passing through, there are more photons
to make the image resulting in less noise.
Along with increasing the kV, one can decrease the exposure time, resulting in image
with less dose to the patient [9].
When a patient is thicker, it requires more photons to create equally good image as for
a standard patient size. But, more photons means higher probability for it to hit tissue.
Thicker patient has more tissue for scattered photons to interact multiple times. Thus
higher kV comes at the price of losing contrast [9].
2.2.6 Influence of Scattered radiation
Reducing the number of scattered photons can be done in two ways:
• Using collimator to narrow the X-ray coming out of the tube
• Using anti-scattering grid
Exposing only the area of interest is the simplest way to lower the number of scattered
photons. It is done by putting a collimator at the exit of the tube.
The need for reducing the scatter radiation comes from the fact that the scattered
photons only attributes to the noise and lowers image quality. The best way to retain
image quality is to use anti-scattering grids.
Anti-scattering grid is a grid of lead strips organized in such a way to absorb most
of the radiation that is scattered. This means it absorbs radiation that is incoming at
an angle greater than : atan(1/GR), where GR is a ratio of the grid1
. The Grid ratio is
1
It is usually manufactured in a form of 8:1, 10:1 or 15:1
22. 20 CHAPTER 2. THEORETICAL BACKGROUND
defined as:
GR = h/b, (2.8)
where h is height of the lead strips and b is the spacing between the strips. The grids
can have parallel or focused lead strips as shown in Figure 2.10.
Figure 2.10: Parallel (left) and focusing (right) anti-scattering grid. Red beam is noted
as primary radiation, and blue beam as scattered radiation.
2.3 Basic Principles of Dual-energy Imaging
Dual energy radiography implies taking two radiographic images on two different mean
beam X-ray energies. The two images are usually referred as "High energy" and "Low
energy" radiographs because of the large difference in X-ray energies.
The result image may emphasize different tissues, distinguishing soft tissue from the
bone.
This diagnosis is possible because different tissues attenuate high and low energy pho-
tons to a different degree. Bone has a higher attenuation coefficient at low beam energies
because it has calcium (higher K-edge energy value, explained in Section 2.1.2). The
separation of bone and tissue is possible because of the different attenuation coefficients.
There are two different ways of acquiring dual-energy images. Single exposure systems
have a single radiograph that is acquired on a CR system using two storage-phosphor
imaging plates with a 1 mm thick copper filter interleaved between the two [21].
The front plate obtains a low kV image, while the back plates obtains the higher kV
image as the copper plate and the front plate have selected out the lower energy photons.
Dual-energy image obtained by this technique has a problem with lower Signal to
Noise Ratio (SNR) [21, 22].
In dual exposure system, two sequential radiographic images are obtained at two
energies [22]. Using Digital Radiography system (DR) with Flat Panel Detector (FPD)
has a higher SNR with its greatest disadvantage being time delay between two exposures
(around 200 ms) [22]. This delay can produce offsets in the alignment of body structures
caused by respiratory, bowel or patient motion.
In this thesis, dual exposure technique was used to obtain the dual-energy images.
2.3.1 Dual-energy physical background
A simple mathematical model assumes that mono-energetic radiation is used and no
scattered radiation is detected, so that the transmitted radiation intensity through a
23. 2.3. BASIC PRINCIPLES OF DUAL-ENERGY IMAGING 21
region of bone and tissue acquired at a low X-ray energy is given by:
ILE
= ILE
0 e−(µLE
st dst+µLE
b db)
, (2.9)
similarly, transmitted radiation intensity at high X-ray energy is given by:
IHE
= IHE
0 e−(µHE
st dst+µHE
b db)
, (2.10)
where µst is linear attenuation coefficient of soft tissue, dst thickness of the soft tissue,
µb linear coefficient of bone and db bone thickness. The superscript LE is an abbreviation
for Low Energy, while HE is abbreviated for High Energy.
Logarithm these equations will give:
ln ILE
= ln ILE
0 + (−µLE
st dst − µLE
b db), (2.11)
ln IHE
= ln IHE
0 + (−µHE
st dst − µHE
b db). (2.12)
For simplicity, the equation is presented as:
SLE
= CLE
+ (−µLE
st dst − µLE
b db), (2.13)
SHE
= CHE
+ (−µHE
st dst − µHE
b db). (2.14)
These two equations are further subtracted by a factor α:
SHE
− αSLE
= CHE
− αCLE
+ dst(αµLE
st − µHE
st ) + db(αµLE
b − µHE
b ) (2.15)
The SHE
presents the information on a high energy image, while SLE
presents infor-
mation in a low energy image.
As one can see that when:
α =
µHE
st
µLE
st
, (2.16)
the value in brackets of the equation 2.15 becomes 0. This means when two images are
subtracted by an appropriate factor, the only visible information is bone image. This
applies when a different α factor is used such that:
α =
µHE
b
µLE
b
. (2.17)
In that case the only information left, is the information about the soft tissue.
In Figure 2.11 one can see the example of a dual-energy subtraction such that there
is only soft tissue in the image.
2.3.2 Additional Filtering
As stated previously, the X-ray tubes have polychromatic spectra that consists of continu-
ous Bremsstrahlung spectrum superimposed with characteristic lines of an anode material
(most frequently Tungsten).
When doing a dual-energy imaging, it is not two distinct photon energies being used,
rather a two energy spectra. The mean energy of both of those "dual" energies is signifi-
cantly lower than the energy of kVpeak. As an example, for a 80 and 140 kVp the mean
energies are 56 and 76 keV, respectively [24].
24. 22 CHAPTER 2. THEORETICAL BACKGROUND
Figure 2.11: High-energy image (left), low-energy image (center), and soft-tissue image
(right) produced by subtraction with an appropriate ratio [23].
In Computed Tomography (CT), usually 80-140 kVp is used because they provide
maximum difference and least overlap between the spectra [24]. In Figure 2.12 one can
see two spectra for 80 kVp and 140 kVp. On high energies, one must use a filter to
additionaly separate spectra.
Figure 2.12: Schematic drawing of the photon energy spectra of 80 kV (—) and the 140 kV
(- - -). The Sn filter increases energy separation by minimizing the overlap of high and
low kVp spectra and reduces dose by blocking low energy photons from the high energy
X-ray tube spectrum [25].
25. 2.3. BASIC PRINCIPLES OF DUAL-ENERGY IMAGING 23
In numerous articles different types of filtering was used. For Dual-kVp2
CT, besides
the standard filtration of 0.9 Ti and 3.5 mm Al, at high energies 0.4 mm Sn was used [26].
2
Dual-kVp scans are realized by two consecutive scans performed at different kVp- and with optimized
mA-settings
27. Chapter 3
Experimental results
The aim of the Thesis is to cover the entire process of X-ray dual imaging on a DR system
as well as image processing and characterization. Thus, this chapter is divided into two
sections: experimental setup and dual-energy image processing.
3.1 Experimental setup
All of the imaging has been conducted in a company Visaris d.o.o incorporated in Belgrade,
Serbia.
3.1.1 DR system details
Samples examined are veal fat, water, veal meat and bone.
Water is held in a polypropylene box. It is considered that this thin transparent box
is not significantly contributing to the X-ray attenuation of water.
The samples are positioned adjacent to each other and the collimation is set to 22 x
22 cm as shown in Figure 3.1.
In Table 3.1 the thicknesses of materials used are presented.
Material Thickness [cm]
Veal Fat 0.88
Veal Meat 0.73
Veal Bone 1.52
Water 2.02
Table 3.1: The thickness of the samples examined.
Detector used is a Thales 3543 EZ portable Flat Panel Detector. To eliminate the
scattering radiation, aluminium anti-scattering grid is used, with a ratio of 10:1 by Soyee.
The tube used is Toshiba 7864X and the CPI 50 kW generator. The images are
acquired using Visaris Imaging console shown in Figure 3.3. The device used is Visaris
Vision C DR system shown in Figure 3.2.
Four samples are imaged in a range 60 kV to 140 kV using a step of 10 kV. From
110 kV to 140 kV additional1
filtering is used: 1 mm Al with 0.1 mm Cu and 2 mm Al
with 0.3 mm Cu.
1
Collimator has its inherent filtration, and in this thesis, Claymount automatic collimator is used. It
has 2 mm Al filter.
25
28. 26 CHAPTER 3. EXPERIMENTAL RESULTS
Figure 3.1: Specimens prepared for X-ray radiation. Fat (upper left), Meat (upper right),
Bone (lower left) and water (lower right).
images/visionc.png
Figure 3.2: The experimental setup on a DR system "Vision C".
3.1.2 verodostojna merenja il tako nesto
While changing the kV, in order to keep the same dose acquired at the detector, one needs
to adjust the appropriate mA s.
29. 3.1. EXPERIMENTAL SETUP 27
The matching values are displayed in Table 3.2 when no additional filtering is used,
for filtering of 1 mm Al with 0.1 mm Cu and 2 mm Al with 0.3 mm Cu.
Tube
voltage [kV]
No additional
filters [mA s]
1 mm Al with 0.1 mm
Cu [mA s]
2 mm Al with 0.3 mm
Cu [mA s]
60 16 - -
70 9 - -
80 5.7 - -
90 4.1 - -
100 3.1 - -
110 2.3 4.6 3.6
120 1.8 3.6 2.8
130 1.6 2.8 2.3
140 1.3 2.3 1.8
Table 3.2: kV and mA s values for keeping the dose the same.
Based on the Table 3.2, the plot kV vs. mA s is shown in Figure 3.4.
The current used is 125 mA for all values of kV.
3.1.3 veryfing the dose with Pixel Value Index (PVI) il tako nesto
For each acquisition, the pixel value is measured on the detector (incoming intensity) and
on each of the samples. The same pixel value measured on the detector confirmed that,
for different values of kV, the dose stayed the same.
The intensity of photons that reach detector is I0, while the intensity of photons that
passes through tissue is represented by Beer-Lambert’s law in equation 2.1.3. For each
type of tissue and on each kV one can obtain the linear attenuation coefficient by solving
Figure 3.3: Visaris Imaging Console.
30. 28 CHAPTER 3. EXPERIMENTAL RESULTS
Figure 3.4: kV-mAs for different filtering values.
Figure 3.5: bolja slika
the equation:
PV I(detector)
PV I(tissue)
=
I0
I0eµtissued
, (3.1)
31. 3.1. EXPERIMENTAL SETUP 29
where PV I is the pixel value index that represents the amount of the photons that
one detector pixel has collected. This value can be measured using the Visaris DR system
and software as shown in Figure 3.5. A large homogenic area of interest is marked and
the PVI value is displayed as a mean value of the marked area. This way, it is verified if
the detector has been saturated.
3.1.4 Obtaining the µ
Solving the equation 3.1.3, one can obtain the values of µ for each tissue at every kV
(Table 3.3, 3.5 and 3.6).
Generator condition [kV] µFat µMeat µBone µWater
60 0.25 0.35 0.70 0.20
70 0.24 0.35 0.70 0.20
80 0.25 0.32 0.60 0.18
90 0.21 0.31 0.53 0.16
100 0.20 0.28 0.49 0.16
110 0.21 0.28 0.44 0.15
120 0.19 0.27 0.43 0.15
130 0.18 0.27 0.41 0.14
140 0.17 0.24 0.40 0.14
Table 3.3: Table of linear attenuation coefficient at different kV without using additional
filtering. Values of µ are expressed in [cm−1
].
Tube voltage [kV] µFat µMeat µBone µWater
110 0.17 0.23 0.35 0.13
120 0.15 0.22 0.35 0.13
130 0.16 0.23 0.33 0.12
140 0.15 0.23 0.32 0.12
Table 3.4: Table of linear attenuation coefficient at different kV with using 1 mm Al with
0.1 mm of Cu additional filtering. Values of µ are expressed in [cm−1
].
Tube voltage [kV] µFat µMeat µBone µWater
110 0.17 0.24 0.33 0.13
120 0.16 0.22 0.32 0.12
130 0.16 0.22 0.32 0.12
140 0.15 0.21 0.31 0.12
Table 3.5: Table of linear attenuation coefficient at different kV using 2 mm Al with
0.3 mm of Cu additional filtering. Values of µ are expressed in [cm−1
].
The values from the tables may be expressed in graphical form, as shown in Figure
3.6.
3.1.5 Choosing the dual-energies
Kako smo se odlucili za 60-120 i 70-130
32. 30 CHAPTER 3. EXPERIMENTAL RESULTS
Figure 3.6: Graphical presentation of µ of different tissues at different kV .
3.1.6 α calculation
The value of α can be calculated for different tissues and for display of either soft or hard
tissue:
α without filters 1 mm Al with 0.1 mm of Cu 2 mm Al with 0.3 mm of Cu
Fat 0.74 0.61 0.64
Meat 0.75 0.63 0.63
Bone 0.61 0.46 0.46
Table 3.6: Value of α for different tissue comination and for display of either soft or hard
image at 60-120 kV. For a certain α the tissue written in a table is subtracted.
3.2 Image processing
The images are acquired in a raw format2
. After reading the data, images are auto cropped
to a collimated area.
The Figure 3.7 shows two images at 60 and 120 kV, without additional filtering, Figure
3.8 with light filtration, while Figure 3.9 has hard filtration.
The values of α for meat "disappearance" for all three types of filtering (none, light,
hard) is calculated to be 0.75, 0.63 and 0.63, respectively (Shown in Table 3.7). This
image is called "hard" image as there is only information of hard tissue left in the image
(bone visible in the image).
2
Raw format refers to the image that has not went through any post-processing. It is the image that
was acquired by the detector.
33. 3.2. IMAGE PROCESSING 31
Figure 3.7: 60 kV (left) and 120 kV (right) image of 4 samples without additional filtering.
Figure 3.8: 60 kV (left) and 120 kV (right) image of 4 samples with light filtering.
34. 32 CHAPTER 3. EXPERIMENTAL RESULTS
Figure 3.9: 60 kV (left) and 120 kV (right) image of 4 samples with hard filtering.
Figure 3.10: Dual-energy subtraction of meat for α = 0.75 for an image without additional
filtering (left), α = 0.67 for light filtering (middle) and α = 0.68 for hard filtering (right).
The values of α for bone "disappearance" for all three types of filtering (none, light,
hard) is calculated to be 0.61, 0.46 and 0.46, respectively (Shown in Table 3.7). This
image is called "soft" image as there is soft tissues visible in the image in comparison
with the bone.
Dual-energy subtraction for bone and meat is shown in Figure 3.11 and 3.10 for α
values close to the calculated one.
Slightly different values of α are shown because it is noted by observation that those
values are giving better tissue subtraction.
The dual-energy subtraction for fat has not been shown as the values of α have similar
results as for meat. It can be seen that, for meat subtraction, the fat and water are also
almost entirely subtracted because the values of mu and further α are fairly similar.
This leads to the conclusion that it will not be possible to separate similar tissues at
this kV.
35. 3.2. IMAGE PROCESSING 33
Figure 3.11: Dual-energy subtraction of bone for α = 0.58 for an image without additional
filtering (left), α = 0.46 for light filtering (middle) and α = 0.46 for hard filtering (right).
α without filters 1 mm Al with 0.1 mm of Cu 2 mm Al with 0.3 mm of Cu
Fat 0.74 0.61 0.64
Meat 0.75 0.63 0.63
Bone 0.61 0.46 0.46
Table 3.7: Value of α for different tissue comination and for display of either soft or hard
image at 60-120 kV. For a certain α the tissue written in a table is subtracted.
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43. Index
CT Computed Tomography, 22
DR (Digital Radiography), 20
FPD Flat Panel Detector, 20
SNR (Signal to Noise Ratio), 20
41
