Download X-ray imaging: Fundamentals and planar imaging

X-ray imaging:
Fundamentals and planar imaging
Mikael Jensen1 and Jens E. Wilhjelm2
1Hevesy
Laboratory, DTU Nutech, 2Biomedical Engineering, DTU Elektro
Technical University of Denmark
(Ver. 4 28/8/14) © 2004 - 2014 by M. Jensen and J. E. Wilhjelm
1 Introduction
X-ray imaging is the most widespread and well-known medical imaging technique. It dates back to
the discovery by Wilhelm Conrad Röntgen in 1895 of a new kind of penetrating radiation coming
from an evacuated glass bulb with positive and negative electrodes. Today, this radiation is known as
short wavelength electromagnetic waves being called X-rays in the English speaking countries, but
“Roengten” rays in many other countries. The X-rays are generated in a special vacuum tube: the Xray tube, which will be the subject of the first subsection. The emanating X-rays can be used to cast
shadows on photographic films or radiation sensitive plates for direct evaluation (the technique of planar X-ray imaging) or the rays can be used to form a series of electronically collected projections,
which are later reconstructed to yield a 2D map (thus, a tomographic image). This is the so-called
CAT or CT technique (see the chapter on CT imaging).
1.1 Characterization of X-ray
X-rays are electromagnetic radiation (photons) with wavelengths, 10 pm <  < 10 nm. They travel
with the speed of light, c0  300 000 km/s and has a frequency of  = c0/ [Hz]. The energy of the
individual photon is E = h [J], where h = 6.62×10-34 Js is Planck’s constant. The energy of an X-ray
is typically measured in electron volts (eV). 1 eV is the energy increase that an electron experiences,
when accelerated over a potential difference of 1 V. Thus, 1 eV = qeV = 1.602×10-19 J, where the
charge of an electron is qe (both qe and V is negative in this context).
Problem 1 Calculate the frequency and energy for monochromatic x-rays with  = 1 nm.
Answer:  = c0/  3×1017 Hz = 300 000 THz. E = h  1.99×10-16 J = 1.24 keV.
Protective lead
Cathode
(heated filament)
I
Glass envelope
Electrons
Anode
+
U
X-rays
Figure 1 X-ray tube showing cathode and anode with electrons accelerated from cathode towards anode. The
tube generates X-rays in all directions, but due to the encapsulation most are lost and only a fraction is used for
imaging.
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Characteristic X-ray emission from Wolfram
Relative intensity at fixed electron current
1200
1 mm Al filter cut-off
1000
V=50kV
V=90kV
V=140kV
800
600
400
200
0
0
20
40
60
80
100
120
140
160
Photon energy (keV)
Figure 2 X-ray emission from Wolfram anaode X-ray tube. Observe that for a given tube voltage, the
higher the energy of the photons, the less there are. And if the number of photons are to increase, then the
tube voltage should increase. Data from [1].
1.2 X-ray generation: The X-ray tube
A typical X-ray tube is depicted in Figure 1. It consists of an evacuated glass bulb with a heated filaments (Danish: glødetråd) as the negative electrode and a heavy metal positive anode. Thermic electrons emitted by the heated filament are accelerated across the gab to the anode. If the voltage between
cathode and anode is U volts and the current in the tube being I amperes, each electron will be hitting
the anode with a kinetic energy of U eV. The power deposited in the anode will be I times U, and the
total energy transferred to the anode in an exposure lasting t seconds will be IUt.
The electrons will be slowed down in the anode material, mainly releasing their energy as heat, but
to a small degree (few percent) the energy is transformed to either Bremsstrahlung or characteristic
X-rays. The Bremsstrahlung originates from the sudden deacceleration and direction changes of the
primary electrons in the field of the anode atoms, the characteristics X-rays originates from the knockout and subsequent level filling of inner electrons in the atoms of the anode material. The highest possible quantum energy of emanating X-rays (measured in eV) will be equal to U. Typical energy
spectra as a function of voltages are shown in Figure 2. Please note that the spectra are all taken at the
same current, only the voltage has been varied. This demonstrates that the total number of X-ray photons are heavily dependent on tube voltage. In addition to the information in Figure 2, a general rule
of thumb says that 15 keV increase in voltage corresponds to a doubling of the photon output. For
practical medical applications, the low energy part of the photons are normally not used but removed
by filtering either inside or just outside the X-ray tube. Normal filter materials are either aluminium
or copper. The thicker the filter and the higher the atomic number of the filter, the greater the cut-off
of low energy photons.
The description of the exposure characteristics of a given X-ray tube will comprise the voltage (in
units of kV), the current (in units of mA), the time of exposure (in units of s) and the degree of filtering
(for example a plate of Al, one mm thick next to a plate of Cu, 0.5 mm thick). As the total number of
photons produced for a given high voltage setting only depends on the product of current and time this
is often stated as a product in units of mAs.
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1.3 Anode material, power dissipation
Heavy elements are normally preferred for anode materials as the high Z-number gives efficient production of the part of the X-ray that originates from Bremsstrahlung. The characteristic X-ray lines,
which add to the total energy spectrum, normally appear in the range 50-70 keV, which ofte is in the
middle of the medical useful energy range.
The thermal load on the anode material both during the short exposures and averaged over time when
performing rapid, multiple exposures heats the anode dramatically. For this reason, normally a high
melting point material is used. Anodes made out of Tungsten (Danish: Wolfram), abbreviated W, are
very common. The area of thermal dissipation can be enlarged by rotating the anode during exposure.
An example of such a device is shown in Figure 3.
2 Typical X-ray system
Figure 4 shows a typical X-ray system. The X-ray is generated by the X-ray tube (Danish: Røntgenrør). Low energy photons are removed by the Al filter, since as they cannot penetrate the object
and contribution to the information on the film, they would only add needless to the dose received by
the object. X-ray radiation outside the image region on the film is removed by the collimator (Danish:
primærblænde). Attenuation (what is measured on the film) and Compton scattering take place at the
object. Only photons moving directly from the source to the film are allowed through the grid at the
bottom (Danish: sekundærblænde).
Figure 3 Rotating anode X-ray tube. "RTM" anode designates a Molybdenum anode mixed with 5% Rhenium
to improve the thermal stability. The metal anode is supported by graphite to improve the total thermal capacity.
Source: Siemens.
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Protective shield of
lead
X-ray tube
Al-filter (removes low
energy radiation)
Collimator
Object
Secondary radiation
(Compton scattering)
Grid (removes Compton scattering)
Screens
Film
Figure 4 Schematic illustration of a typical X-ray system.
Figure 5 provides an interactive illustration of a simple X-ray system allowing for translation of a
simple homogeneous phantom. As the distance is changed, different edge effects can be observed. By
pressing the “Lines” button, the geometry of the edge effect is visualized.
3 Geometrical considerations
Referring to Figure 5 we can define the distance from the X-ray origin (the focus1) to the objects as
FOD. The distance from the focus to the film2 or any other medium of radiation detection can be defined as FFD. Any object will to some degree attenuate the X-ray, and variation in X-ray absorption
across the objects will create a corresponding variation in the radiation impinging on the film. An unavoidable and sometimes desirable geometrical magnification of the image relative to the object can
be deduced from the triangle in Figure 5. The enlargement factor F, can be defined as:3
F = size of film image / size of object = FFD / FOD
(1)
If near to normal picture size and little variation in enlargement is sought for organs having different
depths in the body, the geometrical magnification should be minimized by using a large focus to film
distance (FFD) and a small object to film distance. A small object to film distance also improves image contrast, as blurring by scattering increases with increasing distance between object and film. This
is because the origin of the scatter is mainly inside the object: The longer the scattered radiation is
1. The "focus" is in this context the source of the X-ray photons, as it is the name for the electron spot on the
anode of the X-ray tube.
2. The term "film" is still common language, even though the conventional x-ray film to a large degree has been
replaced by various other imaging plates.
3. Here “size” means any distance e.g. the length of a given object.
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Figure 5 Simple illustration of the geometry of the planar X-ray system corresponding to Figure 4. The Xray image of the green homogeneous box is shown to the right. The focal point is the origin of the X-rays. The
film is identical to the detector. FOD = focus to object distance. FFD = focus to film distance.
allowed to travel between the object and the film, the more this radiation diverges from the true unscattered photons.
Similar geometrical considerations (i.e. similar triangles) can demonstrate that extended size of the
focus will generate blurring on the film.
4 Origins of contrast in the X-ray image
X-rays are attenuated according to the normal linear attenuation law:
I(x) = I0 exp(-x) = I0 exp(-/ x)
(2)
where x is the distance transversed in the material and  is the so-called linear attenuation coefficient
in units of m-1. I0 is the intensity at the entrance to the material (x = 0) and I(x) is the intensity at distance x. In the latter part of (2),  is the density of the material. By giving the attenuation coefficient
in units of / and the thickness in length times density (area weigth, e.g. g/cm2) the attenuation coefficient (now called mass attenuation coefficient) becomes independent on the physical state of the
material. Mass attenuation coefficients for some common tissues are given in Table 1.
The microscopic description of the attenuation comprises photo electric effects and Compton scattering, which are both described in the chapter on nuclear medicine. For the understanding of the X-ray
technique, it suffices to say that the linear attenuation coefficient for human tissues varies approximately as the electron density. Thus, it varies roughly proportional with the physical density (kg/m3).
Air has the lowest density, lung tissue has lower density than fat, fat has lower density than muscle,
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Table 1: Mass attenuation coefficients for typical tissues in /From [2].
µ given in cm2/g
50 keV
100 keV
200 keV
Air =0,0013 g/cm
0,208
0,154
0,122
0,227
0,171
0,137
Adipose tissue =0,95 g/cm
0,212
0,169
0,136
Muscle =1,05 g/cm
0,226
0,169
0,136
3
0,424
0,186
0,131
Lead =11,35 g/cm
8,041
5,549
0,999
3
Water =1,00 g/cm
3
3
Bone =1,92 g/cm
3
3
again having much lower density than the bone mineral of the skeleton. The X-ray attenuation varies
accordingly. X-rays transversing parts of the body having high absorbing material will be much more
attenuated, and the film or radiation capture device will in this region not receive as much radiation.
It should be remembered that the X-ray image is a negative (bright areas correspond to high attenuation) and that the image is a 2D projection of the 3D distribution of attenuation.
An example of a normal X-ray image of the chest (one of the most common medical imaging procedures) is seen in Figure 6. Notes that the most attenuating areas (ribs, vertebral column, heart) appear
white while the lungs with little attenuation appear black on a conventional planar X-ray image.
Problem 2 Does the planar X-ray image have arbitrary units (or, put in other words, are the pixel values relative or absolute)?
Figure 6 Normal chest X-ray image. This image is recorded with a tube voltage of
150 kV to minimize the contribution from bone. Press “?” to try to identify tissues.
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The image information in the planar X-ray is mainly anatomical, actual densitometric measurements
on the film are only performed for quality assurance programs and yields little medical information.
Today, all planar X-ray images are evaluated by a human observer.
5 Film, intensifyer foils and screens
Originally, the radiation was captured by a normal photographic film. In the film, the energetic Xray photons are absorbed in the silver halide (NaB-NaI) crystals, generating very small amounts of
free silver. During film processing, any grain with small amounts of free silver are completely converted to metallic, nontransparent silver, while the remaining unreduced silver halide is removed by
the fixative. X-ray films are of course made to the size necessary for the anatomical situation in question and can be very large. To increase sensitivity and thus lower radiation dose, the photo sensitive
film emulsions are often thicker and occasionally coated on both sides of the film, in contrast to normal photographic film. The silver contents of the films makes X-ray films rather expensive. Any film
has a specific range of optimal sensitivity (exposure range from complete transparency to completely
blackened). Although modern equipment are normally assisted by electronic exposure meters, the
correct choice of film, exposure time, exposure current and high voltage is still left to the judgement
of the X-ray technician.
To improve the sensitivity and thus lower radiation exposure to the patient, the film is often brought
in contact with a sheet of intensifying screen. The screen contains special chemical compounds of the
rare earth elements, that emits visible blue-green light when hit by X-rays or other ionizing radiation.
This permits the use photographic film with thinner emulsions and more normal sensitivity to visible
light. While increasing the sensitivity, the use of intensifying screen on the other hand blurs the images as the registration of X-ray radiation is no longer a direct, but an indicted process.
The patients or the object is not only the source of X-ray absorption but also of X-ray scattering,
mainly due to Compton effect (please see the chapter on Nuclear Medicine). Any part of the patient
exposed to the primary X-ray beam will be a source of secondary, scattered, X-rays. These X-rays will
have lower energy than the original ray, but as no energy discrimination is used in the registration,
also the secondary scattered radiation adds to the blackening of the film. The scattered radiation carries no direct geometrical information about the object and thus only reduces the contrast by increasing the background gray level of the film. Scattered radiation can to some degree be avoided by the
use of special collimators called raster. The raster can be a series of thin, closely lying bars of lead,
only allowing radiation coming from the direction of the focus point to hit the film while other directions are excluded. A typical example of a complete radiography cassette content is shown in Figure 7.
6 Digital radiography and direct capture
During the last years the conventional film based radiography has gradually been replaced by newer
digital techniques. The end points is of course the acquisition and storage of the X-ray image information as computer files. It should be noted that X-ray images are normally of very high resolution
(more than four thousand by four thousand pixels) with large dynamic range (12 to 16 bits). The correct handling and display of such image information without loss or compression is still the subjects
of specialized workstations. The digital X-ray images are normally stored and displayed in so-called
PACS systems (Picture Archiving and Communications System). The image information is either
temporarily captured on phosphor plates for subsequent transfer to digital storage by so-called phosphor plate readers, in function much related to the old X-ray film processors) or by direct, position
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X-ray passing
through the raster
X-ray not passing
through the raster
Raster
IS
Film
Figure 7 X-ray cassette, containing double coated film, intensifying screen (IS) and lead raster.
sensitive electronic X-ray detection devices, the so-called direct capture systems. At present (2003)
the geometrical resolution of the various digital techniques is still somewhat inferior to the best possible film technique. However, the benefits of rapid viewing, interactive image availability, postprocessing and digital transmission often outweighs the reduction in image quality. For special
applications, like breast cancer detection by mammography, the photographic film is still the system
of choice.
7 Analysis of image of phantom
Consider the conventional planar X-ray image of a typical phantom in the course shown in Figure 8.
This X-ray system has a high dynamic range, thus the image has many levels. To include as many
details as possible, this X-ray is shown - very untraditional - in color. Normal X-ray images are shown
in shades of gray.
It is not trivial to analyze such an image, unless one is very careful. First one should remember that
what one see (the contrast in the image) is total attenuation which again is dependent on what kind of
material is present and how much of this material is present in the trajectory between focus and a given
detector. And since the colorbar is representing relative values, both of these observations should be
considered relative to other places in the image.
The analysis goes as follows:
Since the image shows more than the box of the phantom itself, the connectors can be identified.
Those appear dark yellow and since they obviously must attenuate more than the nearby air (which
appears nearly white) it can be concluded, that a high pixel value corresponds to low attenuation (in
contrast to normal X-ray). Be aware of a potential pitfall: if the film is larger than the area that is exposed (as controlled by the collimator at the x-ray tube), then the image has an additional outer frame,
which, obviously, should not be used in the analysis.
Next consider the fiducial markers. Since the lid was in place during recording of the image, we have
two series of materials that are penetrated by the X-rays: one through the markers and another between
the markers:
• agar, acrylic base plate
• acrylic lid, agar, acrylic base plate.
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Vertical (mm)
20
1000
40
900
60
800
80
700
100
600
500
120
400
140
300
160
40
60
80
100
Horizontal (mm)
120
Figure 8 Left: Top photo of phantom 4 from 2009. Right: Planar X-ray image of same phantom. The image
is shown in color to enhance contrast. The two images are sought aligned as well as possible.
Problem 3 Please make a cross-sectional drawing of this, and be sure that the total distance penetrated by the X-ray is the same (for this particular problem, it is assumed that the X-rays hits perpendicular to the surface).
Since air-free agar is mainly water and thus has very much the same density as water and since acrylic has a higher density, the fiducial markers must attenuate less, and thus appear brighter than the surroundings. This is also the case in the image in Figure 8.
Next consider the tube. The tube contains air inside and thus the wall is the most attenuating. But
even if the tube contained water inside, the wall would still be the most attenuating. Thus, the tube
periphery should be darker than the center when looking at the projection image. The shape of the
actual attenuation profile through the tube is considered in Problem 13 in the exam of year 2008.
8 References
[1] Johannes Jensen og Jens Munk: Lærebog i Røntgenfysik, Odense 1973.2.udg. s.121.
[2] Data from http://physics.nist.gov/PhysRefData/XrayMassCoef/cover.html
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