3. 1. X-rays: A brief history
Over one hundred years
ago on November 8 1895,
the German physicist
Wilhelm Conrad Roentgen
(figure 1.1)happened upon
X rays in his laboratory in
Würzburg.
On December 28 1895,
Roentgen announces his
discovery with a scientific
paper, W. C. Roentgen:
About A New Kind of Rays.
1896 Fluoroscopy is
invented in January by
Italian scientist Enrico
Salvioni, while American
inventor Thomas Edison
works on a similar device
Figure 1.1:Wilhelm Conrad Roentgen
3
4. Figure 1.2:The famous radiograph made by Roentgen on December
22, 1895. This is traditionally known as “the first X-ray picture” and
“the radiograph of Mrs. Roentgen’s hand.
4
5.
1901 Roentgen wins the first Nobel Laureate in
Physics prize for his discovery.
1904 Clarence Dally, Thomas Edison’s assistant in Xray research, dies of extreme and repeated X-ray
exposure
1910 Eye goggles and metal shields are commonly
used to shield X-ray users
1919 Dr. Carlos Heuser, an Argentine radiologist, is
the first to use a contrast medium in a living human
circulatory system.
1920–1929 Chest X rays are used to screen for
tuberculosis.
Roentgen dies February 10, 1923.
1927 Portuguese physician, Dr. Egaz Moniz the first to
create images of the circulatory system in the living
brain.
Drs. Evarts Graham and Warren H. Cole of Washington
University,St. Louis, discover in 1923 how to visualize
the gall bladder with X rays by using contrast media.
5
6.
In 1936, the first “tomograph”—an X-ray “slice”
of the body—is presented at a radiology meeting.
The Betatron, a circular electron accelerator, is
developed by Dr. Donald Kerst of the University
of Illinois between 1940–1943.
In 1960, Dr. Robert Egan of the University of
Texas M.D.Anderson Tumor Institute, Houston,
with the support of the U.S. Public Health
Service, publishes the results of an intensive,
three-year study of mammography.
1970–1979 CT, or computed tomography, which
takes X-ray “slices” of the body and images them
on a computer screen, is introduced.
1980s MRI (magnetic resonance imaging; also
referred to as MR)—the marriage of a strong
magnet and a computer—is introduced, instead
of X-ray’s ionizing radiation
6
7.
1980s the development of digital
technologies for X-ray imaging
1995 The introduction of Flat Panel
Detector (FPD)
7
9. 2- How X-ray works
Conventional x-ray radiography produces
images of anatomy that are shadowgrams
based on x-ray absorption.
The x-rays are produced in a region that is
nearly a point source and then are
directed on the anatomy to be imaged.
In medical x-ray imaging, the x-ray
energy typically lies between 5 and 150
keV, with the energy adjusted to the
anatomic thickness and the type of study
being performed.
9
11.
The x-rays emerging from the anatomy are
detected to form a two-dimensional image, where
each point in the image has a brightness related
to the intensity of the x-rays at that point.
Image production relies on the fact that
significant numbers of x-rays penetrate through
the anatomy and that different parts of the
anatomy absorb different amounts of x-rays.
In cases where the anatomy of interest does not
absorb x-rays differently from surrounding
regions, contrast may be increased by
introducing strong x-ray absorbers. For example,
barium is often used to image the gastrointestinal
tract.
11
16. 3.3 Mamography
Figure 3.2 Digital X-ray
Figure 3.3Schematic diagram of a dedicated mammography
machine.
16
17. 3.4 Computed Tomograpgy (CT)
In the mid 1970S, CT gave physicians a whole
new way of seeing.
By eliminating the interfering patterns that
come from over- and underlying bones and
organs, CT provides ample contrast among
the various soft tissues.
CT is routinely used for detailed studies of
abdominal and pelvic organs, the lungs, and
the brain.
CT can image objects down to about 1/3
millimeter
17
18. Figure 3.4 Computed tomography. This view of a slice of bone, eyes,
and brain tissues several millimeters thick displays good soft-tissue
contrast and detail, and low visual noise.
18
20.
3.5 Flouroscopy
Fluoroscopy is radiography's first cousin.
Here, the X rays that pass through and emerge
from the patient are projected onto the front face
of an image intensifier, an electronic vacuumtube device that transforms a life-size pattern of
X-ray shadows into a small, bright optical image.
This visible image can be fed into a film camera;
more commonly, it goes to a television (video)
camera, where it is converted into an electrical
signal and sent to a video monitor for live display.
The image can be recorded on videotape for
subsequent playback and further processing.
20
25. 4- Image Detection
4.1 Screen Film
4.2 X-Ray Image Intensifiers with
Televisions
An x-ray image intensifier detects the xray image and converts it to a small,
bright image of visible light.
This visible image is then transferred by
lenses to a television camera for final
display on a monitor.
25
27. 4.3 Photostimulable phosphors (PSPs)
They were pioneered by Fuji in the 1980s.
In modern hospitals a photostimulable phosphor
plate (PSP plate) is used in place of the
photographic plate.
After the plate is X-rayed, excited electrons in the
phosphor material remain "trapped" in "colour
centres" in the crystal lattice until stimulated by a
laser beam passed over the plate surface.
The light given off during laser stimulation is
collected by a photomultiplier tube and the
resulting signal is converted into a digital image
by computer technology, which gives this process
its common name, computed radiography.
The PSP plate can be reused, and existing X-ray
equipment requires no modification to use them.
27
29.
1.
2.
3.
4.
5.
4.4 Flat Panel Detector (FPD)
The flat electronic detector provides direct
digital registration of X-ray images, without
the intermediate stage of optical or
mechanical scanning.
It has the advantages of :
It has a high DQE, allowing low-dose operation
It has a compact construction for easy integration
It offers a new perspective for live X-ray
examination
High image quality
The integrated flat a-Si detector provides a
significant increase in efficiency, particularly in busy
bucky rooms.
29
31. 5-Flat Panel Detectors (FPDs)
1.
2.
FPDs are classified in two main categories:
Direct FPD's
InDirect FPD's
In all the methods, the charge is accumulated
for a frame period before being read out.
Gamma cameras, in contrast, count each x-ray
photon as it arrives.
That technique is generally not used for x-ray
imaging because the x-ray photon arrival rates
are too high to permit counting
31
32. 5.1 Direct FPDs:
There are 2 types of direct FPDs:
1.
The Intrinsic Method
Figure 5.1: Intrensic Direct FPDs
32
33.
Arriving x-rays are captured by the amorphous silicon
diode where hole-electron pairs are generated.
An applied bias separates the charge to prevent
recombination.
Because a charge pair is generated for about each 5
electron volts of x-ray energy, the signals are high.
Unfortunately, the x-ray absorption of silicon is very low
so the photodiode needs to be 10 to 20 mm thick.
Fabricating such devices of amorphous silicon is not
feasible.
Intrinsic devices have been made from crystalline silicon
but only arrays of one or two lines are practical and even
these are expensive.
33
35.
Photoconductive materials with higher x-ray
absorption than silicon can be coated on an array of
conductive charge collection plates each supplied with
a storage capacitor.
These also produce hole-electron pairs when x-rays
are absorbed but the charge generated must be
stored out of the layer to avoid lateral crosstalk.
The applied field not only separates the charge but
directs it towards the collector plate directly below to
maintain image sharpness.
Currently, the only photoconductor in production,
selenium also has relatively low x-ray absorption and
requires about 50 electron volts to produce a holeelectron pair.
These restrict both the minimum dose needed and the
size of the signal generated.
Other materials with lower energy requirements and
higher x-ray absorption are under development.
35
37.
This method incorporates the usage of a
scintillator.
A scintillator is a compound that absorbs x-rays
and converts the energy to visible light.
A good scintillator yields many light photons for
each incoming x-ray photon; 20 to 50 visible
photons out per 1kV of incoming x-ray energy
are typical.
Scintillators usually consist of a high-atomic
number material, which has high x-ray
absorption, and a low-concentration activator
that provides direct band transitions to facilitate
visible photon emission.
Scintillators may be granular like phosphors or
crystalline like cesium iodide.
37
38.
The FPD consists of a sheet of glass covered with
a thin layer of silicon that is in an amorphous, or
disordered state.
At a microscopic scale, the silicon has been
imprinted with millions of transistors arranged in
a highly ordered array, like the grid on a sheet of
graph paper.
Each of these TFTs is attached to a lightabsorbing photodiode making up an individual
pixel (picture element).
Photons striking the photodiode are converted
into electron-hole pairs.).
Since the number of charge carriers produced will
vary with the intensity of incoming light photons,
an electrical pattern is created that can be swiftly
read and interpreted by a computer to produce a
digital image.
38
40. A.Structure of a phosphor scintillator:
Phosphors are materials which glow when
exposed to x-rays.
For maximum brightness, the phosphors
use in x-ray imaging are made of rareearth oxysulfides doped with other rare
earths.
The most common are gadolinium and
lanthanum oxysulfides doped with
terbium.
These typically emit blue to green light
which is well-matched to film sensitivity.
Various grain sizes and chemical mixtures
are used to produce a variety of resolution
and brightness varieties. In use, these are
mixed with a glue binder and coated on to
plastic sheets.
Tens of electron volts are needed to
produce each visible photon in a phosphor
screen and x-ray absorption is good.
Light scatter can be a problem if the
layers must be thick to stop higherenergy x-rays.
Figure 5.5:
Phosphor
scintillato
rs
40
41. B.Structure of CsI scintillator:
For a better combination of resolution
and brightness, cesium iodide is used.
CsI grows as a dense array of fine
needles under the proper evaporation
conditions.
This produces crystals which act as
light pipes for the visible photons
generated near the input side of the
layer allowing very thick layers to be
used with excellent retention of
resolution.
About 20-25 electron volts are needed
to generate each light photon.
When doped with thallium, CsI emits
at about 550 nm, just at the peak of
the spectral sensitivity of amorphous
silicon.
The combination of CsI and
amorphous silicon has the highest
DQE of all materials in production
today.
Figure 5.6:
CsI
scintillato
rs
41
43. 5.4.1 Scintillator layer
Closely coupled to the array is the X-ray scintillator.
Generally, rare earth screens such as gadolinium
oxysulfide, can be a separate detachable sheet which
is mechanically forced into close contact with the
array.
If a CsI screen is used, this is often directly deposited
on the array, to give the best optical coupling
efficiency.
CsI is used in applications like low-dose fluoroscopy,
where the photon flux is very low.
Figure 5.8 shows a comparison between the
absorption efficiency of CsI and gadolinium oxysulfide.
In addition to its much higher absorption efficiency,
CsI also produces roughly twice the light output of a
gadolinium screen, which results in more than four
times the signal at the photodiode for a given dose.
43
45.
The thickness of the CsI can be greater than that
of a rare earth screen because when CsI is
deposited on the array it grows in a columnar
structure.
The columns tend to act as light pipes, reducing
the amount of light spreading in the scintillator.
So, for example, a 600μm CsI layer can have
resolution similar to a 300μm thick rare earth
screen.
These screens such as gadolinium oxysulfide
have the advantage of much lower cost and
greater flexibility in that the screen can easily be
changed to match the resolution requirements of
the application.
45
46. 5.4.2 TFT/Photodiode layer
The light generated by the scintillator is absorbed by the
photodiodes in the array, creating electrons which are
stored on the capacitance of the photodiode itself.
The peak light absorption efficiency of the photodiodes is in
the green spectrum, at 550nm wavelength.
Both gadolinium oxysulfide and thallium doped cesium
iodide,CsI(Tl), produce their peak light output at this
frequency.
The amorphous- silicon photodiodes are typically the “n-i-p”
type.
This type of amorphous-silicon photodiode has the
advantages of low dark current and a capacitance that is
independent of the accumulated signal.
Compared to crystalline silicon photodiodes like those used
in CMOS imagers, the dark current in amorphous silicon
photodiodes is orders of magnitude less.
So it is not unusual for amorphous silicon flat panel arrays
to be capable of more than ten second integration times at
room temperature.
The fact that the diode capacitance is independent of signal
helps make the detection system linear.
46
49.
Amorphous silicon is not suited to the subsequent signal
processing so every column and row of the array is
brought to the edge of the glass, where it is connected to a
standard crystalline silicon chip by means of a TAB (tape
automated bonding) package.
The chips that need to be directly connected to the array, the
readout chip and the driver chip, are mounted in these TAB
packages.
Figure 5.11:Board-side view of TAB packaged row driver and
custom ASIC readout chips.
49
50.
The TFT/photodiode matrix is normally scanned
progressively, one line at a time from top to bottom.
At the end of each dataline is a charge integrating amplifier
which converts the charge packet to a voltage.
At this point the electronics vary by manufacturer, but
generally there is a programmable gain stage and an
Analog-to-Digital Converter.
One important aspect of the scanning is that the FPD is
continuously collecting the entire incident signal; none is
lost even during the discharge of the pixel.
The FPD is an integrating detector and the integration time
for each pixel is equal to the frame time.
The electronics can be mounted to the side of the array, out
of the beam, as is done in higher energy (MeV) applications
to protect against radiation damage.
But for diagnostic and interventional procedures, to
maintain the best view of the patient, the electronics can
also be mounted behind the array and protected by a thin
layer of lead.
50
51. Flat Panel Operational Advantages
The most obvious advantages of flat panel imagers are size
and weight.
The Image Intensifiers Tubes (IIT) are large and bulky. An
FPD with a 12”x16” active area (20” diagonal) takes up less
than 25% of the volume of a 12” IIT and less than 15% of
that of a 16” IIT.
In addition, the FPD takes the place of not only the IIT, but
also the attached image recording devices, including the
CCD camera, 35mm Cine camera, and the spot film device.
The result is vastly improved access to the patient in
interventional procedures.
In addition to the reduction in size, the weight of the flat
panel imager is 60% less than that of the IIT-based
imaging chain.
Flat panels also are more economical than an IIT of
comparable size considering that we know IIT image quality
deteriorates as a simple consequence of everday X-ray use,
and thus IIT’s have a relatively short service life.
51
52. Figure 5.12:Signal output comparison vs. X-ray dose between
an FPD and an IIT of comparable size.
52
53. FPD Image Quality
FPDs have a very direct, short signal conversion
path, with essentially no optics.
The result is a very flat, uniform “film-like” image
from edge-to-edge.
The ability of flat panel detectors to encompass
multiple X-ray modalities is also a function of
their very large dynamic range.
Figure 5.13 shows the signal response of an FPD
in the full resolution mode, over the dose range
of 1μR to 1.2mR.
Particularly for the amorphous silicon
TFT/photodiode technology, the response to
entrance dose of the FPD is extremely linear.
The response of the imager deviates from the
straight line curve by less than 0.01%.
53
55.
Of equal importance is the Signal-to-Noise (SNR)
behavior versus dose.
The FPD contributes electronic noise to the
image.
However the statistical noise in the X-ray beam is
dominant.
The noise in the beam follows Poisson statistics
(the noise is equal to the square root of the
number of incident X-ray photons)
As shown in Figure 5.14, the SNR of an FPD has a
square root dependence on dose, i.e. is X-ray
quantum limited over a very large range.
This is an indication that the detector contributes
effectively no noise to the image over this dose
range.
55
57. •The DQE as a function of entrance dose (fluoro and cine
range) for an angiographic flat panel imager is shown in Figure
5.15.
•This is typical of the indirect TFT/photodiode technology with
a CsI scintillator.
•Because of the low-loss, high-absorption detection path, the
DQE for indirect CsI-based flat panels is the highest available
and is more than double that of computed radiography, film
screen and CCD based technologies.
•Higher DQE translates directly into better imager quality for a
given dose.
•So with high DQE detection systems, it is possible to get the
same image quality as a low DQE system like screen film at a
fraction of the dose.
57
58. Figure 5.15:DQE vs. spatial frequency over the fluoroscopic and cine
dose range, for a 40x30cm angiographic FPD.
58
59. Examples of Equipments with FPD
Schick technologies intraoral
wireless FPD
Model CDR wireless
59
60. Philips Veradius C-arm with FPD
•1.6k x 1.4k image
•Active image area of 28.7 x 26.5 cm
60
