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Fig. 4.1: Illustration of pulsed and
continuous exposure modes. Note
that the vertical axes are not to scale.
Basic Physics of Digital Radiography/The Image
Receptor
Technological aspectsof digital radiography image receptors are
described in thischapter. These receptors are generally used for
radiography andsome can also be used for fluoroscopy. As a result, in
addition to single-shot radiography, other exposure modes can also be
used with the technology described.The use of continuous X-ray
exposures in fluoroscopygenerally refers to the use of low XRT
currents (i.e. 0.5to 5 mA) - see Figure4.1. The resultant images have a
low image quality but are sufficient for applications such as patient
positioning or monitoring catheterplacement. Its use is also commonly
called Screening, a reference to the days when a sheet of glass coated
with a fluorescent material was used for such imaging. By contrast, the
term fluorographygenerally refers to the use of relatively intense (e.g.
50 to 1000mA), pulsed exposures. Pulses are typically of short duration
and can be applied at a rate of, for example, 1 to 8 per second. The resultant images have a relatively high image
quality and can therefore be used for diagnostic purposes.
Contents
1 Computed Radiography
1.1 Laser Readout
1.2 Imaging Plate
2 Digital Radiography
2.1 Indirect Receptors
2.2 Direct Receptors
3 XII-Video Systems
3.1 X-Ray Image Intensifier
3.2 Video Camera
3.3 Automatic Brightness Control
4 Radiation Dose Management
4.1 Skin Injury4.2 Staff Dose
4.3 Patient Dose
5 References
Computed Radiography
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Fig. 4.2: CR scanning plate.
We have previously described the process of photostimulable luminescence which is exploited in Computed
Radiography, where the absorption of radiation causes electrons to become trapped at intermediate energy levels.
Phosphors having this property are therefore referred to as Storage Phosphors.
A latent X-ray image can be recorded using a plate coated with crystals of barium fluorohalide compounds which
contain trace amounts of europium. The general form of these compounds is BaFlX:Eu, where X can be Cl, Br or
or a mixture thereof. Radiographic information is recorded by elevating electrons to traps in the energy gap and the
number of filled traps is proportional to the amount of radiant energy absorbed. When stimulated by visible
radiation of wavelength around 600 to 680 nm (i.e. red light), the crystals luminesce as the electrons return to theirground state in the range 300-500 nm (blue).
Laser Readout
The plate can be scanned in a raster pattern with a finely focused
laser beam, e.g. He/Ne, and a scanning mirror - see Figure 4.2.
The laser light stimulates the release of electrons from the traps
giving rise to the emission of light which is collected by a light guide
and fed to a photodetector, e.g. a photomultiplier tube. The signalsgenerated by the photodetector as the plate is being scanned are
amplified and digitized by an analogue-to-digital converter (ADC).
The image is in essence built up point by point and line by line to
give a digital image resolution of up to 4k x 4k pixels. Pixel size is
typically 0.1 m. The imaging cycle is completed by flooding the plate with a high intensity sodium discharg
lamp that erases any remnants of the latent image and essentially prepares the plate for reuse - see Figure
4.3.
Fig. 4.3: The CR imaging process.
The spatial resolution of computed radiography is influenced by factors such as the phosphor plate thicknes
the readout time and the diameter of the laser beam, which is typically about 100 m. Note that divergence
through scattering of the laser light in the body of the phosphor layer will broaden the stimulated area to
beyond this diameter.
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Fig. 4.4: Implications of the finite
width of laser beam and phosphor
layer on the spatial resolution in CR.
Note that the nature of the CR imaging process limits its application to single-shot radiographic imaging. In
clinical practice, the process is generally part of a workflow where patient radiographs are recorded as
traditionally in film/screen radiography, but with images now generated with a latent image read-out device
and automatically sent to a quality control workstation for image evaluation, annotation and transfer to a
PACS for reporting.
Imaging Plate
The CR imaging plate is very similar to intensifying screens in its
structure. It consists of tiny phosphor grains of about 5 m
embedded in an organic binder which is coated onto a substrate
material. The plate is turbid and as a result scatters light (both
excitation laser and stimulated light) strongly and isotropically - see
Figure 4.4. Thus, the light diffusion limits the useful thickness of the
phosphor layer. An improvement in X-ray absorption efficiency
can indeed be obtained by increasing the thickness of the
phosphor layer. However,
lateral diffusion of light in the phosphor layer
will increase in proportion to the layer
thickness, impairing spatial resolution; and
sensitivity will not increase all that much when the layer exceeds a certain thickness,
because most of the light stimulated deep in the layer will not reach the surface for
detection.
This sensitivity limitation can be overcome by making the substrate of the imaging plate from a transparent
material and to detect the photo-stimulated luminescence from both the front and the back sides of the
plate[1]. This requires two light-collection systems, but only one laser beam. The improved sensitivity results
since only ~30% of the photostimulated light is collected using the conventional single laser approach. In
addition, sensitivity improvements have been achieved using a structured phosphor, e.g. CsBr:Eu - see the
discussion below - as the imaging plate[2].
A fundamental limitation of CR is the time required to read the latent image. Since the decay time of the phosphor
luminescence is ~0.7 s typically, the readout of a 3,000x3,000 pixel image can take over half a minute to
complete. An improvement can be obtained by line scanning, where a full line of pixels is stimulated and read out
simultaneously instead of single pixels, as described above. This line-scanning approach requires a linear array of
laser light sources, e.g. laser diodes, as well as a linear array of photodetectors as wide as the imaging plate, andgives rise to readout times of less than 10 seconds. Furthermore, the linear scanning mechanism can be built into th
image receptor cassette[3].
Note that clinical radiography with CR plates has been found to generate a range of unique artefacts, which are th
subject of a pictorial review in Cesar et al. (2001)[4].
igital Radiography
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Fig. 4.5: Illustration of an active
matrix array and its associated
electronic circuitry housed inside the
image receptor cassette and
connected to a digital image
processor and image display device.
Considerable research has been conducted in recent years into the
development of flat panel image receptors for digital radiography. This
research has extended from the development of active-matrix liquid-
crystal flat panel displays (AMLCD) for application in, for example,
laptop computers. The underlying technology of AMLCDs is a large area
integrated circuit called an active matrix array which consists of many
millions of identical semiconductor elements deposited on a substrate
material. An intensifying screen or a photoconductor coupled to such an
active matrix array forms the basis of flat panel X-ray image receptors.
These detectors have been applied for both radiography and
fluoroscopy.
Such a receptor is illustrated in Figure 4.5, where the active matrix array
and associated electronic circuitry is mounted in a cassette. Array sizes of
up to 43 cm x 43 cm are available with more than 9 million pixels (pixel
size ~140 m). Operation of the array is controlled by a digital image
processor (see the next chapter), which also stores and displays the
resultant images.
A simplified description of the operation is illustrated in Figure 4.6. Each
pixel of the array has a switch (typically made from a thin film transistor)
which is connected to switching control circuitry in a manner which
allows all switches in a row of the array to be operated simultaneously.
The output from each pixel is connected in columns with individual pre-
amplifiers. All switches are kept in the off position during the X-ray exposure. Following the exposure, the switche
in the first row are turned on and the signal from each pixel is amplified by the pre- amplifiers, digitized using an
ADC and stored in the image memory of the digital image processor. These switches are then turned off and the
switches in the second row are turned on to acquire signals from the second row of pixels. This process is repeate
for the whole array so that an image is acquired in a sequential, line-by-line manner. When the receptor is used forfluoroscopy the array is scanned continuously[5]
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Fig. 4.6: Illustration of a 3x3 pixel region of
an active matrix array, showing the switch
associated with each pixel, the switchingcontrol and pre-amplifier circuitry and
connecting conductors.
Research developments have advanced on two fronts as illustrated in Figure 4.7.
Indirect Receptors
Indirect receptors are based on coupling an intensifying screen to the active matrix array. Phosphors such a
Gd2O2S:Tb and CsI:Tl have been used, and the light produced following X-ray interaction is detected by aarray of pixels consisting of photodetectors - see panel (a) in the figure. Each photodetector generates an
electric charge that is proportional to the amount of light striking it, and this charge is stored until it is read-ou
by the switching control circuitry. The detection process is referred to as indirect since the detected X-rays
are first converted to the light, which is subsequently converted to electric charge.
CsI:Tl has distinct advantages as an intensifying screen because it has favourable K-absorption edges at 33
keV (I) and 36 keV (Cs) and its crystals can be grown in a dense, needle-like (5-10 m in diameter)
structure in a thin layer up to 600 m thick[6]Fluorescent light generated by X-ray absorption can therefore
be guided fiber-optically to the photodetector array without much lateral dispersion. CsI:Tl as a result is a
so-called Structured Phosphor, as opposed to the dispersed powder phosphor obtained using Gd2O2S:T
and similar screens. The packing density of scintillation sites in CsI:Tl is also quite high as a result of its
crystalline structure, whereas a powder screen consists of only about 60% phosphor and 40% binder. This
results in an absorption efficiency for structured phosphors to be about four times higher than that of a
powder screen with similar spatial resolution capabilities.
CsI:Tl fluoresces in the green region of the visible spectrum, where the light absorption efficiency of a-Si is
relatively high.
Indirect receptors have also found application in Computed Tomography. In MDCT, for instance, arrays o
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up to 916 columns by 64 rows of detector elements made from a phosphorescent material are coupled to a
a-Si array of photodiodes. Ultra-Fast Ceramic(UFC) detectors which are manufactured as doped
gadolinium oxysulfide crystals feature a relatively fast response, high detection efficiency and broad dynamic
range. In addition, collimator blades can be mounted directly on the detector array for scatter reduction
purposes.
Fig. 4.7(a): Cross-sectional view of an indirect
flat panel image receptor showing three pixels
mounted on a glass substrate.
Fig. 4.7(b): Cross-sectional view of a direct
flat panel image receptor.
Direct Receptors
Direct receptors are based on coupling a photoconductor to the active matrix array. Photoconductors such
as amorphous selenium (a-Se) have been used, and the electric charge produced following X-ray interactio
is detected by an array of pixels each consisting of an electrode and a capacitor - see panel (b). This charge
is stored in each capacitor until it is read-out by the electronic switching circuitry. The photoconductor
requires a voltage of the order of 5,000 V to be applied, using a surface electrode, so that the charge
produced can be attracted to the pixel electrodes.
It should be noted that the photoconductive properties of amorphous selenium (a-Se) have been known for
many years and have been widely applied in photocopiers, fax machines and laser printers. The direct
imaging technique has evolved from xeroradiography, which was a popular medical imaging technique many
years ago.
Selenium is not as efficient as cesium iodide as an X-ray absorption material because of its lower atomic
number (Z=34). Its K-absorption edge is at 13 keV and its attenuation falls off rapidly with increasing
energy. It can be expected on this basis that rather thick layers of a-Se are required to absorb X-ray
photons with equal efficiency to CsI at higher X-ray energies. Other potential materials such as lead oxide
(PbO), lead iodide (PbI2) and mercuric iodide (HgI
2) exist which have better absorption efficiencies than
selenium for a given layer thickness and may find application in the future[7].
Important design features of flat panel image receptors include the pixel size and the fill factor. Pixel size affects
spatial resolution and sizes of the order of about 100-200 m are typical. The fill factor is the percentage of a pixe
area which is sensitive to the image signal - be it electric charge or light photons. It can never be 100%, given the
need to accommodate conductors (~10 m wide) which input switching signals and which output image signals, a
well as the thin film transistor in each pixel. Typical fill factors are in the range 50-80%.
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Fig. 4.8: Basic elements
of an XII-video system
Flat-panel detectors generally use active matrices of hydrogenated amorphous silicon (a-Si:H+) as readout arrays.
This is the technology developed for laptop displays which makes it possible to manufacture large-area X-ray
imagers with high radiation tolerance. The amorphous silicon layer is formed on a glass substrate, for instance, by
plasma-enhanced deposition of silane (SiH4) gas. The disordered structure of the amorphous layer has the
advantage that X-ray detection can occur without radiation damage to the material. The material can therefore be
repeatedly exposed with comparatively intense X-ray beams. The active matrix array is manufactured from the a-S
layer by photolithography and chemical etching techniques, as widely used in integrated circuit manufacture.
However, a-Si is limited by the minimum size of the structures that can be built into this material. Another possibilit
is a readout array based on complimentary metal-oxide semiconductor (CMOS) technology. CMOS is made usin
crystalline silicon wafers instead of amorphous silicon and is highly developed in the semiconductor industry.
CMOS imagers provide a high fill-factor and very fast readouts. In addition, it is possible to integrate on-chip
electronics (e.g. amplifiers) for each pixel, which can improve performance significantly and may someday provide
X-ray energy discrimination capabilities. Their use in DR however has been limited because of significant electronic
noise.
A variant on the design of indirect detectors is to mount the photodiode/TFT array on the anterior side of the
intensifying screen - so called Irradiation Side Sampling (ISS)[8]
. The benefit here is an improved detectiongeometry between the scintillation sites in the screen and the photodiodes which reduces light attenuation and
blurring effects.
An interesting hybrid design can be obtained using features of both the indirect and direct image receptors. Here a
structured phosphor is coupled to a thin a-Se layer and has a potential advantage when electron amplification is
controlled using programmable applied voltages.
Reduction of electronic noise generated by TFTs and associated circuitry can be achieved using optical readout of
direct detectors. Here two layers of a-Se are used separated by an electron-trapping layer (ETL). One layer is
used to generate electrons when X-rays are absorbed which are trapped by the ETL, and the second thinner layer
is used to read this charge when it is illuminated by a line-scanning array of light-emitting diodes [9].
XII-Video Systems
The basic elements of an X-ray Image Intensifier (XII) and video system are shown
in the Figure 4.8. X-rays emerging from the patient are converted to an optical image
in the XII, which is viewed by a video camera and displayed using a computer
system. These systems are widely used for fluoroscopy and fluorography during
barium-studies, angiography and interventional radiology.
Many systems also feature a mechanism for placing a CR or DR cassette at the XII
input so that radiographic images can be recorded. This device is sometimes referred
to as a Cassette Changerand may allow a number of images to be recorded onto
one image receptor. Units incorporating cassette changers are sometimes referred to
as Radiography/Fluoroscopy (R/F) systems - see Figure 4.9.
X-Ray Image Intensifier
Figure 4.10 illustrates the essential features of an X-ray image intensifier. After the X-ray beam emerges fro
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Fig. 4.9: Typical R/F system
illustrating a triple field XII shown
above the table with video monitor to
the right. The video camera is side
mounted on extreme left of XII. The
receptacle for a CR or DR imager is
shown arrowed. Operator controls
for collimation and XII field sizeselection are also located at this
position. The X-ray tube is mounted
under the table and a second one,
used for overtable w ork , is shown
suspended in the background.
the patient, it enters the XII tube to be absorbed by the input
phosphor and thus create light photons. These photons strike a
photocathode and cause it to emit electrons. The electrons are
accelerated towards the output phosphor by a high electric field
produced between the anode and photocathode. Focusing is
accomplished by the focusing lenses. The electrons hit the output
phosphor and cause large numbers of light photons to be
produced which carry the fluoroscopic image. By this complexmeans (see Figure 4.11), the XII converts an X-ray image to an
electro-optically intensified light image.
Image intensification also results because the output phosphor is so
much smaller than the input phosphor. Output phosphors are
typically about 2 cm in diameter while input phosphors may vary in
diameter from about 15 to 40 cm, depending on the clinical
application. Intensification through image minification can therefore
range from factors of about 56 to 400. Additional intensification is
achieved by the electron acceleration inside the tube, i.e. the FluxGain. The overall brightness gain achieved (i.e. the minification
times the flux gain) can be over 5,000[10].
Fig. 4.10: The X-ray image intensifier (a), with close-
ups of the phosphor/photocathode sandwich (b) and
the output phosphor (c).
The input window is usually made of a thin layer of titanium, carbon fiber or aluminium so as to minimize X-
ray absorption. The input phosphor is usually cesium iodide activated with sodium (CsI:Na). The input
phosphor and photocathode must be in intimate contact with each other in order to preserve the latent imagfidelity - see panel (b). Electrons are produced in direct proportion to the brightness of the input phosphor
and are accelerated through a potential of 25-35 kV before striking the output phosphor. This accelerating
anode is located in the neck of the evacuated tube.
The electrostatic focusing lenses (G1, G2and G3) are at a positive potential with respect to the cathode. Th
photocathode is actually curved (unlike that shown in Figure 4.10) to equalize the path lengths of all electron
so as to minimize image distortion.
The output phosphor is generally made from a phosphor called P20. This phosphor emits green light which
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Fig. 4.11: Illustration of the XII imag
transduction process .
matches well with the sensitivity of video cameras. A very thin layer of aluminium is plated onto this output
phosphor to prevent light from being reflected back through the
tube and activating the photocathode - see panel (c). This layer
also serves as an earth to remove spent electrons from the tube.
The glass envelope is encased in a lead lined canister which
provides a degree of radiation protection from stray X-rays.
Protection from external magnetic fields is also incorporated.
Many XII designs also incorporate a fixed radiographic grid at the
input.
Variable field sizes can be achieved by altering the voltage on the electrostatic lens G3illustrated in Figure
4.10. An increase in this voltage can be used to decrease the area of the input phosphor that can be imaged
on the output phosphor but it also decreases the Minification Gain. An increase in radiation exposure (i.e.
mA) is generally needed to compensate. Such variable field sizes can be useful when fine control is required
in interventional radiology, for example. However, it should be appreciated that the consequent change in
patient exposure is inversely related to the square of the field size when images of equivalent brightness are
needed.
It should be appreciated that even with substantial brightness gains the image on the output phosphor remain
very dull and direct viewing is not practicable. This is one reason why a video camera is generally used to
view the XII output. Furthermore, it should be appreciated that XII images suffer from effects such as
pincushion distortion and vignetting.
Video Camera
The video camera views the XII output through optical coupling consisting a simple system of lenses and an
aperture. The aperture can be used to control the amount of light striking the camera. Alternatively, the vide
camera and XII can be coupled fiber-optically.
There are two general types of video camera used for fluoroscopy. The traditional device is called a
photoconductive camera or Pick-Up Tube (PUT) and is illustrated in the following figure, and the more
recent semiconductor or CCD camera. Both cameras generate images in the form of a video signal, but in
substantially different ways. The operation of these cameras is described below.
Photoconductive Video Camera
The PUT camera uses a photosensor which is called the target. The output image of the XII is
focused onto a target assembly (see Figure 4.12) which is made from a three layer sandwich.
The first layer is a glass face plate which serves only to maintain the vacuum whilst transmitting
the incident light. Next comes the signal plate, which consists of a thin transparent film of
electrically-conductive graphite. A potential of approximately 40 volts is applied to this layer.
The third member of the sandwich is the target which is a thin film of photoconductive material
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Fig. 4.12: Simplified diagram of a photoconductive
video camera.
An important part of the camera is the vacuum tube typically measuring up to 15 cm in length
by up to 3 cm in diameter. It can be surrounded by coils for focusing and deflecting the electro
beam. Alternatively, deflection electrodes called deflectronscan be shaped on the inside wall
of the vacuum tube.
At the opposite end of the tube is the cathode. Through indirect heating, electrons are boiled
from the cathode surface and accelerated by up to 900 volts towards the anode. The anode
extends across the end of the tube as a fine wire mesh which allows most of the electron beam
to pass through it. Once they pass through they decelerate to hit the signal plate. The electron
beam is focused to a tiny point using the focusing coils which extend almost the full length of th
tube. The focused electron beam can be deflected both vertically and horizontally by the
application of suitable voltages to the deflectrons.
Video camera targets are generally scanned in a raster pattern which is similar to that used in
broadcast TV. Here an interlaced raster scan is used to scan 625 horizontal lines of the target
at a rate of 25 images (calledframes) each second, for instance, to generate the perception o
continuous imaging.
Light falling on the photoconductive target changes its conductivity in proportion to the light
intensity and these changes are read by the scanning electron beam. The electric current
generated is conveyed via the transparent conductive film to an external circuit where a video
voltage can be developed.
The choice of photoconductive target has an important impact on the performance of the video
camera. A general-purpose target material used in the vidicon camera is antimony trisulphide
(Sb2S3), while other materials such as lead oxide (PbO), as in the Plumbicon, and cadmium
selenide (CdSe), as in the Chalnicon, can be used for more specialised applications.
Semiconductor Video Camera
These cameras are typically based on charged coupled device (CCD) technology and consist
of a semiconductor chip that is sensitive to light. The chip contains many thousands of electron
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Fig. 4.13: CCD imager showing a
section of the photodiode array along
with vertical and horizontal shift
registers.
Fig. 4.14: The video signal
corresponding to the readout from
one line of a video camera target.
sensors (e.g. photodiodes) each of which reacts to light and generates a signal that varies
depending on the amount of light each receives. CCDs have been developed primarily for the
domestic video camera market and the result is a
compact, lightweight camera with improved
performance relative to photoconductive cameras.
Array sizes of up to 2,028 x 2,028 and greater are
possible.
The device differs critically from the PUT camera in
that a scanning electron beam in an evacuated
environment is not required. The signal is read from
each sensor element at the appropriate rate by purely
electronic means using what are called electronic shift
registers - see Figure 4.13. Sensor signals are first
transferred into the vertical shift registers and then
shifted one line at a time to a horizontal shift register
to be read out to form a video signal. The analogy to
a bucket brigade has often been used to describe thisprocess.
Note that CCD cameras have also found application in the direct viewing of intensifying scree
in radiographic applications[11].
The camera control unit provides the link between the camera and
the display monitor. It houses an oscillator circuit that controls the
electron beam deflections of a PUT camera, for instance, or the
interline transfer in CCD cameras and synchronises the recorded
image with that displayed on a video monitor or stored in a
computer memory. Synchronization is generally achieved using
sync pulses inserted during the blanking interval of each video line
and frame - see Figure 4.14. The main amplifier in the control unit
adjusts the gain of the video amplifier so that the video signal
range, usually 0 to 1 V, can be fully utilised. This function is called
Automatic Gain Control (AGC).
Note that the video image is generally rectangular in shape. The
ratio of the width of the image to its height is called the Aspect Ratio and is typically 4:3. This is not ideally
suited to medical imaging where the video camera is viewing the circular image produced by an XII.
Automatic Brightness Control
The range of body part thicknesses and composition encountered in fluoroscopic imaging is quite large. In a
barium upper GI study , for example, moving from screening the neck to screening the stomach will
substantially increase X-ray attenuation along the way and result in darker images unless the exposure facto
and/or the gain of the video camera are adjusted to compensate. Adjustment of the exposure factors is
generally referred to as Automatic Brightness Control(ABC), while camera-based adjustment is
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provided by the AGC, as described above. ABC is used to keep the brightness of the displayed image at a
constant level during examinations. It involves the adjustment of the kV and mA automatically depending on
the part of the anatomy being examined. This can be achieved using a small photosensor at the XII output,
for instance, which monitors the central portion of intensified images and feeds a signal back to the generato
to adjust the kV, the mA (or both) accordingly.
ABC systems are generally designed to operate between minimum (e.g. 70 kV) and maximum (e.g. 120 kV
kilovoltages. The minimum kV can be used at the start of an exposure sequence, for instance, to prevent low
energy X-rays exposing the patient unnecessary, and is then increased automatically so that a pre-determine
image brightness level is reached. The tube current (mA) can also be adjusted automatically during this
process. Such mA adjustment within the kV range of the ABC is limited by the power rating of the XRT.
When the power limit is reached, in fluoroscopy of the lateral abdomen or at steep-angled cardiac
projections, for instance, further adjustment of exposure factors is no longer possible and the AGC circuitry
can come into play to maintain image brightness. However, images with increased electronic noise generally
result.
The actual kV and mA settings used by ABC systems dictate the contrast displayed in fluoroscopic images
as well as the dose to the patient. A High DoseABC mode can be used which lowers the kV and booststhe mA so that image contrast can be improved at the same image brightness, while a Low Dosemode
increases the kV and lowers the mA to effect a similar outcome. A third intermediate mode can also be
selected on many systems. The heat capacity of the XRT imposes the power limit, which is controlled by th
product, (kV x mA), the ABC system automatically applies. This is dictated by the thickness and
composition of the anatomy being screened. When this product reaches the power limit, in the case of a ver
large patient, for example, the ABC generally maintains constant kV and mA settings. Hence the need for
additional control provided by the camera's AGC. In addition, the AGC control of image brightness happen
almost instantaneously whereas the HV circuitry can take about a second or so to respond to any ABC
detected illuminance changes. The AGC is therefore also of use during this adaptation period so as to
maintain a constant brightness in displayed images.
Radiation Dose Management
An indicator of the radiation exposure required to generate a CR or DR image is provided on many systems. This
indicator can be called the Sensitivity Index(SI), the Log Median(LgM), the Exposure Index(EI) or similar
parameter and can be used to gauge the adequacy of an exposure. Note that these parameters are generally
referenced to exposures generated under specific conditions, e.g. X-ray energy, beam filtration etc., and therefore
can be regarded only as a crude indicator of patient dose.
Dose-Area Product (DAP) - see the discussion below - records can be used to review exposure trends over time
in the clinical environment[12]. One of the aims here is to reduce the phenomenon of Exposure Creep, where
exposures slowly increase over time in the pursuit of images of excellent quality, for example. Another aim here ca
be to review exposure measurements relative to Diagnostic Reference Levels (DRLs).
Retake analysis can also be combined with such exposure reviews so that, for instance, a team of X-ray personne
can work collectively to improve performance. Reasons for retake radiographs can be catalogued into
radiographic exposure errors, including those resulting from patient movement, when too long an exposure is used
for example, and radiographic positioning errors[13].
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Skin Injury
There has been considerable interest in recent years in reducing fluoroscopic doses following numerous
reports of epilation and skin injuries to patients from prolonged interventional procedures[14]. Regulatory
agencies throughout the world have responded to ICRP deliberations[15]and have issued recommendations
to the effect that the person responsible for the apparatus:
Needs to establish standard operating procedures and clinical protocols.Should know typical radiation dose rates for the specific fluoroscopic system.
Must assess the impact of each procedure's protocol on the potential for radiation injury to
patients.
Must modify the protocol to limit the dose.
Should enlist a qualified health physicist to provide assistance in developing and optimising dos
minimisation techniques.
Should give consideration to rotating the tube and the image intensifier through 180 during
prolonged neuroradiological procedures.
Should give consideration to carrying out those cardiology procedures where multiple stenting
necessary over a period of weeks to fractionate the radiation dose.
Should give due consideration when purchasing new equipment to features offered by the
manufacturer that may aid in reducing the patient dose.
Skin injury has been found to be sensitive to factors such as previous high dose exposure, medication,
connective tissue disease and diabetes mellitus. A review of hair and skin effects is given in Balter et al.
(2010)[16]. Single-site skin doses in interventional radiology above 2 Gy, for example, have been found to
cause erythema and epilation and higher doses to cause permanent skin damage.
Mechanisms of patient dose reduction which have been developed include:
Pulsed Fluoroscopy: this is a feature which uses short pulses of radiation of variable duration
and frequency (Fig. 3.1). Such pulsing can be generated by switching the mA in the HV
generator or by controlling the electron beam of the XRT, as in the Grid-Controlled XRT.
Following each exposure pulse, the image is stored in image memory and displayed
continuously until the next pulse to give a contiguous visual effect. Dose rate reductions of the
order of 90% are achievable using this approach although a stroboscopic artefact may result
when imaging fast moving objects such as the heart.
Additional Filtration: the addition of a thin Cu filter (0.1 - 0.3 mm) at the output of the X-raytube can generate substantial dose reductions without a detrimental impact on image quality. A
refinement to this approach is the Region of Interest (ROI) filter which provides little filtration t
the central region of the field of view and substantial filtration in peripheral regions - the
rationale being that high image quality is required in the centre of the field and noisier image da
of lower contrast is tolerable outside of this region for providing general anatomical information
only. This latter mechanism has also been referred to as the X-ray fovea.
Collimation: Since the use of conventional rectangular collimators used with XII-based
fluoroscopy systems (which generate circular images) results in unused exposure, some moder
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systems use circular collimators. These typically have the capability of automatically respondin
to changes in source-to-image distance and in the electronic zoom selection used in the XII. A
complex collimator assembly, based on adjustable multileaf absorbers, results and can include
robotic control. Another feature in this regard is referred to as radiation-free collimator
positioning and involves storing an image in computer memory, stopping the exposure and
adjusting the collimator controls so that the computer generates a graphics display of the
collimated region.
Digital Image Storage: the fact that images are stored in computer memory (see the followin
chapter) allows the possibility of using features such as Last-Image-Hold(LIH), where the
most recently acquired image is displayed continuously without irradiating the patient,
Reference Imaging, where a previously acquired image is displayed on a second video
monitor for comparison purposes, Image Browsing, where multiple images acquired
previously can be displayed, and Fluoroscopic Loop, where an imaging sequence can be
continuously replayed for closer inspection.
Automatic Brightness Control (ABC) can be used in conjunction with exposure pulse control and additiona
filtration to generate what is called Automatic Dose Rate Control(ADRC)[17]. Here, different filters canbe selected automatically for insertion into the X-ray beam depending on the particular exposure factors
encountered when screening commences. For example, some ADRC systems can generate the selection of
a kilovoltage which remains constant during the imaging sequence,
an mA pulse height and width which maintains image brightness, and
the automatic insertion of Cu filters of different thicknesses,
depending of the thickness and composition of the body part being examined. Such systems can be used
with four selectable copper filters, for instance, of thickness 0.2, 0.3, 0.6 and 0.9 mm depending on patient
thickness. Since absorbed doses are greater for larger patient thicknesses, because higher exposure factorsare required to acquire images of adequate brightness, a filter thickness is chosen that hardens the X-ray
beam appropriately and reduces patient dose. This is referred to as Spectral Filtering. The ADRC system
can cause the HV generator to select a constant kV, depending on the Cu filter thickness, and vary the mA
pulse height depending on the anatomy. For example, it may select 60 kV with a 0.9 mm Cu filter for a thin
body part, such as the foot, and 80 kV with a 0.3 Cu filter for a body part greater than 20 cm thick, with th
mA adjusted for each exposure by the generation of pulses of the necessary intensity.
Most fluoroscopy systems also feature an audible alert which sounds after a 'beam-on' time of five minutes.
The accumulated exposure time is generally displayed in real-time on the image display, along with the kV
and mA - and dose rate readings when a Dose-Area Product meter (see below) is installed.
Staff Dose
Three sources of exposure arise for staff when operating radiography and fluoroscopy equipment: the
primary beam used to expose the patient, leakage radiation from the XRT and scattered radiation from the
patient. Lead aprons can be worn and distance exploited to minimise the impact of leakage and scattered
radiation in instances where staff are required to remain in an X-ray room or operating theatre during
exposures. Note however that lead aprons only attenuate X-rays by ~90-95%, depending on their lead
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equivalence. These aprons are generally made from lead-impregnated vinyl, are equivalent to 0.25-0.5 mm
of lead and weigh about 5 kg. Lighter aprons with similar attenuation can be made from composites, such a
tungsten and tin, and may be of value to staff during long interventional procedures. Aprons of 0.5 mm lead
equivalence should be used by staff who remain in close proximity to the patient during exposures. The safe
style is the wrap-around apron, in comparison to backless designs. From a practical perspective, lead
aprons should always be stored vertically on racks since folding them can introduce cracks which reduce
their effectiveness. In addition, they should be radiographed at least annually to check for any such signs of
wear and tear.
In addition, the number of staff in the immediate vicinity of the patient can be kept to a minimum and
warnings given when an exposure is impending. Ideally, staff should be behind a shielded operator's console
area during an exposure. The design of X-ray rooms is such that, in general, adequate shielding and distanc
ensures that doses to any person in the room, behind the console or, indeed, external to the room can be
optimised.
In some interventional cases, the interventionist may find it necessary to have their hands in the primary beam
for parts of the procedure. Such repeated exposure over a number of years may potentially lead eventually
to radiation dermatitis of the hands. Dose optimisation in such circumstances can be achieved by arrangingthe imaging system so that the hands are kept on the beam exit side of the patient. For example, if the XRT
positioned under the patient then the hands manipulating the catheter should be on top of the patient.
Fig. 4.15: Typical occupational skin absorbed dose rates near
fixed fluoroscopic equipment in the absence of protective
aprons and drapes: (a) over-couch and (b) under-couch X-ray
tube.
It should also be noted that the use of over-couch XRTs in fluoroscopy can lead to significantly greater staf
doses than systems which use under-couch XRTs - see Figure 4.15. Increased radiation scattered from the
patient and the lack of shielding provided by the structure of the patient table are the cause. An increased
incidence of lens injuries has been found, for instance, in radiologists who have used over-couch systems
without protective screens[18]. Note, however, that this over-couch argument may not always be valid when
inverted C-arms are used to assist with hand surgery[19].
The elevated dose rates to the trunk and head of the staff member can be noted in panel (a) of the figure.
Radiation dose to the legs of interventionists can also be significant even with the use of under-couch XRTs
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Fig. 4.16: Typical isodose curves (in
Gy/min) for a mobile C-arm
fluoroscopy system.
and with C-arm systems[20][21].
The use of C-arm and similar arrangements in cardiac
catheterisation laboratories, operating theatres and angiography
suites can introduce additional hazards - see Figure 4.16. These
arise because of the ergonomics associated with the use of lead
shielding devices - which are intended for the minimisation of
scatter and leakage at the interventionist's position. It is thereforequite possible for their thyroid and eye lens, for instance, to
receive substantial doses. Thyroid shields and lead spectacles can
be worn and adjustable lead glass shields used for dose
optimisation purposes in these circumstances[22][23]. Shielding
equivalent to 0.25 mm of lead is generally required with glasses
also having side protection. In addition, interventionists can wear a
second dosemeter at the level of their neck to monitor the dose to
their thyroid and eyes. Note that radiation-associated posterior
lens changes have been observed in the eyes of cardiac Cath Lab
staff - both medical and clinical, most not wearing lead glasses -
with an incidence significantly greater than an unexposed control
group[24]and that a study of cataract incidence has been launched for a large sample of interventional
cardiologists[25].
The location of the interventionist during the procedure has a large influence on their hand dose[26]. A ring
dosemeter on the finger proximal to the XRT can be used to monitor such doses. Lead gloves are not
generally worn for ergonomic reasons and also because the additional attenuation generated when shielded
hands are in the beam will increase the automatically controlled exposure factors. Note that particular
additional optimisation techniques can be required in specialities such as urology[27]and endovascular
surgery[28].
Targeted education for all users of specialised X-ray apparatus has been identified as a key optimisation
strategy[29]and has been implemented successfully in, for example, computed tomography[30].
Patient Dose
Good imaging geometry should be used to optimise patient protection in radiography, fluoroscopy and
fluorography. In most situations the image receptor should therefore be moved to the maximum distance frothe XRT and the patient placed as close to the image receptor as possible. In other words, the patient shoul
ideally be moved to the image receptor and not the other way around. Some fluoroscopy systems can do th
automatically so as to maintain a narrow air gap between the patient and image receptor before exposures.
Effective doses to patients from fluoroscopy procedures are considerably higher, as might be expected, tha
in General Radiography - see the table below. The higher doses result from both fluoroscopy-screening
exposures and multiple fluorography exposures. Notice that strongly attenuating parts of the body which
contain a number of radio-sensitivite organs, such as the abdomen, generate the larger doses in the table.
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Mean Effective Doses from Fluoroscopy
Procedures[31][32]
Procedure Effective Dose (mSv)
Cerebral Arteriography 2.5
Nephrostomy 5.5
Barium Meal 8.2
Renal Arteriorography 10.3
Barium Enema 11.7
Biliary Stent Placement 12.5
Enteroclysis 14.0
More recent and extensive data is provided in Mettler et al.(2008)[33]. Note that specific patient shielding
can be of benefit to patient dose in different types of X-ray examinations, e.g. thyroid protection in cerebral
angiography[34].
It can be noted from the table that all doses are higher than that of the average annual background dose (2.4
mSv). From this perspective, the dose from a Cerebral Arteriogram is equivalent to about 1 year of
background exposure - its so-called Background Equivalent Radiation Time (BERT). Note however that
considerable variations can exist in patient dosimetric survey data, by a factor of 10 or more, due to patient
anatomical differences and the quality of the imaging technology, for example. Note also that variations by a
factor of over 100 occur in background exposure in different parts of the world. It is apparent that the BER
is therefore only a very approximate indicator of relative dose.
The higher doses in the above table, and higher still in extended interventional procedures, have prompted
the use of dosimetry equipment to routinely monitor parameters such as the Dose-Area Product (DAP) and
the Peak Skin Dose. DAP (also called the Kerma-Area Product) is a measure of the total energy exiting an
XRT and is generally measured (in Gy cm2) at a location in the beam close to the collimators. The paramete
has the advantage that it is independent of source-to-skin distance, as was described previously, and can be
used to estimate stochastic risk from a procedure. The Peak Skin Dose is a useful indicator of the likelihood
of deterministic effects. Such measurements can be also be used to establish a dose scale to assist
interventional staff in radiation dose management[35].
Implementation of dose monitoring is critical for dose optimization in digital radiography. Staff guidelines as
previously developed for film/screen radiography, should include appropriate collimation, source-to-imagedistance (SID), focal spot size and patient positioning. This information can also enable effective doses to b
subsequently estimated, in situations where DAP meters, for instance, are not available. The use of image
quality indicators for different examinations is an additional step in assisting clinical management of the
balance between dose and image quality[36].
Results of a comprehensive catalogue of dose surveys from 1980-2007 are summarised in Figure 4.17. It
can be seen that effective dose in General Radiography vary by a factor of >1,000, in the range 0.001 to 3
mSv, depending on the area of anatomy irradiated. Radiography/fluoroscopy procedures, as used for bariu
studies, can be seen to generate higher doses (4-8 mSv) as do CT studies (2-16 mSv). Interventional
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Fig. 4.17: Adult effective doses from various X-ray
procedures (adapted from ref. #31).
radiology is seen to generate still higher doses (5-70 mSv). CT and interventional exposures can therefore b
considered at relatively high dose examinations and therefore require greater attention from a dose
management perspective. However, it should be appreciated that these values contain large uncertainties,
which can range by factors of 5-10 or more. Furthermore, the average figures are themselves age- and sex-
averaged and are also subject to a variation of 40% when referred to an individual patient, depending on
variables which include the patient's weight, orientation, X-ray exposure factors and imaging technology
used. It is therefore apparent that they should be considered as indicative values only which should never be
usedas a substitute for the dose to an individual patient from a particular examination. Their use here is soleto provide a comparison for dose management purposes.
Onafinal point, note that patient dosimetry in general
radiography, fluoroscopy, mammography and
computed tomography is reviewed in Huda et al.
(2008)[37].
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(http://www.ncbi.nlm.nih.gov/pubmed/15933078). Radiat
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Advances in computed radiography s ystems and their
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9. Rivetti S, Lanconelli N, Bertolini M, Borasi G, Golinelli P, Acchiappati D & Gallo E, 2009. Physical and psychophys ical
characterization of a novel clinical system for digital mammography (http://www.ncbi.nlm.nih.gov/pubmed/19994524). Me
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http://www.ncbi.nlm.nih.gov/pubmed/16094003http://www.ncbi.nlm.nih.gov/pubmed/19759087http://www.ncbi.nlm.nih.gov/pubmed/20825640http://www.ncbi.nlm.nih.gov/pubmed/20726724http://www.ncbi.nlm.nih.gov/pubmed/20920841http://www.ncbi.nlm.nih.gov/pubmed/17533529http://www.ncbi.nlm.nih.gov/pubmed/17533541http://www.ncbi.nlm.nih.gov/pubmed/12763947http://www.ncbi.nlm.nih.gov/pubmed/11379736http://www.ncbi.nlm.nih.gov/pubmed/9771383http://www.ncbi.nlm.nih.gov/pubmed/17879779http://www.ncbi.nlm.nih.gov/pubmed/20093507http://www.ncbi.nlm.nih.gov/pubmed/11459599http://www.ncbi.nlm.nih.gov/pubmed/8888398http://www.ncbi.nlm.nih.gov/pubmed/18592314http://www.ncbi.nlm.nih.gov/pubmed/15933094http://www.ncbi.nlm.nih.gov/pubmed/9015806http://www.ncbi.nlm.nih.gov/pubmed/16862412http://www.ncbi.nlm.nih.gov/pubmed/199945248/12/2019 Basic Physics of Digital Radiography_The Image Receptor - Wikibooks, Open Books for an Open World
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